Drug Delivery (Handbook of Experimental Pharmacology)

Handbook of Experimental Pharmacology

Band 197

Editor in Chief F.B. Hofmann, Munich Editor J.A. Beavo, Seattle, WA A. Busch, Berlin D. Ganten, Berlin J.-A. Karlsson, Singapore M.C. Michel, Amsterdam CP Page, London W.Rosenthal, Berlin

For more, visit: http://www.springer.com/series/164

Monika Schäfer Discount Editor

medication

Editor Prof. Dr. Monika Schafer-Korting Faculty of Pharmacy, Free University Berlin Koenigin-Luise-Str. 2-4 14195 Berlin Germany[email protected]

ISSN 0171-2004 e-ISSN 1865-0325 ISBN 978-3-642-00476-6 e-ISBN 978-3-642-00477-3 DOI 10.1007/978-3-642-00477-3 Springer Heidelberg Dordrecht London New York USA Library of Congress Control Number: 2009933605 # Springer-Verlag Berlin Heidelberg 2010 This work is protected by copyright. All rights are reserved, whether concerning the material in whole or in part, in particular the right to translate, reprint, re-use illustrations, read, broadcast, reproduce on microfilm or otherwise and store in databases. Reproduction of this publication or parts thereof is only permitted within the framework of the copyright law of September 9, 1965 in the currently valid version and always requires the consent of Springer. Violations are punishable under German copyright law. The use of generic descriptive names, registered names, trademarks, etc., in this publication, even if not expressly stated, does not mean that these names are not protected by relevant laws and regulations and therefore may be used freely. Cover design: Printed on acid-free paper by SPi Publisher Services Springer is part of Springer Science+Business Media (www.springer.com)

foreword

"Handbook of Experimental Pharmacology" is considered "one of the most authoritative and influential series in the field of pharmacology". It aims to provide "a critical and comprehensive discussion of the most important areas of pharmacological research". Apparently, this mission has been going on for quite some time, with the current roll number 197. It all started in the first half of the 20th century by Dr. A. Heffter, who, among other things, was the first president of the German Association of Pharmacologists and Rector of the University of Berlin in 1922. His pharmacological interests include a crystalline glycoside from Strophanthus kombe. Thus, in most books to date, various pharmaceutical active ingredients (principles of action) or groups thereof have been dealt with. More recently, the scope has gotten wider. This is reflected in the accompanying volumes and their themes, namely Volume 196 on Adverse Drug Reactions and Volume 198 on Birth Control. Nonetheless, drug delivery was certainly a topic that probably wasn't covered in the series during the first few decades of the manual. In this case, the question wouldn't be "drugs", although that term seems too broad in the context of a particular volume of a pharmacology handbook. The key is "delivery". Of course, every institution today has to deliver, using the word in the sense of meeting expectations or achieving goals. The latter is actually close to what is meant here. The term "delivery" is used in many situations in general life. In particular, it concerns commercial activities that require the transportation of goods, mainly from producers to consumers. In this context, vehicles have played an important role for centuries and can therefore also be called delivery vehicles. While these vehicles were still horse-drawn carriages in the 19th century, today one can think of fast, self-propelled objects in a narrow sense, such as the fairly small trucks that pretty much dominate highways today.

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foreword

Indeed, even in modern drug treatments, mode (of administration) and precise time frames have become critical to success, at a time when proxy parameters of efficacy are increasingly replaced by actual outcome parameters. The foundational part of this volume's table of contents introduces three terms that may be considered keywords of our time. The first is "aiming". Principles of action absolutely must get where they are supposed to go. The military has known this for thousands of years, and it has recently been emphasized in the field by coining the term "surgical warfare." The idea is to increase the risk/benefit ratio by preventing what soldiers today call "collateral damage." The next term is "nanomedicine." The word reflects a broader term, "nanotechnology." The term and the concepts behind it were inspired by Dr. E. Drexler's 1986 book "The Coming Era of Nanotechnology". This work mirrors earlier work by Ph.D. N. Taniguchi, who defined nanotechnology in 1974. In our context, nanotechnology can provide tiny particles, so-called nanoparticles, as potential carriers for active pharmaceutical ingredients. This approach somewhat mirrors a natural approach of the mammalian body known as protein targeting. Regardless of the starting material for production, we are talking about a vector. A "freight forwarder" is someone who transports goods for a third party, another term is a "carrier". Today, the term "carrier" is used in many different contexts. One of these is an "aircraft carrier," essentially a ship that can launch military fighter jets. In this way, targets beyond the range of the aircraft alone can be reached. The third term is "biosensing". This means using biosensors, biological structures that can recognize specific analytes. One last time, using a military metaphor, we might think of the radar systems that help aircraft carriers and the aircraft they carry carry out their missions. It has now become clear that very specific devices are critical to optimally performing drug delivery tasks—often in order to save and improve quality of life. In fact, there is a wide range to be covered today, from liposomes to medical drug delivery devices. In detail, it is of course important whether drug delivery is considered in the therapeutic context of systemic or topical application. In at least one specific organ, namely the skin, there is a dual-use option: while vehicles may only be needed to optimize topical treatment of skin conditions, transdermal delivery is now also a relevant option in everyday drug therapy. The goal is systemic efficacy. In planning this book, I was struck by the breadth and depth of current drug delivery knowledge. When I finally got the manuscript off to the publisher for printing, I have to admit I was still impressed - and I hope the scientific community concerned feels the same enthusiasm when it reads this new book. This book certainly would not have been possible without the support of 16 distinguished contributors (and, in some cases, their collaborators) from various disciplines or subfields of the life sciences. This only happens because of invitations and

foreword

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The editor-in-chief of the series, my esteemed colleague Dr. Walter Rosenthal from Berlin, to whom I am very grateful. We would also like to thank Ms Susanne Dathe of Springer Heidelberg for her ongoing editorial support. Finally, thanks to Ms. Barbara Brüggener for technical support in editing the manuscript. Berlin, November 2009

M. Schäfer discount

content

first part

basic

Passive and active drug targeting: examples of drug delivery to tumors. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Vladimir P. Torchilin Nanoparticle technology for cancer therapy. . . . . . . . . . . . . . . . . . . . . . . . . . 55 Frank Alexis, Eric M. Pridgen, Robert Langer, and Omid C. Farokhzad Miniature Biosensing and Drug Delivery . . . . . . . . . . . . . . . . . . . . . . . . . 87 Andrea A. Robitzki and Randy Kurz Part Two

equipment

Lipid nanoparticles: effects on changes in bioavailability and pharmacokinetics. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 115 Eliana B. Souto and Rainer H. Muèller Viral vectors for gene transfer: current status of gene therapy. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 143 Regine Heilbronn and Stefan Weger Pulmonary Delivery: Inhaled Drugs. . . . . . . . . . . . . . . . . . . . . 171 Andreas Henning, Stephanie Hein, Marc Schneider, Michael Bur, and Claus-Michael Lehr. Needle-free vaccination. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Reaper 193

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Pharmaceutical polymers: principles, structures and applications of drug delivery systems. . . . . . . . . . . . . . . . . . . . . . . 221 Jayant Khandare and Rainer Haag Mucoadhesive drug delivery systems. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 251 Juliane Hombach and Andreas Bernkop-Schnürch for contraception and gynecology Intrauterine administration of therapy: new approaches. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 267 Dirk Wildemeersch Drug eluting medical implants. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 299 Meital Zilberman, Amir Kraitzer , Orly Grinberg, and Jonathan J. Elsner Part III Systemic Use of Clinical and Preclinical Therapeutics to Enhance Oral Delivery. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 345 Franz Gabor, Christian Fillafer, Lukas Neutsch, Gerda Ratzinger and Michael Wirth Transdermal Drug Delivery. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 399 Richard H. Guy targets the brain—crossing or bypassing the blood-brain barrier. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 411 Heidrun Potschka Part 4 Clinical and preclinical use of topical vehicle therapy in topical treatment of dermatological diseases. . . . . . . . . . . . . . . . . . . . . . 435 Hans Christian Korting and Monika Schäfer-Korting Medical devices for the treatment of eye diseases. . . . . . . . . . . . . . . . . . . . . 469 Index by Tsutomu Yasukawa and Yuichiro Ogura. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 491 Cancellation Notice. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 501

contributor

Frank Alexis Nanomedicine and Biomaterials Laboratory and Department of Anesthesiology, Brigham and Women's Hospital, Harvard Medical School, Boston, MA 02115, USA; MIT Harvard Center of Excellence in Cancer Nanotechnology, Cambridge, MA 02139, USA; Harvard- MIT Division of Health Sciences and Technology, Cambridge, MA 02139, USA Andreas Bernkop-Schnürch Institute for Pharmacy, University of Innsbruck, Josef-Möller-Haus, Innrain 52, 6020 Innsbruck, Austria Michael Bur Biopharmaceuticals and Pharmaceutical Technology, Saarland University , 66123 Saarbrücken, Germany Jonathan J. Elsner Faculty of Engineering, Tel Aviv University, Department of Biomedical Engineering, Tel Aviv 69978, Israel Omid C. Farokhzad Nanomedicine and Biomaterials Laboratory and Department of Anesthesiology, Brigham and Women's Hospital, Harvard Medical School, Boston, MA 02115, USA; MIT-Harvard Center of Excellence for Cancer Nanotechnology, Cambridge, MA 02139 4, 1090 Vienna, Austria xi

twelve

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Orly Grinberg School of Engineering, Department of Biomedical Engineering, Tel Aviv University, Tel Aviv 69978, Israel Richard H. Guy Department of Pharmacy and Pharmacology, University of Bath, Claverton Down, Bath, BA2 7AY, UK Rainer Haag Institute of Chemistry and Biochemistry – Organic Chemistry , Free University of Berlin, Takustr. 3, 14195 Berlin, Germany Regine Heilbronn Institute for Virology, Charité – Universitätsmedizin Berlin, Hindenburgdamm 27, (HBD 27), 12203 Berlin, Germany,[email protected]Stephanie Hein Biopharmaceuticals and Pharmaceutical Technology, Saarland University, 66123 Saarbrucken, Germany Andreas Henning Biopharmaceuticals and Pharmaceutical Technology, Saarland University, 66123 Saarland University Juliane Hombach Innsbruck University, Innrain 52, 6020 Innsbruck, Austria Mark Kendall Australian Institute for Bioengineering and Nanotechnology (AIBN), University of Queensland, Building 75 – Cnr of College and Cooper Road, Brisbane Qld 4072, Australia Jayant Khandare Institute for Chemistry and Biochemistry, Freie Universität Berlin, Takustraße 3, 14195 Berlin, Germany Hans Christian Korting Clinic and Polyclinic for Dermatology and Allergy, Ludwig-Maximilians University of Munich (LMU), Frauenlobstrasse 9-11, 80337 Munich, Germany Amir Kraitzer Faculty of Engineering , Department of Biomedical Engineering, Tel Aviv University, Tel Aviv 69978, Israel

contributor

Thirteen

Randy Kurz Center for Biotechnology and Biomedicine, University of Leipzig, Deutscher Platz 5, 04103 Leipzig, Germany Robert S. Langer MIT-Harvard Center of Excellence for Cancer Nanotechnology, Cambridge, MA 02139, USA; Harvard-MIT Health Sciences and Technology MIT, Cambridge, MA 02139, USA Claus-Michael Lehr Department of Biopharmaceutics and Pharmaceutical Technology, Im Stadtwald, 66123 Saarbrücken, Germany Rainer H. Mu¨ller Department of Pharmaceuticals and Biopharmaceuticals, Free University of Berlin , Kelchstrasse 31, 12169 Berlin, Germany Lukas Neutsch Department of Pharmaceutical Technology and Biopharmaceuticals, University of Vienna, Althanstrasse 14, 1090 Vienna, Austria Yuichiro Ogura Ophthalmology and Vision Science, Graduate School of Medicine, Nagoya City University Department of Veterinary Medicine, 1 Kawasumi , Mizuho-cho, Mizuho-ku, Nagoya, Aichi 467-8601, Japan Heidrun Potschka Faculty of Veterinary Medicine, Ludwig-Maximilians-University, Veterinärstr. 13, 80539 Munich, Germany; Ludwig-Maximilians-Universität Munich, Munich, Germany Eric M. Pridgen MIT-Harvard Center of Excellence in Cancer Nanotechnology, Cambridge, MA 02139, USA; MIT Department of Chemical Engineering, Cambridge, MA 02139, Gerda Ratzinger, USA Department of Pharmaceutical Technology and Biopharmaceuticals, University of Vienna, Althanstrasse 14, 1090 Vienna, Austria Andrea Robitzki Institute of Biochemistry, University of Leipzig, Brüderstrasse 34, 04103 Leipzig

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contributor

Monika Schäfer-Korting Faculty of Pharmacy (Pharmacology and Toxicology), Free University of Berlin, König-Luise-Strasse 2-4, 14195, Berlin, Germany Marc Schneider Pharmaceutical Nanotechnology, Saarland University, Im Stadtwald, 66123 Saarbrücken, Germany Eliana B. Fernando Pessoa University Souto Faculty of Health Sciences, Rua Carlos da Maia, 296, 4200-150 Porto, Portugal Vladimir P. Torchilin Northeastern University Department of Pharmacy and Center for Pharmaceutical Biotechnology and Nanomedicine, 360 Huntington Avenue, Boston, MA 02115 , USA Stefan Weger Institute of Virology, Charité – Universitätsmedizin Berlin, Hindenburgdamm 27, 12203 Berlin, Germany Dirk Wildemeersch CONTREL RESEARCH, Technology Park, University of Gent, 9052 Gent (Zwijnaarde), Belgium Michael Wirth Department of Pharmaceutical Technology and Biopharmaceutics, University Vienna, Althanstrasse 14 , 1090 Vienna, Austria TSUTOMU YASUKAWA Department OphThalmology and Visual Science, Nagoya City University, Graduate School of Medical Sciences, 1 KA WASUMI, Mizuho-Cho, Mizuho-Ku, Nagoya, AICHI 467-8601, JAPAN Meital Zilberman College, biomedical engineering Department, Tel Aviv University, Tel Aviv 69978, Israel

Passive and active drug targeting: drug delivery to tumors with the example of Vladimir P. Torchilin

Table of contents 1 Drug targeting: general considerations. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 2 Concepts of Passive and Active Targets. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6 3 Pharmaceutical excipients: Take liposomes and micelles as examples. . . . . . . . . . . . . . . . . . . . . . . . . . 10 4 Chemistry to deliver drug nanocarriers with multiple functions. . . . . . . 15 Blood lifetime of 5-nanocarriers and their importance in drug delivery. . . . . . . . . . . 17 6 Passive accumulation of liposomes and micelles in tumors. . . . . . . . . . . . . . . . . . . . . . . . . . . . 21 7 Active tumor targeting with drug-loaded liposomes. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 26 8 Active tumor targeting with drug-loaded micelles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 33 9 Conclusion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 35 References. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 36

Summary: Over the past decade, the paradigm of using nanoparticle drug carriers has been well established in pharmaceutical research and clinical settings. The drug carrier is expected to stay in the blood for a long time, accumulate in the pathological sites (tumor, inflammation and infarct area) where blood vessels are affected and leak through the EPR (Enhanced Permeability and Retention) effect, and promote the targeted delivery of specific ligands. Delivery of drugs and drug carriers to hard-to-reach areas. Among the various approaches we have used to target drug delivery systems to desired pathological sites in vivo, two approaches appear to be the most advanced - passive (EPR effect mediated) targeting, which relies The longevity of blood and its accumulation in the blood, pathological sites of compromised vasculature, and active targeting are based on the binding of specific ligands to the surface of drug carriers to recognize and bind pathological cells. Here, we will consider and discuss these two targeting approaches using tumor targeting as an example. Vice President Northeastern University Torchilin Department of Pharmacy and Center for Pharmaceutical Biotechnology and Nanomedicine, 360 Huntington Avenue, Boston, MA, 02115, USA Email:[email protected]

M. Schäfer-Korting (Hrsg.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_1, # Springer-Verlag Berlin Heidelberg 2010

Vice President Tochlin

Key words drug delivery drug targeting tumor liposome polymeric micelles

1 Drug Targeting: General Considerations Drug targeting is generally understood to mean the accumulation of an active pharmaceutical ingredient (API) in a desired area of the body relative to other tissues and organs, mediated by spontaneous or external forces or targeting moieties Increase. For most drugs currently in use, the specificity and activity of a drug to a lesion or individual disease is usually based on the drug's ability to intervene in a local pathological process or defective biological pathway, rather than on its selection within a specific intracellular compartment or cell sexual accumulation or target cells, cells in organs or tissues. The active ingredient of a drug is usually distributed relatively evenly throughout the body, almost independent of the route of administration, and proportional to the local blood flow. Furthermore, in order to reach the site of action, the API must overcome many biological barriers, such as B. other organs, cells and intracellular compartments, where it can be inactivated or adversely affect organs and tissues not involved in the pathological process. Therefore, in order to achieve the desired therapeutic concentration of the active substance in a specific body compartment or tissue, large quantities must be administered (which increases the cost of the treatment), a large proportion of which is used even in the best case, only in normal tissues , whereas cytotoxic and/or antigenic/immunogenic agents may be responsible for many negative side effects. Targeted drugs can address all of these problems. In general, drug targeting is understood as the ability of an active substance to selectively and quantitatively accumulate in a target organ or tissue, regardless of the site and type of administration. Ideally, in this case, the local concentration of the drug at the disease site should be high, while its concentration in other non-target organs and tissues should be below a certain minimum level to prevent adverse side effects. The following advantages of drug targeting are obvious: the dosage regimen can be simplified; the amount of drug required to achieve the therapeutic effect can be significantly reduced, and the cost of treatment can also be significantly reduced; and the drug concentration at the desired site can be significantly increased, while Does not adversely affect non-target compartments. This applies to many diagnostic uses to a large extent. Although the concept of drug targeting proposed by Paul Ehrlich in the early 20th century saw a hypothetical "magic drug" as a two-component entity - the first designed to recognize and bind a target, while the The second aims to provide a therapeutic effect in this target – currently, the whole set of proposed targeting regimens includes many different targeted drug delivery methods. These methods do not necessarily involve the use of specific target entities. In some cases, different physical principles and/or some physiological principles

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Characterization of target regions can be used for successful targeting of drugs and drug carriers. The concept of targeted drugs involves the coordinated interaction of multiple components: active pharmaceutical ingredients, targeting moieties, drug carriers (soluble polymers, microcapsules, microparticles, cells, cytosperms, liposomes and micelles) for each A single targeting moiety is loaded with many drug molecules, and one target. Targets can be identified at different levels: at the level of the whole organ, at the level of certain cells specific to that organ, or at the level of individual components of these cells (cell surface antigens). Molecular-level recognition is undoubtedly the most general form of target recognition, as certain components can be found in any organ or tissue, whereas those components are specific only to that organ or tissue. Numerous drug-targeting approaches have been described to date, allowing the targeted delivery of therapeutic and diagnostic drugs to various tissues and organs. Some of these are discussed in many similarly reviewed publications (see, for example, Francis and Delgado 2000; Gregoriadis 1977; Muzykantov and Torchilin 2003; Torchilin 1995). Attempts have been made to deliver the drug directly to the affected area, to load the drug into specific carriers that respond to specific pH or temperature levels within the pathological range, and even to use specific external forces such as magnetic fields or ultrasound to deliver the drug where it is needed The region where the target is implemented or where the API is published. Many drugs and drug carriers are very effective in tumors, infarcts and areas of inflammation due to their passive accumulation through extravasation through the leaky vasculature (the so-called EPR effect (Enhanced Permeability and Retention)). Drugs and drug carriers can be targeted almost anywhere through the use of specific transport vehicles (specific entities with high specific affinity for target regions). Administration directly to the affected area Intracoronary infusion of thrombolytic enzymes that have been used to treat coronary thrombosis (Chazov et al., 1976), and intraarticular administration of nonhormonal drugs for arthritis (Williams et al. 1996). However, in most cases, it is technically difficult to deliver drugs directly to the affected organ or tissue; moreover, many diseases spread through multiple cells or tissues. All of these limit the applicability of this approach in rare clinical situations. Various endogenous and/or exogenous physical factors have been shown to mediate drug targeting. For example, this approach takes advantage of the pH and temperature differences between normal tissue and pathological regions (tumours, inflammation, etc.), which are characterized by acidosis (lower pH) and hyperthermia. With this in mind, it has been proposed to load various drugs onto pH- or temperature-sensitive drug carriers that can change their properties and The encapsulated drug is released when the area is exposed. The advantage of this approach is that although the drug-loaded carrier is evenly distributed in the bloodstream, the drug is broken down and released only in the targeted area. Furthermore, the target area can be additionally heated from the outside by means of external heat or ultrasound radiation. Cancer drug methotrexate accumulates in mouse tumors when given intravenously, results show

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This was several times faster when temperature-sensitive liposomes were added and external heat applied locally to the tumor area (Weinstein et al., 1979). Drug-loaded pH-sensitive liposomes are also widely used in experiments for the delivery of drugs and genetic material into various damaged tissues (see Budker et al., 1996; Torchilin et al., 1993, just a few of many reviews part). However, in many cases pathological sites do not differ significantly from normal tissue in terms of temperature or local pH, so targeting using pH or temperature differences is not applicable. Purposefully applied external magnetic fields can also be used for targeted drug delivery. In this case, the drug of choice must be combined with a drug carrier with ferromagnetic properties. As a result, drug-loaded ferromagnetic carriers are likely to gather in the region where an external magnetic field is applied. High magnetic field gradients and high blood flow velocities and concomitant high shear strengths in large vessels do not allow this method to work in vessels like the aorta. However, magnetic field-mediated drug accumulation in smaller vessels with slower blood flow and closer to the body surface has been clearly demonstrated (Widder et al., 1983). Dextran-coated iron oxide microparticles were coupled to thrombolytic streptokinase followed by implantation of a small, strong permanent magnet into the tissue, and this formulation was successfully used for targeted thrombolysis of artificially formed thrombi in the experimental canine carotid artery leading to the thrombus Regional blood vessels (Torchilin et al. 1988). Local thrombosis prevention in experimental dogs and rabbits was achieved by intravenous administration of autologous erythrocytes loaded with ferromagnetic colloidal compounds and aspirin by placing strong magnets outside thrombused vessels (Orekhova et al., 1990). There are many examples (McBain et al., 2008; Pauwels and Erba, 2007; Sun et al., 2008) where these drugs target magnetically when loaded into various drug carriers together with magnetosensitive nanoparticles. The tumor is exposed to an external magnetic field concentrated in the tumor. However, magnetic drug delivery has limitations in terms of blood flow velocity in the target area and is nearly impossible in large vessels or "deep" tissues. In this chapter, we focus on EPR effect-based drug targeting (passive targeting) and the use of targeting units (active targeting), with examples from state-of-the-art tumor drug delivery studies.

2 Concepts of Passive and Active Targeting It is now well known that under certain conditions the endothelial lining of vessel walls becomes more permeable than normal. This has been clearly demonstrated in many tumors (Hobbs et al 1998; Jain 1999) and infarcted areas (Palmer et al 1984; Torchilin et al 1992). as a

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Thus, in these regions, macromolecules and even relatively large particles (such as micelles and liposomes, ranging in size from 10 to 500 nm) can leave the vascular bed and accumulate in the interstitial spaces. Given that these large (aggregate) molecules/particles are loaded with a drug, they can bring that drug into areas of increased vascular permeability where the drug can eventually be released from the carrier. Since the limiting size of the permeabilized vasculature varies from case to case (Hobbs et al., 1998; Yuan et al., 1995), the size of drug-loaded particles can be used to control the effectiveness of such spontaneous 'passive' responses. Drug delivery or EPR effects (Maeda 2003; Maeda et al. 2000; Figure 1). This type of targeting requires that the drug delivery system be able to circulate for a long time (i.e. stay in the blood for a long time) to ensure sufficient levels of accumulation in the target. The most common way to keep drug carriers in the blood long enough is to "mask" them by modifying (grafting) their surface with certain water-soluble polymers with good solvation and a flexible backbone, such as polyethylene Diol (PEG) (Klibanov et al. 1990; Torchilin and Trubetskoy 1995). The surface-grafted "protective" polymer effectively prevents (slows down) the opsonization of the drug carrier and its clearance through the reticuloendothelial system. This method is most applicable to liposomes (Lasic and Martin 1995; Lasic and Papahadjopoulos 1998), although it has fairly broad applicability (Torchilin 1998). This is demonstrated by the cancer drug doxorubicin, which is incorporated into long-circulating PEG-coated liposomes and is currently in clinical use

Figure 1 Schematic of the enhanced permeability and retention (EPR) effect or "passive" targeting. (1) Drug-loaded nanocarrier; it cannot extravasate through normal endothelium, only small molecule free drug (4) can cross normal endothelium in either direction to some extent; (2) Gaps between endothelial cells appear in pathological (3) Areas (such as tumors, infarcts, and inflammation) through which nanoparticles can leak and accumulate, resulting in high local drug concentrations

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Very effective in EPR-based tumor therapy and greatly reduces the side effect profile of free doxorubicin (Gabizon 1995, 2001). Long-circulating polymeric micelles (Torchilin 2001) can be used as vehicles for tumor drug delivery with smaller cut-off sizes (Hobbs et al. 1998; Yuan et al. 1995), as demonstrated in mice with Lewis lung cancer Proven (Weissig et al. 1998). Important benefits of prolonged circulation of drugs and drug carriers in the bloodstream include the ability to maintain the desired concentration of the drug or drug carrier in the blood for long periods of time after a single dose; the ability to exploit the EPR effect to accumulate the drug in areas of leaky vessels; and improved The ability of drugs and drug carriers to target areas with limited blood supply and/or low concentrations of the target antigen, where it takes longer for sufficient drug to reach the target area. See below for more information on the importance of drug carrier longevity. However, it is worth mentioning that in many pathological situations, the integrity of the vascular endothelium is unaffected, so the possibility of EPR does not exist. Many of the drug-targeting approaches described so far are not general-purpose. Therefore, it may be technically difficult to administer the drug directly to the affected organ or tissue, or the focus of the disease may become delocalized. Affected areas often show little difference from normal tissue in terms of vascular permeability, temperature, or local pH. Magnetic drug delivery also has limitations related to targeted blood flow rates. The most natural and ubiquitous method of conferring affinity to a nonspecific drug for its target is by linking the drug to another molecule (often called a targeting moiety or carrier molecule) capable of specifically recognizing and binding the target site (Figure 2) .The following substances can be used as targeting entities: antibodies and their fragments, lectins, other proteins, lipoproteins, hormones, charged molecules, monosaccharides, oligosaccharides and polysaccharides and some low molecular weight ligands such as B. folic acid. Monoclonal antibodies against components characteristic of the target organ or tissue are the most commonly used carrier molecules.

Figure 2 Schematic representation of specific ligand-mediated active targeting. Nanocarriers (1) loaded with drugs (2) are modified with moieties for specific ligands (3) that recognize specific binding sites (4) on the cell surface (5). This allows the carrier to adhere to the cell surface and deliver its drug there, or it can be internalized by delivering the drug to the target cell

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Conjugating the drug directly to the targeting moiety seems to be the easiest way to produce targeted drugs. Immunotoxins provide the most illustrative example of this approach (Vitetta et al., 1983). Natural toxins can be "cleaved" into an active part (the toxic part) and a recognition part, the latter being then separated and the former bound to the antibody. As a result, the toxic moiety may be delivered only to those cells expressing the corresponding antigen (usually cancer cells), while antigen-free cells go undetected and are destroyed by the immunotoxin. In this case, however, each individual antibody molecule can only carry one active moiety. Since the toxic part of the toxin/immunotoxin is very active (only the catalytic part of the plant toxin ricin kills the cell after it enters and destroys thousands of ribosomes), immunotoxins can still be used clinically, especially For cancer therapy (Goldmacher et al. 2002; Vitetta et al. 1983). Another example of this type is the binding of different thrombolytic enzymes to antibodies that are specific for different components of the thrombus. Thus, it has been clearly demonstrated in hamsters and baboons that potent thrombosis can be achieved using a conjugate between a single-chain urokinase-type plasminogen activator and a bispecific monoclonal antibody against this activator and fibrin Dissolution (Imura et al., 1992). Data on enzyme-antibody conjugates for thrombolysis and various antibodies for direct delivery of thrombolytic therapy to the site of occlusion are abundant and well-reviewed (Haber 1994; Khaw 2002). Some attempts have also been made to target malignancies, such as human small cell lung cancer (SCLC), using direct drug-antibody conjugates. Antibodies against the proliferative compartment of mammalian squamous cell carcinoma combined with daunomycin resulted in significantly increased potency in mouse models (Ding et al., 1990). The murine monoclonal antibody NCC-LU-243 conjugated to mitomycin C was used for targeted therapy in nude mice transplanted with an antigen-positive human SCLC cell line (Kubota et al., 1992). In general, however, the loading of agents on a single targeting moiety should be much higher than a simple 1:1 ratio for the overall approach to be beneficial and practical. Alternatively, soluble or insoluble supports can be loaded with multiple active moieties, which are then additionally conjugated to targeting moieties, according to a protocol proposed by Ringsdorf in the mid-1970s (Ringsdorf 1975). A variety of reactive and biocompatible soluble polymers are available as soluble carriers, while the family of insoluble carriers includes microcapsules, nanoparticles, liposomes, micelles, and ghosts. Various depot-type systems, such as liposomes or microcapsules, have the following important advantages over other drug carriers: (b) some targeting moieties can carry multiple drug moieties loaded into the depot; (c) controlled size and permeability capabilities. So far, body compartments and conditions that have been shown to be successfully combated through various mechanisms include components of the cardiovascular system (blood pool, vessel walls, lungs, heart), reticuloendothelial system (liver and spleen); the lymphatic system (lymph nodes and lymphatic vessels), tumors, infarction, inflammation, infection and transplantation.

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Parameters that determine drug targeting effectiveness include: size of the target, blood flow through the target, number of binding sites for the target drug/excipient within the target, number and affinity of the targeting moiety. Drug carrier). particles) and the multipoint interaction of drug/drug carrier with the target.

3 Drug carrier: Liposomes and micelles are examples of many drug delivery and drug targeting systems, such as synthetic polymers, microcapsules, cells, cell shadows, lipoproteins, liposomes and micelles (Cohen and Bernstein 1996; Müller 1991) is currently under development or in development. Their use is aimed at minimizing the breakdown of the drug during administration, preventing unwanted side effects, and increasing the bioavailability of the drug and the proportion of the drug that accumulates in pathological areas. To better achieve these goals, all of the listed drug carriers can be made slowly degradable, stimulus-responsive (e.g., sensitive to pH or temperature), and targetable (e.g., Ligand binding of unique components/receptors). Out of interest). In addition, drug carriers are expected to remain in the blood for a longer period of time (Lasic and Martin, 1995; Torchilin and Trubetskoy, 1995) to maintain the desired therapeutic level of the drug in the blood over a longer period of time, thereby slowly accumulating by enhancing Permeability and retention effect (EPR) (Maeda et al. 2000; Palmer et al. 1984) at pathological sites with affected and leaky vessels (tumor, inflammatory and infarcted areas) and facilitates the targeting of specific ligands to release. Delivery of modified drugs and drug carriers to hard-to-reach areas (Torchilin 1998). Pharmaceutical excipients, especially those intended for parenteral administration, are expected to be simple and relatively inexpensive to manufacture, biodegradable, small in particle size, high in loading capacity, long in circulation time, and ideally in the body's desired Pathological site-specific or non-specific accumulation (Gref et al., 1994). The paradigm of using nanoparticle drug carriers to enhance the in vivo efficacy of many drugs, especially anticancer drugs, has been well established in pharmaceutical research and clinical settings over the past decade and does not require further evidence. Nanospheres, nanocapsules, liposomes, micelles, cytosperms, lipoproteins, and several other drug nanocarriers are widely used for experimental (and even clinical) delivery of therapeutic and diagnostic agents (Alonso 2004; Gregoriadis 1988; Müller 1991; Rolland) . 1993). Surface modification of these supports is often used to control their properties in a desired manner while allowing them to perform a variety of different functions. Key outcomes of such modifications include increased lifespan and cycle stability, altered biodistribution, targeting, and increased resistance to

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Stimulus (pH or temperature) and contrast properties. Common surface modifiers (alone or in combination) include: soluble synthetic polymers (to achieve long carrier lifetime); specific ligands such as antibodies, peptides, folic acid, transferrin, sugar moieties (to achieve target tropism); pH- or temperature-sensitive copolymers (to impart stimuli sensitivity); and chelating compounds such as EDTA, DTPA, or deferoxamine (to add diagnostic/contrast moieties to drug carriers). Clearly, different modifiers can be incorporated on the surface of the same nanoparticle drug carrier to confer useful combinations of properties (e.g., longevity and targeting, targeting and stimulus sensitivity, or targeting and contrasting properties) . Liposomes, artificial phospholipid vesicles, can be obtained from lipid dispersions in water in a number of ways. The production of liposomes, their physicochemical properties and possible biomedical applications have been discussed in detail in several monographs (Gregoriadis 2007; Lasic and Martin 1995; Lasic and Papahadjopoulos 1998; Torchilin and Weissig 2003 ; Woodle and Storm 1998). To date, many different methods have been proposed to prepare liposomes of different sizes, structures and size distributions. To increase the stability of liposomes against physiological environmental influences, cholesterol is incorporated into the liposome membrane (sometimes up to 50% mol). The size of liposomes depends on their composition and method of production and can vary from approximately 50 nm to over 1 mm in diameter. MLVs (multilamellar vesicles) range in size from 500 to 5,000 nm and consist of several concentric bilayers. LUVs (Large Unilamellar Vesicles) range from 200 to 800 nm. SUVs (small unilamellar vesicles) are approximately 100 nm (or smaller) in size and consist of a single bilayer (see Figure 3). The encapsulation efficiency of different substances also varies with liposome composition, size, charge and manufacturing method. Encapsulation of 50% or more of material from the aqueous phase into liposomes can be achieved using the reverse-phase evaporation method (Szoka and Papahadjopoulos 1980). Furthermore, various methods have been developed to obtain lyophilized liposomal formulations with good storage stability (Madden et al., 1985). The in vitro release rate of various compounds from liposomes, including medium molecular weight proteins such as lysozyme or insulin, is typically less than 1% per hour (provided the incubation temperature is sufficiently different from the phase transition temperature of a given phospholipid).

Figure 3 Liposomes can vary in size from 50 to 1000 nm. Structure and drug loading: Soluble hydrophilic drugs are trapped in the aqueous interior of liposomes (1), while poorly soluble hydrophobic drugs are located in the liposome membrane (2).

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In vivo, this parameter can vary over a wide range (minutes to hours) and depends on the composition and cholesterol content of the liposome membrane and the location of the liposome in vivo. Liposomes are biocompatible and cause little or no antigenic, pyrogenic, allergic and toxic reactions; they are readily biodegradable; they protect the host from any adverse effects of the encapsulated drug while protecting the entrapped drug Protected against inactivation by physiological media. Last but not least, liposomes are capable of delivering their contents into many cells. Various methods have been developed to deliver liposome contents into the cytoplasm (Connor and Huang 1986). According to one of these approaches, liposomes consist of pH-sensitive components that, after being endocytosed in intact form, fuse with the vacuolar inner membrane and release their contents into the cytoplasm under the effect of a pH decrease in the endosome. In addition, liposomes have been shown to fuse with micropores on the cell surface (eg, as occurs due to ischemia) (Khaw et al., 2001, 1995) and release their contents, including DNA, into the cell cytoplasm. Liposomes surface-modified with TAT peptide (Torchilin and Levchenko 2003) (or other cell-penetrating peptides such as Antp, penetratin, or synthetic polyarginine; see Torchilin 2008 for review) are also capable of transporting their cargo into cells ( Torchilin et al. 2003a). Liposomes have been considered as promising drug carriers for over two decades (Ringsdorf 1975). However, when administered intravenously, simple liposomes are very rapidly (usually within 15-30 minutes) opsonized and sequestered by cells of the reticuloendothelial system (RES) primarily from the liver (Ringsdorf 1975). From this point of view, using targeted liposomes, i. H. Liposomes with a specific affinity for the affected organ or tissue can both increase the effectiveness of liposomal drugs and reduce the loss of liposomes and their content in the RES. To obtain targeted liposomes, various methods have been developed to bind appropriate carriers (antibodies) to the liposome surface. These methods are relatively simple and allow sufficient numbers of antibody molecules to bind to the liposome surface without compromising liposome integrity and antibody affinity and specificity. Currently, more than 100 antibody molecules can be bound to a single 200nm liposome, allowing tight binding of the liposome to a single target. Conventional methods for conjugating antibodies to liposomes include covalent attachment to reactive groups on the liposome membrane and hydrophobic interactions of the protein with the membrane specifically modified with hydrophobic residues (Francis and Delgado 2000; Ringsdorf 1975) . A potentially important problem with liposomes (or other particulate drug carriers) is their inability to reach extravascular targets. Although immunoliposomes have achieved some promising results as drug carriers, the overall approach has been limited by the short circulation life of liposomes and immunoliposomes. Most antibody-modified liposomes still end up in the liver because there is not enough time between the target and the on-target liposome to interact. This is especially true when the target blood supply is reduced (area of ischemia or necrosis). even high liposome

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Affinity for the target does not lead to significant liposome accumulation because only a small amount of liposomes pass the target with the blood, while liposomes remain in circulation. When the concentration of target antigen is very low, even sufficient blood flow (and liposome passage) through the target does not result in good accumulation because there is very little "productive collision" between antigen and immunoliposomes. In both cases, better accumulation can be achieved if the liposomes are allowed to remain in the blood long enough. This is why long-circulating (often pegylated) liposomes have attracted so much attention over the past decade. Micelles represent colloidal dispersions with particle sizes ranging from 5 to 50–100 nm. At a given concentration and temperature, such colloids form spontaneously from amphiphilic or surfactants (surfactants), whose molecules consist of two distinct regions with opposite affinities for the given solvent present ( Mittal and Lindman 1991). At low concentrations, these amphiphilic molecules exist alone as monomers; however, as their concentration increases, aggregation begins at a specific concentration known as the critical micelle concentration (CMC). These aggregates, called micelles, are composed of dozens of amphiphilic molecules and are usually spherical in shape. The hydrophobic segments of amphiphilic molecules form micellar cores capable of dissolving poorly soluble drugs (Lasic 1992). This solubilization phenomenon has been extensively studied and discussed in many publications (see, for example, Attwood and Florence 1983). In an aqueous system, non-polar molecules are dissolved in the micelle core, polar molecules are adsorbed on the micelle surface, and moderately polar substances are distributed in specific intermediate positions along the surfactant molecules (Fig. 4). Polymeric micelles are typically made from amphiphilic block copolymers of hydrophilic PEG and various hydrophobic blocks. Many studies on the formation and properties of polymeric micelles (see eg Gao and Eisenberg 1993; Hunter 1991; Kabanov et al. 1992) have been published. There have been many recent reviews on various aspects of polymeric micelles preparation, physicochemical and biological properties, and possible applications as drug carriers (Adams et al., 2003; Jones and Leroux 1999; Kabanov et al., 2002a,b; Kakizawa and Kataoka, 2002); Kwon 1998, 2003; Lukyanov and Torchilin 2004; Otsuka et al. 2003; Torchilin 2001). In most cases, the amphiphilic monomers contain PEG blocks with a molecular weight of 1 to 15 kDa as the corona-forming block, and the length of the hydrophobic core-forming block is close to or slightly smaller than that of the hydrophilic block (cammas). Etc. 1997). Although some other hydrophilic polymers can be used as the hydrophilic block (Torchilin et al., 1995), PEG is still the first choice for the corona block. Meanwhile, different polymers can be used to build hydrophobic core-forming blocks: propylene oxide (Miller et al. 1997), L-lysine (Katayose and Kataoka 1998), aspartic acid (Harada and Kataoka 1998) , b-benzyl-L-aspartic acid (La et al., 1996), g-benzyl-L-glutamic acid (Jeong et al., 1998), caprolactone (Allen et al., 1998 ) and D,L-lactic acid (Hagan et al., 1996). certainly

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Figure 4 Self-assembly of amphiphilic monomers (e.g. polyethylene glycol-phosphatidylethanolamine conjugate, PEG-PE; see above) with a hydrophobic core (1) and a hydrophilic corona (2) in aqueous media micelles. In water, non-polar molecules are dissolved in the micelle core (3), polar molecules are adsorbed on the micelle surface (4), and moderately polar substances are distributed in specific intermediate positions along the surfactant molecules (5).

In this case, the starting copolymer can be prepared from two hydrophilic blocks, and then one of these blocks can be linked to a hydrophobic agent (e.g. paclitaxel, cisplatin, anthracycline, hydrophobic diagnostic entity, etc.) modification, leading to the formation of amphiphilic micellar copolymers (Katayose and Kataoka 1998; Kwon and Kataoka 1995; Trubetskoy et al. 1997). In some cases, phospholipid residues - short but extremely hydrophobic due to the presence of two long-chain fatty acyl groups - can also be used successfully as hydrophobic nucleating groups (Trubetskoy and Torchilin 1995). Compared to traditional amphiphilic polymer micelles, the use of lipid moieties as hydrophobic blocks to cover hydrophilic polymer chains (e.g. PEG) may provide additional advantages for particle stability due to the presence of two fatty acid acyl groups, which may Significantly contributes to increased hydrophobic interactions between molecular polymer chains in the micelle core. Similar to other PEG-containing amphiphilic block copolymers, diacyl lipid-PEG conjugates (e.g. PEG-phosphatidylethanolamine, PEG-PE) have been found to form very stable micelles in aqueous environments (Klibanov et al. , 1990; Lasic et al., 1991b). ). Their CMC values can be as high as 106 M (Kabanov et al. 2002a; Torchilin 2001), at least 100 times lower than conventional detergents (Rowe et al. Micelles made of these polymers retain their integrity even when highly diluted, such as in blood during therapeutic applications.

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

The high stability of polymeric micelles also allows good retention of encapsulated drugs in solution upon parenteral administration. As with any other drug, long-circulating drug carrier, three targeting mechanisms for micelles can be identified. The first is based on the spontaneous penetration of micelles into the interstitium through leaky vasculature (EPR effect), termed "passive targeting" (Gabizon 1995; Maeda et al. 2000; Palmer et al. 1984). Thus, it has been repeatedly shown that anticancer drugs incorporated into micelles (e.g. doxorubicin, see e.g. Kwon and Kataoka 1999) accumulate much better in tumors than in non-target tissues (e.g. myocardium), thereby Unwanted drug toxicity is minimized. In some cases, it is the small size of micelles that makes them superior to other nanoparticles, including liposomes. The transport efficiency and accumulation of microparticles such as liposomes and/or micelles in the tumor stroma is largely determined by their ability to penetrate the leaky tumor endothelium (Jain 1999; Yuan et al. 1995); see Fig. 1 in Schematic illustration of this phenomenon. Recently, studies have shown that diffusion and accumulation parameters are strongly dependent on the cut-off size of tumor vessel walls, and that the cut-off size varies in different tumors. Thus, the use of PEG-PE micelles for model protein drug delivery to the low-permeability solid murine tumor Lewis lung carcinoma gave the best results compared to other particle carriers (Weissig et al., 1998). The second targeting mechanism is based on the fact that many pathological processes in various tissues and organs are accompanied by local temperature increases and/or acidosis (Helmlinger et al., 1997). Micelles composed of temperature- or pH-sensitive components such as poly(N-isopropylacrylamide) and its copolymers with poly(D,L-lactide) and other blocks can disintegrate in these regions and release incorporated Drugs in micelles (Jones and Leroux 1999). Finally, specific ligands can be attached to the water-exposed ends of the hydrophilic blocks, such as antibodies and/or certain sugar moieties (Rammohan et al., 2001). In this case, the selected antibody or fragment thereof can be chemically linked to the activated, water-exposed free terminus of the hydrophilic block of the micelle-forming polymer to target the micelle without creating a steric site for the antibody. resistance. A relatively simple chemistry similar to that previously developed for liposomes (Torchilin et al., 2001a) and the use of amphiphilic PEG with protein-reactive p-nitrophenylcarbonyl (pNP) groups at the distal end of the liposome -PE can be used for this purpose including hydrophilic PEG blocks.

4 Chemical preparation of drug nanocarriers with various functions The preparation of various functional nanocarriers with controllable properties requires the incorporation of proteins, peptides, polymers, cell-penetrating moieties, reporter groups, and other functional ligands into the carrier. surface (though in some cases).

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In some cases, functional ingredients can be loaded into nanocarriers or distributed within nanocarrier structures: for example, tiny ferromagnetic particles can be loaded into liposomes or polymer nanoparticles, making them magnetic. This binding can be non-covalent, through hydrophobic adsorption of certain intrinsic or specially designed hydrophobic groups in the ligand to bind to or into the surface of the nanocarrier. Thus, amphiphilic polymers or hydrophobically modified proteins can be adsorbed to the hydrophobic surface of polystyrene nanoparticles (Yuan et al., 1995) or incorporated into the phospholipid membrane of liposomes (Torchilin 1998) or the hydrophobic core of micelles (Torchilin 2001). More commonly, attachment occurs through the interaction of reactive groups generated on the support surface with specific groups in the molecule to be attached. In the case of liposomes, the most popular drug delivery system and a practical example of the technique used, the conjugation method, is based on three very efficient and selective main reactions: the reaction between an activated carboxyl group and an amino group, resulting in an amide bond ; pyridyl dithiols react with thiols to form disulfide bonds; reactions between maleimide derivatives and thiols produce thioether bonds (Torchilin and Klibanov 1993). There are some other methods, such as the reaction of p-nitrophenyl carbonyl groups introduced on the surface of nanocarriers with amino groups of different ligands to form carbamate linkages (Torchilin et al. 2001b). A detailed review of numerous coupling methods and protocols for attaching various surface modifiers to drug carriers can be found in Klibanov et al. (2003) and Torchilin et al. (2003c). For example, it has been shown that the carboxyl groups of immunoglobulins can be activated by water-soluble carbodiimides; the activated protein can then be bound to surfaces containing free amino groups, such as PE-containing liposomes (Dunnick et al., 1975). For further attachment of ligands, appropriate reactive groups on the surface of nanocarriers can be premodified using heterobifunctional crosslinkers, such as the popular N-succinimidyl-3(2-pyridyldithio)propionate (SPDP), for the synthesis of PE derivatives , for further coupling to SH-containing proteins (Leserman et al. 1980). Another possibility relies on the reaction of the thiol groups of the ligand (protein) with the maleimide surface (phospholipid molecules in liposomes). This approach (Martin and Papahadjopoulos 1982) is one of the most widely used in research and practical applications today. Various commercially available maleimide reagents can be used to prepare maleimide-bearing phospholipids in a simple one-step process. A variety of high and low molecular weight compounds have been attached to liposomes using pyridyldithiopropionyl-PE or maleimide reagents (Klibanov et al., 2003; Torchilin et al., 2003c). It is also attractive to use free thiol groups located on immunoglobulin Fab fragments. It is believed that these SH groups are located away from the antigen-binding site, thereby allowing the nanocarrier-bound antibody fragment to retain specific interaction with the antigen. Some ligands bear carbohydrate residues that are readily oxidized to generate aldehyde groups that can react with surface amino groups to form Schiff bases (Heath et al. 1980). including nanocarriers (e.g. liposomes)

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Surface-exposed carboxyl groups have been used to attach various ligands (Kung and Redemann 1986). In the case of liposomes, they can be prepared using various techniques and activated directly with water-soluble carbodiimides prior to addition of ligand. The same chemistry can be used to link unmodified proteins and peptides to a variety of nanocarriers, including preformed liposomes containing membrane-incorporating active lipid derivatives such as N-glutaryl PE or glutaryl cardiolipin (Bogdanov et al. 1988; Weissig et al. Gregoriadis 1992; Weissig et al. 1990). Using a four-tailed hydrophobic cardiolipin derivative instead of a two-tailed PE derivative can reduce the number of amino groups involved in the coupling reaction while maintaining the same hydrophobicity. This leads to better retention of the activity of liposome-bound hydrophobins (Niedermann et al. 1991; Weissig et al. 1986). Nobs et al. describe some current methods for attaching various (mainly targeting) ligands to nanocarriers. discussed. (2004). Some specific approaches aim to attach various steric protecting polymers to the surface of nanocarriers (see below). For example, to allow PEG incorporation into liposomal membranes, reactive derivatives of hydrophilic PEG are modified at one end with a hydrophobic moiety (typically PE or residues of long-chain fatty acids attached to PEG hydroxysuccinimide). esters) (Klibanov et al. 1991, 1990). Most protocols use PEG-PE, which must be added to the lipid mixture prior to liposome formation. Alternatively, it has been proposed to synthesize single-end reactive derivatives of PEG that can be coupled to specific reactive groups (such as maleimides) on the surface of prepared liposomes, known as the post-coating method. (Maruyama et al., 1995). Currently, many studies on the preparation and properties of polymer-modified liposomes are well reviewed in several important books (Gregoriadis 1993; Lasic and Barenholz 1996; Lasic and Martin 1995). Spontaneous incorporation of PEG-lipid conjugates from PEG-lipid micelles into liposome membranes has also been shown to be very efficient and does not interfere with vesicles (Sou et al., 2000).

5 Studies of the blood lifetime of nanocarriers and their importance for drug delivery (see Cohen and Bernstein 1996; Lasic and Martin 1995; Moghimi and Szebeni 2003; Torchilin 1996b, 1998; Trubetskoy and Torchilin 1995). Drugs and drug carriers are manufactured and circulated for a number of important reasons. One is to maintain desired levels of agents (therapeutic and diagnostic) in the blood over an extended period of time

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time interval. Long-circulation diagnostics are essential for blood count imaging, which helps assess the current state of blood flow and reveal its irregularities caused by pathological lesions. Blood substitutes represent another important area of use for long-circulating medicines if artificial oxygen carriers should remain in circulation long enough (Winslow et al., 1996). Then, as mentioned above, long-circulating drug-containing microparticles or macromolecular aggregates can slowly accumulate at pathological sites with affected and leaky blood vessels (mainly tumors) (“passive” targeting: Maeda 2001; Maeda et al. 2000 ) and improve or they improve drug delivery in these regions (Gabizon 1995; Maeda 2001; Maeda et al. 2000). In addition, the extended cycle facilitates better targeting of targeted (specific ligand-modified) drugs and drug carriers, allowing them to interact with the target for a longer period of time due to the higher number of passages of targeted drugs (Torchilin 1996b). Goals. Chemical modification of drugs and drug carriers with certain synthetic polymers is the most common way to make drugs and other functions of drug carriers more durable in the body. Hydrophilic polymers have been shown to protect individual molecules and solid particles from interacting with various solutes. This phenomenon is related to the stability of various aqueous dispersions (Molyneux 1984), and in the field of pharmaceuticals, helps to protect active ingredients and drug carriers from adverse interactions with components of the biological environment. The term "steric stabilization" was introduced to describe the phenomenon of polymer-mediated protection (Naper 1983). The most popular and successful approach to obtain long-circulating, biostable nanoparticles is to coat them with certain hydrophilic and flexible polymers, most notably polyethylene glycol (PEG), as first proposed for Liposome-based methods (Allen et al., 1991; Klibanov et al., 1990; Maruyama et al., 1991; Senior et al., 1991). At the biological level, coating nanoparticles with PEG sterically hinders the interaction of blood components with their surface and reduces the binding of plasma proteins to PEG particles, as shown by liposomes (Allen 1994; Chonn et al. 1991 , 1992; Lasic et al. 1991a, Senior et al. 1991, Woodle 1993). This prevents the interaction of the drug carrier with the opsonins and slows their rapid uptake by the RES (Senior 1987; Figure 5). Mechanisms preventing PEG opsonization include surface shielding

Fig. 5 The steric protection mechanism of drug nanocarriers by surface grafted polymers. For example, PEG chains (1) on the surface of liposomes prevent opsonins (2) from attaching to liposomes and allowing them to circulate longer

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Charge, increased surface hydrophilicity (Gabizon and Papahadjopoulos 1992), enhanced repulsive interactions between polymer-coated nanocarriers and blood components (Needham et al. It is also impermeable to other solutes at low polymer concentrations (Gabizon and Papahadjopoulos 1992; Torchilin et al. 1994). Although many polymers have been attempted as steric protectors for nanoparticle drug carriers (Torchilin and Trubetskoy 1995), which will be discussed in more detail below, most studies on long-circulating drugs and drug carriers have used PEG as the With a very attractive combination of properties, steric protective polymers: excellent solubility in aqueous solution and ability to bind many water molecules, high flexibility of its polymer chains, very low toxicity, immunogenicity and antigenicity non-accumulation in RES cells and minimal effect on specific biological properties of modified drugs (Pang 1993; Powell 1980; Yamaoka et al. 1994; Zalipsky 1995). Just as importantly, PEG is not biodegradable and does not subsequently form any toxic metabolites. On the other hand, PEG molecules with a molecular weight below 40 kDa are easily excreted through the kidneys. In fact, PEG is readily available commercially in various molecular weights. PEG is commonly used to modify drugs and drug carriers and has a molecular weight range of 1,000 to 20,000 Da. Single-end reactive (half-helical) PEG derivatives are often used to modify pharmacologically important substances without forming cross-linked aggregates and heterogeneous products. There are many chemical methods to synthesize activated derivatives of PEG and to couple these derivatives to various drugs and drug carriers. These methods and their suitability for solving various drug delivery problems have been extensively reviewed by several authors (Torchilin 2002; Veronese 2001; Zalipsky 1995). Although PEG conjugation chemistry is well developed, there is active work in finding alternative sterically protective polymers. These polymers should be biocompatible, soluble, hydrophilic and have a highly flexible backbone (see Chonn et al. 1992; Lasic et al. 1991a; Maruyama et al. 1994; Takeuchi et al. 1999; Torchilin 1996b; Torchilin al. 2001b, 1995; Torchilin and Trubetskoy 1995; Trubetskoy and Torchilin 1995; Woodle et al. 1994). The most important biological consequences of modifying nanocarriers with protective polymers are a significant increase in their circulation time and a reduction in RES accumulation (in the liver) (Klibanov et al. 1990; Lasic and Martin 1995; Torchilin et al. 1994) . Most importantly, from a clinical standpoint, a variety of long-circulating, relatively small (100-200 nm) liposomes have been shown to efficiently accumulate in many tumors through an "impaired filter" mechanism (Gabizon and Papahadjopoulos 1988; Gabizon 1995) ).; Maeda 2001; Maeda et al. 2000). Thus, PEG-coated and other long-circulating liposomes containing various anticancer drugs such as doxorubicin, arabic Furanosylcytosine, doxorubicin, and vincristine (Allen et al., 1992; Boman et al., 1994; Gabizon et al., 1994). ; Huang et al. 1994). The most successful incorporation of doxorubicin into PEG liposomes has proven to be very good

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Clinical outcomes (Ewer et al 2004; Gabizon 1995; Rose 2005). Allen analyzed the pharmacokinetics of long-circulating nanocarriers using PEG liposomes (Allen et al. 1995b). In general, conjugation of drugs to nanocarriers has significant effects on pharmacokinetics: delayed drug absorption, restricted drug biodistribution, decreased volume of drug biodistribution, delayed drug clearance, and delayed drug metabolism (Hwang 1987). All these effects are due to impaired interstitial penetration of the drug and reduced drug accessibility to the biological environment due to entrapment in the drug carrier. The presence of protective polymers on the carrier surface further alters all these parameters (Klibanov et al., 1990; Senior et al., 1991). Thus, while "simple" liposomes exhibit nonlinear, saturable kinetics, long-circulating liposomes exhibit dose-dependent, non-saturated, and log-linear kinetics (Allen and Hansen 1991; Huang et al. 1992; Mayhew et al. 1992). All pharmacokinetic effects depend on the route of administration of liposomes, their size and composition, and are always less pronounced in sterically protected PEG carriers (Allen et al., 1989; Liu et al., 1991 , 1992; Maruyama et al., 1992). Additional functionality can be added to long-circulating PEGylated drug carriers, allowing the detachment of PEG chains in response to certain local stimuli characteristic of pathological areas, such as B. pH reduction or temperature commonly observed in inflamed and tumor areas raised. The problem is that the stability of PEGylated nanocarriers may not always be favorable for drug delivery. In particular, when drug-containing nanocarriers accumulate in tumors, they may not be able to release the drug to kill tumor cells. When the carrier must be taken up by the cell via the endocytic route, the presence of its surface PEG coating also prevents the contents from leaving the endosome and being released into the cytoplasm. To address these issues, for example in long-circulating liposomes, chemical methods were developed to detach PEG from lipid anchors under desired conditions. In the field of controlled drug release, labile bonds are known to degrade only under the acidic conditions of endocytic vacuoles or acidotic tumor masses. Such linkages can be based, for example, on acid-labile diorthoester chemistry (Guo and Szoka 2001) or vinyl ester chemistry (Boomer and Thompson 1999). The latter reference describes the preparation of PEG lipids that can be cleaved in acidic media. Cysteine-cleavable lipopolymers have also been described (Zalipsky et al., 1999). If the PEG brush is cleaved (e.g., from the liposome surface), membrane destabilization should occur and the liposome contents will be delivered to their target (e.g., by escaping the primary endosome into the cytoplasm). Polymer compositions with pH-sensitive (pH-cleavable) linkages are widely used to create stimuli-responsive drug delivery systems that are stable in the blood stream or normal tissues. However, they acquire the ability to degrade and release entrapped drugs in lower pH body regions or cellular compartments such as tumors, infarcts, areas of inflammation or cytoplasm or endosomes (Roux et al., 2002a, 2004 ; Simoes et al., 2004) . ). As the pH of the "acidic" site drops from the normal physiological value of 7.4 to pH 6 and below, chemical bonds have been used to prepare acidic pH-sensitive carriers, including vinyl esters, diesters ,

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

and hydrazones, which are fairly stable around pH 7.5 but hydrolyze relatively rapidly at pH 6 and below (Guo and Szoka 2001; Kratz et al. 1999; Zhang et al. 2004). To date, a variety of liposomes (Leroux et al. 2001; Roux et al. 2002b) and micelles (Lee et al. 2003a,b; Sudimack et al. 2002) comprising components with such linkages have been described. and various drug-releasing drug conjugates such as doxorubicin (Jones et al., 2003), paclitaxel (Suzawa et al., 2002), doxorubicin (Potineni et al., 2003; Yoo et al. , 2002) and DNA (Cheung et al. 2001; Venugopalan et al. 2002) in acidic cellular compartments (endosomes) and pathological body regions under acidosis. New isolatable PEG conjugates were also reported by Zalipsky et al. describe. (1999), where the separation process is based on the mild thiolysis of the dithiobenzylcarbamate linkage between PEG and an amino-containing substrate such as phosphatidylethanolamine. Serum-stable, long-circulating PEGylated pH-sensitive liposomes have also been prepared using a combination of PEG and pH-sensitive, terminally alkylated copolymers of N-isopropylacrylamide and methacrylic acid on the same liposome (Roux et al., 2004). Attachment of pH-sensitive polymers to the liposome surface can promote liposome destabilization and drug release in the pH-reduced compartment. Extensive in vitro and in vivo experiments have shown great potential to improve drug delivery and targeting efficiency.

6 Passive accumulation of liposomes and micelles in tumors Since it has been repeatedly shown that long circulating liposomes such as macromolecules can accumulate in various pathological regions of the involved vasculature through the EPR effect (Maeda et al. 2001; Yuan et al. 1994). ) Long-circulating polymer (PEG)-coated liposomes have been repeatedly used to deliver drugs into tumors by passive accumulation. An important feature of protective polymers is their flexibility, which allows a relatively small number of surface-grafted polymer molecules to form an impermeable layer on the liposome surface (Torchilin et al., 1994; Torchilin and Trubetskoy, 1995) . Although PEG remains the gold standard for liposome steric protection in passively targeted formulations, attempts are ongoing to identify other polymers that can be used to generate long-circulating liposomes. Torchilin and Trubetskoy (1995) and Woodle (1998) review previous studies on various water-soluble flexible polymers. Recent work describes the use of poly[N-(2-hydroxypropyl)methacrylamide)] (Whiteman et al., 2001), poly-N-vinylpyrrolidone (Torchilin et al., 2001b), and biodegradable Long-circulating liposomes prepared from L-amino acid-based polymers. Lipid conjugates (Metselaar et al., 2003) and polyvinyl alcohol (Takeuchi et al., 2001). As previously described, long-circulating liposomes exhibit dose-dependent, non-saturating, log-linear kinetics and increased bioavailability (Allen and Hansen 1991). studied the relative role of liposome charge and protective polymer molecular size, showing that opsonins have different

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Figure 6 Schematic diagram of the structure of Doxil1-doxorubicin in PEG-coated liposomes

Molecular size may be related to the clearance of liposomes containing differently charged lipids (Levchenko et al., 2002). After PEG-liposomes accumulate at the target site through the EPR effect, PEG is also removably attached to the liposome surface to facilitate cellular uptake of the liposomes (Maeda et al., 2001) and the PEG coating is exposed to topical Under pathological conditions (lower tumor pH). Zalipsky et al. describe new cleavable PEG conjugates. describe. (1999), where the separation process is based on the mild thiolysis of the dithiobenzylcarbamate linkage between PEG and an amino-containing substrate such as PE. Low pH degradable PEG-lipid conjugates based on hydrazone bonds between PEG and lipids have also been described (Kale and Torchilin 2007; Sawant et al. 2006). 1 1 Doxorubicin (Doxil and Caelyx; see schematic diagram in Figure 6) in PEG-coated liposomes has been successfully used to treat solid tumors in patients with metastatic breast cancer and subsequently improved survival (O'Shaughnessy 2003; Perez et al. 2002; Symon et al. 1999). Combination therapy with liposomal doxorubicin and paclitaxel (Schwonzen et al., 2000) or Doxil/Caelyx and carboplatin (Goncalves et al., 2003) targets the same indications. Caelyx is also currently in Phase II studies in patients with head and neck squamous cell carcinoma (Harrington et al., 2001) and ovarian cancer (Johnston and Gore, 2001). Clinical data show that doxorubicin in PEG liposomes is effective against inoperable hepatocellular carcinoma (Schmidinger et al., 2001), cutaneous T-cell lymphoma (Wollina et al., 2003) and sarcoma (Skubitz, 2003) With impressive effect. For a review of the current successful use of Caelyx in ovarian cancer, see Perez-Lopez et al. (2007). However, it should be noted here that recent findings suggest that PEG-liposomes, previously considered biologically inert, can still trigger certain side effects by activating the complement system (Moein Moghimi et al. 2006; Moghimi and Szebeni Year 2003). With the increased pH-sensitive function of liposomal formulations, various methods have been developed to deliver liposomal contents into the cytoplasm (Torchilin 1991). If the liposome consists of pH-sensitive components, after endocytosis, it fuses with the vacuolar inner membrane, destabilizing it and releasing its contents into the cytoplasm (Torchilin et al. al. 1993). Therefore, endosomes become

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Enters the cytoplasm from the outside (Sheff 2004). This approach has been discussed several times in various publications (2004, in Advanced Drug Delivery Reviews #56, a special issue edited by J.C. Leroux devoted to endosomal escape via pH-sensitive drug delivery systems). It is generally believed that in the endosome, low pH and some other factors destabilize the liposome membrane, which in turn interacts with the endosomal membrane, leading to its secondary destabilization and drug release into the cytoplasm. The presence of fusogenic lipids in liposome compositions, such as unsaturated DOPE (dioleoyl-sn-glycero-3-phosphatidylethanolamine), is often required to confer liposome pH sensitivity (Shalaev and Steponkus 1999). Despite reduced pH sensitivity, multifunctional, long-circulating PEGylated DOPE containing pH-sensitive liposomes efficiently releases their contents into the cytoplasm after passive accumulation in tumors (Varga et al., 2000 ). Antisense oligonucleotides (ODNs) have been delivered into cells by anionic pH-sensitive PE-containing liposomes, which are stable in blood but undergo a phase transition at acidic endosomal pH and promote oligonucleotide Acid is released into the cytoplasm (Fattal et al., 2004). More recently, new pH-sensitive liposome additives have been described, including oleyl alcohol (Sudimack et al., 2002) and pH-sensitive morpholine lipids (monostearyl derivatives of morpholine) (Asokan and Cho 2003). In the case of micellar anticancer drug formulations, the passive targeting of micelles to pathological organs or tissues can further enhance the efficacy of micellar-encapsulated drugs. A direct correlation between the circulation lifetime of particulate drug carriers and their ability to reach target sites has been observed several times (Gabizon 1995; Maeda et al. 2001). The results of the blood clearance studies of the various PEG-PE micelles clearly show their longevity: the investigated micellar formulations have a circulating half-life of 1.2 to 2.0 hours in mice, rats and rabbits, depending on the size of the PEG block. Molecular size (Lukyanov). et al. 2002). Increasing the size of the PEG block increases the circulation time of the micelles in the blood, possibly by providing better steric protection against the penetration of opsonins into the hydrophobic micelle core. However, the circulation time of PEG-PE micelles is slightly shorter than that of PEG-coated long-circulating liposomes (Klibanov et al., 1990), which can partly be explained by the fact that micelles, compared with The rate of extravasation was faster and significantly smaller (Weissig et al. 1998). Slow dissociation of micelles under physiological conditions due to continuous clearance of monomers may also play a role by shifting the micelle-monomer equilibrium towards monomer formation (Trubetskoy et al., 1997). Like long-circulating liposomes (Gabizon 1992, 2001; Papahadjopoulos et al. 1991), PEG-PE-based micelles formed from PEG750-PE, PEG2000-PE, and PEG5000-PE efficiently accumulate in tumors through the EPR effect. Notably, micelles prepared with several different PEG-PE conjugates accumulated much higher in tumors compared to non-target tissues (muscle), even in mice with experimental Lewis lung cancer (LLC ), from which they are known to have a relatively small truncated size of vasculature (Hobbs et al. 1998; Weissig et al. 1998).

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In other words, due to their smaller size, micelles may have an additional advantage as a tumor drug delivery system utilizing the EPR effect compared to particle carriers with larger individual particles. A model protein (soybean trypsin inhibitor or STI, MW 21.5 kDa) organized in micelles accumulated more in subcutaneously established mouse Lewis lung carcinoma than the same protein in larger liposomes (Weissig et al., 1998 Year). The accumulation pattern of PEG-PE micelles prepared from all versions of PEG-PE conjugates was characterized by a maximum tumor accumulation time of approximately 3-5 hours. For micelles formed from monomers with relatively large PEG blocks (PEG5000-PE), the maximum total tumor uptake (as AUC) of the injected dose was found 5 h after injection. This can be explained by the fact that these micelles have the longest circulation time and less extravasation into normal tissues than micelles made from smaller PEG-PE conjugates. However, micelles made from PEG-PE conjugates with shorter PEG versions may be more effective carriers of poorly soluble drugs, as they have larger hydrophobic phases compared to hydrophilic ones, and can be more efficiently loaded Medication (based on weight). Similar results were obtained in another mouse tumor model, EL4 T-cell lymphoma (Lukyanov et al., 2002). Some other recent data also clearly show that PEG-PE based micelles spontaneously target other experimental tumors in mice (Torchilin et al. 2003b) as well as damaged cardiac regions in rabbits with experimental myocardial infarction (Lukyanov et al. 2004b) . ). Among the drugs delivered passively targeting micelles is paclitaxel, which has been shown to accumulate in tumors when incorporated into PEG-b-poly(4-phenyl 1-butyrate)-l-micelle Much better than its commercial 1-formulation paclitaxel. Aspartamide conjugate (Hamaguchi et al., 2005). With this formulation, AUC was increased nearly 100-fold, volume of distribution was reduced 15-fold, and drug clearance was significantly reduced, resulting in a 25-fold increase in drug accumulation in C-26 tumors in mice with a corresponding increase in antitumor activity. Several other micellar formulations for passive targeting of paclitaxel have also been tested with varying degrees of success (Hamaguchi et al., 2005; Kim et al., 2004). PEG-b-poly(amino acid)-based micelles loaded with cisplatin (CDDP) have been developed for passive drug delivery into tumors and are in clinical trials (Uchino et al., 2005). Among other micellar formulations used in clinical trials for passive drug targeting, mention can also be made of doxorubicin in PEG-block poly(L-aspartic acid)-doxorubicin conjugate micelles ( These micelles contained free and hydrophobic blocks conjugated to drug) (Matsumura et al., 2004)) and doxorubicin 1-cin in Pluronic micelles (Danson et al. 2004). Another targeting mechanism is based on the fact that many pathological processes in various tissues and organs are accompanied by local temperature increases and/or acidosis (Vutla et al., 1996; Yerushalmi et al., 1994). Therefore, the efficiency of micellar carriers can be further enhanced by subjecting micelles to cleavage at elevated temperatures or lowering the pH at pathological sites, ie. H. By combining the EPR effect with the response to the stimulus.

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

To this end, micelles are prepared from temperature- or pH-sensitive components such as poly(N-isopropylacrylamide) and its copolymers with poly(D,L-lactide) and other blocks, and possess Ability to site disintegration. Drug release incorporated into micelles (Cammas et al. 1997; Chung et al. 1998; Kohori et al. 1998; Kwon and Okano 1999; Meyer et al. 1998). Phthalocyanine-loaded pH-responsive polymeric micelles appear to be promising carriers for photodynamic cancer therapy (Le Garrec et al., 2002), while doxorubicin-loaded polymeric micelles with acid-cleavable linkages can enhance Intracellular drug delivery to tumor cells, thereby increasing efficiency (Yoo et al. 2002). It has been shown that temperature-responsive polymer micelles exhibit increased drug release upon temperature change (Chung et al., 1999). Passively targeting micelles (polymeric micelles) can also exhibit pH sensitivity and the ability to escape from endosomes. Thus, micelles prepared from PEG poly(aspartate-hydrazone-doxorubicin) readily release the drug at the lower pH values typical of endosomes, facilitating its cytoplasmic delivery and toxicity to cancer cells (Bae et al. People, 2005). Alternatively, micelles for intracellular delivery of antisense ODN were prepared from ODN-PEG conjugates complexed with the cationic fusion peptide KALA and enabled higher intracellular delivery of ODN than free ODN (Jeong et al., 2003) . One can also improve the intracellular delivery of drug-loaded micelles by adding in their composition lipid components for membrane destabilizing lipofection. For example, PEG lipid micelles have a net negative charge (Lukyanov et al., 2004b), which may hinder their internalization by cells. On the other hand, it is well known that a net positive charge generally facilitates cellular uptake of various nanoparticles, and after endocytosis, drug/DNA-laden particles can escape from endosomes and enter the cytoplasm through interfering interactions . Cationic lipids with endosomal membranes (Hafez et al 2001). Compensating for this negative charge by adding positively charged lipids to PEG-PE micelles can enhance the uptake of drug-loaded mixed PEG-PE/positively charged lipid micelles by cancer cells. It is also possible that, following increased endocytosis, this micelle leaves the endosome and enters the cytoplasm of cancer cells. In this context, an attempt was made to increase the intracellular delivery of paclitaxel micelles by preparing paclitaxel-containing micelles from a mixture of PEG-PE and 1 Lipofectin lipid (LL), thereby increasing its anticancer activity (Wang et al., 2005 Year). ). Lee et al. also reported multifunctional polymeric micelles that were capable of pH-dependent dissociation and drug release when loaded with doxorubicin and supplemented with biotin as an interaction ligand for cancer cells. describe. (2005). Problems with drug delivery using micelles for passive targeting are often related to excessive drug release from micelles and difficulties in intracellular drug delivery (Aliabadi and Lavasanifar 2006). To minimize drug release from micelles, the drug can be chemically bound to hydrophobic blocks of the micelle-forming components, or the drug-loaded micelles can be additionally chemically cross-linked (Kang et al. 2005; Lavasanifar et al. 2002 ; Shuai et al. 2004; Yuan et al. 2005).

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7 Active Tumor Targeting of Drug-Loaded Liposomes Current development of liposome carriers often involves attempts to combine the properties of long-circulating liposomes and targeting liposomes in one formulation (Abra et al. 2002; Blume et al. 1993; Torchilin et al. 1992) To achieve better selectivity of PEG-coated liposomes, it is advantageous to attach targeting ligands via PEG spacer arms so that the ligands protrude beyond the dense PEG brushes and eliminate The steric hindrance of ligand binding to the target. Various advanced techniques are used for this purpose, and the targeting moiety is usually attached to the protective polymer layer by coupling it to the water-exposed terminus of the activated liposome-grafted polymer molecule (Blume et al. , 1993; Torchilin et al. 2001a). ), see Figure 7. PEG-lipid conjugates for liposomes and other drug nanocarriers for steric protection and for the preparation of polymeric micelles are derived from methoxy-PEG (mPEG) with non-reactive methoxy termini group, so several attempts have been identified to functionalize the PEG tip in PEG-lipid conjugates. To this end, several end-functionalized lipopolymers of the general formula X-PEG-PE were introduced (Zalipsky 1995; Zalipsky et al. 1998), where X contains a reactive functional group and PEG-PE represents a conjugate of PE and polyethylene glycol. An interesting approach for coupling different ligands, such as B. antibodies against liposomes, including PEGylated liposomes, involves the so-called "post-insertion" technique (Ishida et al., 1999). The technology is based on the initial activation of the ligand with any reactive PEG-PE derivatives, followed by the formation of modified ligand-PEG-PE conjugates with preformed drug-loaded simple or PEGylated liposomes. Co-incubation of stable micelles. Finally, the modified ligands spontaneously integrate from their micelles into the more thermodynamically favorable environment of the liposome membrane. In particular, this method was used to prepare Immuno-Doxil via carbonyl modification of p-nitrophenyl

Figure 7 Conjugation of targeting moieties (mainly monoclonal antibodies) to PEGylated drug nanocarriers (liposomes as an example). Although targeting ligands can be immobilized on surfaces together with PEG (1), targeting moieties are usually attached to protective polymers by coupling them to the water-exposed ends of activated liposome-grafted polymer molecules. on the physical layer ( 2 ).

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

(pNP)-PEG-PE modified anticancer 2C5 monoclonal antibody (Elbayoumi and Torchilin 2007; Lukyanov et al. 2004a). Since antibodies are the most diverse and widely used specific ligands for experimental drug-loaded liposome-targeted chemotherapy of various tumors, there are several original papers and reviews on antibody-specific drug-loaded liposomes in cancer ( For reviews see Kontermann 2006; Park et al. 2004; Sapra and Allen 2003); Sofou and Sgouros 2008; Torchilin 1996a, 2000; Vingerhoeds et al. 1994). "First-generation" antibody-modified liposomes have been used to assess certain parameters of their interaction with target cells in vitro (Klibanov et al., 1985) and in vitro and in vivo against certain model and real targets. Liposome targeting. Examples are extracellular matrix antigens or areas of myocardial infarction (Chazov et al., 1981; Torchilin et al., 1985). Importantly, it has been found that modification of antibody-carrying liposomes with PEG (to allow them to circulate for long periods of time) often results in reduced conjugation efficiency because surface-bound antibodies are sterically shielded by liposome-grafted PEG (Klibanov et al. 1991 ; Torchilin et al. 1992). This eventually led to the development of several methods for attaching antibodies to the surface of the PEG layer in PEGylated liposomes. In general, antibody conjugation shortens liposome circulation time, as modified liposomes are increasingly taken up by circulating macrophages or hepatic macrophages via Fc receptors, or liposome-labeled antibodies Molecules are opsonized (Allen et al. 1995a; Kamps and Scherphof 1998). Whole antibodies can also induce complement-mediated and antibody-dependent cytotoxicity (Sapra and Allen 2003). These effects can be minimized by using antibody Fab fragments rather than whole antibodies (Flavell et al. 1997). Although Fab fragments can also accelerate liposome clearance (Maruyama et al., 1997), Fab liposomes generally circulate much longer than fully antibody-modified liposomes (Maruyama et al., 1997). In the case of antibody-modified PEGylated liposomes, even some reduction in circulation time allows sufficiently long circulation for good target accumulation. Of course, care should be taken not to modify PEG liposomes so severely with antibodies that their longevity is severely affected. Interestingly, tumor accumulation of antibody-modified long-circulating liposomes was in some cases comparable to that of non-antibody-conjugated long-circulating liposomes (Moreira et al. 2001; Park et al. 2002, 1997, 2001). However, the therapeutic activity of antibody-targeted liposomes is higher. As described by Kirpotin et al. (2006) using PEGylated liposomes modified or unmodified with anti-HER2 antibodies. Although intratumoral accumulation was similar for both formulations, the antibody-modified formulation was more readily internalized by tumor cells, allowing higher drug doses to enter the cancer cells, ie. H. z Kill cancer cells more effectively. However, in some other cases liposome internalization does not appear to be critical. Thus, Sapra and Allen (2004) showed that PEGylated liposomes were loaded with vincristine or doxorubicin and modified (or unmodified) with antibodies against internalized CD19 antigen or non-internalized CD20 antigen

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The therapeutic effect was shown to depend more on the type of drug used than on its ability to internalize. As expected, the cytotoxicity of target liposomes also depends on the rate of drug release from the liposomes (Allen et al., 2005). An interesting phenomenon was reported by Hosokawa et al. describe. (2003) whose authors showed that while non-targeted doxorubicin-containing liposomes were toxic to a variety of cancer cells to a degree that reflected the sensitivity of the cells to the drug, the cytotoxicity of antibody-targeted liposomes was associated with Surface area antigen density is proportional to what the liposome is targeting. The critical antigen surface concentration is approximately 4104 sites per single cell, beyond which further increases in antigen density are no longer critical. Similar observations were made by Lopes de Menezes et al. production. (1998) and Park et al. (2002). Since cancer cells are often very heterogeneous in the antigens they express, Sapra and Allen (2003) proposed the use of combinations of antibodies against different antigens on a single liposome for better and more uniform targeting of all cells within a tumor, allowing . Alternatively, it is also possible to rely on the "bystander" effect (Sapra and Allen 2003), ie. H. Effects of drugs released from liposomes attached to specific cancer cells on neighboring cancer cells that do not have similar receptors. One antibody that is gaining popularity in the fight against cancer is the monoclonal antibody against HER2, an antigen that is often overexpressed on various cancer cells. Anti-HER2 monoclonal antibodies, including the humanized and currently clinically used Herceptin antibody, have been used to create drug-loaded liposomes (long-circulating liposomes) specific for HER2-positive cancer cells (Kirpotin et al. , 1997, 2006; Park et al., 1995), 1997, 2001; Yang et al 2007). This antibody has been successfully used to deliver doxorubicin in simple and long-circulating liposomes to mice Breast tumor xenografts, thereby significantly increasing the therapeutic activity of the drug. It has been shown that PEGylated liposomes labeled with anti-HER2 antibodies are efficiently endocytosed by HER2-positive cancer cells, leading to better accumulation of the drug (doxorubicin) in the tumor cells and thus better treatment outcome. In the case of antibody-targeted Doxil, more was detected in cancer cells compared to doxorubicin in 1 simple pegylated liposome (Doxil) that normally accumulates in the tumor interstitial space Drug molecules, ie. H. Targeting with antibodies increases internalization of the drug by target cells. Another promising antibody against tumors with drug-loaded liposomes is the monoclonal antibody against the CD19 antigen, which is also frequently overexpressed on various cancer cells. Anti-CD19 antibody-modified liposomes loaded with doxorubicin showed significantly improved targeting and therapeutic efficacy in vivo and in vitro in mice bearing human CD19+-B lymphoma cells (Lopes de Menezes et al., 1998). Similar results were obtained with doxorubicin-loaded liposomes modified with antibodies against the internalizable C19 antigen and against the non-internalizable CD20 antigen (Sapra and Allen 2004). Anti-CD19 antibodies have also been used to target doxorubicin-loaded liposomes with variable drug release rates to experimental tumors (Allen et al., 2005). Recently, successful attempts have been made to target long-circulating liposomes loaded with doxorubicin

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Cancer cells expressing CD19 have the single-chain Fv fragment of the CD19 antibody (Cheng and Allen 2008; Cheng et al. 2007). Since neuroblastoma cells often overexpress the disialoganglioside GD2, antibodies against GD2 and its Fab0 fragment have been proposed to target drug-loaded liposomes to the corresponding tumors (Brignole et al., 2003; Pastorino et al. People, 2006, 2003). Covalent coupling of the Fab0 fragment of an anti-GD2 antibody to doxorubicin-loaded long-circulating liposomes enables increased binding to target cells and higher cytotoxicity in vitro and in vivo, including in nude mice and in metastatic models in human tumor models. GD2-targeted immunoliposomes containing the novel antineoplastic drug fenretinide induce apoptosis in neuroblastoma and melanoma cell lines and have been shown to be effective against neuroblastoma in mice in vitro and in vivo activity (Raffaghello et al., 2003). Combination of doxorubicin-loaded PEGylated liposomes with anti-GD2 and NGR peptides that specifically bind tumor vasculature resulted in enhanced therapeutic effects by acting on tumor cells and tumor vessels (Pastorino et al., 2006). An interesting new target for antitumor drug delivery via targeted liposomes is membrane type 1 matrix metalloproteinase (MT1-MMP), which plays an important role in tumor neovascularization and exists in tumor cells and neovascularization. Angiogenic on endothelial cells and overexpressed. Modification of doxorubicin-loaded long-circulating liposomes with anti-MT1-MMP antibody resulted in increased uptake of target liposomes in MT1-MMP-overexpressing HT1080 fibrosarcoma cells in vitro In vivo more effective inhibition of tumor growth-loaded PEGylated liposomes (Hakeyama et al., 2007). Anti-MM1-MMP antibodies have been shown to enhance endocytosis of drug-loaded liposomes, thereby increasing their cytotoxicity (Atobe et al., 2007). A strong effect of this preparation on tumor endothelial cells was noted. Epidermal growth factor receptor (EGFR) and its variant EGFRvIII can serve as valuable targets for intracellular drug delivery in tumor cells that overexpress these receptors. Conjugation of the Fab0 fragment of the C225 mAb that binds both EGFR and EGFRvIII and the scFv fragment of an EGFR-only mAb to drug-loaded liposomes significantly enhanced the binding of such targeted liposomes to cancer cells-Expression Appropriate recipients, such as U87 glioma cells and A0431 and MDA-MB-468 cancer cells. Better binding leads to increased internalization and increased cytotoxicity (Mamot et al., 2003). In vivo treatment with such targeted drug-loaded liposomes (doxorubicin, epirubicin, and vinorelbine as drugs) always suppressed tumor growth better than treatment with non-targeted liposomal drugs (Mamot et al., 2005). The Fab0 fragment derived from the humanized anti-EGFR monoclonal antibody EMD72000 has been shown to efficiently deliver liposomal drugs intracellularly to colorectal tumor cells (Mamot et al., 2006). The authors of this study also showed that attaching targeting moieties to PEGylated liposomes requires sufficient spacer arm length to overcome possible steric shielding by antibody fragments by sterically protecting the PEG chains. An interesting approach to designing proteins against EGFR

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Pan and Lee (2007) proposed liposomes in which an anti-EFGR antibody (cetuximab or C225) was covalently linked to a folate-binding protein via a thioester bond and then coupled to preformed folate-containing liposomes. Cetuximab liposomes loaded with boron derivatives for boron neutron capture therapy have also been prepared using cholesterol-based anchoring and micellar transfer techniques (Pan et al., 2007). Various extracellular matrix proteins expressed on the surface of cancer cells have also been used as targets for antibody-mediated liposomal drug delivery. Thus, b1 integrin expressed on the surface of human non-small cell lung cancer was targeted by doxorubicin-loaded liposomes modified with the Fab0 fragment of an anti-b1 integrin monoclonal antibody (Sugano et al., 2000 Year). Treatment of SCID mice with lung tumor xenografts with this liposome resulted in significant tumor growth inhibition and also suppressed metastasis compared to all controls. The idea of targeting different antigens, preferably internalizable antigens, by antibody-liposome conjugates on endothelial cells was validated long ago (Trubetskaya et al., 1988). However, this approach has not really received attention until the last few years. Thus, liposomes modified with anti-E-selectin antibodies were successfully internalized by activated endothelial cells in vitro by E-selectin-mediated endocytosis (Asgeirsdottir et al., 2008). Another potential target for antibody-mediated cancer therapy using drug-loaded liposomes is epithelial cell adhesion molecule (EpCAM), which is expressed in many tumors but not in normal cells (Hussain et al., 2007 ). EpCAM-targeting immunoliposomes were generated by covalently binding the humanized scFv fragment of the monoclonal antibody 4D5MOCB to the surface of pegylated doxorubicin-loaded liposomes and were shown to bind to EpCAM-positive cancer cells Binding, internalization and cytotoxicity were significantly improved. Likewise, liposomes conjugated to vascular cell adhesion molecule-1 (VCAM-1) antibodies effectively target activated endothelial cells that overexpress VCAM-1 (Voinea et al., 2005). Liposomes loaded with cytotoxic drugs have also been targeted to ED-B fibronectin using the scFv fragment of the corresponding antibody (Marty and Schhrender 2005). Proliferating endothelial cells are challenged with doxorubicin-loaded liposomes modified with scFv fragments of anti-endocrine antibodies overexpressed on such cells (Volkel et al., 2004). The lipid-based drug carrier is also conjugated to an antibody (or a fragment thereof) against the transferrin receptor (TfR), which is often overexpressed on the surface of various cancer cells. For example, liposome-incorporated maleimide-modified PEG2000-PE molecules of such vectors were modified with the anti-TfR monoclonal antibody OX26 and showed strong binding to TfR-overexpressing cells (Beduneau et al. People, 2007). The same antibody was non-covalently bound to daunomycin-loaded liposomes via an avidin-biotin pair, and the modified liposomes were shown to in multidrug-resistant RBE4 brain capillary endothelial cells in vitro and in vivo There is a good accumulation of (Schnyder et al., 2005). Liposomes loaded with the lipophilic prodrug 5-fluorodeoxyuridine and modified with anti-rat colon cancer monoclonal antibody CC531 showed good binding to target cells (Koning et al., 1999) and a potent intracellular drug

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Delivery compared to all controls (Koning et al. 2002). Antibody CC52 against rat colon adenocarcinoma CC531 binds to pegylated liposomes and causes specific accumulation of liposomes in a rat model of metastatic CC531 tumors (Kamps et al., 2000). Nonpathogenic antinuclear autoantibodies (ANA) are commonly detected in cancer patients and healthy older adults and represent a subclass of natural anticancer antibodies. We have previously demonstrated that certain monoclonal ANAs, such as mAbs 2C5 and 1G3, bind to the surface of many tumor cells but are not recognized by normal cells (Iakoubov et al. 1995a,b; Iakoubov and Torchilin 1997). Some of these monoclonal ANAs have been shown to have nucleosome-restricted specificity, and surface-bound nucleosomes (NSs) of tumor cells have been shown to be their universal molecular targets on the surface of various tumor cells (Iakoubov and Torchilin 1997, 1998 ) ). Since these antibodies can effectively recognize a variety of tumors, they can serve as specific ligands to deliver other drugs and drug carriers to tumors. These antibodies were used to generate drug-loaded, tumor-targeted, long-circulating immunoliposomes (with doxorubicin) and were shown to be compatible with various cancer cells (mouse Lewis lung cancer, 4T1, C26 and human BT-20, MCF- 7. PC3 cells) in vitro (Elbayoumi and Torchilin 2007; Lukyanov et al. 2004a) lead to a significant increase in tumor accumulation and reduced side effects in mouse model tumors, including intracranial human brain U-87 MG tumor xenografts in nude mice and excellent antitumor activity in vivo (Elbayoumi and Torchilin 2006, 2008; Gupta and Torchilin 2007). Doxorubicin-loaded PEGylated liposomes were also modified with the Fab0 fragment of an anti-CD74 antibody by a PEG-based heterobifunctional conjugation reagent and showed significant acceleration and enhancement in human Raji B lymphoma cells in vitro accumulation (Lundberg et al., 2007). Anti-CD166 scFv conjugated to drug-loaded liposomes promotes doxorubicin internalization by multiple prostate cancer cell lines (Du-145, PC3, LNCaP) (Roth et al., 2007). The scFv fragments of antibodies directed against leukemia stem cells and oncogenic molecules involved in the pathogenesis of acute myeloid leukemia have been used to target acute leukemia stem cells (Wang et al. 2008). Using the Fab0 fragment of the OX7 monoclonal antibody against the Thy1.1 antigen, liposomes loaded with doxorubicin were successfully delivered to the kidneys of rats (Tuffin et al., 2005). Since fibroblast activation protein (FAP) is a cell surface antigen expressed by tumor stromal fibroblasts in various cancers, scFvs derived from antibodies that cross-react with human and mouse FAP were used to convert polyethylene Diolated liposomes target tumor stromal cells (Baum et al., 2018). 2007). ). Doxorubicin liposomes modified with a chimeric TNT-3 monoclonal antibody specific for degenerated cells in the tumor necrotic area efficiently targeted the necrotic area of the tumor and showed higher expression in nude mice expressing H460 tumors treatment effect (Pan et al., 2008). Combination of immunoliposomes and endosome-disrupting peptides improves cytosolic delivery of liposomal drugs, increases cytotoxicity and opens new avenues for constructing targeted liposome systems, such as diphtheria toxin A chain as well as pH-dependent fusion The incorporation of the peptide diINF-7 into specific liposomes has been demonstrated to treat ovarian cancer (Mastrobattista et al., 2002).

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Early clinical trials of antibody-targeted drug-loaded liposomes have shown some promising results. For example, doxorubicin-loaded PEGylated liposomes (about 140 nm in size) modified with the F(ab0)2 fragment of the gastric cancer-specific monoclonal antibody GAH were tested in a phase 1 study and showed Similar pharmacokinetic results to Doxil (Matsumura et al., 2004). Since the transferrin (Tf) receptor (TfR) is overexpressed on the surface of some tumor cells, antibodies against TfR, as well as Tf itself, are popular ligands for targeting tumors and liposomes within tumor cells (Hakeyama et al. 2004 ) (although TfR expression in normal cells, especially liver, can compete with tumor targeting by Tf liposomes. Recent studies include conjugation of Tf to PEG on PEGylated liposomes for combined drug delivery Longevity and targeting ability to solid tumors (Ishida et al., 2001). Similar approaches have been used to deliver drugs for photodynamic therapy, including hypericin, to tumors (Derycke and De Witte 2002; Gijsens et al. 2002), and intracellular delivery of cisplatin to gastric tumors (Iinuma et al. 2002. Tf-conjugated doxorubicin-loaded liposomes showed increased binding and toxicity to C6 gliomas (Eavarone et al. 2000). Interestingly, increased expression of TfR was also detected in postischemic brain endothelial cells and used to deliver Tf-modified PEG-liposomes to the ischemic hindbrain of rats (Omori et al., 2003) Tf (Joshee et al., 2002) as well as anti-TfR antibodies (Tan et al., 2003; Xu et al., 2002) have also been used to facilitate gene delivery into cells via cationic liposomes. Tf-mediated Liposome delivery has also been successfully used for brain targeting. Immunoliposomes containing the OX26 monoclonal antibody against rat TfR have been found to localize to the brain microvascular endothelium (Huwyler et al., 1996). Lipids modified with folic acid Somatic targeting of tumors is a very popular approach because the folate receptor (FR) is often overexpressed in many tumor cells. Early studies demonstrated the use of FR endocytosis to convert macromolecules (Leamon and Low 1991) and then lipids into Interest in folic acid grew rapidly following the possibility of plastid (Lee and Low 1994) delivery into living cells to circumvent multidrug resistance - liposome-targeted drug delivery (see Important Information Gabizon et al. 2004 Lu and Low 2002a). Liposomal daunorubicin (Ni et al., 2002) as well as doxorubicin (Pan et al., 2003) and 5-fluorouracil (Gupta et al., 2007) FR delivery to various tumor cells both in vitro and in vivo and showed increased cytotoxicity.Recently, the application of folate-modified doxorubicin-loaded liposomes in the treatment of acute myeloid leukemia has been opposed to the use of all combined with retinoic acid to induce FR (Pan et al., 2002). Folate-targeted liposomes have been proposed as delivery vehicles for boron neutron capture therapy (Stephenson et al., 2003) and have also been used in tumor immunotherapy hapten tumor targeting (Lu and Low 2002b). In the context of gene therapy, folate-directed liposomes have been used for gene targeting of tumor cells (Reddy et al., 2002) and for antisense ODN targeting of tumors (Leamon et al., 2003). The search for new ligands for liposome targeting has mainly focused on specific receptors overexpressed on target cells (especially cancer cells) and some specific receptors

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Specific components of pathological cells. Thus, targeting liposomes to tumors has been achieved through the use of vitamin and growth factor receptors (Drummond et al., 2000). Vasoactive intestinal peptide (VIP) was used to target PEG liposomes with radionuclides to tumor VIP receptors, resulting in enhanced breast cancer suppression in rats (Dagar et al., 2003). PEG-liposomes were also targeted by RGD peptides to tumor vasculature integrins, and because they were loaded with doxorubicin, showed enhanced efficacy against C26 colon cancer in a mouse model (Schiffelers et al., 2003). RGD peptides have also been used to target liposomes to integrins on activated platelets and thus can be used for specific cardiovascular targeting (Lestini et al., 2002) and selective drug delivery to monocytes/middle cells in the brain granulocytes (Qin et al.). . 2007). A similar angiogenic-homing peptide has been used to target drug-loaded liposomes to vascular endothelial cells in experimental treatment of mouse tumors (Asai et al., 2002). Immunoliposomes targeting the epidermal growth factor receptor (EGFR) have been specifically delivered to a variety of tumor cells that overexpress EGFR (Mamot et al., 2003). Long-circulating hyaluronan targeting mitomycin C in liposomes increases its activity against tumors overexpressing hyaluronan receptors (Peer and Margalit 2004). The ability of galactosylated liposomes to concentrate in parenchymal cells has been used to deliver genes to these cells; see Hashida et al. (2001) Available for review. Cisplatin-loaded liposomes that specifically bind chondroitin sulfate, which is overexpressed in many tumor cells, have been used to successfully inhibit tumor growth and metastasis in vivo (Lee et al., 2002). Tumor-selective targeting of PEGylated liposomes was also achieved by grafting these liposomes together with basic fibroblast growth factor-binding peptide (Terada et al., 2007).

8 Active tumor targeting of drug-loaded micelles Like other delivery systems, the drug delivery potential of polymeric micelles can be further enhanced by attaching targeting ligands to the surface of the micelles (Figure 8). Various sugar moieties can be mentioned among these ligands

Figure 8 Like other delivery systems, the drug delivery potential of polymeric micelles can be further enhanced by attaching targeting ligands to the surface of micelles

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(Nagasaki et al. 2001), transferrin (Vinogradov et al. 1999) and folic acid residues (Ota et al. 2002), because many target cells, especially cancer cells, overexpress the corresponding receptors (such as transferrin protein and folate receptors). on their surfaces. Galactose- and lactose-modified PEG-polylactide copolymer micelles have been shown to specifically interact with lectins, thereby mimicking the targeted delivery of micelles to liver sites (Jule et al., 2003; Nagasaki et al., 2001). Transferrin-modified micelles based on PEG and polyethyleneimine, with sizes between 70 and 100 nm, are expected to attack tumors with overexpressed transferrin receptors (Vinogradov et al., 1999). Mixed micelle-like complexes of PEGylated DNA and transferrin-modified polyethyleneimine (Dash et al., 2000; Ogris et al., 1999) have been designed to enhance the expression of the same transferrin receptors. DNA delivery in human cells. A similar targeting approach has been successfully tested with folate-modified micelles (Leamon and Low 2001). Poly(L-histidine)/PEG and poly(L-lactic acid)/PEG block copolymer micelles with folic acid residues on their surface were shown to efficiently deliver doxorubicin to tumor cells in vitro, demonstrating their Potential for treating solid tumors combined with display targeting and pH sensitivity (Lee et al. 2003a). Of all the specific ligands, antibodies offer the broadest range in terms of target diversity and interaction specificity. Several attempts to covalently attach antibodies to surfactants or polymeric micelles (i.e. to create immunomicelles) have been described (Kabanov et al. 1989; Torchilin 2001; Torchilin et al. 2003b; Vinogradov et al. 1999). For example, micelles modified with fatty acid-conjugated Fab fragments of antibodies against brain glial cell antigens (glial fibrillary acid antigen and a2-glycoprotein) and micelles loaded with the antipsychotic drug trifluoperazine in the neck Accumulates progressively in rat brain after intra-arterial administration (Chekhonin et al., 1991; Kabanov et al., 1989). By employing a conjugation technique developed for binding specific ligands to liposomes (Leamon and Low 2001), PEG-PE-based immune cells modified with monoclonal antibodies were generated. The method uses PEG-PE, in which free PEG is end-activated with p-nitrophenylcarbonyl (pNP) groups. The diacyl lipid fragment of this bifunctional PEG derivative is tightly bound into the micelle core, while the pNP group exposed to water is stable at pH below 6 and can be combined with various ligands such as antibodies and their fragments. ) efficiently interacts with amino groups at pH values above 7,5, resulting in stable carbamate linkages (carbamate). All unreacted pNP groups hydrolyze spontaneously at the same pH. To generate immune-targeting micelles, it is sufficient to incubate the antibody to be attached with the drug-loaded micelles at around pH 8.0 (Lee et al. 2003a; Torchilin et al. 2003b). Both the original micelles and the antibody-modified micelles were spherical and uniform in size, about 20 nm. Micelle-bound proteins were quantified using fluorescent labeling or SDS-PAGE (Gao et al., 2003; Torchilin et al., 2003b). It was calculated that 10 to 20 antibody molecules could bind to a single micelle. Antibodies bound to the micellar corona retain their specific binding ability. Mouse blood clearance data showed similar pharmacokinetic profiles of 2C5-modified and normal PEG-PE micelles, confirming the long-circulating nature of primed immune micelles.

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

To specifically increase tumor accumulation of PEG-PE-based micelles, the latter were modified with tumor-specific monoclonal anti-nucleosome antibodies such as mAb 2C5 (Lee et al. 2003a; Torchilin et al. 2003b). Rhodamine-tagged 2C5 immune cells efficiently bound to the surface of several unrelated tumor cell lines: human BT20 cells (breast cancer) and mouse LLC cells (Lewis lung cancer) and EL4 cells (T lymphoma). Paclitaxel-loaded 2C5 immune cells also exhibited the same specificity as "empty" immune cells and efficiently bound various tumor cells. In vivo studies, 111In-labeled 2C5 immune cells accumulated significantly more than normal micelles in LLC tumor female C57BL/6J mice, were able to incorporate more micelles for drug delivery into tumors, and showed significant have a higher ability to inhibit tumor growth (Lee et al. 2003a; Torchilin et al. 2003b). Some other specific ligands (glycoproteins, lipoproteins, carbohydrates, peptides) have also been used to achieve active targeting of polymeric micelles (Jule et al. 2003; Vinogradov et al. 1999; Wakebayashi et al. 2004) . Polymer micelles modified with sugar units (glucose, galactose, mannose, lactose) have been particularly successful (Jule et al., 2003; Nagasaki et al., 2001). Folate-targeted mixed block copolymer micelles consist of folate-PEG-poly(DL-glycolic acid) and folate-free copolymers with one doxorubicin unit per polymer chain (Yoo and Park 2004). This folate-targeted doxorubicin-loaded micelle exhibited better uptake of folate receptor-overexpressing human oral squamous cell carcinoma cells compared to folate-free micelles, and protected against These cells are more cytotoxic. There are many other examples of targeting polymeric micelles for cancer therapeutics (see e.g. Park et al. 2005; Vinogradov et al. 1998, 1999; Xiong et al. 2007). For targeted micelles, the local release of free drug from the micelles in the target organ would lead to increased drug efficacy, while the stability of the micelles on the way to the target organ or tissue should facilitate drug dissolution and reduce toxicity, since unlike non-drug Fewer interactions - target organs.

9 Conclusions In conclusion, it is important to note in this section that there are several clear goals for tumor targeting using antibody-mediated drug-loaded nanocarriers compared to more traditional dosage forms: (1) Such delivery systems should be rapid and effective (2) Such systems should deliver higher amounts of drug to tumors than other delivery systems; (3) Ideally, drugs in nanocarriers should not only accumulate in the tumor stroma, but also Internalized by target cells, resulting in high intracellular drug concentration, thereby avoiding multidrug resistance. To achieve these goals, certain factors should be considered when developing chemotherapy-targeted agents. First, have a goal

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It has been identified that it is present in sufficient numbers (overexpression) on the surface of tumor cells and provides a good opportunity for targeted liposomes to bind tightly to cancer cells (Hosokawa et al., 2003). Second, specific ligands (antibodies or fragments thereof) should be bound to the surface of drug-loaded nanocarriers in a manner that does not impair their specific binding properties (the best choice should be made from the various conjugation methods available, bearing in mind save the following). ). Note that what works for one antibody may not work for another) and the amount is sufficient to allow multisite binding to the target; for pegylated long-circulating vectors, the amount of bound antibody should not be so large that Excessive impact on lifespan (Lukyanov et al. 2004a; Moreira et al. 2002). Third, anticancer drugs targeting antibodies that internalize and facilitate the internalization and incorporation of vectors are highly desirable (Kirpotin et al. 2006; Mamot et al. 2005). Fourth, drug release from carriers within tumors or tumor cells should deliver therapeutic concentrations of the drug at the target and be maintained for a reasonable period of time (several hours) (Allen et al. 2005; Sapra and Allen ). 2004).

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48 years old

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Vice President Tochlin

Weinstein JN, Magin RL, Yatvin MB, Zaharko DS (1979) Liposomes and local hyperthermia: selective delivery of methotrexate to heated tumors. Science 204:188-191 Weissig V, Gregoriadis G (1992) Conjugation of amino-bearing ligands to liposomes. In: Gregoriadis G (ed.) Liposome Technology, Vol. 3. CRC Press, Boca Raton, pp. 231-248 Weissig V, Lasch J, Gregoriadis G (1990) Covalent attachment of peptides to liposome surfaces. Pharmazie 45:849-850 Weissig V, Lasch J, Klibanov AL, Torchilin VP (1986) A novel hydrophobic anchor for binding proteins to liposome membranes. FEBS Lett 202:86-90 Weissig V, Whiteman KR, Torchilin VP (1998) Accumulation of protein-loaded long circulating micelles and liposomes in mouse subcutaneous Lewis lung carcinoma. Pharm Res 15:1552-1556 Whiteman KR, Subr V, Ulbrich K, Torchilin VP (2001) Poly(HPMA)-coated liposomes show prolonged circulation in mice. J Liposome Res 11:153-164 Widder KJ, Marino PA, Morris RM, Senyei AE (1983) Targeting antineoplastic agents using magnetic albumin microspheres. In: Goldberg EP (Ed.) Targeted Drugs. Wiley, New York, pp. 201-230 Williams AS, Camilleri JP, Goodfellow RM, Williams BD (1996) A single intraarticular injection of liposome-conjugated methotrexate inhibits antigen-induced joint inflammation in rats. Br J Rheumatol 35:719-724 Winslow RM, Vandegriff KD, Intaglietta M (1996) Blood substitutes: new challenges. Birkhäuser, Boston Wollina U, Dummer R, Brockmeyer NH, Konrad H, Busch JO, Kaatz M, Knopf B, Koch HJ, Hauschild A (2003) Pegylated liposomal doxorubicin in patients with cutaneous T-cell lymphoma A multicenter study in . Krebs 98:993-1001 Woodle MC (1993) Surface modified liposomes: evaluation and characterization for increased stability and prolonged blood circulation. Chem Phys Lipids 64:249-262 Woodle MC (1998) Control of blood clearance of liposomes by surface grafted polymers. Adv Drug Deliv Rev 32:139-152 Woodle MC, Engbers CM, Zalipsky S (1994) Novel amphiphilic polymer-lipid conjugates form long-circulating liposomes that escape the reticuloendothelial system. Bioconjug Chem 5:493-496 Woodle MC, Storm G (1998) Long-circulating liposomes: old drug, new therapy. Biotechnology Intelligence Unit. Springer, Berlin, p. 301 Xiong on metastatic tumor cells. Biomacromolecules 8:874-884 - Immunoliposomes. Mol Cancer Ther 1:337-346 Yamaoka T, Tabata Y, Ikada Y (1994) Distribution and tissue uptake of polyethylene glycols of different molecular weights following intravenous administration in mice. J Pharm Sci 83:601–606 Yang T, Choi MK, Cui FD, Kim JS, Chung SJ, Shim CK, Kim DD (2007) Preparation and evaluation of paclitaxel-loaded pegylated immunoliposomes. J Control Release 120:169-177 Yerushalmi N, Arad A, Margalit R (1994) Molecular and cellular studies of hyaluronic acid-modified liposomes as bioadhesive vehicles for topical delivery in wound healing. Arch Biochem Biophys 313:267-273 Yoo HS, Lee EA, Park TG (2002) Doxorubicin-coupled biodegradable polymer micelles with acid-cleavable linkages. J Control Release 82:17–27 Yoo HS, Park TG (2004) Folate receptor-targeted delivery of doxorubicin-PEG-folate conjugate stabilized doxorubicin nanoaggregates. J Control Release 100:247-256 Yuan F, Dellian M, Fukumura D, Leunig M, Berk DA, Torchilin VP, Jain RK (1995) Vascular permeability in human tumor xenografts: molecular size and cutoff size dependence sex. Krebs Res 55:3752-3756

Passive and Active Drug Targeting: Examples of Drug Delivery to Tumors

Yuan F, Leunig M, Huang SK, Berk DA, Papahadjopoulos D, Jain RK (1994) Microvascular and interstitial penetration of sterically stabilized (stealth) liposomes in human tumor xenografts. Cancer Res 54:3352-3356 Yuan Copolymer with hydrophobic groups. Langmuir 21:2668-2674 Zalipsky S (1995) The chemistry of polyethylene glycol conjugation to bioactive molecules. Adv Drug Deliv Rev 16:157-182 Zalipsky S, Gittelman J, Mullah N, Qazen MM, Harding JA (1998) Polymer-grafted liposomes with bioactive ligands. In: Gregoriadis G (Ed.) Targeted medicine 6: Strategies for stealth therapeutic systems. Plenum Press, New York, pp. 131-139 Zalipsky S, Qazen M, Walker JA 2nd, Mullah N, Quinn YP, Huang SK (1999) Novel separable poly(ethylene glycol) conjugates: cysteine-cleavable Lipopolymers regenerate natural phospholipids, diacylphosphatidylethanolamines. Bioconjug Chem 10:703-707 Zhang JX, Zalipsky S, Mullah N, Pechar M, Allen TM (2004) Diole containing different types of cleavable lipopolymers Pharmacological properties of acylphosphatidylethanolamine/cholesterol hemisuccinate liposomes. Pharmacodynamic Research 49:185-198

Nanoparticle Technology for Cancer Therapy Frank Alexis, Eric M. Pridgen, Robert Langer and Omid C. Farokhzad

content 1 2

introduce . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 56 Nanoparticle Technology. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 59 2.1 Liposome nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 61 2.2 Polymer-drug conjugate nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 62 2.3 Polymeric nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 63 2.4 Micellar nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 64 2.5 Dendritic polymer nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 65 2.6 Polymersome nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 66 2.7 Protein nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 66 2.8 Biological nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 67 2.9 Inorganic nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 67 2.10 Hybrid Nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 69 3 Strategies for treating cancer with nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 70 3.1 Metastatic cancer. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 70 3.2 Non-targeting nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 71 3.3 Targeting nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 74 4 Summary. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 76 References. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 77

Abstract: Nanoparticles as drug delivery systems enable unique cancer treatments. Over the past two decades, a large number of nanoparticle delivery systems, including organic and inorganic materials, have been developed for cancer therapy. Many liposomal, polymer-drug conjugate, and micellar formulations are state-of-the-art in the clinic, and many more nanoparticle platforms are currently in preclinical development. Recently developed nanoparticles demonstrate the potential complexity of the O.C. Farokhzad (*) Harvard Medical School, Nanomedicine and Biomaterials Laboratory, Brigham and Women's Hospital, 75 Francis St, MRB-5th Floor, Boston, MA, 02115, USA Email:[email protected]M. Schäfer-Korting (Hrsg.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_2, # Springer-Verlag Berlin Heidelberg 2010

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These delivery systems are enhanced by integrating multifunctional capabilities and targeting strategies to increase the effectiveness of these systems against the most difficult cancer challenges, including drug resistance and metastatic disease. In this chapter, we review available preclinical and clinical nanoparticle technology platforms and their implications for cancer therapy. Key words nanoparticle drug delivery targeting metastatic cancer cancer therapy

Abbreviations BBB DSPC DSPE EggPG EPR FDA HPMA HSPC LPS MTD NCI NIR NSCL Cancer PAMAM PDLLA PEG PLA PLA2 PLGA SEM

BBB 1,2-Distearoyl-glycero-3-phosphocholine 1,2-Distearoyl-sn-glycero-3-phosphoethanolamine Egg Yolk Phosphatidylglycerol Increased Permeability and Retention Effect Food and Drug Administration N-(2-Hydroxypropyl)methacrylamide hydrogenated phosphatidyl Choline from Soybean Lecithin Lipopolysaccharide Maximum Tolerated Dose National Cancer Institute Near Infrared Non-Small Cell Lung Cancer Polyamidoamine Poly-DL-Lactic Acid Polyethylene Glycol electron microscope

1 Introduction Nanotechnology is a multidisciplinary field that uses principles from chemistry, biology, physics, and engineering to design and fabricate nanoscale devices (Farokhzad and Langer 2009; Ferrari 2005; Fox 2000; Jiang et al. 2007 2004; Peppas 2004; Sinha et al. 2006; Uchegbu 2006). In its strictest definition, nanotechnology refers to structures with dimensions ranging from 1-100 nm in at least one dimension.

Nanoparticle technology for corona cancer therapy - for influencing biodistribution and half-life in circulation

Payload - including chemotherapy drugs and imaging agents

57 Targeting Ligands – Increased cellular uptake after accumulation in tumor tissue via binding and endocytosis

Core - Affects drug encapsulation and release

Figure 1 Schematic of the physicochemical structure of the nanoparticle platform for drug delivery, including the core, corona, payload, and targeting ligand

However, it more commonly refers to materials developed using top-down or bottom-up techniques with dimensions up to hundreds of nanometers. The resulting nanomaterials exhibit unique properties based on intrinsic properties such as shape and size, as well as functional properties conferred by surface modification (Figure 1). Medicine is likely to be the main beneficiary of nanotechnology advances, and oncology has already begun to benefit from novel nanotechnology (Alexis et al. 2008b; Alexis et al. 2008c; Davis et al. 2008; Euliss et al. 2006 Farokhzad 2008 Farokhzad et al. Langer 2006 Freitas 2005 Jain 2008 Kawasaki and Player 2005 Lanza et al 2006 Levy-Nissenbaum et al 2008 Moghimi et al 2005 Peer et al 2007; Pridgen et al 2007; Riehemann et al 20 09; -Morales et al. 2009a; Venugopal et al. 2008; Zhang et al. 2008b). These benefits include advances in disease detection, imaging and treatment. The National Cancer Institute (NCI) has identified nanotechnology as having the potential to have a paradigm-shifting impact on cancer detection, treatment, and prevention. Strong interest in nanotechnology among researchers in academia and industry has led to increased development of novel nanotechnology platforms for medical applications and a surge in government funding and venture capital investment. The combination of funding and early clinical success provides nanotechnology with the resources and opportunity to address important medical challenges. Early successes in oncology have been catalysts for the application of nanotechnology to cardiovascular disease and other medical problems such as vaccines. One area where nanotechnology could have a significant impact is drug delivery (Farokhzad and Langer 2009; Pridgen et al. 2007). This effect has

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Problems have been felt in bringing multiple nanoscale drug delivery systems into the clinic, although the full potential of these systems is only just beginning to be explored. Nanoscale drug delivery vehicles have demonstrated the ability to encapsulate various therapeutic agents, such as small molecules (hydrophilic and/or hydrophobic), peptides, protein-based drugs, and nucleic acids. By encapsulating these molecules in nanocarriers, drug solubility and stability can be improved, providing an opportunity to re-evaluate potential drugs that were previously overlooked due to poor pharmacokinetics (Langer 1998). Encapsulated molecules can be released from nanocarriers in a controlled manner over time to maintain drug concentrations within a therapeutic window, or release can be triggered by stimuli specific to the delivery site (Moghimi 2006). The surface of nanocarriers can be engineered to increase circulating half-life and affect biodistribution, while attachment of targeting ligands to the surface can lead to enhanced uptake in target tissues (Gref et al. 1994; Moghimi et al. 2001). .The small size allows nanocarriers to cross biological barriers and enable cellular uptake (Brigger et al., 2002). The net result of these properties is to reduce the systemic toxicity of the therapeutic agent while increasing the concentration of the agent in the region of interest, resulting in a higher therapeutic index of the therapeutic agent. In addition to therapeutic drugs, imaging agents can also be incorporated into nanocarriers to improve tumor detection and imaging (Kim et al. 2006; Montet et al. 2006). Finally, nanoparticles can be designed to be multifunctional, capable of targeting diseased tissue, carrying imaging agents for detection, and delivering multiple therapeutic agents for combination therapy (Nasongkla et al., 2006). The multimodal capabilities of nanoparticle delivery systems present opportunities to develop novel drug delivery methods that may lead to alternative or adjuvant therapeutic options for treating disease. In this chapter, we focus on nanoparticle technology (Figure 2), with a particular emphasis on the development of nanocarrier drug delivery systems for cancer therapy. These technologies include polymeric nanoparticles, dendrimers, nanoshells, liposomes, inorganic/metallic nanoparticles, hybrid nanoparticles, micelles, and magnetic and bacterial nanoparticles. Nucleic acid delivery technologies were not considered but are described in detail elsewhere (Chen and Huang 2008; Gao and Huang 2008; Gary et al. 2007; Juliano et al. 2008; Li and Huang 2008b; Luten et al. 2008; Tseng et al. 2009 ). ; Whitehead et al. 2009). This is followed by a discussion of how improvements in the understanding of the tumor microenvironment may affect the development of non-targeting and targeted nanocarriers as anticancer therapeutic delivery vehicles (Bierie and Moses 2006; Bissell and Labarge 2005; Cairns et al. 2006; Fesik 2005). ; Fiedler 1995; Gallon et al. 2007; Overall and Kleifeld 2006; Siklari et al. 2006; Zett 2008). The breakthrough potential of nanoparticle delivery systems is increasingly recognized, and several examples of first-generation nanocarriers have been approved by the FDA for use as targeted nanocarriers in therapeutic and clinical development. Many nanocarrier systems in clinical development have been examined

Nanoparticle Technology for Cancer Therapy

polymer nanoparticles

Liposomes

dendrimer

polymer body

polymer micelles

Inorganic (iron, silica or quantum dot core) Hydrophobic polymer Hydrophilic polymer Lipid

protein carrier

biological nanoparticles

Polymer Drug Conjugates

hybrid nanoparticles

Therapeutic Load Targeting Ligands

Figure 2 Nanoparticle platform for drug delivery. Nanoparticle platforms are characterized by their physicochemical structures, including polymer-drug conjugates, lipid-based nanoparticles, polymer nanoparticles, protein-based nanoparticles, biological nanoparticles, and hybrid nanoparticles

This chapter shows how these systems have transferred to the clinic and what advantages they offer for cancer treatment.

2 Nanoparticle technology The first nanoscale drug delivery system was the lipid vesicle, first described in the 1960s and later called liposome (Bangham et al., 1965). Since then, there have been several important developments that have paved the way for current nanoparticle technology. In 1976, the first controlled-release polymer system for the delivery of macromolecules was demonstrated (Langer and Folkman 1976). The first use of targeted liposomes followed in 1980 (Heath et al., 1980; Leserman et al., 1980). Surface modification of liposomes and polymeric nanoparticles with polyethylene glycol (PEG) in 1990 and 1994, respectively, resulted in increased circulation times or "stealth" properties (Gref et al., 1994; Klibanov et al., 1990 Year). These developments culminated in the approval of Doxil (James 1995a,b), a vesicular delivery system encapsulating doxorubicin that has been shown to be effective in the treatment of several types of cancer (Porche 1996;

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Tejada-Berges et al. 2002). Since then, research has made great strides in developing nanoparticles with multifunctional and "smart" properties, such as the ability of B. to respond to their environment to enable more effective drug delivery strategies. There are currently 70 clinical trials evaluating nanoparticle carriers, 208 evaluating drug conjugates, and 361 clinical trials evaluating vesicle-based carriers (http://www.clinicaltrials.gov). Clinical trials include combination therapies and treatments by different routes of administration, such as pulmonary and oral. Nanoparticle technologies for cancer therapy include polymeric nanoparticles (Moghimi 2006; Pridgen et al. 2007), vesicle-based carriers such as liposomes (Kaneda 2000; Torchilin 2005), micelles (Fan 2008; Liggins and Burt 2002 ; Matsumura 2008), dendrimers (Florence et al. Hussain 2001, Lee et al. 2005, McCarthy et al. 2005, Najlah and D'Emanuele 2007), polymer conjugates (Greco and Vicent 2008, Li and Wallace 2008, Thanou and Duncan 2003), protein carriers (Hawkins et al. 2008; Wang and Uludag 2008), inorganic nanoparticles (Murakami and Tsuchida 2008) and bacterial nanocarriers. The various delivery systems discussed below allow the development of nanoparticles with a wide variety of shapes, sizes, and assemblies, which can thus be tailored for specific applications. However, the main goal in developing drug delivery systems is to achieve more effective therapy through drug control

tumor

Advantages of Using Nanoparticles to Treat Cancer: Tumors. Selective accumulation at the tumor site increases drug concentration in the tumor due to the EPR effect. Lower drug concentration in healthy tissue minimizes side effects of chemotherapy

Higher maximum tolerated dose for nanoparticle drugs

hoch

Organic Drug Concentration. Poor distribution of free medicines

Drug distribution encapsulated in nanoparticles

Figure 3 Advantages of using nanoparticles as a drug delivery system for cancer therapy compared to free drug

Nanoparticle Technology for Cancer Therapy

Concentrations within the therapeutic window, reduction in cytotoxic effects and improvement in patient compliance (Figure 3). This enables effective treatment cycles to be maintained while reducing damage to healthy cells and minimizing recovery periods.

2.1

Liposome-nanoparticle

Lipids form nanoparticle vesicles through self-assembly of amphiphilic lipids and excipients. Lipids form bilayers based on hydrophobic interactions in continuous parallel packing, with the hydrophilic headgroups oriented toward the aqueous environment. Hydrophilic molecules can be encapsulated in the inner aqueous phase, while hydrophobic molecules can be transported in the hydrophobic domain of the lipid bilayer. The physicochemical properties of liposomes can be precisely altered by simply mixing commercially available lipid molecules to control surface charge, functionality, and size. This offers significant advantages over other supports that require more controlled synthesis steps and additional chemical modifications. In general, lipids used to make vesicle formulations are present in the human body and are FDA-approved, such as DSPE (1,2-distearoylglycero-3-phosphoethanolamine), HSPC (hydrogenated phospholipid from soybean lecithin Acylcholine) and EggPG (egg yolk). Phosphatidylglycerol) and DSPC (1,2-Distearoyl-glycero-3-phosphocholine). Any of these lipids can be obtained with or without PEG, which can be used to modify the surface of the resulting liposomes. Doxil, a polyethylene glycol liposome, is clinically used to treat a variety of cancers and is a breakthrough in liposomal drug delivery systems. Doxil consists of a PEGylated surface (2 kDa PEG chains) and is loaded with doxorubicin by drug diffusion based on an ammonium salt gradient. Using this method, stable drug entrapment in crystalline form with reduced leakage can be achieved over a long period of time. Doxil liposomes have a size of approximately 100 nm, a surface charge of approximately 10 mV, and a long-term storage stability of approximately 2 years at approximately 4 °C. Recently published by Aphios Corp. The developed nanobodies (small liposomes, 200 nm) are generally larger spherical shapes, whereas nanoparticles are smaller (1%) (Kreuter 1994). Due to their large endothelial gap of approximately 150 nm, distribution in vivo is also affected by the extravasation of nanoparticles from the peripheral capillary walls of these organs. Therefore, passive targeting severely limits the use of nanoparticles in API delivery other than those that fall within the context of renewable energy (Wolburg and Lippoldt 2002). To overcome these limitations, nanoparticles are often surface-modified with hydrophilic molecules (e.g., surfactants and hydrophilic polymers or proteins) to prevent recognition by the mononuclear phagocytic system (MPS). Furthermore, negative surfaces are widely believed to activate the complement system and coagulation factors (Moghimi et al., 2001). In addition to particle size reduction, API biodistribution is altered, enhancing the systemic temporal circulation of the vector and its deposition in non-RES organs (Kreuter 2001). In fact, one of the current approaches to achieve site-specific delivery is to bypass normal physiological defenses by reducing particle size, thereby persisting in the systemic circulation for a longer period of time.

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It is also known that the size and deformability of nanoparticles are very important for their clearance through the sinusoidal spleen in humans and rats, viz. H. To avoid splenic filtration in the endothelial-intercellular space (IES) of the sinus wall, nanoparticles must be sufficiently small and deformable (Moghimi et al., 1993, 1991). It has been reported that the size of engineered long-cycling particles should ideally not exceed 200 nm (Groom 1987). Otherwise, nanoparticles must be sufficiently deformable to bypass IES filtration. Alternatively, long-circulating rigid particles with a size greater than 200 nm could act as splenogenic agents and be removed later if non-rigid (Moghimi et al., 1991). When the SLN is below 200 nm, systemic circulation increases and thus the time the drug remains in contact with the target site increases. SLN has been suggested as a suitable system for parenteral administration of hydrophilic APIs such as dimimazine, as well as other BCS class IV APIs such as paclitaxel, vinblastine, camptothecin, etoposide, and cyclosporine (Cavalli et al. 2000a; Chen et al. 2001; Yang et al. 1999a, b). Due to its lipophilicity, SLN can be rapidly absorbed by the RES, which may lead to treatment failure due to insufficient API plasma concentration. Stereostabilization is also an option, as it forms a dense conformational cloud around the particle, reducing opsonization and phagocytosis and uptake by neutrophils. The result would be an increase in the systemic half-life of the drug. An example of steric stabilization is the stabilization of lipid nanoparticles provided by polyethylene glycol (PEG) molecules. PEG is a hydrophilic, electrically neutral polymer with high chain flexibility. The absence of functional groups prevents physicochemical interactions with the biological environment. PEG molecules with a molecular weight between 2,000 and 5,000 kDa are generally required to inhibit adsorption of plasma proteins, and those that form a thicker hydrophilic layer around the particle also contribute to reduced hepatic clearance (Chen et al., 2001). To increase the selectivity of SLNs for specific targets, ligands or homing devices (surface epitopes or receptors that specifically bind target sites) can be coupled to their surface. Cancer cells are known to overexpress specific receptors. B. Folate receptor (overexpressed in cancer cells of epithelial origin) and low-density lipoprotein (LDL) receptor (ie, the B16 melanoma cell line showed higher LDL receptor expression). ) and peptide receptors (eg, somatostatin, vasoactive intestinal peptide, gastrin-related peptide, cholecystokinin, gonadotropin-releasing hormone). Therefore, attaching appropriate ligands for these specific receptors to the SLN surface could improve selectivity (Pardridge 2007b).

4.4

brain positioning

Interest in APIs targeting the brain has grown over the past decade (Blasi et al. 2007; Goppert and Müller 2005; Kreuter 2001; Pardridge 2005, 2007b,c,d,e). Lack of understanding of central nervous system physiology

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The nervous system (CNS) is one of the limiting factors in the development of effective APIs and appropriate API delivery systems for targeting and delivery in the brain (Pardridge 2003, 2007a,c,d). The specific blood-brain barrier (BBB) strictly regulates the exchange between the peripheral blood flow and the CSF circulatory system. Thus, these physiological features of the brain's microvasculature severely limit the amount of API that can enter the brain upon systemic administration. In fact, more than 98% of potential new CNS-active drugs fail to cross the BBB (Pardridge 2007a). Drug molecules with high lipophilicity and a molecular weight below 500 Da can pass through the blood-brain barrier. Several strategies have been tried to efficiently achieve the delivery and deposition of APIs in the CNS (Badruddoja and Black 2006; Johanson et al. 2005; Vyas et al. 2005), in particular the use of API delivery systems (Tiwari and Amiji 2006). One possibility for entry into the brain is receptor-mediated transport, since the BBB expresses receptors for endogenous macromolecules (e.g., insulin, transferrin, leptin, apoE, thiamine) on the luminal side. Receptor-mediated transport of these molecules can be used for specific delivery to the brain (Cornford and Hyman 1999). APIs are delivered directly to the brain by conjugating drugs or carriers such as liposomes and nanoparticles to specific ligands (peptides) or peptidomimetic monoclonal antibodies (Pardridge 2003). These monoclonal antibodies act as a Trojan horse, delivering the nanoparticles to the brain. Use of peptidomimetic antibodies capable of binding BBB transcytosis receptors, i. H. Pegylated immune nanoparticles targeting the brain have also been proposed. The entrapped API can be delivered into the brain parenchyma without altering the BBB permeability (Harris and Chess 2003). However, some transporters present in the blood-brain barrier, such as P-glycoprotein, can also limit the delivery of APIs in the brain and prevent the accumulation of various drugs including APIs in the brain (Stouch and Gudmundsson 2002). To overcome this limitation, well-established pharmaceutical surfactants have been proposed to inhibit P-glycoprotein (Batrakova et al., 1999; Miller et al., 1999). Polymeric nanoparticles are thought to be particularly useful in overcoming the blood-brain barrier (Garcia-Garcia et al., 2005; Müller and Keck 2004b), and when the nanoparticles are coated with polysorbate 80 (Tween 80), the blood-brain barrier appears to very high (Goppert and Muüller 2005); Koziara et al. 2003). SLNs have also been tested for brain targeting (Garcia-Garcia et al. 2005; Goppert and Müller 2005; Müller and Keck 2004a,b). The potential advantages of SLNs over polymeric nanoparticles for brain targeting are based on their lower cytotoxicity, higher API loading capacity, and better production scalability. Surfactant coating technologies developed for brain targeting have been transferred to SLNs and related vehicles with relatively great success. Goppert and Müller developed polysorbate-surfaced SLNs that can deliver a variety of drugs to the brain. These studies also showed that ApoC and ApoCII adsorbed on the SLN surface inhibited receptor-mediated binding of b-VLDL (expressing ApoE on the particle surface) to the LDL receptor (Goppert and Müller 2005). The authors highlight the benefit of adsorbing high ApoE/ApoCII ratios on particles for brain targeting. In addition, they also found

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This stealth SLN containing polysorbate 80 adsorbed the least amount of ApoCII on the particle surface. Pathfinder Technology, i. H. Differential protein adsorption, using plasma proteins adsorbed to the surface of intravenously injected SLN to target it. ApoE is one such unit for SLN targeting of BBB endothelial cells (Müller and Keck 2004a). Zara and colleagues developed SLNs and PEG-coated SLNs containing increasing amounts of this cloak to deliver doxorubicin to the brain following intravenous administration. management (Kaur et al., 2008). When the stealth compound was increased, the concentration of doxorubicin in the brain increased. The concentration of doxorubicin in rabbit brain ranged from 27.5 ng g1 for non-camouflaged SLN to 242.0 ng g1 for camouflaged SLN (with PEG molecules on the surface). Thor et al. Improved interaction with brain endothelial cells and higher intracellular accumulation of sterically stabilized liposomes conjugated to cationized albumin were reported compared to bovine serum albumin nanoparticles (Thole et al., 2002 ). Positively charged albumin nanoparticles have been taken up into brain endothelial cells via the caveolae-mediated endocytic pathway. The effect of SLN surface charge on brain output was also examined after administration of etoposide-loaded tripalmitine SLN. Brain levels were compared to etoposide solution. Positively charged etoposide-loaded SLN reached the highest brain concentration (0.07% injected dose/g), compared to negatively charged etoposide-loaded SLN (0.02%) and etoposide solution (0 .01%), Significantly more than uptake (Reddy and Venkateswarlu 2004). Furthermore, nitrendipine-loaded SLN was composed of various acylglycerols (tripalmitine, trimyristine, tristearic acid), surfactants (soy lecithin, porin Loxamer 188) and charge regulators (dicetyl phosphate, stearylamine), the purpose is to increase the systemic half-life of API compared with intravenous injection. Mode of administration compared to traditional nitrendipine suspension (Manjunath and Venkateswarlu 2006). It was found that the SLN formulation was absorbed to a greater extent by the brain and maintained high API levels for 6 hours, whereas the nitrendipine suspension only reached these levels for 3 hours. SLNs from tripalmitine, tripalmitine dicetyl phosphate and tripalmitine stearylamine demonstrated 3.2, 7.3 and 9.1-fold increases in Cmax compared to API suspensions, respectively. Similar results were also reported for SLN loaded with 30,50-dioctanoyl-5-fluoro-20-deoxyuridine (Wang et al., 2002). Stearic acid-based SLNs loaded with camptothecin were administered intravenously. managed. In mice (1.3 mg kg1), this resulted in a significantly longer residence time of the drug in the body compared to camptothecin solution (Yang et al., 1999a). A 5-fold increase in plasma AUC and a 10-fold increase in brain AUC were observed when the camptothecin dose was increased from 1.3 to 3.3 mg/kg1. In addition to the advantages of SLNs in enhancing brain drug uptake, the extremely low brain cytotoxicity of SLNs makes these vectors very attractive candidates for brain delivery (Müller et al., 1997b). It needs to be emphasized that the toxicity of SLN is not only related to the type of lipid but also to the surfactant used to stabilize the particles in aqueous dispersion. The most common surfactant used to target nanoparticles in the brain is polysorbate 80. Interestingly, free polysorbate 80 was more toxic than bound (Koziara et al., 2006), which was attributed to

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The fact that this surfactant is incorporated into the SLN matrix rather than being adsorbed and thus released minimally also reduces toxicity.

5 Conclusions and Outlook This chapter provides an overview of recent achievements in the modification of API pharmacokinetic parameters and bioavailability using lipid nanoparticles (SLNs and NLCs). These carriers are made of bioenvironmentally compatible materials. SLNs and NLCs have been used for oral, dermal, pulmonary and parenteral administration. Clearly, the in vivo behavior of these nanoparticles and the resulting therapeutic potential depend on their physicochemical properties as well as the route of administration. The type of lipid nanoparticle system (SLN vs. NLC) should be strictly chosen according to the route of administration, e.g. B. NLC is unlikely to be used for brain delivery. However, both systems can be used to reduce API toxicity. The pharmaceutical industry is interested in developing a delivery system that is versatile enough for multiple routes of administration. Variations in carrier surface properties (charge, hydrophilicity) and matrix composition may be required to minimize or overcome limitations associated with more traditional colloidal carriers (e.g., liposomes, polymeric nanoparticles, nanoemulsions). SLNs and NLCs can be designed according to the physicochemical properties of the API molecule as well as the route of administration and target/delivery purpose.

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Viral Vectors for Gene Transfer: The Current State of Gene Therapy Regine Heilbronn and Stefan Weger

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Gene therapy: definitions and state-of-the-art. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 145 AAVs for gene therapy carrier. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 147 2.1 Overview of attributes. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 147 2.2 AAV structure. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 148 2.3 AAV life cycle. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 148 2.4 Cellular receptors used by AAV. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 149 2.5 AAV vector production. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 149 2.6 Persistence and safety of AAV vectors. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 154 2.7 AAV Segmentation Vectors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 155 2.8 Dimeric, self-complementary (sc) AAV vectors. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 155 2.9 Cell targeting strategies for AAV. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 156 2.10 Future directions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 157 3 Retroviral vectors for gene therapy. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 157 3.1 Retrovirus structure and life cycle. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 158 3.2 Cancer Design and development of retroviral vectors. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 158 3.3 Self-inactivating (SIN) oncoretroviral vectors. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 159 3.4 Design and development of lentiviral vectors. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 161 3.5 Safety of retroviral integration. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 162 3.6 Production and stability of retroviral vectors. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 163 3.7 Purification and amplification of retroviral vectors. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 164 3.8 Vector quantization and quality assessment. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 165 3.9 Future directions for retroviral vector development. . . . . . . . . . . . . . . . . . . . . . . . . . . . . 166 4 Outlook. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 167 References. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 168

R. Heilbronn (*) Institute for Virology, Charité – Universitätsmedizin Berlin, Hindenburgdamm 27 12203 Berlin, Germany Email:[email protected]

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Abstract: In recent years, gene therapy for the correction of congenital or acquired diseases has become increasingly important. Children with severe combined immunodeficiency (SCID) have been successfully treated using retroviral vectors for gene transfer. Encouraging improvements in vision have been reported for a genetic eye disorder (LCA) that causes early childhood blindness. In these experiments, adeno-associated viral vectors (AAV) were used for gene transfer. This chapter provides an overview of the design and delivery of viral vectors for the delivery of therapeutic genes to target cells or tissues. Includes construction and production of retroviral, lentiviral, and AAV vectors. The focus is on production methods suitable for scale-up and further processing of biopharmaceuticals. Quality control measures and biosafety considerations for the use of vectors in clinical trials are discussed. Key words viral vector gene therapy adeno-associated virus retrovirus lentivirus production method safety aspects

Abkürzungen AAV AAVS1 Ad ADA cDNA CMV CNS cPPT CsCl DOC EIAV ELISA FGFR GFP GMP HIV HSV ITR lacZ LCA LMO2 MoMLV PCR PIC RCV

Adeno-Associated Virus Adeno-Associated Virus Integration Site 1 Adenosine Deaminase Complementary DNA Cytomegalovirus Central Nervous System Equine Infectious Anemia Central Polypurine Tract Cesium Chloride Virus Enzyme Immunoassay Fibroblast-Wachtum Factor Green Fluorescent Protein GMP human immunodeficiency virus herpes simplex virus inverted terminal repeat b-galactosidase hepatic amaurosis LIM domain-only 2 (rhomboid 1) Moloney murine leukemia virus polymerase chain reaction pre-integration complex replication competent virus

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Rev Response Element Self-complementing AAV Severe combined immunodeficiency SDS-polyacrylamide gel electrophoresis Spodoptera frugiperda cell line Self-inactivation Unique domain 30 virus-associated RNA Viral protein Vesicular stomatitis virus

1 Gene therapy: definition and prior art The original concept of gene therapy is to supplement or replace the function of a defective gene by introducing a functional copy of the gene into target cells for protein expression. Variations on this original concept have led to the development of gene therapy approaches to induce, harness, or modulate gene expression through various DNA or RNA molecules. Expresses full-length genes that result in the expression of specific proteins, but also expresses their truncated versions, complementing or disrupting gene expression. In addition, antisense RNA, shRNA, etc. regulate and interfere with the expression of RNA molecules, resulting in more complex regulation of gene expression at the transcriptional or post-transcriptional level. Given this widening range of possible applications, gene therapy should be defined today as a drug delivery modality in which the drug is a nucleic acid, either DNA or RNA, delivered to target cells by a specific drug formulation that must meet these dual goals. The active ingredient DNA or RNA must be protected so that it can safely reach the site of action. Unprotected naked DNA or RNA tends to degrade. In addition, the protective shield must effectively attach to the cell membrane and allow cellular entry, and then be safely transported and efficiently delivered to the intracellular site of action, usually the nucleus. Strategies that viruses have evolved over millions of years are used to solve problems related to protecting DNA or RNA molecules and ensuring efficient and safe transport across cell membranes. The viral nucleic acid is protected by a core protein encoded by the virus itself. Many viruses carry an additional envelope consisting of a cell-derived lipid bilayer with embedded viral proteins. The three-dimensional structure and binding properties of exposed viral outer capsid or envelope proteins ensure binding to cellular receptors. This initiates the entry of the virus into the cell. Due to the evolution of viral entry efficiency, most gene therapy approaches are based on engineered human or animal viruses carrying foreign genes of interest.

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Gene therapy uses pharmaceutical DNA or RNA to treat a wide range of diseases, covering all major disease groups. This is reflected in the number of current clinical trials of gene therapy in 2008: 65% for the treatment of cancer, 9% for cardiovascular disease, 8% for monogenic diseases, 7% for infectious diseases, 2% for neurological and ophthalmic diseases and the remaining few percent cover various other applications. Gene therapy is widely considered an option for treating otherwise incurable and life-threatening diseases. Gene therapy has been met with great promise and painful failure since it was first attempted more than 15 years ago. Despite its many pitfalls, impressive and enduring success has been documented, for example in children with X-linked SCID, an inherited form of severe combined immunodeficiency (SCID). These children can only survive in an absolutely sterile environment. Retroviral vector-mediated gene exchange of defective genes in transduced T cells resulted in complete recovery of immune function in most patients (Cavazzana-Calvo et al., 2000; Gaspar et al., 2004). Unfortunately, three to five years after successful treatment, four of the 20 children initially treated developed leukemia. In all cases, retroviral LTR-mediated activation of nearby cell oncogenes was detected. Currently, 18 of the original 20 children are alive and have recovered immunity (Kohn and Candotti 2009). Recently, successful long-term follow-up of retrovirus-mediated gene therapy for ADA forms of SCID has been reported. All treated patients were alive and 8 out of 10 patients maintained excellent and sustained immune reconstitution eight years after initial treatment with no signs of side effects (Aiuti et al., 2009). The year 2008 witnessed yet another impressive success in gene therapy for different diseases and target tissues, achieved with different types of vectors derived from adeno-associated virus (AAV). A gene therapy for Leber congenital amaurosis, a monogenic eye disease that causes blindness in early childhood, has led to dramatic improvements in eyesight in treated adolescents with advanced disease. Continued improvements offer hope that early childhood treatment before disease onset can eventually prevent disease progression and preserve vision. This has been demonstrated previously in a canine model of the disease where treated eyes continued to heal over eight years without any evidence of side effects (Bainbridge et al 2008; Hauswirth et al 2008; Maguire et al 2008). AAV vectors are particularly effective in neuronal and retinal gene delivery, and they result in no appreciable decline in long-term gene expression after months and years of available follow-up, setting the tone for other central nervous system and ocular diseases. Here comes hope (Wilson et al., 2017). 2008). ). A variety of gene therapy vectors are currently being used in gene therapy trials based on various human or animal viruses that are capable of incorporating and delivering a gene of interest into specific target cells. Although shielding methods have improved over the years to effectively protect naked or plasmid DNA from premature degradation, viruses are generally superior, and most gene therapy protocols use recombinant viral vectors. The clinical success stories of gene therapy described above are based on retroviral vectors and adeno-associated virus (AAV)-based vectors for the hematopoietic system.

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Other viral vectors in use today are mainly based on adenoviruses, but also on other viruses such as vaccinia/variola virus or herpes virus-based vectors. All of these vectors are highly immunogenic. Due to their high transgenic capacity and easy production of vectors, adenoviruses were widely used in clinical trials until, worryingly, one patient died after exposure to high doses of adenoviruses, which appeared to result in over- or under-stimulation of the immune system. As long as the inherent problem of the high immunogenicity of these vectors remains unresolved, their manufacture and application will remain strictly experimental and academic. Therefore, this chapter discusses the drug design and drug delivery of adeno-associated viral vectors and retroviral/lentiviral vectors, whose safety profile is more defined, the bioreactor scale-up method has been advanced, and most importantly, the results of clinical studies are finally obtained confirmed. available so that the early promise of gene therapy can be realized.

2 AAV vectors for gene therapy AAV vectors are derived from adeno-associated virus, a small single-stranded DNA virus that was originally described as a contaminant of adenoviral preparations, hence the name (Muzyczka and Berns 2001).

2.1

Property Overview

AAV vectors incorporate a range of popular properties that hastened their widespread use, although AAVs can be up to 4.5 kb in size, offering limited transgenic capacity to traditional AAV vectors containing single-stranded DNA. Well-documented benefits of AAV vectors include: AAV capsids exhibit very low immunogenicity while maintaining high chemical and physical stability, even at temperatures as high as 60 °C for prolonged periods of time. This allows stringent virus purification and concentration methods, a major advantage of bioreactor-scale production required for clinical use. AAV vector transduction results in sustained transgene expression in post-mitotic non-dividing cells, as shown in brain, retina, muscle, and liver. As a result, long-term gain-of-function was achieved in mice, dogs, and primates, thereby curing the underlying disease. Last but not least, the growing number of AAV subtypes and variants offers unique opportunities to specifically target selected tissues and cell types. Unfortunately, the production of AAV vectors has long been laborious. In parallel, suitable scale-up protocols have been developed for the production of AAV vectors on a clinical scale. Ultimately, bioreactor fabrication will be required to meet clinical AAV vector gene therapy needs.

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AAV structure

AAV is a small non-enveloped single-stranded DNA virus enclosed by a 20 nm diameter icosahedral capsid composed of the AAV structural proteins VP1, VP2 and VP3 in a ratio of 1:1:10. At least 12 different serotypes have been described to date, all of which are of human and nonhuman primate origin. The human serotype AAV-2 was considered the prototype strain and was used for most early studies in deciphering the molecular properties of AAV. AAV-2 is also the strain used in the original development of AAV as a vector system for mammalian gene transfer. AAV-2 consists of a 4.7 kb linear single-stranded DNA genome with two open reading frames rep and cap flanked by 145 bp inverted terminal repeats (ITRs). These repeats contain origins of DNA replication, packaging signals, and also mediate chromosomal integration. The ITR is the only AAV element retained in AAV vector construction. The reading frame on the left side of the genome, called rep, encodes four overlapping nonstructural proteins. Two of these, Rep78 and a C-terminal splice variant called Rep68, are expressed from the AAV p5 promoter. In addition, N-terminally truncated versions of the two proteins, called Rep52 and Rep40, respectively, are expressed from the internal p19 promoter. Rep78/68 is required for most steps of the AAV life cycle, including regulation of gene expression, chromosomal integration, and AAV DNA replication (Muzyczka and Berns 2001). Rep78/68 initiates AAV DNA replication at the hairpin ITR, which is also an important step in AAV vector production. The cap gene is expressed from the p40 promoter and encodes three capsid proteins VP1, VP2 and VP3. These arise from a combination of differential splicing and translation initiation.

2.3

AAV life cycle

The typical bipartite AAV life cycle has been well characterized in cell culture (Muzyczka and Berns 2001). Efficient AAV replication requires co-infection with an unrelated helper virus, such as adenovirus or herpes simplex virus. In the absence of a helper virus, AAV integrates into the host cell genome with a high preference for a specific region on human chromosome 19, called AAVS1. The AAV Rep78/68 protein appears to mediate the site-specificity of chromosomal integration by forming a ternary complex with the Rep-binding site on the AAV ITR and a cognate DNA element in the human genome that promotes tight integration. An intact copy of the integrated AAV genome can be rescued and amplified by superinfection with a helper virus, resulting in the release of progeny AAV by productive infection. Until recently, the in vivo mode of AAV infection remained largely unknown. Recently, the AAV-2 genome has been detected in human tissue samples by PCR analysis, and it has been shown that the AAV genome persists as circular nuclear episomes.

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Cellular receptors used by AAV

AAV infects a variety of cell types, as reflected in its interaction with ubiquitous cell surface receptors. Various glycan residues on the cell surface act as primary receptors. AAV-2 and AAV-3 are heparan sulfate proteoglycans. AAV-4 uses O-linked and AAV-5 N-linked sialic acid as primary receptors. The latter also binds to AAV-6 and AAV-1. Various co-receptors that facilitate cell entry through receptor-mediated endocytosis, such as B. Also described are fibroblast growth factor receptor 1 (FGFR-1), hepatocyte growth factor receptor, laminin receptor Integrins and various isoforms. For AAV-5, the platelet growth factor receptor appears to act as a co-receptor. So far, no specific cellular receptors have been assigned for other AAV serotypes. Currently, serotype-specific variations in cell tropism appear to reflect their interactions with specific receptor combinations.

2.5

AAV-vector production

Basically, AAV vectors are designed such that only inverted terminal repeats (ITRs) are retained as cis-acting DNA elements required for AAV vector replication and packaging. The AAV coding region with the corresponding promoter can be replaced with the foreign gene of interest. Up to 4.5 kb of transgenic DNA can be incorporated. These include promoters and other regulatory elements. For a more complete overview of the principles of AAV vector design, the reader is referred to a number of excellent reviews (Snyder and Flotte 2002; Grimm and Kay 2003; Choi et al. 2007). For the use of AAV vectors in clinical applications as medicinal products, various validated AAV vector production protocols have been developed and adapted according to Good Manufacturing Practice (GMP) standards. The suitability of a particular protocol for intended preclinical or clinical use depends on the scale of manufacturing, as described below. 2.5.1

Production of AAV vectors by co-transfection of packaging plasmids

The production of AAV vectors is made more difficult because the cell lines must not only provide the missing AAV genes rep and cap, but also the necessary helper virus functions. Currently, these requirements are met by transfection of plasmids with the well-characterized adenoviral (Ad) helper genes VA RNA, E2A and E4, and the AAV Rep and Cap genes, either on two separate plasmids or on a single helper . Build up the mix. AAV vector plasmids containing transgenes between AAV ITRs were co-transfected into 293 cells, a human cell line that constitutively expresses the remaining required Ad helper genes E1A and E1B (Figure 1). This results in the amplification and packaging of AAV vectors carrying the foreign DNA (Grimm et al. 1998; Xiao et al. 1998). plasmid transfection

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represent

build

AdV Accessibility ITR Representative

build

Wild-type AAV genome

Agency for International Trade Negotiations

genetically modified

Agency for International Trade Negotiations

recombinant AAV genome

Helper plasmid 293 cells (E1A/E1B)

Recombinant AAV (rAAV)-iodixanol-gradient

Affinity chromatography

raw extract

Highly purified rAAV

Figure 1 Generation and purification of AAV vectors. To generate recombinant AAV vectors (rAAV), the AAV rep and cap genes are replaced with the desired transgene, including a heterologous promoter and additional selection regulatory sequences. This leaves only the 145 bp inverted terminal repeats (ITRs) at both ends of the vector as the only AAV-derived DNA sequence. The Rep and Cap proteins required for AAV vector genome amplification and packaging are provided in trans by a transfected helper plasmid that also contains the required adenovirus type 5-derived helper functions E2A, E4ORF6, and VA RNA. Adenoviruses E1A and E1B were provided by 293 cells for rAAV packaging. Preparation of crude cell lysates and purification of supernatants containing rAAV vectors by iodixanol density centrifugation followed by affinity or ion exchange chromatography

This technique eliminates the problem of contamination of AAV vector preparations with infectious helper adenoviruses previously encountered after use of wild-type adenoviruses to provide helper functions. In addition, AAV vector titers were significantly increased compared to conventional adenovirus infection protocols. Sustained cell viability and prolonged viral gene expression contribute to this effect. Therefore, one of two or three plasmid co-transfection protocols represents the method of choice for laboratory-scale production of AAV vectors. Clinical applications require significant scale-up, which is difficult to achieve with protocols based on a DNA transfection step. 2.5.2

Scaling up AAV vector production

The need for bioreactor-scale manufacturing processes of AAV vectors for clinical use has led to some small but cumulative advances and conceptual changes in methodology, as described in more detail below.

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AAV packaging cell lines. In terms of reproducibility and ease of use, an ideal large-scale production method would be based on stable cell lines that consistently express in sufficient quantities all components required for the packaging process. Stable HeLa, 293, or A549-derived cell lines expressing AAV rep/cap genes have been extensively tested for bioreactor-scale production of AAV vectors. Unfortunately, most of these cell lines proved to be unstable. This appears to be due to the inherent cytotoxicity of the Rep component, which is enhanced by adenovirus E1A. Elaborate systems have been developed to control cytotoxicity through modulation of E1A and Rep gene expression. However, AAV vector yields remain low and difficult to control as the number of cell passages increases. It is generally accepted that AAV vector yields eventually decline after passage 40 to 50, further limiting the use of producer cell lines for bioreactor-based AAV vector production. Recombinant helper virus strains for AAV production. The inherent instability of AAV-Rep/Cap-containing cell lines has stimulated research into alternative concepts for large-scale AAV packaging. An obvious and elegant concept seems to be to incorporate the required AAV components rep and cap into a recombinant "super helper virus" based on adenovirus or herpes simplex virus (HSV). It was hypothesized that infection of cells with a helper virus expressing rep/cap would allow adequate regulation of all components required for AAV vector production. A major advance of this method is the ability to introduce all components of AAV vector production through an infection step rather than DNA transfection. The infection protocol is superior for scaling up bioreactors because many basic production principles are already established for vaccine production. Initially, many laboratories focused on constructing recombinant adenovirus-based vectors expressing the AAV rep/cap genes under the control of their cognate promoters. Unfortunately, most adenoviral recombinants generated were found to be unstable to the AAV Rep component, which was quickly deleted or otherwise inactivated. This was not entirely unexpected, as Rep has previously been shown to significantly impair adenoviral replication. HSV acts as an alternative helper virus to facilitate productive AAV infection with comparable efficiency to adenovirus. In contrast to recombinant adenoviruses, HSV-based recombinants readily tolerate integration of the AAV rep/cap gene cassette (Heilbronn et al., 2003). HSV DNA replication was not significantly affected by simultaneous Rep expression and AAV DNA replication, which seems to explain the observed stability of HSV recombinants expressing AAV rep/cap. The first wild-type HSV expressing recombinant rep/cap was further improved by introducing mutations in essential HSV genes not required for AAV vector production (Conway et al., 1999). This extends the time frame of the AAV production process and ensures that progeny HSV will not arise as contamination of AAV vector stocks. HSV expressing recombinant rep/cap can be used to infect cell lines carrying an integrated AAV vector genome that is rescued after infection with a recombinant helper virus (Toublanc et al., 2004). Alternatively, co-infection with a second recombinant herpesvirus carrying the AAV vector backbone and transgene can be used to prime the AAV vector

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Production. The latter design has proven to be more versatile, as only HSV components and not cell lines need to be reengineered to generate new AAV vectors. Furthermore, variations in infection dose and infection mode allow for fine-tuning of the production process. AAV vector yields have been reported to increase to 10,000 viral units per cell using recombinant HSV infection methods (Kang et al., 2009). Another infection-based approach for the large-scale production of AAV vectors uses combinations of recombinant baculoviruses to infect insect cells (Urabe et al., 2002). SF9 suspension cells were co-infected with three baculovirus vectors expressing AAV rep, AAV cap, and recombinant AAV vector backbone. This system has been shown to deliver large quantities of AAV-2-based vectors in bioreactors. However, the design of the individual components must be carefully retuned to accommodate alternate AAV serotypes and new vector backbones. Whether the adaptability of baculovirus-based production systems remains competitive with HSV-based production systems remains to be seen.

2.5.3

AAV vector purification

A fortunate circumstance for the production and scale-up of AAV vectors is the high stability of AAV virions, which allows stringent purification protocols. Dramatic improvements in AAV vector purification over the years have increased the burst size per cell in transfection-based systems from less than 1 to 100 vector transduction units per cell. The introduction of the infection-based production method described above increased the yield of 5,000-10,000 transduction units per cell. Purification of AAV vectors is not only necessary to remove residual cellular and viral contaminants, but also increases vector potency and increases the ratio of AAV infectious units to AAV particles. Traditionally, AAV vectors have been purified by cesium chloride gradient centrifugation. Unfortunately, AAV infectivity is lost with continuous exposure to CsCl and cannot be restored by dialysis or alternative purification steps to remove CsCl. Therefore, alternative methods were investigated as the first cleaning step. The method of choice today is the combination of AAV on an isotonic iodine density gradient (Zolotukhin et al. 1999). This protocol has been shown to effectively eliminate cell-derived contamination and empty AAV particles, resulting in a more favorable ratio of infectious to non-infectious vector particles. First, crude cell extracts are produced by deoxycholate (DOC)-mediated cell lysis, freezing and thawing, a combination of both, or an alternative method of mechanical cell lysis. Crude cell extracts are then treated with benzonase to digest cell-derived and unencapsulated viral nucleic acids. Extracts were loaded onto an Iodixonal step gradient for subsequent ultracentrifugation (Figure 2). Banded AAV virions were collected at the interface of the 54% and 40% iodixanol fractions and were sufficiently pure for subsequent heparan sulfate or ion-exchange chromatography steps. The choice of chromatography matrix is determined by the AAV serotype to be produced. The AAV vector is then dialyzed against a buffer of your choice. Alternatively, ultrafiltration can be used to further implement

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AdV auxiliary functions AAV-2 代表

AAV-2-Kappe

AAV-2 helper plasmid

AAV-1-Kappe

AAV-2-Kappe

AAV-5-Kappe

AAV-6-Kappe

AAV-8-Kappe

AAV-9-Kappe

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rAAV-2

rAAV-2/5

rAAV-2/6

rAAV-2/8

rAAV-2/9

ZNS, skull musk deer

ZNS

retina, airway epithelium

skeletal muscle

Lebel

Cardiac muscle, skeletal muscle

Figure 2 Sero(pseudo)type AAV vectors. To generate an AAV-pseudotyped vector containing a viral capsid derived from another AAV serotype, the AAV-2-derived vector backbone and rep gene of AAV-2 were retained. Therefore, the amplification of rAAV still depends on the AAV-2 Rep gene and AAV-2 ITR. Only the AAV-2-derived cap gene is replaced with the cap gene of another AAV serotype, such that the vector genome is cross-packaged into the capsid of the selected serotype. AAV serotype vectors have increased transduction efficiency of certain tissues or target cells compared to first-generation AAV-2 vectors, or have abnormal distribution patterns in target tissues with mixed cellular phenotypes

Concentration exchange buffer selection. A detailed protocol for the described cleaning protocol can be found in the current overview (Zolotukhin 2005). Vectors produced according to the latter protocol are highly concentrated and extremely pure, meeting the requirements for nontoxic, long-term gene expression in the brain, eye, and other tissues.

2.5.4

Quantification of AAV vector production

To quantify and compare the yield of AAV vectors, quantitative PCR determination of AAV genome copy number is the method of choice. While the amount of infectious virus is usually the preferred information, assessing the infectivity of a particular AAV vector preparation can be difficult. It is highly dependent on the readout of the transgenic product. Furthermore, even with sensitive detection of a specific transgene, the detected viral titer is critically dependent on the precise conditions used for transduction detection. These include seeding density of cells, growth rate and availability of suitable recipients in the test

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cell line. The increasing use of tissue-specific promoters and AAV pseudotyped capsids further limits the comparability of infectious AAV titers. Therefore, to compare not only the titers of AAV vectors from different batches, but also capsids of different transgenes and AAV serotypes, quantitative PCR determination of DNA copy number serves as the gold standard. Reliable and comparable results require AAV vector preparations that are sufficiently purified and free of contaminating DNA from the originally transferred packaging plasmid or helper virus. Therefore, a DNase or Benzonase step must be performed in the purification protocol. In any case, the so-called "DNA titer" of the purified AAV vector stock does not reflect the amount of infectious virus in a 1:1 ratio due to the excess of defective non-infectious particles. Extensive purification of prototype AAV-2 vectors using iodixon gradient centrifugation and chromatography typically results in a ratio of infectious virions to DNA-containing virions between 1:50 and 1:100. For crude cell extracts or partially purified vectors, this ratio may be lower. The detectability of infectious particles is limited by the availability of the required cellular receptors and the efficiency of intracellular transduction, including nuclear entry, envelope, and second-strand DNA synthesis. To evaluate the efficiency of the AAV vector production process, the burst size i was determined. H. The number of infection or cell transduction units per initially infected cell is a good measure. In summary, biotechnological optimization of AAV production protocols in recent years has resulted in an increase in the production efficiency of AAV vectors from an initial 1-10 transducing units/cell to more than 5,000-10,000 transducing units/cell, with more advanced HSV-based Or baculovirus-based production protocols followed by complex scale-up purification protocols (Urabe et al. 2002; Zolotukhin 2005; Kang et al. 2009).

2.6

Persistence and safety of AAV vectors

As explained at the outset, AAV vectors are characterized by stable long-term gene expression in quiescent or post-mitotic cells without a significant decrease in transgene expression. This is an ideal gene expression profile to complement defective genes. However, safety concerns regarding AAV persistence arise. Numerous studies have shown that wild-type AAV-2 integrates into the host cell genome with a preference for a specific region on human chromosome 19q13.42. The specificity of AAV wild-type integration is mediated by AAV Rep, which is absent in AAV vectors. AAV vectors have been shown to persist as nuclear episomes, often as large tandem DNA molecules that rarely integrate into the host cell genome. No site preference was observed in the occasional integration events described after selection in cultured cells. The predominantly episomal state of AAV vectors in quiescent cells ensures a high level of safety for clinical use (Miller et al., 2005).

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AAV splitting vector

AAV vectors have a more limited packaging capacity than adenovirus- or retrovirus/lentivirus-based vectors. Gene cassettes up to 4.5 kb including promoters and other regulatory elements represent maximum transgenic capacity. This is sufficient for many genes, especially when they are expressed as cDNA. Occasionally, however, expression of larger gene cassettes is unavoidable. To meet these requirements, so-called AAV split vectors were developed that split the transgene of interest into two parts. The first AAV vector carries the promoter and coding region at the N-terminus of the transgene product, followed by a splice donor site. The second AAV vector carries the coding region for the C-terminal portion of the transgene, preceded by a splice acceptor site. Both vectors are individually packaged. After co-infection of target cells with the two vectors, recombination events lead to vector multimerization, which involves head-to-tail concatenation of the vector genomes. In these vectors, the two exons of the shared gene are separated only by a splice site with an AAV ITR in between. Proper splicing of the pre-mRNA leads to translation of the precise fusion transgene. In animal models using this technique, transgene levels sufficient to correct muscle genetic defects were detected (Lai et al., 2005). The split vector technique is not as efficient as the traditional single AAV vector approach due to the need to infect both vectors simultaneously, but it does open up future prospects.

2.8

Dimeric, self-complementary (sc) AAV vector

A limitation of the application of AAV vectors is the relatively slow onset of transgene expression. This is because AAV vectors contain a single-stranded DNA genome that cannot be transcribed immediately after entering the nucleus. The host cell's DNA replication machinery is required to generate the double-stranded DNA template required for mRNA transcription. However, in quiescent or post-mitotic cells, the preferred target for AAV transduction, the host cell's DNA replication machinery is largely inactive. As a result, the onset of gene expression is delayed for days to weeks until a plateau of transgene expression is reached. To circumvent these problems, so-called "dimeric" or self-complementary (sc) AAV vectors packaged as pseudo-double-stranded genomes were developed. This is achieved by making a small deletion in the so-called terminal resolution site (trs) of one of the two ITRs so that only one ITR is correctly resolved during DNA replication and packaging of the AAV vector. As a result, dimeric genomes were generated in a head-to-head conformation joined by mutant ITRs. These can fold back into a double-stranded form, providing a suitable template for transcription immediately following transduction with the AAV vector. scAAV vectors show onset of gene expression within hours and reach their plateau levels within days after AAV vector-resting cell transduction. Increased efficiency comes with reduced package size:

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The packaging limit of scAAV vectors is 2.15 kb. This is sufficient for very small proteins and peptides, and is ideal for expressing small regulatory RNAs, such as those used for RNA interference. To meet the challenge of reduced transgenic capacity, small and optimized promoters and other regulatory elements must be employed. Despite these limitations, the favorable expression kinetics of scAAV vectors have accelerated the use of AAV vectors and stimulated new developments in broader and even more customized gene therapy applications (Choi et al., 2007).

2.9

Cellular Targeting Strategies for AAV

As noted above, a variety of naturally occurring AAV serotypes are available and AAV vectors derived from them have been constructed. The production of AAV variant serotype vectors has gained momentum with the advent of packaging of AAV-2-based vector genomes into variant AAV capsids (so-called AAV pseudotyped vectors). Technically, this is achieved by co-transfecting a helper plasmid containing the AAV-2 rep gene together with the cap gene of the variant AAV serotype (Figure 2). AAV-2 Rep is required for DNA replication of the AAV-2-based ITRs of the AAV vector backbone. AAV2 Rep packages AAV-2-based single-stranded vector DNA into preformed capsids formed from variant AAV serotype capsid proteins. Packaging plasmids for variant serotype capsids have been developed for virtually all naturally occurring serotypes known today and their many derivatives (Grimm et al., 2003). This speeds up the search for the best AAV serotype to target specific tissues or cell types in vivo. The nomenclature of the pseudotyped vector reflects the origin of the AAV ITR in the vector construct and corresponding rep gene on the one hand, and the serotype of the capsid protein on the other. For example, AAV-2/1 vectors contain AAV-2 serotype ITRs and AAV-1 serotype capsids. AAV 2/1 vectors exhibit an excellent targeted gene expression profile in the CNS, often resulting in higher levels of gene expression compared to traditional AAV 2/2 vectors. Other pseudotyped vectors such as AAV-2/8 or AAV-2/9 have been shown to produce higher and broader gene expression in the central nervous system as well as in liver, skeletal and cardiac muscles (Figure 2). ). Depending on the therapeutic goal, centralized or comprehensive and uniform gene expression is required. An example of the former is Parkinson's disease, and an example of the latter is α1-antitrypsin deficiency, a monogenic disorder that causes loss of systemic function. AAV-5 serotype vectors have been shown to be extremely effective in transducing retinal and airway epithelial cells. AAV-8 is well suited for gene transduction in the liver, and AAV-9 transduces cardiac and skeletal muscle. Studies describing the distribution of vectors of various serotypes according to site and mode of application are currently emerging at a high rate. A current overview can be found in Li et al. (2008). The repertoire of AAV capsids has been further refined and expanded by techniques leading to the construction of targeting chimeric vectors

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Evolutionary methods (Muller et al., 2003; Perabo et al., 2003). The technology uses a vector backbone containing a capsid protein library generated by fragmenting and shuffling known capsid protein repertoires. These are used to select chimeric AAV pseudotypes optimized for binding to selected target cells. These techniques have led to the isolation of AAV Cap gene variants, enabling the targeting of derived vectors to cardiac endothelial cells and various other cells and tissues previously resistant to AAV transduction. Analysis of chimeric Cap genes revealed that, in most cases, the preferred cell surface binding of the parental virus, such as the heparan sulfate proteoglycan of B. AAV-2, was an alternative receptor.

2.10

Future direction

At present, the development of AAV vector technology mainly has two directions. On the one hand, upscaling protocols for GMP-compliant bioreactor production are being refined to ensure efficient and cost-effective AAV vector production for clinical trials and the production of AAV vectors as pharmaceutical products. On the other hand, AAV vectors are being developed to improve cell targeting and cell entry based on rational approaches. The improvement of specificity is often accompanied by the decline of infection efficiency, which is an urgent problem to be solved in future clinical applications. An important step in this direction was recently reported: a detailed study of the fate of the vector after AAV entry revealed that cell surface bound AAV vector is not sufficient to indicate how much virus reaches the nucleus for gene expression. Studies with proteasome inhibitors have shown that most AAV capsids entering cells undergo ubiquitination and degradation and do not reach the nucleus. Based on the crystal structure of the AAV2 capsid, putative tyrosines in the exposed capsid domain were identified as preferred target sites for ubiquitination. These individual amino acids are exchanged by mutation, which then leads to a dramatic increase in intracellular AAV vector stability and up to a 100-fold increase in transgene expression (Zhong et al., 2008). Comparable amino acid exchanges in other serotypes showed the same effect (Petrs-Silva et al., 2009). These and other continuing developments in AAV vector technology will make it easier to manufacture AAV vectors as drugs. Some recent clinical studies have achieved notable success and will further accelerate the development of this field.

3 Retroviral Vectors for Gene Therapy The retroviral family of enveloped RNA viruses includes six distinct genera found in avian and mammalian species (Goff 2001). These are divided into simple retroviruses, also known as oncogenic retroviruses, and so-called complex retroviruses

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These include lentiviruses and spumaviruses. While oncogenic retroviruses contain only the minimum essential genes gag, pol, and env, complex retroviruses encode additional regulatory and accessory proteins required to fine-tune viral replication and persistence. Retrovirus-derived vectors used in clinical gene therapy are derived from oncogenic retrovirus or lentivirus families.

3.1

Retrovirus structure and life cycle

All members of the retrovirus family carry the so-called gag, pol and env genes. The gag gene encodes capsid, matrix, linker and nucleocapsid proteins. These are components of the retroviral core, closely related to both copies of the viral RNA genome. Gag-derived proteins are produced by proteolytic cleavage of the gag precursor protein. The pol gene encodes reverse transcriptase, integrase, and retroviral protease. Typically, the gag-pol precursor protein is used as a template for further processing. The env gene encodes the envelope glycoprotein required for viral attachment and entry. The retroviral life cycle involves cell entry, reverse transcription of the RNA genome into complementary (c)DNA, and integration of the cDNA into the host cell genome. Transcription of full-length progeny RNA and mRNA encoding regulatory and structural proteins begins from the integrated proviral DNA copy. Accumulation of new virus particles occurs at the plasma membrane. The cis-acting sequences required as regulatory signals are contained within or near long terminal repeats (LTRs) flanking the retroviral genome.

3.2

Design and development of oncoretroviral vectors

All viral regulatory proteins involved in cell entry, reverse transcription and integration are incorporated into the virus particle. Therefore, when developing oncogenic retroviral vectors, the entire coding region of the viral protein can be removed to accommodate transgenes of interest up to 8 kb. Only the packaging signal, viral LTR, and adjacent elements necessary for reverse transcription and integration are retained as viral sequences in the resulting recombinant vector construct. The first generation of oncogenic retroviral vectors retained the entire LTR, including the retroviral enhancer and promoter sequences. Some of these can also be replaced by various heterologous enhancer/promoter combinations, such as inducible or tissue-specific elements. So-called packaging or producer cell lines are used to produce oncogenic retroviral vectors expressing the desired retroviral proteins from gag, pol and env. Packaging cell lines have been developed to prevent unwanted homologous recombination between the transgene-containing vector and the packaging construct

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are produced by expressing gag/pol and env from different chromosomal loci. This reduces the risk of homologous recombination leading to replication-competent retroviruses (Figure 3A).

3.3

Self-inactivating (SIN) Oncoretroviral Vectors

The use of first-generation oncogenic retroviral vectors has been associated with many problems. Although the packaging genes were placed on two different plasmids and cell clones containing the genes at different genomic loci were selected (see section 3.2), the vector backbone for expressing the transgene was not completely free of redundant DNA sequences. Occasionally, but inevitably, this leads to the production of replication-competent viruses through homologous recombination. Another perturbation is the inherent tendency of retroviruses to integrate into active chromosomal loci. This repeatedly leads to unwanted and uncontrolled LTR-mediated gene activation in host cells. Numerous side effects and adverse events due to uncontrolled activation of oncogenes have been identified in animal models and clinical studies. This led to the development of several cases of childhood leukemia whose SCID had been successfully treated many years earlier. In all reported cases, retroviral vector-mediated activation of nearby oncogenes was demonstrated. Therefore, in order to advance retroviral vector-mediated gene therapy more safely, measures to control the risk of insertional mutations are imperative. Finally, a self-inactivating (SIN) oncoretroviral vector was developed, the concept and design of which are described below. The first step in developing oncogenic retroviral SIN vectors was to minimize sequence homology between the vector and helper constructs while increasing full-length RNA transcription for more efficient vector packaging. A non-retroviral promoter was chosen to replace the U3 promoter region within 50 LTR. The U3 region normally directs the transcription of the full-length RNA for viral packaging. Replacing the U3 region within the 50 LTRs with a heterologous promoter, such as a strong cytomegalovirus (CMV) promoter, greatly increased transcript levels at inherently low LTRs. Hybridization of the CMV/LTR promoter combination resulted in increased vector titers. When a retroviral vector infects a target cell, reverse transcription and second-strand DNA synthesis involve copying the U3 region from the 30th LTR to the 50th LTR. This typically results in recovery and repair of 50 LTR (Figure 3B). To avoid correction of the LTR, a retroviral SIN vector was generated with a comparable deletion in the U3 region of the 30-LTR as in the 50-LTR. When the deleted U3 region of the 30-LTR is copied to the 50-LTR, it replaces and deletes the CMV promoter. Due to this deletion, the 50 LTR remains transcriptionally inactive after integration and cannot be rescued again. The transgene is expressed from an internal heterologous promoter of your choice without further involvement of the LTR. The SIN concept minimizes two risks of oncogenic retroviral vectors: the risk of insertional mutagenesis caused by functional retroviral LTRs and the risk of vector mobilization.

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Figure 3 Retroviral vectors. (a) In an oncogenic retroviral vector, the viral genes gag, pol, and env are replaced by the transgene of interest. The retroviral LTR and adjacent packaging signal c. The Gag/Pol and Env proteins are provided in trans by constructs integrated at different locations in the genome of the packaging cell line host cell. (b) In a self-inactivating (SIN) vector, the U3 region of 50 LTR is replaced by a heterologous enhancer, such as that of human cytomegalovirus (CMV), and the U3 region of 30 LTR is partially deleted. The CMV/LTR hybrid promoter drives transcription of the vector genome for vector packaging. After vector entry into target cells and reverse transcription, the partially deleted 30 LTR is replicated into a 50 LTR without promoter and enhancer sequences. This results in deletion of the CMV promoter element and transcriptionally inactivates the 50 LTR. Internal heterologous promoters mediate transgene expression from vectors integrated into the host cell genome. Defective LTRs are no longer able to activate cellular genes adjacent to the integration site

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Design and development of lentiviral vectors

Originally, retroviral vectors were derived from simple oncogenic retroviruses such as Moloney murine leukemia virus (MoMLV), a member of the gamma retrovirus family. Unfortunately, these viruses are unable to infect non-dividing differentiated cells. This property has been attributed to the inability of retroviral preintegration complexes to enter the nucleus unless the nuclear envelope is disrupted during mitosis (Miller et al., 1990). The preference for dividing cells could be exploited in cancer therapy or in the treatment of rapidly proliferating cells, such as those of the hematopoietic lineage. Because most target cells in the human body are dormant or slowly dividing, the range of application of conventional (cancer) retroviral vectors remains limited. This deficiency has spurred interest in the development of integrating retroviral vectors based on lentiviruses such as equine infectious anemia virus (EIAV) or primate lentiviruses such as human immunodeficiency virus (HIV). As a wealth of data on their molecular biology became available, research quickly focused on HIV-based lentiviral vectors. HIV can infect non-dividing cells in the G0 or G1 phase of the cell cycle. The HIV preintegration complex (PIC) has the ability to penetrate the intact nuclear envelope. In addition to the typical retroviral genes gag, pol, and env, HIV-1 synthesizes six accessory proteins, Tat, Rev, Nef, Vif, Vpr, and Vpu, that regulate viral transcription, replication, and persistence. Among them, Rev is the only conserved accessory protein in lentiviral vector design. Rev is required for nuclear export of unspliced and singly spliced HIV RNA. Despite their considerable complexity, the development of lentiviral vectors largely mimics the principles established for vectors derived from conventional (onco)retroviruses. In the vector backbone, all viral sequences were removed except for the essential cis-acting sequence of the LTR and the packaging signal located in the adjacent untranslated region. In addition, the rev-responsive element (RRE) was preserved to ensure Rev binding (Figure 4). Rev protein is provided in a separate packaged structure. Two separate packaging constructs are used for the gag/pol and env genes. To overcome the host range limitations of the glycoprotein encoded by HIV-1 env, lentiviral vectors can be pseudotyped with the vesicular stomatitis virus (VSV-G) glycoprotein, which has a broad host range. VSV-G encoding env is expressed by replacing the HIV env gene in the appropriate packaging construct (Figure 4). Similar to the concept established for oncogenic retroviral vectors, the expression of viral proteins is driven by heterologous promoters (such as the CMV promoter) to ensure high levels of protein expression and minimize the gap between the lentiviral vector genome and packaging constructs. sequence homology. Since then, many improvements to the original lentiviral vector design have increased the robustness of the vector. The U3 portion of the 50 LTR can be replaced by the CMV promoter, resulting in a CMV/LTR hybrid promoter. This improves the production efficiency of lentiviral vectors and, more importantly,

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Figure 4 Lentiviral vectors. As with oncogenic retroviral vectors, the viral genes gag, pol, and env are replaced with the transgene of interest. The LTR and packaging signal c are retained for vector amplification and packaging. Among lentiviral accessory proteins, only the rev response element (RRE) remains as an additional cis-acting sequence required for nuclear export of unspliced and single-spliced viral RNA in the presence of Rev protein. Rev is represented by one of three wrapping structures. The other two encode Gag/Pol and Env proteins, respectively. For pseudotyping, the lentiviral Env protein is often replaced by the vesicular stomatitis virus (VSV-G) glycoprotein, which targets multiple cell types

Makes the LTR independent of transactivation of the viral Tat protein. To improve nuclear import of proviral DNA, HIV's central polypurine region (cPPT) was added to the vector as a cis-acting sequence. The improved biosafety concept described above for SIN vectors for oncogenic retroviral vectors has been adapted for lentiviral vector design. A large deletion in the transcriptional activation sequence of the U3 portion of the 30-LTR results in post-retrotranscriptional inactivation of the 50-LTR in target cells.

3.5

Retroviral Integrated Security

The first successful clinical trial of gene therapy resulted in a cure for SCID type X1. Unfortunately, a few years later, a case of leukemia apparently caused by a retrovirus led to a major setback for the field. In the first Paris study, 4 in 10 children developed T-cell leukemia about three years after initial treatment and cure of the disease. Among these, chromosomal integration of retroviral vectors was found

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Proximate to the promoter region of the LMO2 proto-oncogene. This results in an upregulation of the LMO2 gene product, leading to uncontrolled stimulation of cell division (Hacein-Bey-Abina et al., 2003). These disturbing findings accelerated the development of SIN retroviral vectors (see 3.3) and led to intensive studies of the genome-wide distribution patterns of oncogenic retroviral integration sites. These screens revealed that the transcription initiation region is the preferred target site for integration of MoMLV-based oncoretroviral vectors. Integration sites for HIV-based lentiviral vectors also favor genomic transcription units, but are more evenly distributed throughout the gene (Mitchell et al. 2004; Fischer and Cavazzana-Calvo 2005). The vector used in the first clinical trials was a first-generation oncogenic retroviral vector. Extensive animal studies have demonstrated a significantly improved safety profile of self-inactivating (SIN) oncoretroviral vectors. Cell culture-based high-throughput assays have now been developed to assess the safety of retroviral vector integration (Modlich et al., 2006). In addition, lentiviral vectors are being investigated for transduction of hematopoietic stem cells, but the risk of insertional mutagenesis needs to be noted. A recent study documented a high incidence of liver cancer in mice infected neonatally or in utero with an EIAV-based lentiviral vector (Themis et al., 2005). Another major concern with the safety of retroviral vector formulations involves vector mobilization and the emergence of replication competent viruses (RCV). These typically occur when using first generation retroviral vectors. The SIN design of oncogenic retroviral and lentiviral vectors is thought to eliminate the risk of vector mobilization and homologous recombination leading to RCV generation. Whether SIN vectors deliver on their promise of improved safety remains to be seen in ongoing clinical trials.

3.6

Production and stability of retroviral vectors

A major obstacle associated with the use of oncogenic retroviral and lentiviral vectors is the low titer viral stocks produced by currently available production systems (Rodrigues et al., 2007). This is exacerbated by the inherent instability of retroviral vector particles released into the cell supernatant. Using packaging cell lines or transient transfection of plasmid-helper constructs, cell supernatants with maximal vector titers in the range of 105-107 infectious particles per mL can be obtained. While these titers may be sufficient for a variety of ex vivo applications, most gene therapy applications require further concentration of the supernatant. Furthermore, cell supernatants, regardless of their concentration levels, are inherently impure and require further purification. For cell supernatants containing unpurified retroviruses, adverse side effects such as cytotoxicity, inflammatory or immune responses, and loss of vector transduction efficiency have been reported due to the presence of transduction inhibitors.

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Purification and amplification of retroviral vectors

An important prerequisite for purification of retroviral vectors is knowledge of the optimal conditions for maintaining their stability (Segura et al., 2006). After budding from the producer cell line, retroviral particles rapidly lose infectivity at 37°C. Their half-life in culture medium is approximately eight hours. Thus, at harvest, 24-48 hours after production begins, only a fraction of the initially produced retroviral particles remains infectious. This, together with the production of defective viral particles, results in a typical ratio of total particles to infectious retroviral particles of over 100:1. Therefore, cleaning protocols must be brief and usually performed at 4 °C. In addition, the inherent instability of retroviral particles outside the pH range of 5.5-8.0 and their sensitivity to high salt concentrations must be considered. Based on these limitations, retroviral vector purification protocols based on centrifugation-based methods, membrane separation steps and chromatography, or a combination thereof have been developed. Baseline titers tend to be low due to the release of retroviral particles into the production medium. On the other hand, there is no need to destroy the producer cells, which reduces possible contamination by cell debris. Low-speed centrifugation or microfiltration through membranes with a pore size of 0.45 mm can be used to separate viral particles from isolated producer cells and cell debris. Virus particles can then be pelleted by ultracentrifugation for vector concentration. Compared to AAV non-enveloped vectors, the utility of ultracentrifugation for retroviral purification is limited for several reasons. Most importantly, both cesium chloride and sucrose, commonly used in density equilibrium centrifugation, lead to a significant loss of infectivity (Powell et al., 2000). This applies especially to non-pseudotyped retroviral vectors with wild-type Env. These retain the authentic surface domains of wild-type glycoproteins, which are easily lost during ultracentrifugation. Furthermore, ultracentrifugation-based methods are time-consuming and difficult to scale up. The ultrafiltration process appears to be an attractive alternative. They are fast, robust and can be used for the concentration of vector preparations and simultaneous dialysis to adjust the buffer conditions required for downstream chromatographic purification. At laboratory scale, centrifugation-dependent ultrafiltration units with molecular weight cut-offs in the 20-500 kDa range or stirred vessels with high recovery have been used. Membrane geometries commonly used for scaling are tangential flow devices such as flat cassettes or hollow fibers (Rodrigues et al., 2007). Conditions for further purification of vector preparations by chromatographic methods are very stringent because retroviral vectors are sensitive to the methods used for desorption from various chromatographic matrices. For example, affinity adsorbents for conjugated antibodies to viral Env proteins have not been described for retroviral vectors, apparently due to the harsh conditions required to disrupt the antigen-antibody complex. However, the purification of biotinylated

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Retroviral particles on streptavidin-biotin affinity columns and particles with histidine-tagged Env proteins on metal affinity columns have been described. Standard reagents used to elute proteins from their respective affinity matrices (guanidine hydrochloride/urea or imidazole/EDTA) readily inactivate retroviral vectors. Therefore, use an alternative desorbent or immediately remove the desorbent by dialysis. Anion exchange chromatography desorbed at high salt concentration (1 M NaCl) has been successfully used to purify oncogenic retroviral and lentiviral vectors. Size exclusion chromatography suffers from low throughput as an alternative strategy but is commonly used as the final polishing step in retroviral vector purification (Rodrigues et al. 2007).

3.8

Vector quantization and quality assessment

The quality of purified retroviral vector preparations can be described in terms of dose, potency, purity, and safety criteria (Rodrigues et al. 2007). Potency represents the ability of target cells to express the transgene following retroviral transduction and can only be measured by assays designed specifically for the transgene and target cells. Parameters describing vector doses are total particle number, biological activity as measured by infectious particle content, and their protein content. Total particles can be quantified using several methods with different diagnostic value. A sensitive and rapid assay measures carrier RNA molecules by real-time quantitative PCR after reverse transcription. This approach often underestimates the number of total viral particles because empty particles are not considered. Quantification of viral antigens such as p24 Gag protein of HIV and p30 Gag protein of oncoretroviral vectors by standard ELISA assay or measurement of reverse transcriptase activity can be used to calculate the total number of particles. In contrast, electron microscopy of negative-stained viral preparations is a tedious and impractical method for routine quantification, but can serve as an important tool for assessing the integrity of purified vector preparations (Segura et al., 2005). A commonly used method for detecting infectious, transducible viral particles is based on determining the expression of marker genes (such as GFP or lacZ) or antibiotic resistance genes in transduced cell lines. Most developing retroviral vectors carry one of these marker genes. However, these assays often underestimate actual titers. Due to differential diffusion and decay of virions in the infection medium, only a fraction of initially active virions in vector strains ultimately transduce target cells (Andreadis et al., 1997). Furthermore, viral titers are mainly dependent on the parameters used for transduction testing, such as seeding density, growth rate, and receptor availability of the B. target cell line. In addition, changes in incubation time, incubation volume, and polybrene concentration (Andreadis and Palsson 1997) can affect titer measurements based on transgene expression. choose,

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As a measure of biological activity, quantification of the proviral genome inserted into the target cell genome (Sastry et al., 2002) or the level of transgene mRNA by real-time PCR (Lizee et al., 2003) can be used. Usually , a combination of at least two of the above methods should be used, both to determine the amount of active, transcriptionally competent virus and to measure total viral particles in the vector preparation. Major contaminants in retroviral vector preparations include proteins, DNA, and transduction inhibitors of various chemical constituents. Protein contaminants come primarily from serum used to supplement growth media or excreted into the media by producer cells. One difficulty in identifying protein contaminants is the inherent property of retroviruses to incorporate host cell proteins during budding between the viral envelope lipid bilayer and the internal nuclear structure. Protein contamination is usually determined by SDS-PAGE followed by silver staining (Segura et al., 2005). Contaminating DNA from producer cells or transfected packaging plasmids can be detected by standard DNA hybridization techniques using plasmid or host cell DNA probes. Tolerable limits for DNA contamination vary depending on the vector being manufactured and the intended application. As a guideline for clinical trials, the FDA recommends an upper limit of 10 ng of foreign DNA per vector dose. To minimize DNA contamination, a DNase digestion step should be included at the beginning of the purification process. A major safety concern with retroviral vector formulations is contamination with replication competent viruses (RCV). Methods for detecting RCV in retroviral preparations are technically demanding and time-consuming. They are based on multi-passage expansion of RCV in permissive cell lines combined with detection by cell-based assays or more sensitive PCR-based techniques. For more detailed information, the reader is referred to FDA (2001).

3.9

Future directions for retroviral vectors

Vectors based on classical oncoretroviruses and lentiviruses have emerged as powerful tools for gene therapy of various disease entities. X-linked and ADA forms of SCID have been successfully treated in clinical trials using classical oncogenic retroviral vectors. Reported shortcomings in leukemia development have been shown to be due to activation of oncogenes through LTR-mediated activation, a well-documented problem with early vector design. The improved safety and efficiency of the SIN vector concept has been demonstrated in extensive animal studies, making it safe for clinical studies. It is hoped that SIN vectors will also address the safety concerns associated with generating replication-competent viruses. The concept of SIN vectors can be expanded by adding chromatin insulators or matrix attachment regions to prevent vectors from transactivating genes in adjacent cells. To further reduce the risk of insertional mutagenesis, lentiviral-based vectors,

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Foamyviruses and non-murine retroviruses are being explored because they have been shown to integrate into larger regions of the genome (Kohn and Candotti 2009). Last but not least, to circumvent the problem of insertional mutagenesis, episomal integrase-deficient vector variants were investigated. Since highly proliferative cells and tissues rapidly lose these vectors, the goal of episomal vector persistence is focused on long-term gene expression in post-mitotic cells (Philpott and Thrasher 2007), which has been achieved by AAV vectors with excellent long-term performance result. As retroviral vectors continue to advance to phase III clinical trials, the need to address production-related safety issues has pushed research on retroviral vectors to a higher technical level. The required larger vector doses require improved production methods and downstream processing. Two main directions are being explored: production of vectors in serum-free media to reduce the amount of serum proteins previously shown to elicit an unacceptable immune response, and replacement of adherent cells with suspension culture-based production methods. Suspension culture combines ease of scale-up with reduced contamination by proteoglycans derived from secreted extracellular matrix proteins (Merten 2004). The combination of ongoing retroviral vector development will significantly improve the genetic safety of retroviral vector applications and the convenience and safety of bioreactor-scale production for clinical use.

4 Outlook The basis of gene therapy is safe and permanent gene transfer and safe gene expression. Because chromosomal integration of gene therapy vectors has an inherent risk of insertional mutagenesis, in situ gene repair methods have been extensively studied. Two directions are being explored: one is to locate the vector to a supposedly safe site in the human genome; the other is to directly target the defective gene sequence to repair specific genetic defects. To this end, DNA sequence-specific designer zinc finger nucleases are being developed. These are associated with appropriate endonucleases that induce double-strand breaks near the intended target site for gene repair. Simultaneously, a repair matrix containing homologous sequences surrounding an intact copy of the target sequence is transferred into the cell as a repair template to induce homologous recombination at sites of DNA strand breaks. Gene transfer of the repair matrix can be accomplished via non-integrating viral vectors such as AAV or via non-integrating retroviral or lentiviral vectors. Although the efficiency of the described DNA repair process is still relatively low, the technology has improved rapidly in recent years. Stem cell gene therapy animal models of the hematopoietic system and other rapidly dividing cell types are being studied extensively (Bohne and Cathomen 2008). In situ gene repair utilizes cellular gene repair mechanisms to perform templated gene repair. A prerequisite is actively proliferating target cells. However, most cell types in the human body cycle slowly or do not divide at all, and are therefore unsuitable for the described method of gene repair

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more than. To enable expression of full copies of the mutated gene in these differentiated cells, episomal persistent vectors such as AAV are used, resulting in long-lived and safe gene expression profiles.

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Perabo L, Buning H, Kofler DM, Ried MU, Girod A, Wendtner CM, Enssle J, Hallek M (2003) In vitro selection of viral vectors with improved tropism: Adeno-associated virus display. Mol Ther 8:151–157 Petrs-Silva H, Dinculescu A, Li Q, Min SH, Chiodo V, Pang JJ, Zhong L, Zolotukhin S, Srivastava A, Lewin AS, Hauswirth WW (2009) Efficient transduction of tyrosine Mouse retinas of mutant AAV serotype vectors. Mol Ther 17:463-471 Philpott NJ, Thrasher AJ (2007) Gene therapy using non-integrating lentiviral vectors. Hum Gene Ther 18:483–489 Powell SK, Kaloss MA, Pinkstaff A, McKee R, Burimski I, Pensiero M, Otto E, Stemmer WP, Soong NW (2000) Retroviruses bred by DNA shuffling for improved stability and processing Yield. Nat Biotechnol 18:1279–1282 Rodrigues T, Carrondo MJ, Alves PM, Cruz PE (2007) Purification of retroviral vectors for clinical use: biological implications and technical challenges. J Biotechnol 127:520–541 Sastry L, Johnson T, Hobson MJ, Smucker B, Cornetta K (2002) Lentiviral vector titration: a comparison of DNA, RNA, and marker expression methods. Gene Ther 9:1155–1162 Segura MM, Kamen A, Trudel P, Garnier A (2005) A novel purification strategy for retroviral gene therapy vectors using heparin affinity chromatography. Biotechnol Bioeng 90:391–404 Segura MM, Kamen A, Garnier A (2006) Downstream processing of oncoretroviral and lentiviral gene therapy vectors. Biotechnol Adv 24:321-337 Snyder RO, Fleet TR (2002) Production of clinical-grade recombinant adeno-associated viral vectors. Curr Opin Biotechnol 13:418-423 Themis M, Waddington SN, Schmidt M, von Kalle C, Wang Y, Al-Allaf F, Gregory LG, Nivsarkar M, Themis M, Holder MV, Buckley SM, Dighe N, Ruthe AT, Mistry A, Bigger B, Rahim A, Nguyen TH, Trono D, Thrasher AJ, Coutelle C (2005) Tumor initiation following delivery of non-primate lentiviral gene therapy vectors to fetal and neonatal mice. Mol Ther 12:763–771 Toublanc E, Benraiss A, Bonnin D, Blouin V, Brument N, Cartier N, Epstein AL, Moullier P, Salvetti A (2004) Identification of recombinant adeno-associated replication-defective herpes simplex virus using stable production Particle assembly of cell line virus type 2 (rAAV2). J Gene Med 6:555-564 Urabe M, Ding C, Kotin RM (2002) Insect cells as factories for the production of adeno-associated virus type 2 vectors. Hum Gene Ther 13:1935–1943 Wilson JM, Gansbacher B, Berns KI, Bosch F, Kay MA, Naldini L, Wei YQ (2008) Good news for the field of clinical gene transfer. Hum Gene Ther 19:429-430 Xiao J Virol 72:2224-2232 Zhong L, Li B, Mah CS, Govindasamy L, Agbandje-McKenna M, Cooper M, Herzog RW, Zolotukhin I, Warrington KH Jr, Weigel-Van Aken KA, Hobbs JA, Zolotukhin S, Muzyczka N, Srivastava A (2008) Next-generation adeno-associated virus 2 vectors: point mutations at tyrosine lead to efficient transduction at low doses. Proc Natl Acad Sci USA 105:7827-7832 Zolotukhin S (2005) Production of recombinant adeno-associated virus vectors. Hum Gene Ther 16:551-557 Zolotukhin S, Byrne BJ, Mason E, Zolotukhin I, Potter M, Chesnut K, Summerford C, Samulski RJ, Muzyczka N (1999) Purification of recombinant adeno-associated virus using a new method improves infectious titers degree and output. Gene Therapy 6:973-985

Pulmonary Administration: Inhaled Drugs Andreas Henning, Stephanie Hein, Marc Schneider, Michael Bur, and Claus-Michael Lehr

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introduce . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 173 Principles of Aerosol Transport. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 174 2.1 Inhalation therapy. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 174 2.2 Lung structure. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 175 2.3 Aerosol deposition. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 176 2.4 Lung clearance. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 177 3 Methods of Pulmonary Drug Delivery. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 179 3.1 Asthma/COPD. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 179 3.2 Immunosuppressants. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 180 3.3 Vaccines. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 181 3.4 Anti-infectives. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 182 3.5 Lung gene therapy. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 183 3.6 Lung Cancer Treatment. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 185 4 Future prospects. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 186 References. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 187

Abstract: Humans have inhaled substances for medicinal and other reasons for thousands of years, leading in particular to the cultural manifestations of smoking tobacco and opium. The concept of pulmonary drug delivery, including inhalation devices and drug formulations, has been and continues to be developed over time. State-of-the-art instruments can even adapt the inhalation process to the patient's breathing pattern, allowing for individualized drug application. Pulmonary drug delivery offers promising advantages over "classical" drug delivery via oral or transdermal routes, as is the case

centimeter. Lehr (*) Biopharmaceuticals and Pharmaceutical Technology, University of Saarland, P.O. PO Box 15 11 50, 66041 Saarbrücken, Germany Email:[email protected]

M. Schäfer-Korting (Hrsg.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_6, # Springer-Verlag Berlin Heidelberg 2010

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This is reflected in the growing interest and number of marketed products for inhalational treatments. However, efficient clearance mechanisms in the lung still limit the utility of many therapeutic concepts. Therefore, current research and development in the field of pulmonary drug delivery is aimed at overcoming and controlling drug release from the intended target. Some of the most promising future drug delivery concepts are presented and discussed here to give readers an insight into this emerging field of medicine. Key words aerosol deposition mucociliary clearance macrophage clearance telomerase inhibition inhalation vaccination

Abkürzungen 2-OMR A549 APCs ATD AUC BALT Calu-3 CF CFTR CLIJ COPD DOTAP DPI DPLC DPPC DSPC DSPE FDA ICRP MC MDI PEG PEI PLGA PTEN siRNA SLIT SLM TAT WHO

Antisense Oligonucleotide 20-O-Methyl RNA Lung Cancer Cell Line (CCL-185; ATCC) Antigen Presenting Cell Area Under the Curve for Anti-TB Drugs Bronchoalveolar Lymphoid Tissue Human Epithelioid Lung Cancer Cell Line (HTB-55; ATCC ATCC) N-[1-(2,3-dioleoyloxy)]-N,N,N-trimethyl-N-[1-(2,3-dioleoyloxy)]-N,N,N-trimethyl Propane Methyl Sulfate Dry Powder Inhaler Dipropionate Dilauroyl Phosphatidyl Choline Dipalmitoyl Phosphatidyl Choline Distearoyl Phosphatidyl Choline Distearoyl Phosphate Diethanolamine United States Patents Food and Drug Administration International Radiation Defense Committee Mucociliary Clearance Metered-Metered Inhaler Polyethylene Glycol Polyethyleneimine Polylactide-Co-Glycolide Phosphatase and Tensin Homologue: Tumor Suppressor Gene Small Interfering RNA Sustained Release Lipid Inhalation Targeting Solid Lipid particle human immunodeficiency virus 1 transactivator protein World Health Organization

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1 Introduction Thousands of years ago, humans used the airway to administer drugs. Ancient Egyptian physicians used the energy of hot stones to vaporize alkaloids in plants to induce patients to inhale the active ingredients (Sanders 2007). Indian and Native American shamans knew datura for its anti-asthmatic properties when the leaves were smoked in a pipe or simply burned in a small room (Dessanges 2001). In addition, smoking opium, which contains powerful pain-killing alkaloids derived from poppies, has a long tradition in Chinese culture, although the medical aspects were not the main focus of this case. It is therefore not surprising that inhalation techniques, as well as knowledge about therapeutic inhalation in general, have and will continue to develop over time. To make inhalation more effective, the ancient simple whistle was quickly replaced by more sophisticated inhalation instruments. The first illustration of an inhaler appears in Christopher Bennet's Theatri Tabidorum (Bennet 1654) in 1654. However, the term "inhaler" was introduced in 1778 by the English physician John Mudge, who suggested inhaling opium vapors to treat coughs (Mudge 1778). Along with continued technical optimization, a better understanding of the principles of pulmonary drug delivery has also led to a clearer and more precise vocabulary in the field. According to Aiache, R. Whitlaw and E. Gray Patterson defined the term "aerosol" in 1932 in terms of "aer" (air) and "sol" (solution) (Aiache 1990). Until now, the terms "nebula", "micronebula", "fog" and "smoke" have been used imprecisely and often confusingly. From today's perspective, this is astonishing, as defining a vocabulary is one of the basic requirements in every field of modern science. The drive to improve drug inhalation, especially in the 19th century, was therapeutic wasting or "lung wasting", whereas today the focus is more on treating asthma, chronic obstructive pulmonary disease (COPD) and cystic fibrosis (CF) .The number of approved inhalation products has steadily increased in recent decades and several new therapeutic approaches have emerged, such as vaccination via the pulmonary route (Bivas-Benita et al., 2005; Lu and Hickey, 2007) and highly specific Lobar targeting (Selting et al. 2008) has shown promising results. While there have been some excellent reviews and textbooks on pulmonary drug delivery (Bechtold-Peters and Luessen 2007; Patton and Byron 2007), this chapter focuses on cutting-edge technology that goes beyond the well-known devices and formulated doses used in metered inhalation device (MDI), dry powder inhaler (DPI) or nebulizer. Using examples, the principles of pulmonary drug delivery and current developments and concepts in this emerging field are explained and discussed. Additional references enable the well-intentioned reader to find detailed information in a wide range of publications, reviews, etc.

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books. Finally, a prospective outlook should reflect the growing demand and future developments in inhalational therapies.

2 Aerosol delivery principle 2.1

inhalation therapy

Pulmonary drug delivery has received increasing attention due to the lung's unique barrier properties (Figure 1) and enormous surface area (~140 m2). Today, inhalation therapy adequately treats a wide variety of respiratory disorders (Groneberg et al., 2003). For many asthma and COPD patients, drug inhalation is even suitable as the only form of therapeutic intervention. In fact, this offers important advantages in terms of patient compliance and the overall benefit of treatment. Depending on the current state and general severity of the disease, one or more daily inhalations of the drug are required. It is especially important to note that correct drug application, in this case inhalation maneuvers, leads to better therapeutic outcomes and thus lower application rates (Serra-Batlles et al 2002; Welch et al 2004; Booker 2005 ). . Therefore, any patient receiving inhaled drug therapy should receive comprehensive training to optimize drug deposition in the airways. the bottom line

Figure 1 Alveolar air-blood barrier (TEM image). The unique barrier properties of the lung are related to the extremely thin air-blood barrier in the alveolar region. TEM image (cross-section) showing epithelial cells (ep), endothelial cells (en), basement membrane (bm), and erythrocytes (ery) within alveolar capillaries. Courtesy of Prof. Dr. Peter Gehr, Institute of Anatomy, University of Bern, Bern, Switzerland

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The importance of controlled suction operation can be easily understood from the morphological structure.

2.2

lung structure

The lungs are composed of at least two distinct regions: the central conducting airway and the peripheral respiratory zone. More specifically, the trachea, main bronchi, and conducting bronchioles make up the conducting airways, while the respiratory bronchioles, alveolar ducts, and of course the alveoli make up the respiratory zone of the lung (Figure 2). Based on this tree-like branching pattern, the structure of the lung is often divided into distinct airway generations, starting with the trachea as generation 0 and ending with the alveoli as generation 23 (Albertine et al., 2000). As the term implies, the function of the conducting airway is primarily limited to the substantial airflow during active inspiration and expiration. The essential and specialized gas exchange functions of the lungs are performed entirely in the airways

Figure 2 Schematic diagram of the human airway structure. The human airway structure presents symmetrical branching from the trachea to the alveolar region. The upper airway is dominated by columnar epithelium composed of ciliated and goblet cells, whereas the alveolar area forms squamous stratified epithelium of type I and type II alveolar cells. Schema modified from Weibel (1963) and Forbes and Ehrhardt (2005)

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district. Whenever it comes to pulmonary drug delivery, these dimensional facts must be carefully considered for controlled and efficient drug deposition in the lung.

2.3

Aerosol separation

Within the respiratory system, inhaled material is deposited in different lung regions, mainly depending on the aerodynamic particle diameter and the overall inhalation maneuver (Fig. 3). Differences between depositional sites occur through three different main mechanisms: Brownian motion, deposition, and impact (Heyder 1981). For particles smaller than 1 mm, Brownian motion is the key mechanism for deposition primarily in the alveolar region. Particles 1-5 mm can penetrate into the tracheo-bronchial region and be deposited there, whereas particles larger than 5 mm are deposited mainly in the oropharyngeal airways by impact (Heyder et al., 1986; Oberdorster et al., 2006) . 2005). Interestingly, ultrafine particles with a size of 0.005-0.2 mm are effectively deposited deep in the lungs, whereas most particles with a size of 0.2-1.0 mm are exhaled (Dolovich et al., 2000; Heyder and Svartengren, year 2002). However, this size range is not currently used for aerosolized drugs due to the lack of appropriate formulation technologies capable of producing ultrafine drug particles, and due to the inherent limitations of the drug dose that ultrafine particles can deliver in a reasonable aerosol volume or inhalation time. . As mentioned earlier, incorrect pulmonary application inevitably results in significantly less effective therapeutic drugs. However, this is the case for most inhaled drugs

Figure 3 Regional deposition of particles in the lung. Particles are deposited in different areas of the lungs depending on particle size and overall inhalation maneuver (Denmark EPA, Report No. 12352008).

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Aim to reach the alveolar region, as the very thin epithelium (0.1-0.5 mm) and large alveolar surface (~140 m2) provide excellent conditions for drug uptake. Overall, the alveolar surface accounts for more than 95% of the total lung surface area (Weibel 1979). In this setting, a defined and controlled inhalation maneuver is critical for regional responses and the usual expected alveolar deposition of the drug. The decisive parameters for regional particle deposition are inspiratory flow rate and total volume (Bennett et al., 1999; Brand et al., 2000; Brown and Bennett, 2004). High flow rates (1,000 ml s1) were associated with an increased proportion of particle deposition in the central airways, whereas low flow rates (200 ml s1) increased the likelihood of alveolar deposition (Scheuch et al. 2007). Thus, an important paradigm for aerosol administration can be defined as "slow and steady wins the game" (Dhand 2005). To date, the effectiveness of many commercially available inhalation products is still compromised by the deposition of the majority of the drug in the airways (Hochrainer et al., 2005; Pitcairn et al., 2005). It is important to note that several more effective techniques are available and promise to improve future inhalation therapy (Scheuch and Siekmeier 2007; Scheuch and Fischer 2008). However, the increasing use of these emerging technologies is currently limited by the economic barriers faced by most new and expensive concepts. All that said, most of the technical work has been done on the pulmonary drug delivery device, and now it's time to hand the ball over to the "science team" to figure out what happens when the drug is administered in a controlled manner.

2.4

lung space

Parallel to the different morphological structures, there are also marked differences in the clearance mechanisms of the conducting airway and alveolar regions. Deposition in conductive airways typically occurs due to particles adhering to the mucus layer on the airway surface. In general, mucosal barriers exist in different parts of the human body, such as B. the lining of the gastrointestinal tract, nasal epithelium, airways, etc. In all cases, the mucosal barrier acts as a protective space, designed to protect the body from external influences, such as bacteria, viruses or fungal spores. These pathogens usually interact with the sticky mucus layer and are thus immobilized and eliminated by clearance mechanisms present at specific body sites. Airway mucus is a viscous gel composed of highly glycosylated mucin proteins called mucins (Desseyn et al 2000; Thornton et al 2008). After being trapped by the airway mucus, the particles are subsequently cleared from the site of deposition as the mucus continues to move proximally to the tip of the endotracheal tube (Figure 4). Eventually, trapped material and mucus are swallowed and undergo further biochemical processing in the gastrointestinal tract. This clearance mechanism, also known as mucociliary clearance (MC), significantly limits the practicality of many inhalational therapies: drug particles are quickly cleared after the first inhalation and deposition

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Figure 4 Top view (SEM image) of airway surface epithelium. Airway epithelium is characterized by a densely ciliated surface with mucus-producing goblet cells but no cilia

Figure 5 Top view (SEM image) of the alveolar epithelium. In the alveolar region, pathogens and inorganic particles are normally cleared by alveolar macrophages that patrol the alveolar surface (bottom left). Courtesy of Lawrence Berkeley National Laboratory

Clean (4–6 mm/min!) and thus remove from the site of therapeutic effect (ICRP 1994; Bailey et al 2007). Drug particles that have successfully deposited on the alveolar surface cannot be removed by MCs because this lung region lacks ciliated cells. Alveolar macrophages (Figure 5) are a key mechanism against inhaled pathogens, particles, etc. (Sibille and Reynolds 1990; Moeller et al. 2005). Adult macrophages derived from monocytes in the bone marrow migrate to and patrol the alveolar surface. Gaiser et al. were able to show that ultrafine TiO2 particles deposited in the lungs of rats were phagocytized by alveolar macrophages within hours (Geiser et al., 2005). Subsequently, alveolar macrophages undergo MC or active migration as they ascend into the airways

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Enter the alveolar lymphatics to further participate in the immune regulation process.

3 Pulmonary drug delivery methods 3.1

Asthma/COPD

There are a number of inhaled products on the market that are used to treat asthma and COPD. Most of these formulations contain drug particles mixed with lactose as the carrier material and are administered via a dry powder inhaler (DPI) or by aerosolizing the drug dissolved/dispersed in a propellant using a pressurized metered dose inhaler (MDI). Although these formulations produce effective pulmonary deposition, they are not designed to provide sustained or controlled release. When such systems exhibit prolonged drug efficacy, as in the case of salmeterol and formoterol, this effect is due to the pharmacological half-life of the drug rather than the delivery system or formulation design. In fact, particle technology in early conventional formulations was designed to improve aerodynamic properties and thus increase the deposition rate of aerosol particles. However, it was quickly discovered that longer drug residence times significantly improved therapeutic outcomes, leading to the development of new particle delivery systems for the treatment of asthma, COPD and other localized lung diseases. Some of these new approaches aim to find sustained-release formulations that reduce the frequency of dosing in patients and increase bioavailability in the lungs. Arya et al. (2006) coated budesonide particles with a very thin film of polylactic acid using pulsed laser ablation. Higher lung AUC values were observed for coated budesonide following intratracheal administration of coated and uncoated budesonide to neonatal rats. In addition, budesonide had reduced systemic exposure compared to uncoated budesonide. Garber et al. Beclomethasone-loaded micelles were successfully prepared using PEG-DSPE, a block copolymer of polyethylene oxide and distearoylphosphatidylethanolamine (Gaber et al., 2006). Lyophilized beclomethasone-loaded polymer micelles exhibited high encapsulation efficiency (>96%) and sustained release over 6 days in in vitro drug release studies. Another interesting approach is the incorporation of albuterol acetone into solid lipid particles (SLM) (Jaspart et al., 2007). SLM has previously been shown to exhibit physicochemical stability in rats and lack acute in vivo toxicity (Sanna et al., 2004). Subsequent in vitro release studies showed that albuterol acetone SLM exhibited slower drug release than pure albuterol acetone. Therefore, SLMs hold promise for sustained pulmonary drug delivery, reducing the dose required for effective treatment.

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In addition, liposomes have also been considered as vehicle vehicles for pulmonary delivery of anti-inflammatory drugs. Surrey et al. The distribution of 99m Tc-labeled beclomethasone dipropionate dilauroylphosphatidylcholine (DPLC) and dipalmitoylphosphatidylcholine (DPPC) liposomes in healthy volunteers was studied (Saari et al., 1999 ). They found that the clearance of DPPC liposomes was slower than that of DPLC liposomes, most likely due to the different phase transition temperatures. Interestingly, for both formulations, approximately 80% of the deposited radioactivity remained in the lungs 24 hours after inhalation. Learoyd et al. Chitosan-based particles of terbutaline sulfate were produced in which chitosan modified drug release (Learoyd et al., 2008). The use of low molecular weight (310 kDa) chitosan and its mixtures resulted in a longer duration of release of terbutaline from high molecular weight formulations.

3.2

immunosuppressants

Pulmonary drug delivery will be assessed for topical use of immunosuppressants in lung transplant patients. To date, intravenous and oral tacrolimus formulations are available but poorly tolerated by many patients. Hineswat et al. Nanostructured aggregates containing amorphous (lactose-containing) or crystalline tacrolimus nanoparticles were generated by ultrafast freezing (Sinswat et al., 2008). These aggregates could be successfully delivered by nebulization and showed high rates of drug absorption in the lungs of mice. Cyclosporin A, another commonly used immunosuppressant, is hydrophobic and thus has problematic physicochemical properties. Previous aerosol formulations were based on ethanol and propylene glycol solutions, which were inevitably highly irritating to the lungs of animals and humans (O'Riordan et al., 1995; Iacono et al., 1997). In parallel to beclomethasone, liposomal encapsulation of cyclosporine A led to an optimized drug formulation that exhibited efficient absorption rates in lung tissue and reduced drug side effects (Gilbert et al. 1997; Letsou et al. 1999) . In a recent approach, Chiou et al. Cyclosporin A powder was prepared using the Confined Liquid Impingement Jet (CLIJ) technique followed by spray drying (Chiou et al., 2008). They optimized the technique to obtain particles suitable for lung delivery of proteins. Taken together, all of these advanced formulation technologies for cyclosporin A are expected to reduce systemic plasma levels, thereby reducing deleterious toxicity to other organs such as kidneys and liver. From the discussion above, it is clear that most approaches are aimed at (1) improving bioavailability, (2) controlling the release profile, or (3) reducing the frequency of dosing of the drug. All of these trials aim to improve patient compliance and the overall effectiveness of treatment. However, a major problem limiting any therapeutic inhalation strategy remains unresolved: To date, there are no techniques to inhibit or circumvent efficient clearance mechanisms

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inside the airway. Of course, the potential for such new drugs and/or carrier technologies for pulmonary delivery is enormous. Emphasized here is that new delivery systems that prevent clearance from the lung and build a robust drug pool are urgently needed and must be taken into account in future pulmonary drug delivery.

3.3

vaccination

Lung vaccination is simple, rapid and noninvasive, making it an effective strategy in the fight against infectious diseases, especially in developing countries. Furthermore, this route of vaccination allows for mass vaccination campaigns without the need for medical personnel. A number of lung vaccines are in development for various infectious diseases such as influenza (Amorij et al., 2007; Garmise et al., 2007), measles (LiCalsi et al., 2001; de Swart et al., 2007 ; Burger et al., 2008) and diphtheria (Amidi et al., 2007) and hepatitis (Lu and Hickey, 2007). Lung vaccination can trigger local and systemic immune responses (Hobson et al. 2003). As mentioned previously, the lung mucosa and underlying epithelial layer serve as the main physical barrier in the lung. Underneath the epithelium is a collection of immune cells such as B. antigen-presenting cells (APCs) and bronchoalveolar lymphoid tissue (BALT), which are usually only present in children and the elderly, but can be induced by local infection (Tschernig and Pabst 2000). Localized activation of the immune response in the lungs has the advantage of being able to fight pathogens directly at the point of entry. To date, several formulations for intranasal administration of influenza vaccines have been tested and shown to elicit modest systemic immune responses (Read et al., 2005; Garmise et al., 2007). Smith et al. Inactivated or subunitized influenza virus vaccines were encapsulated in spray-dried microparticles containing dipalmitoylphosphatidylcholine (DPPC) and distearoylphosphatidylcholine (DSPC) and administered intratracheally to small Mice and rats (Smith et al., 2003). The formulation showed increased local bioavailability of BALT, increased antigen load of APC, IgG antibodies and local and systemic T cell responses. In another approach, an inulin-stabilized influenza vaccine powder was produced by spray freeze-drying and administered to the lungs of mice (Amorij et al., 2007). The vaccine powder formulation resulted in higher levels of IgG and IgA compared with conventional intramuscular influenza vaccine, suggesting that the modified vaccine can increase local and systemic antibody production. Like flu, measles is another airborne infectious disease. Several research groups have reported significantly higher immune responses in humans following live attenuated measles vaccine administered as a wet-mist aerosol compared to injectable vaccine formulations (Bennett et al., 2002; Dilraj et al., 2007) . However, the vaccine is very sensitive to temperature and maintenance

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Extending the cold chain from industry to "patients" or end users can significantly increase the prices of these products. To date, several research groups have attempted to develop dry powder vaccines with greater stability. Deswater et al. Rhesus monkeys were given two different powdered measles vaccines, although the vaccination was less effective than intramuscular or aerosolized vaccination (de Swart et al., 2007). Therefore, further work is needed to develop measles vaccine formulations with acceptable properties for administration via DPI, resulting in increased serum antibody levels.

3.4

Anti-infective

One-third of the world's population is infected with Mycobacterium tuberculosis, the causative agent of tuberculosis (WHO 2008). Tuberculosis treatment is a major challenge because Mycobacterium tuberculosis invades and multiplies in macrophages. So far, anti-tuberculosis drugs are administered orally, and the administration time is relatively long. Thus, side effects and high dosing frequency can lead to patient compliance issues and discontinuation of drug therapy. In this context, specific macrophage targeting can improve the overall treatment regimen by reducing systemic exposure, lowering total drug dose, and reducing treatment side effects. Furthermore, in this case, a specific targeting strategy is required that can both target the drug to infected macrophages and ensure long-term release of the drug after administration. Pandey et al. Biodegradable polylactide-co-glycolide (PLGA) nanoparticles containing three anti-tuberculosis drugs (ATDs), rifampicin, isoniazid, and pyrazinamide, were prepared and aerosolized nanoparticles were administered in infected guinea pigs (Pandey et al., 2003). The experimental results showed that the bioavailability of the three drugs was improved compared with intravenous administration. Furthermore, inhalation and ATD-loaded nanoparticles resulted in drug concentrations in the lungs exceeding therapeutic concentrations for 11 days. Sharma et al. Attempts were made to increase ATD bioavailability by using ATD to create bioadhesive wheat germ agglutinin-coated PLGA nanoparticles (Sharma et al., 2004). Wheat germ agglutinin was used because it is known to bind to alveolar epithelial cells (Brueck et al., 2001). Here, in parallel with Pandey's data, the results showed that plasma concentrations of aerosolized ATD remained in the therapeutic range for about 15 days. As mentioned earlier, liposomes are excellent for delivery to the respiratory tract. The similarity of liposomal compounds to natural surfactants prevents them from becoming irritating once deposited in the lungs. Zaru et al. Various rifampicin-loaded liposomes were designed and reported reduced toxicity to alveolar epithelial cells (A549) compared to free drug (Zaru et al. 2007). Today, "stealth liposomes", i. H. Steric-stabilized liposomes prevent rapid clearance through the reticuloendothelial system (Allen and Hansen 1991) and have been used in intravenous cancer therapy (eg Caelyx1/Doxil1). Using the "stealth" concept, Deol et al. Stealth liposomes developed for pulmonary drug delivery

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Increased affinity for mouse lung tissue by surface modification with O-stearylamylopectin. Encapsulated drugs, namely isoniazid and rifampin, were significantly less toxic to peritoneal macrophages of infected mice compared to free drugs (Deol and Khuller 1997; Deol et al. 1997). Another targeting strategy utilizes the mannose receptor expressed on alveolar macrophages by mannosylation of liposomes. Wijagkanalan et al. reported effective targeting of mannosylated liposomes to alveolar macrophages following intratracheal instillation in rats (Wijagkanalan et al., 2008), as did Chono et al. when they used ciprofloxacin-loaded mannosylated liposomes against parasitic infections in lung cells (Chono et al., 2008). Other groups have developed microspheres as delivery systems for anti-infective drugs. Takenaga et al. showed that rifampicin-loaded lipid microspheres could be delivered to alveolar macrophages in vitro and in vivo with reduced side effects in the liver (Takenaga et al., 2008). Hirota et al. The phagocytic activity of alveolar macrophages on different sizes of rifampicin-containing PLGA microspheres was investigated (Hirota et al., 2007). Interestingly, they found that 3 mm particles were the most effective for drug delivery to alveolar macrophages. Capreomycin, used to treat multidrug-resistant tuberculosis, has shown severe side effects after intravenous administration. García Contreras et al. macroporous particles of capreomycin sulfate were developed and administered to the airways of guinea pigs and reported a decrease in both inflammation and bacterial load in lung tissue (Garcia-Contreras et al., 2007). However, the new approach to delivering anti-infective drugs to the respiratory tract is not limited to tuberculosis treatment. Tobramycin is an anti-infective drug used against Pseudomonas aeruginosa, a pathogen commonly found in the lungs of CF patients. Piercer et al. Inhalable lipid-coated tobramycin particles were formulated and reported improved drug deposition due to reduced particle agglomeration (Pilcer et al., 2006). Furthermore, moxifloxacin was successfully loaded into chitosan microspheres, which were subsequently cross-linked with glutaraldehyde. In vitro testing of microspheres in a Calu-3 cell culture model showed promising and delayed absorption of moxifloxacin compared to free drug (Ventura et al., 2008).

3.5

lung gene therapy

Cystic fibrosis (CF) is caused by various mutations in the gene encoding cystic fibrosis transmembrane conductance regulator (CFTR), a chloride ion channel that regulates the flow of water and ions through epithelial cells to important. Pathophysiologically, CF is characterized by abnormal mucus production, inflammation, and chronic bacterial infection in the airways (Sueblinvong et al., 2007). The treatment of CF by gene therapy is an exciting field, as replacement of the defective CFTR gene with a functional gene transfer vector leads to normalization of mucus production and reduction of inflammation. Although CFTR gene transfer is one of the

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Other genetic diseases such as alpha-1 antitrypsin deficiency (Cruz et al., 2007) or hemophilia (Murphy and High, 2008) are prime targets for gene therapy, where gene therapy can have a positive effect on the etiological pathology. Influence. Furthermore, recent developments suggest that the treatment of various types of cancer will become one of the largest areas of gene therapy in the near future (Eager et al. 2007). Nebulization of naked plasmid DNA results in low transfection efficiency and poor DNA stability (Lentz et al., 2006). Therefore, DNA delivery to the lung mucosal surface must be delivered via a vector system that protects the DNA from enzymatic degradation, improves long-term protein expression, and increases transfection efficiency. There are two different types of DNA vectors, viral and non-viral, and both options have distinct advantages and disadvantages. Viral vectors used for DNA delivery exhibit high efficiency in gene transfer, although they have been modified to eliminate their inherent pathogenicity. However, one of the major disadvantages of DNA delivery via viral vectors compared to non-viral vectors is their immunogenic potential, which severely limits the options for multiple-dose therapy (El-Aneed 2004). Therefore, the next section focuses on non-viral vectors. Non-viral vectors allow for the administration of multiple doses but exhibit lower gene transfer efficiencies compared to viral vectors. Most carriers in use today are positively charged to allow complexation and adsorption of negatively charged DNA through electrostatic interactions. Nonviral vectors must be biocompatible, nontoxic and capable of delivering DNA across various cellular barriers to the nucleus. For these reasons, liposomes and polymeric particles are well suited as vehicles for DNA delivery, and they can be easily and relatively inexpensively manufactured. Chitosan is a very popular polymer for gene delivery and has been used by many research groups (Bivas-Benita et al. 2004; Ko¨ping-Ho¨gga˚rd et al. 2004; Howard et al. 2006; Issa et al. 2006; Li and Birchall 2006) because of its mucoadhesive properties (Lehr et al. 1992). Lee et al. developed lipid/polycation-enriching plasmid DNA-chitosan particles and was able to show that chitosan-modified powders had higher in vitro deposition than unmodified powders and increased reporter gene expression levels (Li and Birchall 2006). Likewise, Tahara and others. Cellular uptake of PLGA nanospheres with a chitosan-modified surface was improved compared to traditional carriers (Tahara et al., 2007). Another polymer used for pulmonary gene delivery is polyethyleneimine (PEI) (Kleemann et al. 2004; Chen et al. 2007; Tagalakis et al. 2008). Kleiman et al. A TAT-PEG-PEI conjugate was developed to deliver plasmid DNA and reported improved DNA protection and higher in vivo transfection efficiency compared to unmodified PEI (Kleemann et al., 2005). Cationic lipids such as Lipofectin have also been used as vehicles for gene delivery. Batley et al. Lipofectin polymer micelles (poly(p-dioxanone-lactide)-block poly(ethylene glycol)) carrying the tumor suppressor gene PTEN were administered to C57BL/6 melanoma mice. Based on the experimental results, they observed a significant improvement in PTEN gene expression in the lungs, with no evidence of cytotoxicity or acute inflammation, and a significant increase in survival time (Bhattarai et al., 2007).

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In the context of gene therapy, it must be noted that many of these particle-based targeting concepts are also used to deliver antisense DNA/RNA or siRNA to the lung. Similar to plasmid DNA, these smaller nucleotide sequences must be formulated with a delivery system that protects the antisense nucleotides and increases overall transfection efficiency.

3.6

lung cancer treatment

Anticancer drugs are often administered into the systemic circulation when high plasma levels may be required to provide adequate drug concentrations at the site of action. To date, lung cancer drugs have been administered systemically, causing severe side effects on healthy organs such as the liver, heart, and kidneys due to the drug's inherent cytotoxicity, which is also its mechanism of action. With regard to lung cancer, pulmonary drug delivery offers the opportunity to achieve more potent local action and even sustained release in the lung, while reducing unintentional systemic exposure of anticancer drugs. Various approaches have been used to target different anticancer drugs to the lung and are discussed below. Hitzman et al. Nebulized lipid-coated nanoparticles loaded with 5-fluorouracil were administered to hamsters with lung squamous cell carcinoma (Hitzman et al., 2006c). Previous in vitro studies confirmed the sustained release properties of the lipid-coated nanoparticles used (Hitzman et al., 2006a,b). In animal studies, serum 5-fluorouracil levels were much lower compared to the lung, indicating potent local exposure and sustained release properties. Injectable paclitaxel-loaded albumin nanoparticles were approved by the FDA in 2005 for the treatment of breast cancer (Gradishar et al., 2005), and although no studies have been published on their effectiveness in the lung, the potential of this technology for inhalation therapy Worth investigating. In another approach, the toxic effects of cisplatin are reduced by slow-release lipid inhalation (SLIT). SLIT-cisplatin is a lipid vesicle-encapsulated cisplatin dispersion that releases 50% of the dose immediately, while the other 50% is retained in liposomes for sustained release (Perkins et al., 2005; Wittgen et al. People, 2007). Although this phase I study showed that SLIT-cisplatin dosing is feasible and safe, the deposition efficiency (10-15%) needs to be optimized. Given the unique structure of the lung, cell-specific targeting systems have the potential to further improve the treatment of respiratory cancers. For example, lectin-functionalized liposomes specifically bind the tumor-derived cell line A549 (Abu-Dahab et al., 2001; Brueck et al., 2001), and thus can serve as an effective targeting system. Abdahab et al. The effect of nebulization on the stability of lectin-functionalized liposomes and their binding to A549 cells was successfully investigated. A more specific target may also be the transferrin receptor, which is overexpressed in many human tumor cells. Anabousi et al. Examined the uptake level and cytotoxicity of transferrin-conjugated liposomes and showed increased uptake

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As well as increasing the cytotoxicity of this specific targeting approach (Anabousi et al. 2006a). Pegylation of these liposomes increases stability and will allow promising aerosolization experiments to prove this concept in vivo (Anabousi et al. 2006b). Last but not least, telomerase is an interesting emerging target for cancer therapy since this enzyme is present in most human cancers (Hiyama et al., 1995; Shay and Wright, 2006) . Inhibition of this enzyme may represent a new approach to lung cancer therapy, except that certain telomerase inhibitors such as the antisense oligonucleotide 20-O-methyl-RNA (2-OMR) require specialized delivery systems to deliver in target cells play a biological role. Beisner et al. In various formulations containing DOTAP (N-[1-(2,3-dioleoyloxy)]-N,N,N-trimethylpropane methylsulfate), cationic lipids or a mixture of DOTAP and cholesterol The telomerase inhibitor was administered to A549 cells in a liposomal formulation (Beisner et al., 2008). Thus, these formulations enhanced transfection of A549 cells and effectively inhibited telomerase. Nafi et al. Chitosan-coated PLGA nanoparticles were recently developed, originally designed for the delivery of plasmid DNA (Ravi Kumar et al., 2004), as carriers of 2-OMR antisense oligonucleotides (Nafee et al., 2007) . The cationic surface modification of chitosan enables PLGA particles to form nanocomplexes with nucleotide-based drugs. This allows the nanocomposite structure to protect these molecules from premature degradation and facilitate their cellular uptake. According to this concept, Taetz et al. 2-OMR was delivered to A549 cells using cationic chitosan/PLGA nanoparticles, and enhanced uptake of the 2OMR nanocomplex in A549 cells, potent telomerase inhibition, and terminal Kernels are significantly shortened (Taetz et al., 2009). Clearly, this nanotechnology-based delivery system represents an interesting new platform for the safe and effective delivery of telomerase inhibitors in the context of lung cancer and potentially other cancers.

4 Future Outlook Despite the wide variety of delivery devices and medications, pulmonologists agree that the standard of care is inadequate in terms of drug duration and regional targeting within the respiratory system. In this chapter, we would like to give an overview of some upcoming and promising concepts in modern inhalation therapy. The technological foundations are in place for the production of various advanced drug delivery systems (such as nanoparticles, liposomes, and macroporous particles), and efficient pulmonary deposition—even under pathophysiological conditions—seldom limits therapeutic utility. However, inhalation therapy remains problematic and raises several unanswered questions regarding drug/particle clearance from the lung. However, more work is needed to elucidate and control what happens after drug deposition in the airways. Since the drug/particle must be able to escape the clearance mechanisms of the lung, i.e. macrophages and MCs, creating a local depot for prolonged drug action still appears to be rather difficult.

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Furthermore, the efficacy of any therapeutic intervention must be assessed in terms of its likely effect on these clearance mechanisms, which were originally designed by evolution to protect the body from invading pathogens. Lung diseases such as cystic fibrosis and cancer can affect tissue properties in several ways, for example by altering the properties of airway mucus, causing chronic inflammatory processes or creating airless areas of the lung. Modifications in morphology and physiology re-emphasize the need for multifunctional concepts capable of treating a variety of diseases. Promising progress is being made in at least one of these areas, the penetration of mucus by particulate drug carriers (Lai et al., 2007). Translating innovative drug delivery technologies into marketed medicines is a rather slow process due to the (understandably) heightened focus on safety, complex regulatory hurdles in the drug approval process, and other characteristics of the pharmaceutical market. Sometimes backtracking may occur for economic rather than scientific reasons, as in the case of recently inhaled insulin. Nonetheless, the progress made in pulmonary drug delivery in recent years has been impressive, and we believe new therapies will continue to be developed, even if the road ahead is long and bumpy.

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Needle-Free Vaccinations Mark A.F. Kendall

content 1 2 3

introduce . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 194 Targeting of skin and mucosal cells: principles of immunology. . . . . . . . . . . . . . . . . . . . . . 195 Development of physical methods for targeted therapy of skin and mucosal cells. . . . . 197 3.1 Mechanical properties of the SC barrier. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 198 3.2 Biological methods. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 198 3.3 Physical cell targeting methods. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 199 4 genes Gun particle delivery. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 201 4.1 The functional principle of the gene gun. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 201 4.2 Develop a hand-held gene gun device for clinical use. . . . . . . . . . . . . . . . . . . . . 202 4.3 Ballistic particle delivery to the skin. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 207 4.4 Clinical outcomes and commercial applications. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 212 5 Conclusions. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 215 References. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 215

Summary A major obstacle to millions of deaths from infectious diseases each year is our limited ability to deliver vaccines to optimal locations in the body. In particular, there is a lack of efficient methods for delivering vaccines to the skin and outer mucosal layers—sites of immunological, physical, and practical advantages that traditional delivery methods cannot reach. This chapter examines the challenges of physical delivery methods, most of which are needle-free. We study the structural and immunogenic properties of the skin in the context of the physical cell targeting requirements of the living epidermis and review selected current physical cell targeting technologies developed to meet these requirements: needles and syringes, diffusion patches, Liquid jet syringes and microneedle arrays/patches. MSc Kendall Australian Institute of Bioengineering and Nanotechnology (AIBN), The University of Queensland, Building 75 – Cnr of College and Cooper Road The University of Queensland Brisbane, Brisbane, QLD4072, Australia Email:[email protected]

M. Schäfer-Korting (Hrsg.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_7, # Springer-Verlag Berlin Heidelberg 2010

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We then focus on the delivery of biolistic particles: we first analyze the development of these systems to meet demanding clinical needs. We then examine the interaction of biolistic devices with the skin, focusing on the mechanistic interactions of ballistic impact and cell death, and finally we discuss current clinical results of DNA vaccines, a key application of engineered delivery devices. Key words gene gun, DNA vaccine, drug delivery, gene gun, immunotherapy, Langerhans cells, skin mechanical properties, skin microneedle vaccine

Thumbnail A APC (APCs) CST CHMP D dDCs DEM DGV GMT HIV PIV RH SC VE V vi,ver rt s

Particle Cross-Sectional Area Antigen Presenting Cell Profile Shock Tube Human Drug Panel Particle Resistivity Dermal Dendritic Cell Discrete Element Modeling Doppler Global Velocity Measurement Geometric Mean Titer Human Immunodeficiency Virus Particle Image Velocity Measurement Relative Humidity Stratum Corneum Active Epidermal Particles Velocity Activity Limit Skin Target Density Target Yield Stress

1 Introduction Vaccines are most commonly used with needles and syringes, which were first invented in 1853. Needles and syringes are valid but unpopular, and there are billions of needlestick injuries or needle reuse risks resulting in iatrogenic illness each year. Additionally, needles and syringes do not optimally deliver vaccine components to antigen-presenting cells (APCs), which themselves respond to the combination of antigen and adjuvant (innate immune stimulation) that makes the vaccine successful. Provides a safe and effective way to deliver vaccines to immune-sensitive dendritic cells in the skin (and mucous membranes) has the potential to do just that

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Improving strategies for treating serious diseases. Applying physical methods to achieve this presents unique technical challenges for physically delivering vaccines to these cells. In this chapter, the physiological, immunological, and material properties of the skin are discussed in the context of physical cell targeting requirements. Check for a viable epidermis (as an example). Selected cell-targeting technologies developed specifically to meet these requirements are briefly described. The working principles of these methods are described and their usefulness for non-invasive targeting of live epidermal cells and DNA vaccination against severe disease is discussed. We then focused on one of these needle-free approaches called biolistics, which ballistically delivers millions of biomolecule-coated microparticles to the outer layers of the skin. The architecture of these devices is presented, starting with early prototypes and then examining more advanced systems configured for clinical use. Theoretical and experimental analyzes of the ballistic particle impact process, including studies of the induction of cell death, were subsequently performed. Finally, the results of applying this technique to a major human clinical trial are presented.

2 Attacking Skin and Mucosa Cells: Immunological Principles Why are outer skin cells an important target for disease treatment? The answer comes from looking at the structure of the skin, as shown here. 1 and 2. Human skin can be divided into several layers: the outer stratum corneum (SC, 10-20 mm deep), the active epidermis (VE, 50-100 mm) and the dermis (1-2 mm). (Givens et al., 1993; Fuchs and Raghavan, 2002). The SC is an effective physical barrier to dead cells in "solid" structures (Menton and Eisen 1971; Nemes and Steinert 1999). The underlying VE is composed of cells such as immune-sensitive Langerhans cells, keratinocytes, stem cells, and melanocytes (Fuchs and Raghavan 2002). Unlike the underlying dermis, the VE is devoid of blood vessels and sensory nerve endings—important features of a painless delivery site with minimal trauma. In VE, the skin has developed a highly efficient immune function, usually with a large number of Langerhans cells (500-1,000 cells mm-2) (Berman et al 1983; Chen et al 1985; Stenn et al 1992 ). It is the first line of defense against many pathogens (Babiuk et al. 2000). In particular, Langerhans cells (shown in Figure 2) are extremely efficient APCs responsible for the uptake and processing of foreign substances to generate an effective immune response. These cells have been reported to be 1,000-fold more efficient than keratinocytes, fibroblasts, and myoblasts in eliciting various immune responses (McKinney and Streilein 1989; Banchereau and Steinman 1998; Timares et al. 1998; Chen et al. et al. 2002)).) There is less information about APCs in the underlying dermis. A new population of dermal dendritic cells (dDCs) expressing langerin (originally thought to be...).

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Figure 1 Schematic of the structure of mammalian skin (a), the epidermis of mammalian skin (b), and the corresponding bilayer approximation of the epidermis for the theoretical penetration model (c). Penetration case A indicates particle delivery to the stratum corneum (dsc), while in case B the stratum corneum is completely penetrated (tsc) and the final particle location is within the active epidermis (dve). The impact velocity is vi and the living epidermis enters with velocity vi,ve. Adapted from Kendall et al. (2004b)

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（A）

Langerhans cells ~10µm

(b) Cuticle 10-20 µm

Vibrant epidermis

（C）

(Four)

50-100 microns

Leather 1-2mm

(e) Figure 2. Schematic of a cross-section of skin with Langerhans cells. Five physical cell targeting methods are also shown. (a) A small-bore needle and a half-section of a syringe; (b) the diffusion path of the stain; (c) the penetration of the liquid jet injector; (d) the hole of the micro-syringe; (e) the distribution of microparticles after gene gun injection. Kendall (2006)

(Exclusive marker of Langerhans cells) has been reported to exist (Poulin et al., 2007) and have a distinct immune function in the skin (Nagao et al., 2009). Thus, by directly targeting specific Langerhans cell or dermal APC populations, post-vaccination immune responses can be modulated. For simplicity, this chapter focuses on delivery methods that target vaccines to the most defined population of cutaneous APCs in the epidermis (Langerhans cells). Efficient in situ (in vivo) targeting of Langerhans cells and other epidermal cells with polynucleotides or antigens will open new applications in disease control (Chen et al., 2002), including targeting important viruses/diseases such as human immunodeficiency virus (HIV) vaccination. ) and cancer.

3 Development of physical methods for skin and mucosal cell targeting Within the VE, the location of Langerhans cells—as delivery targets for immunotherapy—is precisely defined as: l l

Vertical placement of consistent suprasubsal locations (Hoath and Leahy 2002); Horizontal spatial distribution evenly distributed across the skin (Numahara et al., 2001);

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Consists of 2% of the total epidermal cell mass (human skin) (Bauer et al., 2001).

Despite their recognized potential, VEs were not considered possible cellular targets until recently with the advent of new biological and physical techniques. The challenge is to efficiently penetrate the SC and pinpoint the cells of interest.

3.1

Mechanical properties of the SC barrier

The SC is a semipermeable barrier that is difficult to breach minimally invasively to attack the underlying living epidermal cells due to its variable mechanical properties. Mechanistically, SCs are classified as bioviscoelastic solids and exhibit highly variable properties. Obvious differences include large variations in thickness and composition, depending on skin location and individual age (Hopewell 1990). However, more subtle but equally important variations in SC properties must be considered when configuring the target method. For example, the mechanical fracture stress of SCs is strongly influenced by the ambient humidity or moisture content (Wildnauer et al. 1971; Christensen et al. 1977; Rawlings et al. 1995; Dobrev 1996; Nicolopoulos et al. 1998) - relative humidity Im at Over a range of 0% to 100%, this resulted in a decrease in the fracture stress of resected human SCs from 22.5 MPa to 3.2 MPa (Kendall et al. 2004b). Likewise, an increase in ambient temperature results in an order of magnitude decrease in SC fracture stress (Papir et al., 1975). Recently, in permeation studies using small probes (2 and 5 mm in diameter) attached to NANO indenters (Kendall et al., 2007), we found that the main mechanical properties of SC and the underlying VE complexity and variation bigger. Specifically: l

Lift

Both the storage modulus and mechanical fracture stress were significantly reduced by SC (Fig. 3a, 3b); at a certain depth within SC and VE, reducing the probe size significantly increased the storage modulus (Fig. 3a).

These and other sources of variability in the mechanical properties of SCs present challenges to configure methods to disrupt SCs in a minimally invasive manner and efficiently deliver vaccines to underlying cells.

3.2

biological method

Although this chapter focuses on physical methods targeting epidermal cells, it is also important to emphasize biological methods. A robust biological approach to delivering biomolecules to epidermal (and other) cells in vivo exploits the evolved function of viruses for cell delivery. in gene transfer

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Figure 3. Mechanical properties (mean standard deviation) as a function of displacement, obtained with a microprobe inserted into a mouse ear. (a) Memory modules with 5 and 2 mm microprobes. (b) Tension using a 2 mm microprobe. Adapted from Kendall et al. (2007)

Researchers have used genetically engineered viruses for DNA vaccinations and gene therapy for serious diseases, with encouraging results. However, viral gene delivery has been hampered by safety concerns, limited DNA carrying capacity, production and packaging issues, and high cost (Lu et al. 1997; Tang et al. 1997).

3.3

physical cell targeting

Alternatively, many physical techniques are being developed. They may be able to overcome some of the limitations of biological approaches by using a needle-free approach

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Mechanisms to breach the SC barrier to facilitate direct drug and vaccine access to epidermal cells. Figure 2 schematically shows important physical targeting methods in relation to the scale of typical skin and the Langerhans cell layer of interest. Needles and syringes. The most common physical delivery methods, small needles and syringes, are shown in half section in Figure 2a. Although this method easily disrupts SCs, it is practically impossible to achieve precise localization of VEs enriched in Langerhans cells. Thus, needles and syringes are used for intradermal or intramuscular injections. This inefficient method of indirectly targeting dendritic cells with DNA resulted in a modest immune response (Mumper and Ledebur 2001). Other disadvantages of needles and syringes include the risk of needlestick injuries (WHO 1999) and needle phobia (Givens et al 1993). Diffusion/osmotic release. The least invasive method of disrupting the SC is likely to be penetration through it, driven by diffusion from a patch attached to the skin (Fig. 2b) (Glenn et al., 2003). However, the current prevailing opinion is that this mode of delivery is best for smaller biomolecules (20,000, especially 40,000), resulting in longer plasma residence times (Duncan 2003). Systemic toxicity of cancer drugs can be detected by PEGylated conjugates (Minko et al., 2007) and PEGylated adenosine deaminase (Adagen1, Enzon, Bridgewater, NJ) and L-asparaginase (Oncaspar1 , Medac, Wedel, Germany) were marketed in the early 1990s for tumor-specific targeting. In addition, PEGylated interferon alpha and G-CSF are approved drugs, and PEG-camptothecin (Prothecan) has entered phase II clinical trials (Duncan 2003).

2.3.3

Dextran

Dextran is a naturally occurring macromolecule consisting of linear units covalently linked to (1!60)-pyranoses branched at the a-(1!40) position (Figure 10). Polydextrose biopolymers are characterized by an α-(1!6) linkage to a hydroxylated cyclohexyl moiety and are commonly produced by the enzymes of certain Leuconostoc or Streptococcus strains. Dextran contains multiple hydroxyl groups for bioconjugation and is available in various molecular weight ranges. Dextran has the most compact structure of all polymers and is another ideal candidate for prodrug formation due to the following properties: water solubility, nontoxicity, and highly stable glycosidic linkages. Dextran is FDA-approved as a plasma expander; it has also been investigated as a vehicle for passive targeting of tumors and inflamed areas based on the EPR effect. Methylprednisolone prodrugs have been prepared by dextran conjugation using succinic acid as a spacer (Mehvar et al., 2000), and colon-specific delivery of dextran-conjugated glucocorticoids was studied ( McLeod et al., 1993).

2.3.4

HPMA-copolymer

N-(2-hydroxypropyl)methacrylamide (HPMA) homopolymer was developed in the 1980s as a plasma expander (Kopecek and Bazilova 1973). Hydrophilic HPMA

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J. Khandare 和 R. Haag

Figure 10 Chemical structure of dextran

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Copolymers increase the water solubility of APIs and have been shown to be biocompatible, non-immunogenic, and nontoxic (Rihova et al. 1989; David et al. 2002). The distribution of HPMA in vivo is well characterized. It has great potential as a targeted delivery vehicle, eg. B. Controlled release of antisense oligonucleotides and small molecules demonstrated (Wang et al. 1998). The HPMA-doxorubicin conjugate (Figure 11) is less toxic than the free drug and can accumulate in solid tumors. Conjugates were synthesized as follows: N-succinimidyl-3-(2pyridyldisulfanyl)propanoate was used to incorporate the Pyridyldisulfanyl group into poly(HPMA) hydrazide for subsequent conjugation to the modified antibody. Anticancer drugs are then covalently attached to the remaining hydrazide groups via acid-labile hydrazone linkages (Ulbrich et al., 2003). Finally, human immunoglobulin G was coupled to the HPMA polymer, which was modified with 2-iminosulfane by replacing the 2-pyridylsulfanyl group of the polymer with the -SH group of the antibody.

2.3.5

poly(anhydride)

Poly(anhydrides) (Table 1) were specifically designed and developed for drug delivery applications. These polymers are aromatic poly(anhydrides) based on the monomers p-(carboxyphenoxy)propane and p-(carboxyphenoxy)hexane and aliphatic poly(anhydrides) based on sebacic acid (Uhrich et al., 1999 Year). Poly(anhydrides) are prepared by melt polycondensation, starting with dicarboxylic acids, and forming prepolymers from mixed anhydrides and acetic anhydride (Uhrich et al., 1999). Poly(anhydrides) undergo hydrolytic degradation to form water-soluble degradation products that dissolve in aqueous environments. They are also susceptible to surface erosion due to the high water instability of the surface anhydride bonds

Pharmaceutical polymer CH3 C

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Figure 11 Poly(N-(2-hydroxypropyl)methacrylamide-doxorubicin antibody bioconjugate. Poly-HPMA hydrazide with N-succinimidyl 3-(2-pyridyldithio ) propionate modification to introduce pyridyldithiol groups and subsequent conjugation to the modified antibody (Ab ) (modified from Ulbrich et al., 2003)

and hydrophobicity, preventing water from entering the clumps (Langer and Chasin 1990; Lee and Chu 2008). Polymers have been widely used to incorporate small molecules and proteins such as insulin, enzymes, and growth factors due to matrix degradation and erosion (Chasin et al., 1990). A sebacic acid/p-(carboxyphenoxy)propane copolymer intracranial device improves the efficacy of nitrosourea antineoplastic agents in patients with fatal brain tumors (Brem et al., 1995). Poly(anhydrides) have satisfactory biocompatibility (Laurencin et al., 1990).

2.4

Current Developments: Dendrimers

In the early 1980s, polymer science research introduced multifunctional nanoscale dendrimers such as B. dendrimers and hyperbranched polymers. The latter are highly branched macromolecules (with 50-75% branching) with a polydispersity index (PDI) typically in the range of 1.5-2.0, but with a well-defined chemical structure. Dendrimers, named after the Greek word "dendron" for tree, can be

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Chemically designed and synthesized to have precise structural properties (Newcome et al. 1994; Tomalia and Fre'chet 2005). These multifunctional polymers are synthesized from monomeric units, gradually adding new branches until a unified tree-like structure is formed. Dendrimers are interesting drug delivery vehicles because they are nanoscale, chemically defined, and multifunctional (Tajarobi et al. 2001; Lee et al. 2005; Majoros et al. 2005). Dendrimers (Figure 12) are unique nanosystems because of their monodispersity (PDI ~1.0), nanometer size (1–10 nm), low viscosity, multifunctional end groups, high solubility, and biophase Capacitance. Various reports of dendrite binding to drugs are provided in the literature, including methotrexate, camptothecin, and paclitaxel. In addition, folate residues, antibodies, and hormones can be attached to the surface of dendrimers for potential tumor cell specificity and targeting (Stiribara et al. 2002; Haag and Kratz 2006; Zhou et al. 2006) .

Figure 12 PAMAM dendrimers (4th generation) for biological applications

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Figure 13 Dendritic polyglycerol as a novel highly biocompatible polymer scaffold for delivery and multivalency

Dendritic polyglycerols are aliphatic polyethers with multiple functional hydroxyl end groups (Figure 13). They can be prepared either as perfect dendrimers by stepwise synthesis (Haag et al. 2000) or as hyperbranched polymers at the kilogram scale (Sunder et al. 1999). Since dendritic polyglycerols are synthesized in a controlled manner to obtain specific molecular weights and narrow molecular polydispersities (~1.7) (Sunder et al. 2000), they have been evaluated for various biomedical applications, as nonpolar Delivery vehicles for APIs (Turk et al. 2007; Quadir et al. 2008), heparin analogs (Turk et al. 2004) and multivalent selectin ligands (Papp et al. 2008).

3 Specific aspects of polymers in oral drug delivery In the next section, we describe specific drug delivery systems for oral administration. For information on transdermal and pulmonary drug delivery systems, see Guy and Lehr's chapter (this volume).

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A challenging aspect in the development of CRSs is to provide zero-order extended kinetics of total drug release without time lag or burst effects (Hamdan et al., 2008). Such a release profile can optimize the frequency of dosing in drug therapy. For oral therapy, coated pellets or granules in capsules and compressed tablets are inexpensive; however, advanced matrix systems, reservoir systems, and osmotic pump systems are currently proving to be the most effective. In addition, the gastrointestinal coating form is said to protect the active ingredient from gastric pH and enzymatic degradation, and prevent nausea and vomiting caused by the active ingredient irritating the gastric mucosa. Many drugs also require protective coatings to improve patient compliance by masking unpleasant tastes and odors. However, it should be noted that due to the chemical nature of the polymer used as the outer coating, the pharmaceutical dosage form may irritate the gastrointestinal mucosa.

3.1

modified release dosage form

Modified release dosage forms determine the rate and/or location of drug release through specific formulation design and/or manufacturing methods, including extended release, delayed release and pulsed release. These can be single unit systems (such as tablets) or multiple unit systems (coated pellets or granules in capsules). The latter is preferred due to predictable gastric emptying, lack of dose-dose risk, and superior and less variable bioavailability (Roy and Shahiwala 2009). Sustained-release dosage forms allow at least a two-fold reduction in the frequency of administration compared to conventional dosage forms administered by the same route. Sustained-release formulations release the charged drug after a lag time, ie. H. After a period of "no drug release". This type can be used to deliver drugs to the colon (Chorasia and Jain 2003), for example B. in the topical treatment of ulcerative colitis and Crohn's disease. Delivery of anti-inflammatory agents and immunosuppressants to damaged areas of the gut improves efficacy and reduces systemic side effects. As one of the most important forms of targeted drug delivery, drug delivery in the colon is currently receiving increasing attention. Pulse drug delivery systems (PDDS) are dosage forms with modified drug release and continuous API release (Haus 2007; Gazzaniga et al. 2008). Pulse-release formulations can also tailor API plasma levels to clinical needs by delivering a single dose after a programmed lag period, e.g. B. Suppression of morning asthma attacks. Pulse therapy can also enhance the therapeutic effects of testosterone, glucocorticoids, and antihistamines due to circadian fluctuations (Dittigen et al., 2000). Therefore, pulsatile drug delivery should be of great benefit in certain clinical situations. Gastric retention (GR) delivery systems delay delivery of the API to the gut and are ideal for APIs that are only absorbed in a limited area of the small intestine and therefore have a very limited "absorption window", such as B. due to the saturable uptake mechanism.

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Gastrointestinal coating

The advantages of enteric coating polymers are good storage stability of drugs, pH-dependent drug release, protection of gastric juice-sensitive drugs with subsequent increase in drug efficacy, and protection of gastric mucosa from aggressive drugs. Gastrointestinal and colon targeting is possible. Table 3 summarizes the drug delivery sites for drug absorption at pH 5.5-7.0 with reference to Eudragit polymers. Prozac Weekly™ capsules (Eli Lilly, Bad Homburg, Germany) are an extended-release formulation containing enteric-coated pellets of fluoxetine hydrochloride (90 mg fluoxetine). Bioavailability following a single oral 40 mg dose reaches a maximum plasma concentration of fluoxetine of 15-55 ng·mL1 after 6-8 hours (FDA 2006a).

3.3

matrix system

As the most commonly used controlled release system, the matrix system includes tablets and granules. This system is preferred because of its availability and ease of manufacture and therefore low cost. Typically, it is formulated as a once-daily dosage form containing the active ingredient uniformly dissolved or dispersed (Boldhane 2008). Solidifying matrices that are soluble or dispersible in solvents allow controlled release. This also applies to active substances with high water solubility which have to be used in higher doses. Using traditional delivery systems, these drugs are available almost immediately, and toxic plasma levels early after ingestion quickly fall below therapeutic levels because the drug is rapidly eliminated from the organism (Figure 1). In contrast, matrix systems can maintain constant plasma levels (Mura et al., 2003). Importantly, matrix systems can also be formulated to avoid interactions between active and dietary ingredients, and advanced systems can be employed regardless of dietary pattern. Sustained release can be achieved through the use of erodible monolithic materials that deliver the active ingredient to the lower gastrointestinal tract (FDA 2006b). Controlled release is achieved through swelling and erosion of the monolith, followed by sustained release of the API throughout the gastrointestinal tract. Table 3 Dissolution of Polymer at Drug Administration Site, Polymer Grade Eudragit1, Grade and Form Duodenum pH 5.5 and above L30 (powder) D-55 (30% aqueous dispersion) Jejunum pH 6.0 Above L100 L (powder) S12.5 (12.5% organic solution) ileum pH 7.0 above S100 (powder) S12.5 (12.5% organic solution) FS30D (30% aqueous dispersion) Colonic discharge pH 7.0 above FS30D (30% aqueous dispersion)

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Multilayer matrix tablets are being developed that could allow multiple release kinetics of a single active ingredient or a combination of two (or more) active ingredients with the same or different physicochemical properties (Boldhane 2008). Burst release through transient dissolution of the monolithic layer provides the loading dose required to achieve active plasma concentrations. The second layer then slowly releases the API in a controlled manner as described above. Physicochemical barriers can also avoid incompatibility between two drugs, excipients, and drug-excipient interactions by providing physical separation. A well-known example of this interaction is the Millard reaction, which occurs when tablets are compressed. Using the commercially available multilayer matrix system Cipro XR1 (500-1,000 mg per day; Bayer Pharma AG, Leverkusen, Germany), the dissolution of ciprofloxacin from the tablet was prolonged by a barrier layer separating the core from the dissolution medium. freed. Approximately 35% of the dose is contained in the immediate-release composition, while the remaining 65% is embedded in the sustained-release matrix. In the case of Cipro XR, the maximum plasma concentration of ciprofloxacin is reached between 1 and 4 hours after administration. Bioavailability approaches that of immediate-release ciprofloxacin when the same dose is used (FDA 2004). Multilayer matrices can also be used to make repeated action products. One layer of a tablet or the outer layer of a compressed dragee provides the initial dose by disintegrating rapidly in the stomach. The inner tablet consists of ingredients that are insoluble in gastric juices but release the active ingredient in the intestine. A degradable barrier layer can further improve the release profile. A commercially available multilayer matrix system is PAXIL CR1 (GlaxoSmithKline, Munich, Germany), which contains the psychotropic drug paroxetine hydrochloride.

3.4

reservoir system

For matrix systems, it may not be possible to eliminate the adverse burst effects of all drugs, which can result in higher initial drug delivery and reduced useful lifetime of the device, thus requiring more advanced systems. In polymeric reservoir systems, there is a defined resistance to drug diffusion from the reservoir to the sink (Colombo et al., 2000), the driving force being the concentration gradient of the API molecule between the reservoir and the sink. Such systems are capable of exhibiting a linear release pattern. The reservoir tablet system consists of a semi-permeable barrier that participates in the release of the drug from the core of the tablet. Common methods include coating of tablets or multiparticulates, microencapsulation of drug granules, and compression coating of tablets. Reservoir systems have several advantages and disadvantages. Benefits include linear release of the API, independent of the solubility of the drug in the medium and the pH and ionic strength of the medium. Drawbacks include complex manufacturing process and speed

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Large quantities of drug are transported from the reservoir, the drug loading capacity is small, and eventually the active substance is released incompletely. The complexity of the design, manufacturing process, and need for specialized equipment can lead to an imbalance in the cost-benefit ratio of these systems.

3.5

Osmotic pump system

Osmotic pump systems consist of tablets sealed by a semipermeable membrane with openings. Primary osmotic pumps typically consist of a single-layer core containing the API (usually water-soluble) encased in a semipermeable membrane with one or more laser-drilled holes. This is the basic single chamber osmotic pump system. In the gastrointestinal tract, water is drawn through the membranes in a controlled manner, leading to gradual dissolution of the active ingredient. The increased pressure causes the resulting API solution to flow out through the orifice at the same rate as water flows through the membrane. Examples of products that use this system are Acutrim (phenylpropanolamine hydrochloride, Norvartis, Nuremberg, Germany), Efidac (pseudoephedrine hydrochloride, Hogil, White Plains, New York, USA) and Volmax (salbutamol sulfate, G, Darmstadt, Germany; Santus and Baker 1995). There is also a second type of osmotic pump system, the two-chamber osmotic pump system, also known as the push-pull system, which is usually used when the active ingredient has limited water solubility . In this system, a semipermeable membrane surrounds a bilayer core, one containing the drug and the second containing the water-swellable osmotic agent (Figure 14). As water flows through the rate-controlling membrane into the core, the osmotic agent expands, forcing the dissolved API out through the laser-drilled holes. Examples of products that use a push-pull system are Procardia XL (nifedipine, Pfizer, Berlin, Germany) and Minipress XL (prazosin hydrochloride, Pfizer, Berlin, Germany). The weight of the semipermeable membrane (measured by weight as increased membrane thickness) slows the rate of drug release. According to FDA regulations, it is important for applicants developing GI dosage forms to determine and carefully control the weight of the semipermeable membrane during the coating process. This consideration is important in preapproval inspections (FDA 2008a). Among various pulsating conveying systems, stand-alone pulsating systems are very popular and consist of the following subcategories: l

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Capsule-based system (Pulsincap1, Scherer International Corporation, MI, USA), osmotic system (Port1 system, Therapeutic Systems Research Laboratories, Ann Arbor, MI),

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Push Layer Drug Release Advanced Push Layer Drug Layer

Figure 14. Multi-layer skeleton push layer composed of polymer plasticizer

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Soluble or erodible membrane (chronotropic) delivery systems consisting of a core containing a drug depot coated with a hydrophilic polymer (eg, HPMC) and a reservoir system with a rupturable coating (eg, soft gelatin or HPMC) .

The length of the expandable hydrogel plug and its point of insertion into the capsule controls the lag time of the Pulsincap1 system. Upon contact with the dissolving liquid, the hydrogel (e.g. HPMC, PMMA, PVA, PEO) swells and the plug sealing the drug contents in the capsule is pushed outward, releasing the active ingredient rapidly. The length of the plug and its point of insertion into the capsule controls the lag phase (Arora et al. 2006). Pulsincap1 has been reported to be well tolerated in human volunteers (Takada 1997; Hebden et al. 1999; Ross et al. 2000). The Port1 osmotic system consists of capsules coated with a semipermeable membrane. Capsules contain an insoluble but osmotically active compound and API. The entry of water through the semi-permeable membrane increases the pressure in the capsule and pops the stopper after a delay time. The Port1 system is designed to deliver methylphenidate to children with ADHD to avoid taking a second dose during the day. Alternatively, pulsatile release (three pulses) of methylphenidate from tablet-formed dosage units is possible. The coating is based on corrodible polymers (Arora et al., 2006). Pulsys1 (MiddleBrook Pharmaceuticals, formerly Advancis Pharmaceutical Corp.) is a once-daily delivery system that delivers three regular pulses of amoxicillin for optimal bactericidal effect. Preclinical studies have demonstrated excellent efficacy (Arora et al., 2006; Roy and Shahiwala, 2009); results of comparative clinical trials with standard of care are awaited. This product is commercially available (Moxatag™ Tablets). Probably the most reliable gastric emptying formulations are granules because they distribute freely in the GI tract. The particles are coated with rupturable polymers or polymers that change their permeability. In the latter method for diltiazem delivery, the rate of delivery is controlled by the thickness of the coating. Membrane thickness also controls the rate of drug delivery from rupturable polymers (Arora et al., 2006; Roy and Shahiwala, 2009). a control and

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Extended-release formulations of verapamil hydrochloride and propranolol hydrochloride are currently marketed. The release rate is independent of pH, diet, and gastrointestinal motility (Innopran1 XL tablets, MiddleBrook Pharmaceuticals, Inc.).

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Mucoadhesive Drug Delivery Systems Juliane Hombach and Andreas Bernkop-Schnürch

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introduce . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 252 Mucoadhesion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 252 2.1 Slimes. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 252 2.2 Theories of mucoadhesion and types of attachment. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 253 2.3 Mucoadhesion test. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 254 2.4 Factors affecting mucoadhesion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 257 3 Mucoadhesive polymers and derivatives. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 258 3.1 anionic polymer. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 258 3.2 Cationic polymers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 258 3.3 Nonionic polymers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 260 3.4 Amphiphilic polymers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 260 3.5 Polymer derivatives. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 260 4 Drug delivery system. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 261 4.1 Nasal. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 261 4.2 Buccal. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 261 4.3 Vagina. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 262 4.4 Eyepiece. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 262 4.5 Oral. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 263 5 Conclusions. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 264 References. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 265

Summary: Drug absorption is often limited by short contact times and rapid leaching between the formulation and the absorbing membrane. However, by using mucoadhesive polymers, the residence time of the dosage form on the mucosa can be significantly increased. In this chapter, the components of mucus, the different mucoadhesive theories and the types of association between mucus and mucoadhesives, Mucoadhesion A. Bernkop-Schnürch (*) Institute of Pharmacy, University of Innsbruck, Innrain 52, 6020 Innsbruck, Austria Email:[email protected]

M. Schäfer-Korting (Hrsg.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_9, #Springer-Verlag Berlin Heidelberg 2010

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Tests and factors affecting mucoadhesion are described. In addition, various mucoadhesive polymers are described and various mucoadhesive delivery systems are outlined. Key words mucus mucoadhesion mucoadhesive test mucoadhesive polymer mucoadhesive delivery system

1 Introduction Drug absorption is usually limited by the residence time of the drug at the site of absorption. For example, for ocular administration, the drug solution is excreted through tears within a few minutes after administration. Therefore, mucoadhesion is an important strategy to increase the residence time of drug delivery systems in the mucosa. Mucoadhesive dosage forms can be used to optimize systemic and local delivery by maintaining close contact with the site of absorption or site of action, resulting in high local drug concentrations and high flux through the absorbing tissue. Furthermore, the mucoadhesive material itself may also be a therapeutic agent for tissue protection (eg, in the case of gastric ulcers) or for lubrication (in the eyes or vagina).

2 Mucoadhesion In pharmaceutical science, mucoadhesion is defined as a condition in which a material and mucus or mucous membranes are held together by interfacial forces for prolonged periods of time (Gu et al., 1988).

2.1

mucus

In some body cavities, monolayer (eg, stomach, intestine, bronchi) or multilayer (eg, vagina and cornea) epithelial cells are hydrated by the presence of a mucus gel layer. Simple epithelium contains goblet cells that secrete mucus to the epithelial surface, while stratified epithelium contains specialized glands such as salivary glands that secrete mucus to the epithelial surface. Mucus gels are composed of mucin glycoproteins, lipids, inorganic salts and up to 83% water. Mucin glycoproteins are the most important structural components because their cysteine-rich subdomains can form intramolecular and/or intermolecular disulfide bonds. Mucus is negatively charged due to the presence of sialic acid and sulfate. The mucus layer is constantly being released and eroded

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Enzymatic and mechanical challenges of luminal surfaces. The thickness of this mucus layer varies widely between different mucosal surfaces, ranging from less than 1 mm in the oral cavity (Sonju et al., 1974) to 450 mm in the stomach (Allen et al., 1990). Mucus has different functions, which are manifested more or less in different parts of the body. First, it provides a diffusion barrier for xenobiotics, especially drugs. It also protects the gastric mucosa from intraluminal hydrochloric acid. Finally, as with all organs, mucus keeps the mucous membranes moist.

2.2

Mucoadhesion Theory and Types of Attachment

Many attempts have been made to explain the phenomenon of mucoadhesion. Because many parameters affect mucoadhesion, as detailed below, various theories of adhesion have been proposed, but no generally accepted theory has yet been proposed. However, two basic steps are generally accepted. In step I, the contact phase, there is intimate contact between the mucoadhesive layer and the mucogel layer. In the second step, the consolidation phase, the adhesive bond is strengthened and strengthened, thus ensuring a long-lasting bond (Wu 1982).

2.2.1

chemical bond

Bonds between mucus and mucoadhesive molecules can be formed in a number of ways (listed below in order of decreasing strength): 1. Covalent bonds between thiopolymers and cysteine-rich subdomains of the mucus layer , such as disulfide bonds 2. Ionic bonds, such as the interaction between cationic polymers and sialic acid moieties of mucus. 3. Hydrogen bonding occurs in the presence of hydrophilic functional groups such as hydroxyl, carboxyl, or amino groups. 4. Van der Waals bonds between dipoles. 5. Hydrophobic bonds based on non-polar groups in aqueous solution.

2.2.2

Adhesion theory

Several theories describe the basic mechanism of adhesion (Ahuja et al. 1997 and Smart 2005). Electron theory proposes that an electric double layer forms at the interface due to electron transfer when adhered surfaces with different electronic structures come into contact.

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Adsorption theory suggests that the binder binds due to covalent and/or hydrogen bonding and van der Waals forces. Wetting theory is mainly used in liquid systems. It takes into account the ability of a liquid to spread across a mucosal surface and calculates the contact angle and energy required to separate the two phases. According to the theory of diffusion or interpenetration, mucoadhesives penetrate sufficiently deeply into mucus glycoproteins, resulting in a strong semi-permanent adhesive bond. The penetration depth of polymer chains depends on the diffusion coefficient and contact time. Chain flexibility is a key parameter that favors polymer interpenetration. Fracture theory describes the force required to separate the two surfaces involved after adhesion. Typically, however, separation does not occur at the interface, but usually at the weakest point, the cohesion of one of the links. Mechanistic theories postulate that liquid adhesives will embed in irregularities of rough surfaces. The theory of mucus dehydration proposes that dehydration of the mucus gel layer increases its cohesion and promotes the retention of the adhesion system. Mucus glycoproteins may also be transported with water into mucoadhesive polymers, resulting in interpenetration. Physical entanglement of polymer chains may also explain mucoadhesive and cohesive properties.

2.3

mucoadhesion test

Various in vitro methods, such as visual, tensile, and rheological testing, are established methods for determining adhesive performance and are described in detail below. In addition, spectroscopic methods can detect chain interpenetration by 13C nuclear magnetic resonance spectroscopy (13C NMR) (Kerr et al., 1990) or attenuated total reflection Fourier transform infrared spectroscopy (ATR-FTIR) (Jabbari et al., 1993) or hydrogen bond formation. .a new method called the BIACORE1 system is based on a chip on which the tested polymer is immobilized and mucin particles flow through while the Mucoadhesion (Takeuchi et al., 2005).

2.3.1

visual test

Rotating the cylinder (Figure 1a) appears to be an appropriate method to assess the duration of attachment to the mucosa and the cohesion of the mucoadhesive polymer. Specifically, tablets containing the test polymers are applied to freshly excised mucosa (eg, intestinal, oral or nasal mucosa of pigs, vaginal mucosa of cattle).

Mucoadhesive drug delivery system Figure 1. Mucoadhesive test setup. (a) Rotary cylinder. (b) Irrigation channels. (c) Tensile studies (MDF, maximum peel force; TWA, total work of adhesion)

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artificial liquid, rotating cylinder

Mucous membrane

b Test formulation: mucosal buffer

TWA/MDF

Mucous membrane

Mucosa) were stretched on stainless steel cylinders (Apparatus 4-Zylinder, USP). The cylinder is then placed in a USP dissolution apparatus containing an artificial liquid suitable for mucous membranes at 37 °C with a stirring speed of 125 rpm. Separation, disintegration and/or erosion times of the test tablets were determined visually. In the flush channel method (Fig. 1b), freshly excised mucosa is spread over an inclined channel with a mucus gel layer on top and placed in a channel

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37°C constant temperature room. After application of the test material to the mucosa, the residence time of the mucoadhesive polymer is determined visually or by fluorimetry with a fluorescently labeled material (Rango Rao and Buri) by flushing with a suitable artificial fluid at a constant flow rate . 1989).

2.3.2

Stretching test

The stretch study (Figure 1c) is one of the most established and widely used (eg Grabovac et al., 2005; Mortazawi and Smart, 1993) in vitro systems for testing mucoadhesion. A flat disc of the test polymer is attached to freshly excised mucosa. After the test disc has been in contact with the mucosa for a certain time, the mucosa is pulled away from the disc at a certain velocity (mm s 1). Total work of adhesion (TWA) represents the area under the force-displacement curve and determines the maximum peel force (MDF). Tensile studies can also be performed on hydrated polymers to minimize hydration-induced adhesion effects (Ch'ng et al., 1985).

2.3.3

Rheological test

Due to the chain interpenetration of mucoadhesives and mucin macromolecules, there are physical entanglements, conformational changes, and chemical interactions that lead to changes in rheological behavior. The resulting synergistic increase in viscosity can be assessed by mixing mucoadhesives with mucus and measuring viscosity by classical rotational viscometers at specific shear rates or by dynamic oscillation measurements (Hassan and Gallo 1990), providing information on the structure More information on polymer-mucin networks.

2.3.4

in vivo method

In vivo methods for assessing mucoadhesion are few. Typically, the adhesion time of the dosage form in vivo is determined visually or by g-scintigraphy. After oral administration of mucoadhesive material, the extent of remaining particles can be assessed visually after a period of time. This technique can be used to assess mucoadhesion of various tissues in animal studies, but only in the human oral cavity. In contrast, the g-scintigraphy approach appears to have no tissue limitations. The radionuclides most commonly used for imaging include 99mTc, 111mIn, 113m In, and 81mKr, and have been used to study gastric emptying in humans (Khosla and Davis 1987) and gastrointestinal (GI) transit times (Harris et al 1990) . Arbos et al. (2002) described another approach, applying fluorescently labeled nanoparticles to rats, which were sacrificed at specific times.

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The gastrointestinal tract was removed, divided into different regions, and fluorescent markers were extracted from different intestinal regions and quantified by fluorescence spectrophotometry.

2.4

Factors Affecting Mucoadhesion

Many parameters affect mucoadhesion. The nature of the polymer and the nature of the environment, including mucus physiology, affect the degree of mucus adhesion.

2.4.1

aggregation factor

The optimal molecular weight for maximal mucoadhesion depends on the type of polymer. In general, low molecular weight polymers interpenetrate more easily, while high molecular weight polymers favor entanglement. Furthermore, the higher the polymer concentration, the stronger the mucoadhesion in solid dosage forms (Duchene et al., 1988). The flexibility of polymer chains is an important feature for interpenetration and entanglement. Cross-linked polymers exhibit lower chain mobility, with reduced chain lengths to effectively penetrate mucus, thereby reducing mucus adhesion. The three-dimensional structure of the polymer is also important. Dextran with a helical conformation is able to shield the adhesive groups, so a much higher molecular weight is required for the same bond strength than linear polymers.

2.4.2

envirnmental factor

The pH affects the charge of the polymer and the charge of the mucosal surface due to the differential dissociation of the functional groups of the polymer and the amino acid backbone. The degree of hydration of a polymer depends on its chemical structure and pH, which are described in more detail below. Initially applying higher pressure for contact increases the depth of interpenetration, and longer initial contact time between the delivery system and mucus increases mucoadhesive strength. Another important characteristic is the swelling behavior, which is specific to each polymer but also depends on the concentration and presence of water. On the one hand, the swollen polymer chains unravel and promote interpenetration, but on the other hand, excessive swelling leads to a decrease in the cohesive properties of the polymer.

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2.4.3

mucus physiology

Natural mucin turnover limits the residence time of mucoadhesives on the mucus layer. Mucin turnover varies across mucosal surfaces and across individuals. Human mucus turnover time is estimated to be 12-24 hours (Forstner 1978; Allen et al 1998). Mucus viscosity and thickness should also be considered. A thicker mucus layer provides more available groups for interactions and a deeper layer for polymer chain entanglement; however, high mucus viscosity makes entanglement difficult. In addition, diseases such as gastric ulcer, bacterial or fungal infection or inflammation can change the physicochemical properties of the mucous membrane.

3 Mucoadhesive polymers and derivatives Mucoadhesive polymers can be differentiated according to their origin (e.g. natural-synthetic), the type of mucosa to which they are primarily applied (e.g. eye-buccal), their chemical structure (e.g. cellulose derivatives- polyacrylate) or its binding mechanism (e.g. covalent-non-covalent). They can also be classified into anionic, cationic, nonionic and amphiphilic polymers according to their surface charge, which is important for the adhesion mechanism. Important representatives of various polymer groups are listed in Table 1.

3.1

anionic polymer

The carbonic acid groups and to a lesser extent the sulfate and sulfonate moieties are responsible for the adhesion of the anionic polymer to the mucus gel layer. Carboxyl groups are capable of forming hydrogen bonds with hydroxyl groups of oligosaccharide side chains on mucins. However, a disadvantage of anionic mucoadhesive polymers is their incompatibility with multivalent cations such as Ca2+, Mg2+ and Fe3+. In the presence of such cations, these polymers can precipitate and/or coagulate (Valenta et al., 1998), leading to a severe loss of their adhesive properties. Furthermore, the swelling of anionic polymers is pH dependent. The higher the pH, the greater the swelling, which at higher pH leads to deterioration of the cohesive properties of the polymer, making the drug delivery system no longer mucoadhesive.

3.2

cationic polymer

Strong mucoadhesion of cationic polymers is based on ionic interactions between these polymers and anionic substructures such as sialic acid moieties

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Table 1 Chemical structures of mucoadhesive polymers

Note H

red lotus

alginate

–

red lotus

–

disaster

oxygen

Anionic

COOH COOH COOH

carbomer

Anionic, cross-linked with sucrose COOH

hyaluronic acid

CH2OH O

disaster

hydrogen nitrogen

CH3

CH2OCH2COONa

Sodium Carboxymethyl Cellulose (NaCMC)

disaster

O CH2OCH2COONa

carbon dioxide

pectin

disaster

Hydroxide

oxygen

oh oh oh

oh oh

oxygen

Anionic

Anion, 0.3-1.0 carboxymethyl groups per glucose unit

anion, R = OH or methyl H

COOR

COOH COOH COOH

polycarbophil

Anionic, cross-linked with divinyl glycol CH2OH CH2OH H

Chitosan

Europe

Hydroxide

disaster

Hydroxypropyl Cellulose O

Hydroxypropylmethylcellulose

Cationic primary amino groups can be partially acetylated

disaster

Cationic

methane or methane

Ammonia

Ammonia nitrogen

Grape

disaster

Ammonia

Polyhemolysin NH

Ammonia

oh or

disaster

Nonionic, R=H or hydroxypropyl

CH2OR

Nonionic, R=H or Methoxy or Hydroxypropyl (continued)

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260 Table 1 (continued) Polymer Poly(ethylene oxide)

Chemical structure OH

non-ionic

Poly(vinyl alcohol)OH

poly(vinylpyrrolidone)

Comment

disaster

oh oh

disaster

non-ionic

Mucus gel layer. For example, chitosan is the most important representative of this group because it also has permeability-increasing properties (Artursson et al. 1994; Lueßen et al. 1997) and is available in large quantities at reasonable prices. Their swelling at lower pH values is improved compared to anionic polymers.

3.3

nonionic polymer

In general, nonionic polymers are less viscous than anionic and cationic polymers. Their adhesion originates from the interpenetration of the polymers and the subsequent entanglement of the polymer chains. However, mucoadhesion of nonionic polymers is neither pH-dependent nor affected by electrolytes.

3.4

amphiphilic polymer

Amphiphilic polymers have both anionic and cationic substructures. Thus, mucoadhesion is caused by hydrogen bonding of carboxylic acid moieties and ionic interactions with negatively charged mucosal surfaces. However, the combination of these two properties leads to reduced mucoadhesive properties compared to simply charged polymers (Lueßen et al., 1996). On the other hand, since the cationic and anionic moieties in the polymer are stabilized by ionic interactions, the cohesion of the delivery system can be greatly enhanced.

3.5

Polymer derivatives

A new generation of mucoadhesive polymers are thiolated polymers or so-called thiomers, which are polymers with attached thiol side chains

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(Bernkop-Schnürch et al., 1999). These novel polymers are designed to form covalent disulfide bonds with cysteine-rich mucus glycoprotein subdomains based on thiol/disulfide exchange reactions and/or oxidation processes. By covalently binding to mucus, the mucoadhesive properties are greatly enhanced, while the polymer still exhibits cohesive properties.

4 drug delivery system 4.1

nasal

The nose is not only an area for topical drug delivery, but also for systemic drug delivery. The advantage of nasal administration is that the nasal surface area is large, the epithelial cells of the nasal cavity are thin, porous, and rich in blood vessels, which can ensure a high absorption rate, and can quickly transport absorbed substances directly into the bloodstream of the whole body, thereby avoiding drug metabolism in the liver. When absorbed in the olfactory region, they bypass the tight blood-brain barrier and go directly to the central nervous system. In addition, epithelial cells have lower enzymatic activity than the gastrointestinal tract, thus enabling higher bioavailability of active pharmaceutical ingredients (APIs) such as peptides and proteins. However, nasal administration also has certain limitations. Due to the large amount of interference with the normal function of the nose, only limited amounts can be administered intranasally. In addition to hydrophilic drugs, macromolecular drugs are also poorly absorbed. In addition, production of approximately 1.5–2 L of nasal mucus per day and ciliary beating at a frequency of 20 Hz results in a mucociliary clearance rate of 6 mm/min (Proctor 1977). There is a need for mucoadhesive formulations to remain in place in the nasal cavity longer and increase absorption that would not otherwise occur. This has been tested by several research groups using different polymers and active substances such as antibiotics and proteins (Ugwoke et al., 2005). Some mucoadhesive polymers, such as chitosan and polyacrylic acid, also increase the permeability of epithelial cells and exhibit enzyme inhibitory activity (Dyer et al., 2002; Lueßen et al., 1994).

4.2

North Cal

Oral administration has two distinct therapeutic goals: local treatment of the oral mucosa (eg, antifungal, antiviral, local anesthetic, or corticosteroids) or systemic treatment (eg, proteins, peptides, or oligonucleotides). The oral mucosa has many advantages over other drug delivery routes. It has a rich blood supply that flows directly into the jugular vein, thereby protecting the drug from first-pass hepatic metabolism and gastrointestinal enzymatic degradation (Park and Robinson 1985).

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Waterproof back layer API layer without mucoadhesive Mucoadhesive layer (without API)

Figure 2 Example of a patch system for oral drug delivery

An alternative to traditional dosage forms such as oral gels, liquids or lozenges are plaster systems. For a successful oral patch delivery system, a bioadhesive that retains the drug in the oral cavity and prolongs contact with the mucosa, a carrier that releases the API adequately under oral conditions, and a drug that overcomes the lack of oral mucosal strategic penetration. In addition, it is usually covered with an impermeable backing to prevent drug release and drug loss in saliva (if necessary) and to increase patient comfort. A schematic example of a patch system is shown in Figure 2.

4.3

vaginal

In addition to locally acting drugs such as antifungal, antibacterial, antiviral, and anti-inflammatory agents, estrogens, and spermicides, the vagina also provides an effective platform for systemic drug delivery due to its abundant blood supply and large surface area. Promising site (Vermani and Garg 2000). In addition to avoiding hepatic first-pass metabolism and reducing gastrointestinal and hepatic side effects, it has demonstrated good permeability to a variety of compounds, including peptides and proteins (Muranishi et al., 1993 ), making the vaginal route an alternative to the parenteral route for drugs such as bromocriptine, oxytocin, calcitonin, human growth hormone, and steroids used for replacement therapy or contraception. However, despite all these advantages, the vaginal route for systemic drug delivery is sex-specific and varies widely in epithelial thickness, vaginal fluid, and cervical mucus (volume, viscosity, pH) as they depend on the age, hormones, etc. Other limitations of currently available vaginal delivery systems are leakage and short dwell times, resulting in poor patient compliance. Robinson and Bologna (1994) overcame this problem using a polycarbophil-based mucoadhesive gel that was reported to remain in the vaginal cavity for 3–4 days and used to deliver drugs such as progesterone or nonphenol- 9 vectors.

4.4

lens

The ocular bioavailability of conventional ophthalmic formulations such as aqueous solutions and ointments typically ranges from 2-10%. mayor

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Problems with ocular drug delivery are the small penetration area, the presence of lipophilic corneal epithelium as an absorption barrier, and the short contact time. Contact time is shortened by loss of instillation solution, lacrimation, lacrimation, evaporation of tear fluid, biotransformation, and protein binding of some drugs (Lee and Robinson 1986). From 50 mL of eye drops instilled in the eye, approximately 20-30 mL is lost by spillage, with a continuous loss of 2 mL per blink (Maurice and Mishima 1984). Therefore, a small number of APIs are only available for ingestion for a few seconds. Goblet cells on the surface of the conjunctiva secrete mucin, which is distributed on the ocular surface by blinking. As drug delivery systems, mucoadhesive polymers can interact with mucus and increase the residence time of drugs, thereby increasing their bioavailability. For example, Mengi and Deshpande (1992) compared polyacrylic acid (PAA) (Carbopol 940) and poloxamer hydrogels with eye drops as delivery systems for flurbiprofen in rabbits. However, both gel formulations showed sustained effects, and the authors hypothesized that the superior results with the PAA vehicle were due to mucoadhesion. More recently, Mansour et al. (2008) developed a poloxamer-based in situ gel formulation of ciprofloxacin hydrochloride, which showed Offers controlled release, mucoadhesive properties, and improved ocular bioavailability compared to conventional commercially available eye drops.

4.5

oral

The residence time of drugs and their delivery systems in the GI tract varies widely due to nutritional status and gut motility. The intimate contact of the delivery system with the mucosa increases residence time and thus higher local drug concentrations enhance local therapy and enhance absorption. The first part of the gastrointestinal tract is the esophagus, which has no surface mucus. Therefore, adhesion occurs directly on epithelial cells. Esophageal bioadhesions are common when the dosage form is supine or ingested with little or no water. For example, in esophageal cancer, fungal infections, movement disorders, or gastroesophageal diseases, the use of bioadhesive topical drug delivery formulations may increase the contact time with epithelial cells. Alginate has a long retention time on porcine esophageal tissue, is a potential drug carrier, and provides a barrier that protects the underlying epithelium from gastric reflux (Batchelor et al., 2004). Some mucoadhesive gastric systems such as tablets, granules, pills, and granules have been shown to have longer residence times in the stomach of rats and dogs (Preda and Leucuta 2003; Hosny and Al-Meshal 1994). The small intestine is the site of absorption for many drugs. However, high locomotor activity, relatively short transit times, mucus turnover, and enzymatic degradation limit drug absorption. Mucoadhesive polymers can result in closer and longer contact with the absorbent membrane, thereby increasing absorption.

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Furosemide plasma concentration [µg/ml]

0,25

0,2

0,15

0,1

0,05

0 0

6 times [h]

Figure 3 Plasma levels of furosemide after oral administration of 10 mg of non-adherent (open squares) or adhesive (closed squares) microspheres to fasting volunteers (mean SD, n=10). Adapted from Akiyama et al. (1998)

When both adherent and nonadherent microspheres containing furosemide were administered to humans, plasma concentrations and absorption of furosemide were higher in adherent microspheres compared to nonadherent microspheres (Fig. 3) (Akiyama et al., 1998). Mucus adhesion may be more successful in the colon than in the stomach or small intestine due to a thicker mucus layer, less disruptive colonic motility, and lower mucus turnover. Varum et al. (2008) gave an overview of various studies on colonic mucosal adhesion in animals.

5 Conclusions The advantages of mucoadhesive drug delivery systems are enormous. Mucoadhesive polymers can increase residence time and enhance contact of the dosage form with the mucosa. The development of multifunctional polymers also exhibits enzyme inhibition, permeation enhancement, and controlled release properties, making them interesting for both topical and systemic administration. Although there are already many formulations based on mucoadhesive polymers on the market, the number of delivery systems taking advantage of these advantages will surely increase in the future.

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References Ahuja A, Khar RK, Ali J (1997) Mucoadhesive drug delivery systems. Drug Dev Ind Pharm 23:489-515 Akiyama Y, Nagahara N, Nara E, Kitano M, Iwasa S, Yamamoto I, Azuma J, Ogawa Y (1998) Furosemide and riboflavin-based pharmacokinetic evaluators Oromucosa-adhesive microspheres, compounds with limited site absorption from the gastrointestinal tract. J Pharm Pharmacol 50:159–166 Allen A, Cunliffe WJ, Pearson JP, Venables CW (1990) Adhesive gastric mucosal barrier and changes in gastric ulcer disease in humans. J Intern Med 228:83-90 Allen A, Hutton DA, Pearson JP, Sellers LA (1998) Mucus and mucosa. In M O'Connor (ed.) Ciba Foundation Symposium 1984, Vol. 109, Wiley, New York Arbos P, Arangoa MA, Campanero MA, Irache JM (2002) Bioadhesive properties of protein-coated PVM/MA nanoparticles Quantify. Int J Pharm 242:129-136 Artursson P, Lindmark T, Davis SS, Illum L (1994) Effect of chitosan on permeability of intestinal epithelial cell (Caco-2) monolayer. Pharm Res 11:1358–1361 Batchelor HK, Tang M, Dettmar PW, Hampson FC, Jolliffeb IG, Craig DQM (2004) Feasibility of a bioadhesive drug delivery system targeting esophageal tissue. Eur J Pharm Biopharm 57:295–298 Bernkop-Schrechch A, Schwarz V, Steininger S (1999) Polymers containing thiol groups: a new generation of mucoadhesive polymers? Pharm Res 16:876–881 Ch'ng HS, Park H, Kelly P, Robinson JR (1985) Bioadhesive polymers as a platform for oral controlled drug delivery II: Some swellable, water-insoluble bioadhesive polymers synthesis and evaluation. J Pharm Sci 74:399-405 Duchene D, Touchard F, Peppas NA (1988) Pharmaceutical and medical aspects of bioadhesive systems for drug delivery. Drug Dev Ind Pharm 14:283-318 Dyer AM, Hinchcliffe M, Watts P, Castile J, Jabbal-Gill I, Nankervis R, Smith A, Illum L (2002) Insulin administered intranasally using a novel chitosan-based formulation: Simple chitosan formulations and chitosan nanoparticles were compared in two animal models. Pharm Res 19:998-1008 Forstner JF (1978) Intestinal mucins in health and disease. Digest 17: 234–263 Grabovac V, Guggi D, Bernkop-Schnürch A (2005) Comparison of mucoadhesive properties of different polymers. Adv Drug Deliv Rev 57:1713-1723 Gu JM, Robinson JR, Leung SHS (1988) Binding of acrylic polymers to mucin/epithelial surfaces: structure-property relationships. Crit Rev The Drug Carr Syst 5:21-67 Harris D, Fell JT, Sharma HL, Taylor DC (1990) Gastrointestinal transport of potential bioadhesive formulations in humans: a scintigraphic study. J Control Release 12:45-53 Hassan EE, Gallo JM (1990) A simple rheological method for the in vitro assessment of bioadhesive binding strength of mucin polymers. Pharm Res 7:491-495 Hosny EA, Al-Meshal MA (1994) In vivo evaluation of bioadhesives containing indomethacin tablets. Drug Dev Ind Pharm 20:2715-2720 Jabbari E, Wisniewski N, Peppas NA (1993) Detection of mucoadhesion by chain interpenetration at poly(acrylic acid)/mucin interfaces using ATR-FTIR spectroscopy. J Control Release 26:99-108 Kerr LJ, Kellaway IW, Rowlands C, Parr GD (1990) The effect of poly(acrylic acid) on the rheology of glycoprotein gels. Proc Int Symp Contr Rel Bioact Mat 17:122-123 Khosla L, Davis SS (1987) Effect of polycarbophil on gastric emptying of pellets. J Pharm Pharmacol 39:47–49 Lee VHL, Robinson JR (1986) Topical ocular drug delivery: current developments and future challenges. J Ocular Pharmacol 2:67-108 Luessen HL, de Leeuw BJ, Langemeyer MW, de Boer AG, Verhoef JC, Junginger HE (1996) Mucoadhesive polymers VI in oral peptide drug delivery. Carbomer and chitosan improve intestinal absorption of the active peptide buserelin in vivo. Pharmaceutical Research 13:1668-1672

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Lueßen HL, Lehr CM, Rentel CO, Noach ABJ, de Boer AG, Verhoef JC, Junginger HE (1994) Bioadhesive polymers for oral peptide drugs. J Control Release 29:329-338 Lueßen HL, Rentel CO, Kotze' AF, Lehr CM, de Boer AG, Verhoef JC, Junginger HE (1997) Mucoadhesive polymers in oral peptide delivery IV. Polycarbophil And chitosan is an effective enhancer of peptide transport across the intestinal mucosa in vitro. J Control Release 45:15–23 Mansour M, Mansour S, Mortada ND, Abd Elhady SS (2008) In situ formation of ocular poloxamer-based ciprofloxacin hydrochloride gels. Drug Dev Ind Pharm 34:744-752 Maurice DM, Mishima S (1984) Ophthalmic pharmacokinetics. In: Sears MC (Ed.) Ophthalmic Pharmacology. Springer, Berlin, pp. 19-116 Mengi S, Deshpande SG (1992) Development and evaluation of flurbiprofen hydrogels for breakdown of the blood/water barrier. STP Pharma 2:118-124 Mortazawi SA, Smart JD (1993) In vitro assessment of mucoadhesion using tensile and shear stress. J Pharm Pharmacol 45 (suppl): 1108 Muranishi S, Yamamoto A, Okada H (1993) Rectal and vaginal absorption of peptides and proteins. Pharm Biotech 4:199-227 Park H, Robinson JR (1985) Physicochemical properties of water-insoluble polymers important for mucin/epithelial adhesion. J Control Release 2:47-57 Preda M, Leucuta SE (2003) Oxyprenol-loaded bioadhesive microspheres: fabrication and in vitro/in vivo characterization. J Microencapsul 20:777–789 Proctor DF (1977) The upper airway: I. Nasal physiology and lung defenses. Am Rev Respir Disord 115:97–129 Rango Rao KV, Buri P (1989) A novel in situ method for testing bioadhesion of polymeric and coated microparticles. Int J Pharm 52:265-270 Robinson JR, Bologna WJ (1994) Treatment of the vagina and reproductive system with bioadhesive polymers. J Control Release 28:87–94 Smart SD (2005) Principles and underlying mechanisms of mucoadhesion. Adv Drug Deliv Rev 57:1556-1568 Sonju T, Cristensen TB, Kornstad L, Rolla G (1974) Electron microscopy, carbohydrate analysis and bioactivity of proteins adsorbed to tooth surfaces in vivo over two hours. Caries Res 8:113–122 Takeuchi H, Thongborisute J, Matsui Y, Sugihara H, Yamamoto H, Kawashima Y (2005) Novel mucoadhesion tests of polymers and polymer-coated particles to develop optimal mucoadhesion drug delivery system. Adv Drug Deliv Rev 57:1583–1594 Ugwoke MI, Agu RU, Verbeke N, Kinget R (2005) Nasal mucoadhesive drug delivery: background, applications, trends and future prospects. Adv Drug Deliv Rev 57:1640-1665 Valenta C, Christen B, Bernkop-Schnuchch A (1998) Chitosan-EDTA conjugate: a new polymer for topical gels. J Pharm Pharmacol 50:445-452 Varum FJO, McConnell EL, Sousa JJS, Veiga F, Basit AW (2008) Mucoadhesion and the gastrointestinal tract. Crit Rev The Drug Carrier Syst 25:207–258 Vermani K, Garg S (2000) The scope and potential of vaginal drug delivery. Pharm Sci Technol Today 3:359-364 Wu S (1982) Adhesive bond formation. Polymer interfaces and adhesion. Marcel Dekker Inc, New York, pp. 359-447

Intrauterine Drug Administration for Contraception and Gynecological Therapy: New Approaches Dirk Wildemeersch

content 1 2

introduce . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 268 Frameless IUDs and Systems. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 270 2.1 Frameless intrauterine IUD and system development. . . . . . . . . . . . . . . . . . . . . 270 2.2 Frameless Copper Release IUD. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 272 2.3 Frameless Intrauterine Levonorgestrel System (LNG-IUS) . . . . . . . . . . . . . . . . . . . . . 277 2.4 Description. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 277 2.5 Clinical manifestations of frameless LNG-IUS. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 278 2.6 Frameless LNG-IUS in Estrogen Replacement Therapy (ERT) Acceptability and endometrial safety in women. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 279 2.7 Effect on menstrual blood loss. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 279 2.8 Contraceptive effect. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 280 2.9 Effects of frameless LNG-IUS in women with primary or secondary dysmenorrhea. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 280 Frame type levonorgestrel intrauterine release system. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 281 3.1 Development. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 281 3.2 Framework standard levonorgestrel intrauterine release system. . . . . . . . . . . . . . 282 3.3 Framework criteria Clinical presentation of LNG-IUS. . . . . . . . . . . . . . . . . . . . . . . . . . 283 3.4 Frame-type ultra-thin levonorgestrel intrauterine sustained-release system. . . . . . . . . . . . . . . . . . . 286 Conclusion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 287 4.1 Long-acting contraceptive methods should be used to prevent unwanted pregnancy. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 287 4.2 Long-term intrauterine contraception as an alternative to irreversible female sterilization. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 288 4.3 Safer contraceptive methods. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 288 4.4 Pain control with intracervical anesthesia during IUD/IUD insertion. . . . . . . . . . . . . . . . 289 4.5 Intrauterine hormonal, amenorrhea contraception for all women. . . . . . . . . . . . . . . 289 4.6 Intrauterine hormonal contraception can avoid the need for hysterectomy. . . . 290

D. Wildemeersch Contrel Research Technology Park, Ghent University, 9052 Ghent (Zwijnaarde), Belgium Email:[email protected]M. Schäfer-Korting (Hrsg.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_10, # Springer-Verlag Berlin Heidelberg 2010

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Reduces risk of heart disease, stroke, dementia and Alzheimer's after menopause. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 291 4.8 Future. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 292 References. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 293

Summary This chapter presents the development of novel intrauterine drug products aimed at providing improved methods of prevention and treatment of gynecological disorders, improvement of contraceptive methods, and increased levels of safety, user acceptance, compliance, and quality of life for women. The rimless IUD system was developed to improve the performance and acceptability of existing IUDs and potentially address the major problems encountered with traditional IUDs (eg, discharge, abnormal or excessive bleeding, and pain) . However, the performance of frameless devices is dependent on proper anchoring of the device, which requires specialized technical skills not required for deployment of traditional IUDs. Furthermore, the current research paves the way for new developments. The frameless copper and LNG eluting IUD/IUS and the framed LNG-IUS are the first of a series of innovative developments in this field. Novel compounds such as progesterone antagonists and selective progesterone receptor modulators (SPRMs) can be incorporated into polymeric drug delivery platforms for use in the uterus, cervix, vagina, or subcutaneously. Current and newer hormone-releasing intrauterine systems may also be available for bleeding-free contraception in HIV-positive (HIV+) women. It is hoped that this work will help increase the use of intrauterine contraception globally and address common health issues women face in non-surgical ways. Key words Intrauterine contraception Frameless Copper IUD Frameless LNG-IUS Frameless LNG-IUS GyneFix FibroPlant Femilis Intrauterine treatment

1 Foreword On July 29, 2005, a World Health Organization (WHO) International Agency for Research on Cancer (IACR) task force concluded that estrogen-progestogen-combined oral contraceptives (OCs) and estrogen-progestin-combined menopausal therapy are carcinogenic in in humans (Cogliano et al. 2005a, b). This widely circulated claim is based on long-term epidemiological studies showing a slightly increased risk of breast cancer in current and recent OC users. The study also showed that the risk after 10 years of stopping OC use appeared to be similar to that of people who had never used it. The risk of cervical cancer and hepatocellular carcinoma also increased with prolonged use of co-administered oral drug products. This is

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This is of major public health importance considering that 100 million women worldwide are currently using combined hormonal contraceptives—approximately 10% of all women of reproductive age. On the other hand, however, oral contraceptives have many non-contraceptive benefits, including a reduced risk of ovarian, endometrial, and colon cancer. The task force concluded that while both positive and adverse side effects of combined hormone therapy, other than cancer, have been documented, a rigorous risk-benefit analysis would help put the various effects in perspective and assess the overall public health impact. Another disadvantage of OC and intercourse-related methods is poor compliance. The effectiveness of oral contraceptives and barrier methods depends on their correct and consistent use. Inappropriate contraceptive use still results in many unintended pregnancies. It is estimated that at least 30% of pregnancies are unplanned (National Institute for Clinical Excellence (NICE) 2005). The typical use-related failure rate of the pill is 5% during the first year of use (Trussell 1998). In contrast, the effectiveness of long-acting, reversible contraceptive methods such as the intrauterine device (IUD) and intrauterine system (IUS)1 does not depend on daily consistency. In light of the above observations, the search for safe, effective, and convenient methods of contraception should not stop, and alternative methods that reduce risks for women should be actively promoted. In addition, long-acting methods should be developed to maximize contraceptive effectiveness. Several long-acting birth control methods have been shown to be safe and minimize the risk of unwanted pregnancy. These include: l l l l l

Copper IUD, progestin-only IUD, progestin-only injectable contraceptive, progestin-only subcutaneous implant, combined estrogen-progestin vaginal ring.

Intrauterine methods are a safe alternative to OC. With nearly 160 million users worldwide, the IUD is the second most commonly used contraceptive method after sterilization. In countries where the IUD is widely used, its popularity is largely due to its effectiveness and long duration of action. In addition, a recent meta-analysis found that use of a copper IUD may reduce the risk of endometrial cancer (Being et al., 2008). Publications have also suggested that copper-containing IUDs may also be associated with a lower risk of invasive cervical cancer (Lassise et al., 1991; Parazzini et al., 1992). Long-term copper IUD use is associated with reduced risk, suggesting that copper IUD use may be protective against the development of invasive cervical cancer. However, this should be interpreted with caution. Intrauterine contraception is also currently the most cost-effective reversible method of contraception (World Health Organization 2002). new generation,

The term "IUD" is used for copper-releasing intrauterine methods, while the term "intrauterine system" stands for hormone-releasing intrauterine methods.

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The miniature copper and hormone-releasing IUD described in this chapter is suitable for all women of childbearing potential, including young women who have not yet given birth and women who are pregnant at the same time, for convenient and effective contraception. Problems such as B. Menorrhagia. Contrel Research is a research institute in Ghent, Belgium, established to manage clinical research and to develop and research innovative drug delivery technologies aimed at improving quality of life through the prevention and treatment of gynecological diseases and the development of high probability contraceptive methods to improve safety levels, user adoption, and compliance. The development of the Frameless Copper IUD started in 1985, the Frameless Copper IUD was approved in the European Union in 1995 and is currently sold in Europe and China. The frameless levonorgestrel extended-release IUD, which has been in development since 1997, and the framed T-shaped levonorgestrel IUD, which has been in development since 2002, are currently in the final stages of clinical trials stage. The first LNG-IUS (Femilis1, see below) was recently approved for commercialization in Mexico. The purpose of this chapter is to provide an overview of the clinical aspects of these devices and systems to prevent fertilization and implantation. 2 It was concluded that the IUD and IUS are particularly attractive because they have the advantage of being primarily locally effective, preventing serious adverse events. After the first few months, they have less effect on the menstrual cycle. New developments in intrauterine birth control technology have made possible smaller frameless devices. They may be ideal for younger women because they are small, potent, and well tolerated. Unlike OC, they are true "install and forget". In use, they are significantly more effective than birth control pills in this age group. Furthermore, they have long-lasting and reversible effects. So the rewards are substantial. Given the current state of unwanted pregnancy, it should be used more frequently as first choice, in combination with condoms if necessary.

2 Frameless IUDs and systems 2.1

Development of frameless IUDs and systems

The size and shape of the uterine cavity vary widely, and the uterus changes in size and volume during the menstrual cycle (Hasson 1984; Kurz 1984). These changes are most noticeable during menstruation. Therefore, it is unreasonable to expect that a standard-sized IUD/IUD will fit into a uterine cavity that varies in size and volume from woman to woman and from time to time within the same woman (Figure 1). Clinical experience shows that the two are incompatible

IUDs and IUDs work primarily by preventing fertilization.

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Figure 1 Uterine cavity of different sizes and shapes. (a) Width differences; (b) Length differences; (c) Examples of functional changes and incompatibilities

The IUD/IUD and the uterine cavity can cause partial or total discharge, pain, unwanted pregnancy, and abnormal or heavy uterine bleeding leading to device removal. The Lippes ring, developed in the 1960s, had a high discontinuation rate due to its large surface area and size-related side effects. Hence, less endometrial deformation due to the use of copper as an effective antifertility agent and the T-shaped design (Zipper et al. 1971). If the width of the uterine cavity is too small, side effects and complications may occur. The cross-arm of a standard T-shaped IUD is usually too long for many uterine cavities, since the average width of most uterine cavities is usually smaller than the width of the IUD itself. If the uterine cavity is much longer than the IUD, the device will remain partially or fully in the fundus isthmus of the uterus and trigger uterine activity that facilitates expulsion and causes pain. The most important factor in reducing IUD side effects is eliminating deformation of the uterine cavity (Howard Tatum, inventor of the T-shaped IUD). Although incompatibility issues and the effect of the TCu380A IUD on menstrual bleeding are significantly reduced compared to the Lippes ring, there is still room for improvement due to the prevalence of bleeding, pain, and discharge (Xiao 1995). For these reasons, the frameless copper-releasing GyneFix1 IUD and the frameless FibroPlant1 LNG-IUS were developed.

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Instructions for the Frameless Copper-Releasing IUD

The standard rimless IUD consists of six copper sheaths or four copper sheaths (small size), each 5 mm long and 2.2 mm in diameter, threaded over a length of polypropylene suture (Figures 2 and 3). By crimping the upper and lower sleeves onto the threads, the sleeves are prevented from slipping off the material. The proximal end of the suture is knotted and inserted and secured in the fundus with a specially designed insertion tool (see insertion procedure below). The total effective copper surface (both internal and external) is 330 mm2 for the standard version and 200 mm2 for the small version. Since this implant has no plastic body, it is completely flexible. It should be noted that rimless IUDs that consist only of a copper sheath are distinguished by their effective copper surface area compared to conventional copper IUDs, which have a coil wrapped around the shaft. copper wire. The nominal and effective copper surfaces are the same only for bushings that are not attached to the plastic frame. When using copper wire or copper sleeving attached to a plastic body, the portion of the wire or sleeving that is in contact with the plastic body is not valid and should not be counted as part of the effective surface area (Kosonen 1981; Wagner 1999). The effective copper surface of the TCu200 IUD is only 120 mm2

飞哥。 2 (a) GyneFix1 330 IUP； (b) GyneFix1 200 IUP

Figure 3 Anchor tissue slices. In this case, there is no or very limited foreign body reaction

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The TCu380A IUD is 252 mm2. The large effective copper surface of the rimless IUD explains its high effectiveness (see below). The safety of the material and the response of the implant system to myometrial tissue at the polypropylene anchor site were evaluated in 14 women using a rating system developed by Sewell (Sewell et al., 1955; Coppens et al., 1989) safety). ). The time between onset and hysterectomy varied from 1 day to 4 years (Fig. 2). One third of the specimens showed no histologic reaction of the myometrium. The remaining two-thirds were mild to moderate reactions, and the diameter of the two obvious uterine inflammatory reactions did not exceed 1mm. No tissue response was observed in the other two specimens collected 4 years after insertion. In none of the cases examined, grafted endometrial tissue could be observed in the adjacent myometrium. The study confirmed the safety of the material and the safety of the implant system.

2.2.2

effectiveness

The efficacy of the standard frameless IUD has been evaluated in a large, long-term international multicentre randomized and nonrandomized comparative study involving 15,000 woman-years of experience in parous and nulliparous women. Data from these studies show that the standard IUD is very effective. Outages ranged from 0.0/100 users to 2.5/100 users (cumulative rate) over 1-9 years of use. Effectiveness was confirmed in a randomized comparative clinical study by WHO (1995). The failure rate was slightly lower than the TCu380A IUD (0.4/100-3.2/100 users) and similar to the levonorgestrel-releasing IUS (Sturridge and Guillebaud 1996). In a randomized comparative study conducted by WHO (Meirik et al., 2009). In a three-year multicenter study of approximately 400 women using a small frameless IUD, the failure rate was also low (Cao et al., 2004).

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Safety and Side Effects

bleeding. The most common reason to stop using a copper IUD is increased menstrual blood loss (MBL). The extent to which the MBL increases depends primarily on the size of the device. For larger types of non-medicated IUDs, such as the Lippes ring, the blood loss is about 70-80ml per month, which is about twice the normal menstrual blood loss. For smaller copper devices, such as the copper T series, the volume of excessive bleeding is less (50–60 mL) (Guillebaud et al. 1976).

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Clinical studies have shown that standard frameless IUDs have lower MBL than TCu380A (Andrade et al., 1987). For the frameless compact IUD, studies using visual assessment techniques using charts showed no increase in MBL after the first few months. This is attributed to the very small size of the small frameless version (Wildemeersch and Rowe 2004a). The chart method does not provide accurate flow rates in milliliters, but in practice its sensitivity and specificity are quite high, outperforming subjective MBL assessments in women. For assessing treatment effects, the visual assessment technique is convenient compared to quantitative methods because it eliminates the need for women to submit sanitary gowns to the laboratory (Janssen et al., 1995). Figure 4 shows the size differences between these different helices. The GyneFix1 200 IUD has one-third the surface area of the TCu380A and one-sixth the surface area of the Lippes Loop IUD. The reduced size of the Frameless Mini IUD minimizes the risk of menorrhagia and anemia. Women with an average MBL of 66 ml had significantly lower hemoglobin levels over 12 menstrual cycles using an IUD (Guillebaud et al., 1976). About 10% of women with an IUD may be at risk for secondary anemia, especially those who bleed more than 80ml per menstrual cycle. It has been postulated that the risk of iron deficiency increases with as little as 40 ml of blood loss (Jacobs and Butler 1965). pain. Because of the small size and flexibility of frameless IUDs, both clinical research and clinical practice experience have shown that frameless IUDs cause little pain (Wildemeersch 2003; Dou et al 2001; D'Souza et al 2003). This shows promise in primiparous women with small uterine cavities for whom standard IUDs are often poorly suited. With conventional framed IUDs, incompatibility between the device and the endometrium causes myometrium to expand. Depending on the degree of intolerance, severe cramping pain can lead to abnormal bleeding and partial or total expulsion of the IUD.

Figure 4 Three generations of IUDs (from left to right): Lippes Loop (1960), TCu380A (1980), GyneFix 200 (2000)

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Insertion, Eviction and Removal

Since frameless spirals are a new device, more familiarity with the onboarding process may be required, depending on the supplier's capabilities. GyneFix1 can be used by physicians, midwives, nurses, or other healthcare providers after they have received appropriate family planning training and have been trained in general practice of IUD placement. Insertion failures are rare if the insertion instructions are followed closely. A relevant medical history should be obtained prior to insertion to identify conditions that may influence the choice of an IUD as a method of contraception. The physical examination should include a pelvic examination and, if indicated, a "Pap smear" and appropriate testing for other forms of genital disease. Pregnancy should be ruled out before insertion. IUD providers should be aware that attention to pain relief during IUD insertion can significantly increase IUD adoption rates. If a woman is afraid, local intracervical anesthesia or local/regional anesthesia should be considered. Administration of misoprostol prior to IUD insertion can also be used to dilate the cervical canal (Saay et al. 2007). Spontaneous expulsion occurs in less than 1% of rimless IUDs over a 5-year period after correct placement. Pull-out studies evaluating the force required to retrieve an anchored IUD confirmed the reliability of the anchoring concept (Batar and Wildemeersch 2004; Wildemeersch 2004). A long-term multicentre clinical study using the current GyneFix1 insertion device demonstrated low expulsion rates (including insertion failure - see below) in both parous and nulliparous women, ranging from 0.5/100 to 3.0/100 in the first three years Lie to users. Compared to expulsion rates of 2.7/100 to 7.4/100 users with TCu380A IUDs (Cao et al., 2004; Wu et al., 2003). When applied to a rimless IUD, the term "failed insertion" has a broader meaning, including failure to implant the node into the myometrium fundus. If the node is not implanted, the device will remain in the uterine cavity but will not attach to the uterine wall as intended. This results in the rimless IUD being ejected days or weeks after attempted insertion. GyneFix1 can be removed from the uterine cavity by pulling the wire.

2.2.5

perforation

Commonly quoted conventional helical perforation rates are 1/1,000 to 3/1,000 insertions. However, the true incidence of this complication is between 0.0/1,000 and 8.7/1,000 insertions and is directly related to the skill of the person performing the insertion (Tatum and Connell 1989). One of the main causes of perforation is that the size and orientation of the uterus cannot be determined by careful examination of the pelvis. This is especially important when there is significant anteflexion, retroflexion, or lateral deviation of the uterus and the axis has not been previously straightened

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Insert using tenaculum traction. To date, perforation diagnosed at or after insertion, or translocation of a frameless IUD, has not been documented in large international multicenter clinical studies. However, a perforation rate of 1-2/1000 insertions was observed in post-marketing studies, indicating the importance of induction training. A recent study by Van Houdenhoven et al. In 2006 the estimated incidence of uterine perforation associated with the introduction of LNG-IUS (Mirena1, Bayer Schering, Berlin, Germany) was 2.6 per 1,000 introductions. Complete or partial perforation of the myometrium with a hysterometer or insertion tube increases the risk of partial or complete placement of the IUD/IUS in the abdominal cavity. Onset in breastfeeding women, even beyond 6 weeks postpartum, has been shown to be an important risk factor. A shrinking uterus (caused by long-term use of extended-release contraceptives) is also a risk factor because it causes the myometrium to thin.

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Special Uses of Frameless Copper IUDs

Emergency contraception. In 1976, the copper-containing IUD was shown to be highly effective in emergency contraception (Lippes et al., 1976). They have three main advantages over emergency hormonal oral contraceptives: (1) copper-containing IUDs are more effective, with a pregnancy rate of less than 0.1% (Trussell and Ellertson 1995), whereas emergency hormonal-only The birth control pill has a pregnancy rate of only 1% (Post-Ovulation Task Force). Fertility Regulation Methods 1998). (2) A copper IUD can be inserted at least 5 days after unprotected intercourse or up to 5 days after the earliest estimated date of ovulation (Webb 1997). In this case, the copper coil prevents implantation. With prolonged use, copper ions can trigger reactions that would normally prevent fertilization (Mishell 1998). (3) Once inserted, the IUD provides continuous contraception for 5 years or more. In a randomized study, the rimless IUD was compared with the TCu380S for emergency contraception. The results showed that while the actual insertion of the GyneFix1 Frameless IUD may have been more painful, it was less painful over the next 30 days. This significantly reduces the likelihood of requesting removal after 6 weeks due to pain with the rimless IUD. No pregnancy was reported in this study (D'Souza et al., 2003). The overall high retention rate (>80%) of all emergency IUDs at 6 weeks favors IUD insertion after unprotected intercourse, which is also supported by the superior efficacy of emergency contraception compared with oral hormonal methods One finds out. Although the standard rimless IUD was used in this study rather than the more appropriate small IUD, the latter should be preferred for emergency use because of its more acceptable bleeding profile. Contraception immediately after miscarriage. Women with Immediate IUD Insertion After Miscarriage Have Lower Pregnancy and Recurrent Pregnancy Rates

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Abortions were more common among women planning to have an IUD inserted at follow-up (Reeves et al., 2007). Therefore, the rimless IUD may provide a useful new option for the prevention of recurrent miscarriage. In limited clinical studies, insertion of an IUD immediately after pregnancy termination did not expel until 13 weeks' gestation (Bata'r et al 1998; Gbolade 1999). However, in a multicentre clinical study of 212 post-abortion women in China during the first 6 months of follow-up, four 'early' miscarriages pooled at one center were reported (unpublished data). This finding is in contrast to WHO (World Health Organization 1983) reporting rates of early pregnancy loss for framed IUDs (Lippes Loop, TCu220C, and Copper 7) ranging from 5/100 to 14/100 at 2 years. Contraception immediately after delivery. Since 1984, Belgium, Hungary, and China have tested an anchoring system insertion and fixation technique for the suspension of IUD/IUS in the uterine cavity immediately after delivery. Various anchor types were tested in pilot and multicenter tests. These studies suggest that the Immediate Postplacental Insertion and Fixation (IPPIF) technique is safe and does not increase the risk of perforation or infection. It was concluded that a frameless fixed IUD and introducer could be developed as a practical postplacental contraceptive for general use (Wildemeersch et al., 1986; Van Kets et al., 1991, 1993).

2.2.7

life

Extended IUD use is important to IUD users because it is economical and reduces some of the health risks associated with frequent removal and replacement. GyneFix1 330 IUD lifetime calculations, based on weight and surface area measurements of remote devices that have been in utero for up to 12 years, show that copper release appears to be constant for up to 12 years and is approximately 36% of the copper was released. Therefore, this IUD has a service life of at least 10 years.

2.3

Frameless Levonorgestrel Intrauterine System (LNG-IUS)

2.4

describe

The Intrauterine System (IUS) FibroPlant1-LNG is an anchored levonorgestrel (LNG) releasing device. It is a multi-component system consisting of nonabsorbable sutures with a knot at the proximal end (Fig. 5). Connected to it is a fiber delivery system measuring 3 cm long by 1.2 mm wide (FibroPlant1 14) or 3.5 cm long by 1.6 mm wide (FibroPlant1 20), releasing 14 mg or 20 mg of LNG per day. The fiber consists of an LNG ethylene vinyl acetate (EVA) core and an EVA membrane for speed control. Both systems have a service life of 5 years.

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Figure 5 (a) FibroPlant1 LNG-IUS. (b) Fiber cross-section showing drug-containing inner core and outer rate-controlling membrane

Figure 6 FibroPlant1 LNG-IUS vaginal ultrasound

The fiber is attached to the anchor suture and a stainless steel clip is placed 1 cm from the anchor knot. The anchoring node is implanted in the fundus muscle layer using the same introducer as the frameless copper-releasing IUD, securing the implant in the uterine cavity. The frameless LNG-IUS is clearly visible on ultrasound (Fig. 6); the metal clip improves the visibility of the system in X-ray images. The visibility of the IUS makes it possible to verify its correct position in the uterine cavity during insertion and subsequent examinations. Because the system is frameless, it is very flexible and can fit into any size and shape cavity, unlike framed IUSs which sometimes do not fit properly into the uterine cavity.

2.5

Clinical manifestations of frameless LNG-IUS

Two levonorgestrel-releasing IUS were evaluated in the following conditions requiring endometrial suppression:l

Hormone replacement therapy: To assess the acceptability and endometrial safety of continuous parenteral estrogen plus intrauterine levonorgestrel in postmenopausal women (3-year study).

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Menorrhagia: Effects on menstrual bleeding in women with normal menses; in women with idiopathic menorrhagia; and in women with menorrhagia associated with uterine fibroids. Contraception: Evaluated the effectiveness and acceptability of contraception in parous and nulliparous women for up to 5 years. Dysmenorrhea: The effects of a frameless LNG-IUS were studied in women with primary and secondary dysmenorrhea.

2.6

Acceptability and endometrial safety of frameless LNG-IUS in women using estrogen replacement therapy (ERT).

To measure acceptance of the FibroPlant1 LNG-IUS, women who had used oral or parenteral ERT and had worn an IUS for at least 3 years were asked if they would like to continue combination therapy and if they would accept an extension of the device. At the time of the study analysis, 150 postmenopausal women aged 33-78 were using a low-dose frameless LNG IUS. The number of women who decided to continue with the approach was 142 (94.6%). No serious adverse events such as pelvic inflammatory disease or uterine perforation were observed, and no miscarriages were recorded. Endometrial safety was assessed by endometrial biopsy and transvaginal ultrasonography (TVU). In a subset of 101 consecutive postmenopausal women who underwent endometrial biopsy three years after using the protocol, the histological findings were severe endometrial suppression, characterized by glandular atrophy and stromal decidua change. There was a good correlation between histological findings and endometrial thickness, which was thin with and without menorrhagia in all low-dose frameless LNG-IUS-LNG users (n=98) ( 90%). The only exceptions were one woman with submucosal fibroids and another woman with large polyps. After two years of use, amenorrhea occurred in up to 80% of women; ferritin levels improved significantly in all women tested.

2.8

contraceptive effect

304 insertions were performed using a frameless LNG-IUS delivering 20 mg LNG per day, 14.1% of which occurred in nulliparous women. The mean age of all women was 34.7 years (range 15-48). The total observation period was 11,299 woman-months, and the follow-up period was up to five years. Only one pregnancy occurred after the unnoticed discharge of LNG-IUS. During the first year, two expulsions and two uterine perforations occurred during insertion. These devices can be easily removed laparoscopically. The cumulative overall rate of use-related discontinuations over five years was 23.6, most of which were due to spotting/bleeding problems (n=24), pelvic pain (n=12), most unrelated to the IUD, and some was because of emotional disturbance (n ¼ 5). 16 removals were requested due to the desire to conceive (Wildemeersch and Andrade 2009, submitted).

2.9

Effects of Frameless LNG-IUS on Women with Primary or Secondary Dysmenorrhea

A non-comparative pilot study was conducted in 18 women aged 16-52 (Wildemeersch et al., 2001); 8 with primary dysmenorrhea and 10 with secondary dysmenorrhea; four trials were conducted in non-pregnant women. two insertions; twelve women complained of profuse bleeding; three women had significant fibroids (3-6 cm in diameter), and three women had suspected adenomyosis. The trial period ranges from 3 to 33 months. All women reported significantly less pain or no pain and a significant reduction in bleeding, which started in as little as one month after the frameless LNG-IUS implantation. The only exception is women with multiple fibroids. She reported significantly less bleeding, but the relief of menstrual cramps was not as pronounced as the other women in the study. All women continued to use the method.

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3 Frame type levonorgestrel intrauterine sustained release system 3.1

develop

Because the T-shaped IUD has been used for decades, healthcare providers are familiar with its insertion and fitting, so minimal training is required. Therefore, combining drug delivery technology with traditional IUD frames is an attractive option for lay providers (eg, nurses, midwives, general practitioners) and those who use IUDs infrequently. Although the insertion procedure of the T-coil is relatively simple, further simplification is needed to increase the use of intrauterine contraception. The new T-shaped LNG-IUSs Femilis1 and Femilis1 Slim have been developed with this aim in mind. With a simplified insertion process, the end of the cross-arm remains on the outside of the insertion tube, rather than folding into it like a traditional T-shaped IUD. This allows the transverse arms to deploy immediately after entering the uterine cavity. This insertion process appears to be easier than the traditional T-shaped copper IUD and Mirena1 insertion technique. It is also safer because spreading out the cross-arms does not allow the narrow protruding element to press against the fundus of the uterus, which can happen with traditional T-shaped IUDs and the Mirena1 LNG-IUS with the cross-arms folded during insertion. IUD/IUD time tube. This minimizes the risk of perforation (Fig. 7). As uterine perforation is a potentially serious adverse event, all efforts to avoid this complication are of course welcome.

Fig. 7 Simplified insertion procedure for Femilis1 LNG-IUS. Step 1: Place the loaded inserter against the cervix; Step 2: Push the loaded introducer into the uterine cavity to the fundus and insert the Femilis1 LNG-IUS; Step 3: Remove the insertion tube ( while twisting) and cut the suture

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3.2.1

D. Wildemeersch

Standard framework description for the levonorgestrel intrauterine release system

The Femilis1 (standard) consists of a fiber delivery system 3 cm long and 2.4 mm wide (Fig. 8). The inert carrier made of EVA copolymer contains 60 mg of LNG and is covered with a rate controlling membrane also made of EVA. The drug chamber releases about 20 mg of LNG per day. The polyethylene crossarm contains 22% barium sulfate, making it radiopaque. Single tail is made of 00 gauge polypropylene. The framed LNG-IUS can be used in parous and nulliparous women, whereas the Femilis1 Slim can only be used in postmenopausal women with atrophied uterus. It is worth noting that the cross-arm of the standard frame LNG-IUS (Femilis1) is 4 mm shorter than that of the standard copper T-IUD and LNG-IUS Mirena1 (Fig. 8). This smaller size optimizes compatibility with the uterine cavity of most women, especially nulliparous women (Kurz 1984; Hasson 1984). The short transverse arms also prevent contact with the vaginal wall during insertion. The Femilis1 Slim (Fig. 8) is similar to the standard Femilis1 LNG-IUS, but is 2.8 cm longer instead of 3 cm. The drug delivery compartment is only 2.0mm wide instead of 2.4mm, and contains 40mg of LNG instead of 60mg. The drug compartment, which releases about 20 mg of LNG per day, is equipped with a thin and highly flexible cross-arm attached to the upper part of the drug delivery rod. The overall length of the wishbone is 24mm, compared to 28mm on the standard version. The polyethylene crossarm contains 22% barium sulfate, making it radiopaque. Single tail is made of 00 gauge polypropylene.

Figure 8 Left to right: Comparison of ParaGard1, Mirena1, Femilis1 and Femilis1 Slim

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283

Standard frame LNG IUS clinical performance, easy and safe insertion and contraceptive performance

In the first contraceptive study using the standard framework LNG-IUS (Wildemeersch 2006), the main focus was the simplicity and safety of the new insertion technique. Published in 2006, this paper is the first report to use the LNG-IUS framework. The push-in insertion technique is considered simple and safe. Easy insertion was reported for almost all women (97.9%). 24.7% of women had no insertion pain and 67.7% had mild pain. In terms of pain, there was no statistical difference between the parous and nulliparous groups. It was concluded that framed LNGIUS is an effective contraceptive and is easy to insert. The simple and safe insertion procedure may benefit lay providers such as nurses, midwives, general practitioners and people who do not insert IUDs on a regular basis. In the second study (Wildemeersch et al., 2009), the focus was on contraceptive effectiveness. Among women with a mean age of 35.7 years (range 17-48), 280 implantations were performed, of which 60% were parous and 40% nulliparous. The study included 24 women with uterine disorders (eg, fibroids, menorrhagia). Cumulative total termination mortality was determined. The entire observation period was 8,028 female months. The LNG-IUS was easily inserted and no perforation occurred in 95.7% of cases. No pregnancies were observed and there was only one deportation (0.4/100 women after five years). The cumulative overall discontinuation rate related to use over five years was 14.7/100. Nine amputations were due to pain, six of which occurred in nulliparous women. Four women requested IUS removal due to bleeding problems. Fourteen were for "other" medical reasons, the most common being mood disorders (five cases), and 12 were for non-medical reasons. Because you want to conceive, please move fifteen. Twelve of these women conceived within a year, all without incident. The LNG-IUS with frame is well received by primiparous and multiparous women alike. Most women with heavy menstrual bleeding prior to onset, whether fibroid-related or not, report significantly less bleeding, very little bleeding, or no bleeding at all. This study demonstrates that a framed LNG-IUS releasing 20 mg of LNG per day is a highly effective and well-tolerated contraceptive in both parous and nulliparous women.

3.3.2

reserve

Results of clinical studies show that Femilis1 LNG-IUS is well preserved even in parous and nulliparous uteri, which can be explained by the strong inhibitory effect of the hormone levonorgestrel on the endometrium, decreased menstrual flow and suppressed intrauterine Peristaltic activity of the membranous uterus. Earlier

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Studies have shown that the efficacy of intrauterine levonorgestrel injections for dysmenorrhea is due to inhibition of prostaglandin synthesis (Lumsden et al., 1983). This may also explain why, in the few cases where the IUS was moved down, partial or full evictions did not occur. Downward displacement does not require replacement of the IUS unless the patient is in pain or the shaft of the IUS can be seen during speculum examination. For conventional IUDs, such as the TCu380A IUD, expulsion rates of 2.7 to 7.4/100 users or more have been observed during the first year and subsequent years (Population Reports 1995). However, partial and complete ovulation rates have been reported to be even higher in nulliparous women - as high as 17/100 users in the first year of use (Petersen et al., 1990). The downward shift in copper IUDs significantly reduced the contraceptive effectiveness of the IUDs compared to the LNG IUS (Kaivola 1990; Anteby et al 1993). In the Population Council study, the expulsion rate of the levonorgestrel-releasing IUS (11.7 per 100 users) was higher than that of the copper-releasing IUD (8.3 per 100 users) (Sivin et al., 1990) . In the European study, the eviction rates were 5.8 and 6.7 per 100 users, respectively. Another important factor that may have contributed to the low expulsion rate of the Femilis1 LNG-IUS was related to the ease of insertion and the instantaneous deployment of the transverse arms when pushing the IUS into the uterine cavity. This minimizes the risk of incorrect positioning. The very short time between folding, passage through the cervix, and subsequent deployment also contributes to the good retention of the Femilis1 LNG-IUS. For the traditional IUD and the Mirena1 LNG-IUS, the transverse arm is retracted during insertion of the insertion tube. This results in a slower deployment of the plastic beams. When the IUD/IUD is then pushed out of the tube, the cross-arm may not fully expand as it is pushed halfway into the uterine cavity, so the cross-arm may not rest properly on the fundus. If the uterine cavity is narrow and the span of the transverse arm of the IUD or IUS is significantly greater than the transverse fundal diameter, or if the insertion procedure is not strictly followed, push the IUD/IUS out of the tube instead of retracting it as suggested. This may result in poor positioning and Partial or full expulsion due to uterine contractions. The shorter transverse arms of the Femilis1 better fit the small uterine cavity and thus also facilitate proper positioning of the IUS.

3.3.3

Effects on Menstrual Blood Loss

Menstrual blood loss was assessed in 60 women, 50%, using a visual assessment technique (Boudrez et al., 2004).

4.7

Reduced risk of heart disease, stroke, dementia and Alzheimer's in postmenopausal women

After ovarian cessation, when estrogen therapy can be started, LNG-IUS provides endometrial protection as long as hormone therapy is needed as a transition to menopause (Sitruk-Ware 2007). However, since the WHI study was published, many physicians remain confused about the benefits and risks of HRT. For the reasons discussed earlier, transdermal estrogen and intrauterine progestin or progestin administration is probably the safest and most acceptable route for women in utero, resulting in high patient compliance and greatest benefit in postmenopausal women. It is considered by several experts (e.g. J. Manson, T. Mikkola, F. Naftolin, personal communication) to be the "way forward" because this therapy can be considered an almost pure Estrogen therapy circulation is low. Transdermal estrogen administration has been shown to be significantly more effective than oral estrogen administration in controlling menopausal symptoms (Pratapkumar 2006). More recently, emerging evidence suggests that the age at which HRT is initiated is important. A time window can be defined. Women in the 50-59 age group in the estrogen-only group of the WHI study were asked to participate in an additional study—

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WHI-CACS (Coronary Artery Calcification Study), which examines the extent of coronary artery calcification measured using ultrafast coronary CT. Coronary artery calcification is part of the process of atherosclerosis and is closely related to the results of coronary angiography. The results of WHI-CACS were very encouraging, as women randomized to the GHI estrogen group had significantly lower calcification scores than women in the placebo group (Manson et al., 2007). Effects of all severities were noted, estrogen users were 20-30% less likely to be classified as having a mild to moderate calcification score (below 100) and 50% less likely to be classified as having a calcification score More than 100 is late. This study confirms what has actually been known for years based on animal data and observational studies of women. Estrogens have a wide range of well-documented beneficial metabolic and vascular effects. "If treatment is initiated early in menopause, it reduces the rate of atherosclerosis accumulation and reduces the risk of coronary events" (Pines et al 2007; Vitale et al 2008; Karim 2008; International Menopause Society (2008) ); Mikola 2008). According to Henderson, there may also be a critical time window for reducing the incidence of dementia and Alzheimer's disease (Henderson 2008). In addition, early postmenopausal initiation of low-dose ERT does not appear to increase stroke risk compared with overall risk in the WHI study (Grodstein 2008).

4.8

future

New indications for LNG-IUS in combination with other hormones were found. For example, the combination of LNG-IUS with a 100 mg daily transdermal estradiol-releasing patch has been shown to be very effective in treating the physical and psychological symptoms of severe premenstrual syndrome (PMS) (Royal College of Obstetricians and Gynecologists 2007). Furthermore, in the future, the intrauterine route of administration of hormones may be used not only for the treatment of gynecological diseases, but also for their prevention (Fraser 2007). The frameless and framed IUD/IUD platforms will be used for the development of other drug delivery systems in the future. Some of these developments have already been designed and provided proof of feasibility. In the near future, intrauterine, subcutaneous and intravaginal drug delivery systems for sustained-release progesterone antagonists (PAs) and selective progesterone receptor modulators (SPRMs) are likely to be developed. These systems could treat or cure conditions such as fibroids and endometriosis/adenomyosis, as they may be more effective than current drug delivery systems, including LNG-IUS (Maruo et al., 2007). SPRM can also be used in postmenopausal women (Sitruk-Ware 2008).

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Additionally, a dual-chamber delivery system could be developed that releases PA or SPRMs as well as effective fungicides. The combination of topically administered contraceptives and microbicidal compounds is very useful for bloodless contraception and prevention of HIV transmission in HIV-positive women. This dual system may also be useful for HIV-positive women to protect them from transmission by male partners, especially in areas of high HIV prevalence. The dual-chamber IUD system can also be used in combination with other viricidal delivery systems, such as the vaginal ring. Due to high local concentrations of antiviral drugs, drug resistance as well as toxic effects, side effects and poor compliance leading to discontinuation of treatment are avoidable. In addition, intrauterine administration of hormones or antihormones can prevent menstrual blood loss, which is an important cause of viral shedding.

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National Institute for Clinical Excellence (NICE) (2005). Long-term contraception, reversible contraception? Effective and appropriate use of long-acting reversible contraceptives. www.nice.org.uk. National Institute for Health and Clinical Excellence (NICE). heavy menstruation. http://www.rcog.org.uk/news/rcog-release-variations-treating-heavy-menstrual-bleeding-found Palma S, Perez-Iglesias R, Prieto D, Pardo R, Llorca J, Delgado-Rodriguez M (2008) Iron supplementation but not folic acid supplementation reduces the risk of low birth weight in non-anemic pregnant women: a case-control study. J Epidemiol Community Health 62:120–124 Parazzini F, La Vecchia C, Negri E (1992) Intrauterine device use and risk of invasive cervical cancer. Int J Epidemiol 21:1030–1031 Petersen KR, Brooks L, Jacobson B, Skouby SO (1990) Intrauterine devices in nulliparous women. Proc 1st Congr Eur Soc Contraception, Paris, France Peterson HD (2008) Sterilization. Obstet Gynecol 111:189–203 Pines A, Sturdee DW, MacLennan AH, Schneider HPG, Burger H, Fenton A (2007) Heart of the WHI study: Time to revise hormone therapy guidelines. Climacteric 10:267-269 Population Reports (1995) Johns Hopkins University, Baltimore. Spiral - updated. Series B, No. 6 Prager S, Darney PD (2007) Intrauterine levonorgestrel systems in nulliparous women. Contraception 75:S12-S15 Pratapkumar AV (2006) Comparison of transdermal and oral HRT for control of menopausal symptoms. Int J Fertil Womens Med 51:64–69 Qi X, Zhao W, Duan Y, Li Y (2008) Intrauterine introduction of levonorgestrel in two infertile patients with complex atypical endometrial hyperplasia Successful pregnancy after sustained release system. Gynecol Obstet Invest 65:266–268 Reeves MF, Smith KJ, Creinin MD (2007) Contraceptive effectiveness of immediate versus delayed insertion of an intrauterine device after abortion: a decision analysis. Obstet Gynecol 109:1286–1294 Rietzschel E, de Buyzere M, de Backer D, Bekaert S, Segers P, Cassiman P (2007) Asklepios Investigators Abstract 3614: Contraceptive use and increased prevalence of carotid and femoral plaques : Asklepios population data. Edition 116:II_820 Rizkalla HF, Higgins M, Kelehan P, O'Herlihy C (2008) Pathologic findings associated with the Mirena intrauterine system during hysterectomy. Int J Gynecol Pathol 27:74-78 Royal College of Obstetricians and Gynecologists (2007) Management of Premenstrual Syndrome. Green Top Guideline #48 (http://www.rcog.org.uk/) Saay I, Aronsson A, Marions L, Stephansson O, Gemzell-Danielsson K (2007) Sublingual administration before IUD insertion Misoprostol for cervical priming devices in nulliparous women: a randomized controlled trial. Hum Reprod 22:2647-2652 Sewell WR, Wiland J, Craver BN (1955) A new method for comparing three sheep gut sutures with bovine gut sutures. Surg Gynecol Obstet 4:483–494 Sitruk-Ware R (2007) Intrauterine levonorgestrel system for perimenopausal and postmenopausal women. Contraception 75:S155–S160 Sitruk-Ware R (2008) Progesterone receptor modulators: clinical applications. 12th World Congress on Menopause, Madrid (Spain), Abstract p. 19 Sivin I, Shaaban M, Odlind V, Olsson SE, Diaz S, Pavez M, Alvarez F, Brache V, Diaz J (1990) Intrauterine devices Gyne T 380 and Gyne T 380 Slimline. Contraception 42:379-389 Sturridge F, Guillebaud J (1996) Risk-benefit assessment of the levonorgestrel-releasing intrauterine system. Drug Safety 15:430–440 Suhonen S, Haukkamaa M, Jakobsson T, Rauramo I (2004) Clinical performance of levonorgestrel intrauterine release system and oral contraceptives in young nulliparous women: a comparative study. Contraception 69:407–412 Postovulatory Fertility Regulation Methods Working Group (1998) Randomized controlled trial of levonorgestrel versus the Yuzpe regimen, a combined oral contraceptive for emergency contraception. The Lancet 352:428-433

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Tatum HJ, Connell EB (1989) The intrauterine device. In: Flilshie M, Guillebaud J (eds.) Contraception: Science and practice. Butterworths, London, pp. 144-171 North American Menopause Society (NAMS) (2003) The role of progestogens in hormone replacement therapy in postmenopausal women: The North American Menopause Society opinion. Menopause 10:113-132 Thiery M (1985) Pain relief during intrauterine device insertion and removal: a simplified paracervical block technique. Adv Contracept 1:167-170 Trussell J (1998) Contraceptive efficacy. In: Hatcher RA, Trussell J, Stewart F, Cates W, Stewart GK, Kowal D, Guest F (eds) Contraceptive techniques, 17th revised edition. Ardent Media, New York, NY Trussell J, Ellertson C (1995) Efficacy of emergency contraception. Birth Control Rev 4:8-11 UNDP, UNFPA and WHO Special Program for Research, Development and Research Training in Human Reproduction. World Bank: IUD Study Group (1995) TCu380A IUD and frameless "FlexiGard", preliminary three-year data from an international multicentre study. Contraception 52:77–83 Van Kets H, Parewijck W, Kleinhout J, Osler M, Zighelboim I, Tatum H (1991) Clinical experience with the Gyne-T intrauterine device in the postpartum period. Fertil Steril 4:197–205 Van Kets H, Parewijck W, Van der Pas H et al (1993) Immediate placement and fixation of the CuFix implant system in the postpartum period. Contraception 48:349–357 van Van Houdenhoven K, Kaam KJAF, van Grotheest AC, Salemans TH, Dunselman GA (2006) Uterine perforation in women using a levonorgestrel-releasing intrauterine system. Contraception 73:257–260 Varma R, Soneja H, Bhatia K, Ganesan R, Rollason T, Clark TJ, Gupta JK (2008) Intrauterine levonorgestrel-releasing (LNG-IUS) system for the treatment of endometrial hyperplasia. effectiveness. A long-term follow-up study. Eur J Obstet Gynecol Reprod Biol 139:169–175 Vitale C, Mercuro G, Cerquetani E, Marazzi G, Patrizi R, Pelliccia F, Volterrani M, Fini M, Collins P, Rosano GM (2008) Postmenopausal effects acute and estrogenic Chronic effects on endothelial function. Arterioscler thromb Vasc Biol 28:348-352 Wagner H (1999) Intrauterine contraception: past, present and future. In: Rabe T, Runnebaum B (eds.) Birth control: updates and trends. Springer, Berlin, pp. 151–171 Webb AMC (1997) Intrauterine contraceptives and antiprogestins as emergency contraception. Eur J Contracept Reprod Health Care 2:243–246 Westhoff CL, Heartwell S, Edwards S, Zieman M, Stuart G, Cwiak C, Davis A, Robilotto T, Cushman L, Kalmuss D (2007) Oral contraceptive withdrawal: causes Side effects? Am J Obstet Gynecol 196(412):e1–e6 Discussion 412.e6-7 Wildemeersch D (2004) The force required to remove a frameless 0 suture anchor system: a comparison of premenopausal and postmenopausal women. Contraception 69:513–515 Wildemeersch D (2006) Brief safety presentation, contraceptive performance and effects on menstrual blood loss of the new LNG-IUS Femilis for use in pares and nulliparous women. 9th European Congress on Social Contraception. Wildemeersch D, Bata´r I, Affandi B, Andrade ATL, Wu S, Hu J, Cao X (2003) "Frameless" intrauterine systems for long-term, reversible contraception: a review of 15 years of clinical experience, Istanbul, Turkey. J Obstet Gynaecol Res 29:160–169 Wildemeersch D, Rowe PJ (2004a) Menstruation in Belgian women using a frameless copper-releasing IUD with a copper surface of 200 mm and a copper-levonorgestrel-releasing intrauterine system Assessment of blood loss. Contraception 70:169–172 Wildemeersch D, Rowe PJ (2004b) Evaluation of menstrual blood loss in women with idiopathic menorrhagia using a frameless levonorgestrel-releasing intrauterine system. Contraception 70:165–168 Wildemeersch D, Janssens D, Andrade A (2009) The Femilis1 LNG-IUS: contraceptive performance - an interim analysis. Eur J Contracept Reprod Health Care 14:1-8

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Wildemeersch D, Rowe PJ (2005) Evaluation of menstrual blood loss in Belgian women using the novel T-type levonorgestrel intrauterine system. Contraception 71:470–473 Wildemeersch D, Schacht E (2001) Treatment of menorrhagia with a novel "frameless" intrauterine levonorgestrel-releasing drug delivery system, a pilot study. Eur J Contracept Reprod Health Care 6:93-101 Wildemeersch D, Schacht E (2002) Effect of a novel "frameless" intrauterine levonorgestrel-releasing drug delivery system on menstrual blood loss in women with uterine fibroids, a pilot study. Eur J Obstet Gynecol Reprod Biol 102:74-79 Wildemeersch D, Defoort P, Thiery M, Tatum H (1986) A new insertion fixation instrument and technique for the immediate retroplacental introduction of an intrauterine device. Ann Meet Am Fertility Soc, Toronto Wildemeersch D, Schacht E, Wildemeersch P (2001) Treatment of primary and secondary dysmenorrhea with a novel "frameless" intrauterine levonorgestrel-releasing drug delivery system: a pilot study. Eur J Contracept Reprod Health Care 6:192–198 Wildemeersch D, Schacht E, Wildemeersch P (2003) Performance and acceptability of intrauterine-released levonorgestrel for hormone replacement therapy, contraception, and postmenopausal women. Micro-delivery system. Maturitas 44:237–245 Wildemeersch D, Janssens D, Schacht E, Pylyser K, De Wever N (2005a) Intrauterine levonorgestrel administration via a frameless system combined with systemic estrogen: in perimenopausal and Acceptability and endometrial safety after three years of use in postmenopausal women. Gynecol Endocrinol 20:336–342 Wildemeersch D, Janssens D, Vrijens M, Weyers S (2005b) Ease of introduction, contraceptive efficacy and safety of a novel T-type levonorgestrel intrauterine release system. Contraception 71:465–469 Wildemeersch D, Janssens D, Weyers S (2005c) Continuous combined parenteral estrogen replacement and intrauterine progestin administration: ideal HST combination? Maturitas 51:207–214 Wildemeersch D, Pylyser K, De Wever N, Dhont M (2007a) Treatment of atypical and atypical endometrial hyperplasia with the levonorgestrel intrauterine system: long-term follow-up. Maturitas 57:210–213 Wildemeersch D, Pylyser K, De Wever N, Pauwels P (2007b) Combination of transdermal estrogen and intrauterine levonorgestrel for 5 years in postmenopausal women after hormone replacement Membrane security. Maturitas 57:205–209 World Health Organization (1983) Intrauterine devices: their role in family planning. Geneva, (WHO Offset Publication No. 75) p. 53 World Health Organization (2002) Prog Reproduct Health Res http://www.who.int/reproductive Health/en/ Wu S, Hu J, Wildemeersch D (2003) Performance Frameless GyneFix and TCu380A intrauterine devices in a three-year, multicenter, randomized comparative study of parous women. contraception. 61:91–89 Xiao S (1997) The role of scientific research and scientific research institutes in China's family planning programs. In: Proc Beijing Intern Sympos Fertility Regulation, 1995. Nat Inst Child Health and Human Develop Publication No. 97-4118 Zipper JA, Tatum HJ, Medel M, Pastene L, Rivera M (1971) Contraception through the use of intrauterine metals. I. Copper as a complement to the "T" fixture. Copper "T" in utero. Am J Obstet Gynecol 109:771-774

Improving Word of Mouth Franz Gabor, Christian Fillafer, Lukas Neutsch, Gerda Ratzinger and Michael Wirth

content 1

introduce . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1.1 Physicochemical considerations. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1.2 Physiological conditions of the gastrointestinal tract. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Improving formulation parameters for oral administration . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Dimensions. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 pH-dependent drug delivery. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3 Swelling. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4 Osmotic pressure. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.5 Density. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.6 Enzyme-mediated release in the colon. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.7 Biological detection. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.8 Absorption enhancers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Future Outlook. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .refer to. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

346 348 349 357 357 365 369 372 374 377 380 384 389 389

Summary An estimated 90% of all pharmaceuticals are oral formulations and their market share is still increasing due to significant benefits to patients, the pharmaceutical industry and healthcare systems. However, oral administration is one of the most difficult routes considering the physicochemical requirements of drugs and biopharmaceutical aspects such as physiological conditions. By considering solubility, permeability and residence time in the gastrointestinal environment as key parameters, various properties of the drug and its delivery system can be tuned, such as size, pH, density, diffusion, swelling, adhesion, degradation and penetration To adapt to the improvement of oral administration. Future developments will focus on further improving patient compliance and the feasibility of oral biotech drugs. F. Gabor (*) Institute of Pharmaceutical Technology and Biopharmaceuticals, University of Vienna, Althanstraße 14, A-1090, Vienna, Austria

M. Schäfer-Korting (Hrsg.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_12, # Springer-Verlag Berlin Heidelberg 2010

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Key words G/I physiology absorption enhancer pH dependent release enzyme-mediated release floating delivery system hydrogel microparticles nanoparticles biorecognition targeting osmotic pump gastric retention delayed release

Abkürzungen AAL API BCS EGF EOP EPAS FAE FDA FDDS GALT GI GRAS HBS HEC HPMC IBD LEA MC OPV PEG PVP SFL SOTS UEA WGA

Aleuria aurantia Lectin Active Pharmaceutical Ingredients Biological Classification System Epidermal Growth Factor Element Osmotic Pump Aqueous Evaporation Precipitation Hair Follicle-Associated Epithelium Food and Drug Delivery Floating Drug Delivery System Gut-Associated Lymphoid Tissue Gastrointestinal Recognized Safe Hydrodynamic Balance System Hydroxyethylcellulose Hydroxypropyl Inflammatory Bowel Disease Tomato esculentum-lectin Methylcellulose Oral Poliovirus Vaccine Polyethylene Glycol Polyvinylpyrrolidone Spray, Frozen Liquid Sandwich Osmotic Tablet System Ulex europaeus Lectin Wheat Germ Agglutinin

1 Introduction Despite tremendous innovations in the field of drug delivery and gaining detailed knowledge of promising alternative routes of administration, an estimated 90% of all drugs are in oral form and oral drug delivery systems account for more than half of the drug delivery market . The oral drug delivery market was a $35 billion industry in 2008 and is expected to grow by $20 billion

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10% annual increase until at least 2012. Among liquid and semisolid formulations, tablets remain the drug of choice for oral administration with distinct advantages. Oral administration is the most convenient form of treatment for the patient. Painless compared to injections, convenient compared to enemas, self-administered without training, and easy to use, at least at first glance. In practice, however, some basic principles must be followed during use to ensure adequate oral intake even with skilled drug delivery systems: first, the head should be raised to avoid narrowing of the pharynx, thereby reducing patency. Second, oral dosage forms should be swallowed with at least 100 mL of water in the bottle. Otherwise, the preparation may get stuck in the esophagus, mainly in the upper third. When 12.5 mm tablets were swallowed without water in the supine position, only two of the 20 volunteers observed the orogastric passage. Increasing the water volume to 100 mL resulted in a 92% passaging success rate (Gallo et al., 1996). In the case of esophageal adhesions, repeated drinking of even larger volumes of water does not help to separate the preparation. This is especially important when taking viscous medications such as bisphosphonates, tetracyclines, penicillins, NSAIDs, and theophylline. Third, even in bedridden patients, an upright position or at least a 45-degree tilt is beneficial to ensure a rapid transition to the stomach. In addition, some tablets may not be divided into portions for successful absorption (van Santen et al 2002; Hussar 2000). In the case of enteric-coated and dry-coated tablets, the protective layer is destroyed and osmotically compressed tablets lose their release properties. As a result, tablets containing microgram quantities of pharmaceutically active ingredients split in half lack dose uniformity, and tablet fragments initially weighing less than 50 mg are difficult to handle. In addition to technical issues, toxicological concerns for nursing staff must also be taken into account when crushing tablets, especially those containing highly potent substances, such as hormones and their antagonists, cytostatics, antivirals, and immunosuppressants. In addition, extended-release oral dosage forms improve patient comfort and compliance by reducing dosing frequency. In addition to being a convenience for patients, oral medicines can also be profitable for the pharmaceutical industry. Today, oral formulations are produced in high volumes and with short production times through skilled automation of the manufacturing process. High-quality orally administered products are reliably and reproducibly manufactured under Good Manufacturing Practice (GMP) regulations. After development costs are paid for, the ease of preparation results in a mass-produced product that is cheap and profitable. Finally, healthcare systems could easily benefit from the ease of use of oral drug delivery systems. Especially given the growing geriatric population and the resulting rise in healthcare costs, orally administered products enable cost-effective treatment without the need for qualified medical intervention. Therefore, an oral dosage form appears to be an ideal formulation for treatment. However, from a biopharmaceutical point of view, oral drug delivery is one of the most challenging drug delivery routes. as the main task

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The gastrointestinal (GI) tract is responsible for the absorption of nutrients and has a large absorptive surface area and adequate blood supply to aid in the absorption of drugs. At the same time, however, there are highly effective life-saving systems and methods of breaking down and inactivating compounds that are harmful to the body. Therefore, improving oral absorption by adjusting formulation factors is a balancing act that considers GI physiology as well as the physicochemical properties of active pharmaceutical ingredients.

1.1

physical chemistry considerations

Although reports of oral uptake of nanoparticles disprove the paradigm that only solutes can be absorbed, for most formulations it is generally accepted that solute molecules of active pharmaceutical ingredients (APIs) are transported across membranes, with passive diffusion largely superior to active diffusion in terms of drug molecules (Florence 1997; Kararli 1989). An ideal absorption process requires maximum water solubility for rapid dissolution in body fluids and diffusion according to Fick's law. At the same time, minimal lipophilicity is a prerequisite for penetration through the phospholipid bilayer of the cell membrane. To estimate the lipid solubility of an API, the equilibrium distribution of the API between the hydrophilic and lipophilic phases is typically used, followed by calculation of log P for non-dissociable APIs or log D for acidic or basic API molecules. Due to pH changes during gastrointestinal transit, only log D is meaningful when ionizable drugs are administered orally. As a rule of thumb, for a drug to be fully absorbed, its log P value should be between -1 and 3.4. To calculate absorption parameters in drug discovery and early development stages, the so-called "rule of five" has been proposed to predict drug-like properties of substances using solubility and permeability parameters. Thus, when there are more than 5 hydrogen bond donors (-OH and -NH groups), more than 10 hydrogen bond acceptors, and a molecular weight greater than 500, the calculated log P is greater than 5 (Lipinski et al., 2001). Only APIs that are substrates for transporters are exempt from this regulation. Given that almost one-third of the drug substances listed in the United States Pharmacopeia (USP25) are poorly or insoluble in water, chemical and/or technical means are required to ensure drug absorption (Langguth et al., 2004). Solubility optimization may involve the use of more soluble salts, eg the water solubility of tetracycline is increased six-fold by the formation of hydrochloride, or nine-fold in the case of phosphate. It should be noted that the presence of excess chloride ions in gastric acid reduces solubility, leading to salting out of the drug in vivo (Thomas and Rubino 1996). A study evaluating seven salt forms of protein kinase inhibitors found that mesylate monohydrate was five times more soluble, more stable, and 2.6 times more bioavailable than oral beagle hydrochloride, and are easier to process (Engel et al., 2018). 2000). therefore,

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The solubility of weakly acidic drugs can be increased by forming salts with cations. Although the amount of counterion is usually too small to cause adverse effects, it should be taken into account that potassium may be irritating to the gastrointestinal tract, magnesium is laxative, and calcium is constipating. When drugs exist in different crystalline forms, they can have different solubilities. Among the different polymorphs, the metastable state with the highest free energy has the highest solubility and thus greater uptake than the stable polymorph. For example, the hydrophobic oral antidiabetic drug tolbutamide can occur in four polymorphic forms with different conformations and motility of the toluene and n-butyl moieties (Kimura et al., 1999). In the dissolution test, the metastable polymorph with the highest solubility transformed into another metastable polymorph with almost the same solubility within 3 hours. According to Beagle studies, the latter metastable monoclinic polymorph is 2.5 times more bioavailable than the stable orthorhombic form. Since this metastable form of tolbutamide requires at least 60°C and 75% relative humidity to convert to the poorly soluble orthorhombic form, it is not expected to be exposed during processing, storage, or dissolution of the final oral drug product. problem appear. Another metastable state of a drug is the amorphous form, which requires less energy to dissolve due to the lack of a crystal lattice and is therefore more bioavailable than the crystalline form. Amorphous forms of drugs can even emerge during pharmaceutical processing such as milling, spray drying, lyophilization, granulation, compression, and film coating. To avoid spontaneous crystallization of amorphous drugs, pharmaceutical formulations should be stored at temperatures below the glass transition temperature of 50°C (Yoshioka et al., 1994). For indomethacin, which has a melting point of approximately 50 °C, co-precipitation with 5% PVP has been reported to remain amorphous up to 70 °C, allowing storage at ambient temperatures (Yoshioka et al., 1995 ). Furthermore, concomitant drugs for exercise or pH adjustment, the use of surfactants or complexation for solubilization can enhance solubility as a prerequisite for absorption (see Section 2.8). In addition to the physicochemical properties of drugs, the physiology of the gastrointestinal tract determines the absorption properties of active pharmaceutical ingredients from oral drug delivery systems.

1.2

Physiological conditions of the gastrointestinal tract

When a formulation is administered orally, it encounters a wide variety of environmental conditions: friendly or hostile environments, short or long-term residence, exposure to light and even severe stimuli. Clearly, these conditions greatly affect the uptake of the agent. These environmental conditions vary by region throughout the GI tract, reflecting the anatomy of the stomach and intestine. See Table 1 for some characteristics.

200

5 5 2,5– 5,0 –

–

500 5.000 3.000

8.500

13 (1–44)

– – –

105 (45–319)

7,0

Fixed pH

11 (2–97)

– – – –

5,5–6,5 6,1–7,1 7,0–8,0

– 6.0–7.0 686b (534–2.1 (male) 859) 2.8 (female) 54 (20–5.5–8.0 156)

Inlet volume (ml)a –

15–48 standard

>5 minutes 1–2 standard. 2–3 standard

3–4 standard

3.5 seconds fast for 2 hours, eat for 2-8 hours

Residence time of a single oral dosage form –

Cecum – 0.05 7.0 – – – – – Colon – 0.25 5.0 3,000 – – 8.0 – Rectum 0.15 – 2.5 500 – – 7.0 – Average and extreme values in parentheses; b represents filled volume, not just liquid. Sources: Feldman and Barnett (1991), Rouge et al. (1996), Chawla et al. (2003), Schiller et al. (2005)

0,3

0,2–0,3 0,1 2,5 60 3,5 60

Colon 1.5

Duodenum Jejunum Ileum

small intestine 7.0

Table 1 Some physiological properties of the gastrointestinal tract. The maximum length of the liquid part. Suction diameter (m2). 2.5 - - Stomach 0.2 0.1 15 1,500 45 (13-72)

Various bacterial enzymes – – –

Galle, Peptidasen, Lipasen, Amylase – – –

– pepsin, lipase,

Amylase

enzyme

– 1011 –

1011

10 105 107

–

– 10

Microorganisms (number per gram of content) -

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351

Mouth and Esophagus

During administration, the oral dosage form comes into contact with saliva, albeit for a brief period. In humans, 0.5-1.5 liters of saliva are secreted per day, which corresponds to an average pH of 6.4 between the extremes of 5.8 and 7.1. Ptyalin in saliva contains 6 g l-1 of dry matter and digests starch, while mucin lubricates solids. With the exception of specific oral bioadhesives or sublingual formulations containing readily absorbable lipophilic agents, the contact time with the oral epithelium is too short for appreciable absorption (Smart 2005). When swallowed, the dosage form usually passes rapidly through the esophagus. This muscular tube is strongly folded in the transverse and longitudinal directions, is 2 cm in diameter, and connects the pharynx with the gastric cardia. Approximately 300 glands are embedded in the submucosa and secrete mucus onto the squamous epithelium, providing a moist environment. Under the control of the medulla oblongata, a normal adult swallows 600 to 700 times a day, but at night, the frequency is reduced to one-tenth. Based on magnetically labeled monitoring of a capsule 16.1 mm long and 5.5 mm in diameter, esophageal transit times were in the range of less than 10 s and further decreased to about 1.4 s with increasing co-intake of water under optimal conditions (Weitschies et al., 2005). Although the upright position accelerates passage compared with the supine position, esophageal transit time does not appear to be related to esophageal propulsion motor activity. Although the submucosa contains large blood vessels, short exposure times often impede drug uptake.

1.2.2

magan

After hopefully unimpeded passage through the cardia (the area 5-6 cm above the stomach with a pressure of 15-40 mmHg and no "splash"), the drug delivery system enters the stomach. The mucosa of this highly extensible organ contains approximately 3.5 million gastric pits, each with four glands, corresponding to a density of approximately 100 gastric pits per square millimeter of surface. The four secretory cells that line these pits secrete mucus and as much as 1.5 liters of gastric juice per day. Gastric juice mainly contains pepsin, hydrochloric acid, gastrin, intrinsic factor, 1.7 g·L-1 soluble mucin and other compounds. This acidic environment helps dissolve the essential active ingredient, which is the first step in absorption and action; on the other hand, peptides or other acid-labile drugs may denature and/or proteolytically degrade, thereby losing their potency. At higher magnification, the mucosa consists of a single layer of columnar epithelium, 20–40 mm in height, covered by a viscoelastic mucus layer with an average height of 140 mm. The average turnover time of the mucus layer is 4-5 hours due to bicarbonate secretion, creating a pH gradient from luminal pH 2 to neutral pH at the cell surface. This converts the ionization of the basic drug into a non-dissociated drug

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A shape that facilitates diffusion across cell membranes. In contrast, the mucus layer represents a diffusion barrier preventing absorption, even for small molecules. In general, there are two main parameters that affect the fate of a drug delivery system in the stomach and thus absorption: pH and gastric residence time (see Table 1). Although gastric pH as determined by the Heidelberg capsule ranges between 1.3 and 2.1 in fasting people, it fluctuates widely after meals (Dressman 1986). Foods such as dairy products can buffer and neutralize stomach acid, but the pH can become even slightly alkaline, such as B. This can lead to rupture of the gastric lining and subsequent premature release of acid-sensitive or irritating drug substances. Since gastric emptying is also highly dependent on the fed or fasted state, it determines the time it takes for the formulation to reach the small intestine and thus is a key parameter for reaching the site of absorption. The "digestive interphase migratory motor complex" occurs every two hours in the fasted state or after eating, as soon as the digestible contents leave the stomach. It begins with a delay phase of 30-45 minutes (first phase) of inactivity. Irregular mixed contractions lasting about 30 minutes (second stage) are followed by intense circular contraction waves, so-called housekeeping waves, and sometimes (hopefully faint) borborygmi (third stage). These waves sweep all indigestible material, including pills, through the open pylorus into the small intestine, where the surge slowly subsides. After 5-15 minutes of peak activity, there is a period of rest (Phase IV) in which locomotor activity is negligible. When eating, the "Antrum Grinder" mixes, grinds, pushes, and empties material smaller than 1-2mm from the stomach by repetitive and irregular contractions at a rate of about 3 times per minute. Thus, in the fasted state, unit dosage forms larger than 4mm remain in the stomach until the next housekeeping wave arrives, which may be anywhere from a few minutes to two hours. When fed, these larger whole preparations undergo antral grinding and exit the stomach after 2-8 hours. This time frame depends largely on the composition of the diet. In general, drug absorption takes longer (40 minutes) with a high-fiber diet than with a no-fiber diet (30 minutes) when administered as a solid dosage form. In comparison: liquids have a very short delay time of 8-9 minutes (Walter et al., 1989). In addition, fats and oils increase gastric residence time because they inhibit antral contraction but at the same time increase pyloric contraction (Houghton et al., 1990). However, the uninterrupted feeding didn't give the housekeeper time, at least in the evening, to clean up Dawanzi's stomach. A problem associated with dangerous side effects observed in obese patients is "dose dumping". Once they stop eating for a few hours, all the collected tablets are immediately released into the small intestine, triggering severe overdose symptoms. In contrast, multiple units smaller than 2 mm leave the stomach within 10-40 minutes in the fasted state because they are processed like fluid. During charging, the residence time of the particles depends on their density. With a water-like density, the particles leave the stomach quickly in the central fluid stream. At higher densities, they settle in the greater curvature of the stomach, escape antral grinding, and are emptied after meals, causing gastric retention lasting more than 40 minutes. In general, drugs are not absorbed efficiently in the stomach.

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small intestine

When passing through the pylorus, the more or less decomposed dosage form reaches the small intestine, which has an average diameter of 3-4 cm. Due to the folding of the mucosa, the surface of the epithelial cells and the apical membrane of the epithelial cells are larger than a tennis court, which facilitates the absorption of nutrients and drug molecules. Despite similar histology, the absorptive capacity and secretions of the duodenum, jejunum, and ileum are quite different. In humans, the absorptive villi are short and wide in the duodenum, taller and most numerous in the jejunum, and become smaller and smaller in the ileum. The mucosal area per cm of serosal length is estimated to be approximately 98 cm in the jejunum, reducing to approximately 20 cm in the lower ileum (Kompella and Lee 2001). In addition, the intestinal epithelium is less hyperpermeable to small molecules. This emphasizes the excellent absorptive capacity of the duodenum, which may be offset by the rapid duodenal passage of seconds to minutes (Schiller et al. 2005). In addition, high blood perfusion supports absorption, maintaining a drug concentration gradient between the intestinal lumen and plasma. Almost one-third of the cardiac output goes through the gastrointestinal organs and 10% supplies the small intestine at a flow rate of 500 ml/min. Interestingly, the villi receive 60% of the blood flow (Washington et al. 2002). There are also various secretions that aid or hinder dissolution, integrity, and absorption. Approximately 0.7-2.5 liters of isotonic pancreatic juice, pH 7.0-7.7, is excreted daily into the duodenum. It contains several different enzymes, including proteases and lipases as well as amylases, that activate prodrugs, degrade protein drugs and increase osmolarity by cleaving polysaccharides into low molecular weight oligosaccharides. In addition, bile secretion supports absorption. Up to 600 ml of hepatic bile is produced daily, concentrated and collected in the gallbladder, which has a volume of 50-65 ml. While doing chores or eating, bile is released into the duodenum. Lipids contain natural emulsifiers such as bile salts, lecithin, and cholesterol and are therefore emulsified into small droplets that are broken down by lipase. The resulting free fatty acids form mixed micelles with bile acids, likely accumulating on the cell surface where the fatty acids are taken up. Bile salts are absorbed by the ileum to close the enterohepatic circulation. It has been proposed that lipophilic drugs are absorbed through this fat absorption route (Florence and Attwood 2006). To neutralize the acidic chyme and set an optimal pH for proteolytic activity, the Brunner glands in the duodenal crypts secrete bicarbonate in addition to mucus. Since pH is on a logarithmic scale, this change has a dramatic effect on ionization, water solubility, and absorption efficiency. Although different pH ranges are reported for the anatomical regions of the small intestine, it is generally accepted that the pH of the duodenum is mildly acidic to neutral and remains more or less at pH 7.5 until descending through the ileocecal junction to pH 5.5 (Lui et al., 1986; Hardy et al., 1987).

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In addition to this absorption barrier consisting of pH and enzymes, another factor to consider is the mucus barrier. Brunner's glands and goblet cells secrete mucus granules that swell when they come in contact with water. In this way, a viscous aqueous layer with an average thickness of 30-100 mm to 192 mm forms next to the absorbed enterocytes. This so-called unstirred water layer acts as a diffusion barrier and can sometimes be the rate-limiting step of absorption (Allen 1984). Interestingly, pH measurements using microelectrodes revealed the presence of an acidic microclimate in the pH 4.5-6.0 range near the cell surface (Washington et al., 2002). While this could lead to weak base malabsorption, this hypothesis has been experimentally disproved. Another driving force for absorption is convective water flow. The proximal small intestine absorbs approximately 8 L of water per day at a flow rate of 50 mL min-1 (Schiller et al., 2005), probably due to slow passive water transport due to osmotic pressure differences between blood and luminal contents due to Differences in hydrostatic pressure due to contraction and rapid active water transport through aquaporins (King et al., 2004). There is an experimentally confirmed phenomenon that the absorption of acidic and basic drugs increases simultaneously with water intake, which contributes to successful absorption in the intestinal tract. This so-called solvent resistance can be simply explained by the fact that removal of water concentrates the API, thereby facilitating diffusion in the mucosal epithelium (Florence and Attwood 2006). As demonstrated by magnetic resonance imaging of fasting subjects, the fluid volume was not evenly distributed throughout the small intestine, but separate pockets were found, especially in the distal region, which were filled with approximately 12 ml of water. After meals, this volume decreases from 12 ml to 4 ml while the amount increases (Schiller et al. 2005). This observation raises questions about the impact of such subdivision on impaired drug absorption in non-deagglomerated dosage forms, resulting in intra- and inter-subject variability in plasma levels. Residence time is another key factor affecting the degree of absorption. After a meal, the motility pattern of the small intestine consists of approximately 2 cm longitudinal contractions, which mix chyme and expose it to the absorptive surface, and peristaltic contractions, which push food toward the large intestine in cycles of 1–3 contractions followed by 5–40 seconds of inactivity. In the fasted state, the "interphase migratory motor complex" travels from the stomach to the small intestine, but with reduced intensity. In contrast to the stomach, residence times for solutions, pellets, and individual units are in the same range of approximately four hours and are not affected by food (Davis et al., 1987). Microparticles have a longer and more reproducible residence time than tablets, resulting in more predictable plasma levels. Unlike the stomach, meals do not affect the residence time of the formulation in the small intestine, but can significantly affect the absorption of dissolved API. Binding of the API to dietary proteins, adsorption to food, or competition of the API with dietary compounds for carrier proteins can reduce the fraction absorbed. In addition, viscous chyme can limit the entry of dissolved API into the absorptive epithelium. On the contrary, bioavailability

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Drugs with a high first-pass effect are enhanced in the presence of food. Although the absorption process itself is not affected, the fraction absorbed increases due to the saturation of metabolic brush border and liver enzymes and increased meal blood flow (Melander 1987). In the small and large intestines, absorption takes place not only in the blood, but also in the lymph vessels. The open lymphatic capillaries in the villi are involved in fat absorption; in addition, there is the local immune system. Gut-associated lymphoid tissue (GALT) is composed of single and clustered specialized epithelial cells called M cells. These oval or oblong collections of M cells are called Peyer's patches and are usually located at the mesenteric border of the bowel. They range in size from a few millimeters to an extreme of 28 cm in adults and make up 25% of the gastrointestinal mucosa. The number of patches increases towards the distal end of the colon. The use of this route for particle uptake and vaccination is described below. In addition, absorption is not uniform throughout the gastrointestinal tract and there are local differences in drug absorption. These "windows of absorption" are attributed to regional differences in pH, mucus composition and thickness, surface area, and enzyme activity. Based on human studies, e.g. piretanide absorbed from the stomach and duodenum (Brockmeier et al., 1986), ciprofloxacin (Harder et al., 1990), captopril (Hu and Amidon, 1988) and metoprolol (Jobin et al., 1985) from the gut and diltiazem from the gut of the colon in small animals.

1.2.4

colon

Entry of dosage forms into the colon is controlled by the ileocecal portion, which prevents reflux of colonic contents, but also allows retention of larger dosage forms in the small intestine. Colonic residence times of 2-20 hours have been reported after large single-dose administration to young and elderly volunteers (Metcalf et al., 1987). In general, APIs are less well absorbed in the colon than in the small intestine. The absorptive area is only 0.25% of the area of the small intestine due to the absence of villi and its shortened length, although the diameter of the small intestine can reach 8.5 cm in the cecum and decrease to 2.5 cm in the sigmoid segment (Read et al. 1980). For comparison: The small intestine has an absorbable area of approximately 1,700 cm2 per centimeter of bowel segment, and the large intestine approximately 20 cm2. One of the main functions of the colon is to reabsorb water and electrolytes, which also limits drug dissolution. The large intestine absorbs almost 90% of the water that enters it (Debongie and Philips 1978). Ingest up to 4 L of water per day at a rate of 2.7 mL/min to form semi-solid stools that are 60-85% water. The aqueous environment for absorption is provided by water in the chyme and irregularly distributed and infrequent pockets of water of 1-2 mL in volume in the transverse colon. Their number increases after meals, but their volume remains largely the same. Since 90% of oral capsules do not come into contact with liquid

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The impact of these pockets on drug absorption in solid dosage forms remains to be elucidated (Schiller et al., 2005). Another function of the colon affecting absorption is mixing and lubricating the contents of the cavity. Goblet cells located in many crypts secrete mucus, which reduces friction but also forms a layer of mucus that coats the absorbing enterocytes. In this way, the mucus layer acts as a diffusion barrier, similar to an 8% polyacrylamide gel (Smith et al., 1986). The mucus layer is part of the microclimate adjacent to the enterocytes and maintains a pH of 6.8, in contrast to the varying pH in the main phase (McNeil et al., 1987). Colonic motility involves segmental contractions, primarily in the proximal colon, that mix the contents and increase contact with the mucosa. Antiperistaltic contractions push the luminal contents towards the ileum and thus contribute significantly to prolonging the residence time of drugs and formulations in the colon. Finally, the distal colon is emptied with massive movements of intense sustained contractions 3 to 4 times per day. Interestingly, eight subjects showed increased colonic motility within 4 minutes of drinking caffeinated or decaffeinated coffee, but six volunteers did not respond (Brown et al., 1990). Overall, the interaction of different movement patterns resulted in a longer dwell time of 24 to 26 hours in children and young adults, with a further increase to 110 hours in older adults (Kirwan and Smith, 1974). In healthy volunteers, tablets and capsules have an average residence time in the large intestine of 20 to 30 hours, thus allowing sufficient time for absorption even at low rates (Kompella and Lee 2001). In addition, small particles pass through the large intestine more slowly than large particles (Hardy et al. 1985). However, increasing fiber content shortens residence time and may negate any absorption benefits. Despite lower blood flow (8-75 mL/min) compared to the small intestine, delivery in the large intestine opens up great promise for peptide and protein delivery, as protease activity is 20 to 60-fold lower than in the small intestine (Washington et al. 2002). In addition, there is a neutral luminal pH ranging from 6.4 in the ascending colon to 6.6 in the transverse colon to 7.0 in the descending colon (Evans et al. 1988). However, the pH in the large intestine fluctuates widely, and depending on the protein content of the food, the pH may increase to 8.0 or decrease in the case of a high-fiber diet, thereby affecting drug absorption. A unique feature of the colon is the over 400 different bacterial flora, including aerobic bacteria mainly in the proximal colon and anaerobic bacteria in the cecum. Similar to a fermentation chamber, multiple reactions such as hydrolysis, aromatization, reduction, esterification, decarboxylation, and deamination are possible, which affect the amount of drug absorbed but are also used as a strategy for colon-targeting drugs (Edwards 1997 ). When cellulose is digested by bacterial cellulase, volatile compounds are formed and these gas bubbles reduce the contact of the active substance with the mucous membranes. In addition, short-chain fatty acids are formed during carbohydrate fermentation, which temporarily lowers luminal pH prior to absorption (Lipton et al., 1988).

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2 Improving Formulation Parameters for Oral Drug Delivery In order to save costs by accelerating drug development, 20 years of research have concluded that the absorption properties of a drug can be predicted by its solubility as a measure of its physicochemical properties and permeability, taking into account to the physiological conditions of the digestive tract. These two parameters form the basis of the Biopharmaceutics Classification System (BCS), which consists of four classes that predict the biopharmaceutical quality of drugs (Amidon et al., 1995; Lennernas and Abrahamsson, 2005). For class I drugs with high solubility and permeability, no problem is expected, whereas for class II drugs with low solubility and high permeability, solubility is the rate-limiting step and thus needs to be increased. On the other hand, Class III drugs with high solubility and low permeability require some effort to improve absorption. Finally, Class IV drugs have both low solubility and low permeability, resulting in limited bioavailability, thus requiring both increased solubility and absorption. According to this classification, there is a toolbox of different techniques that can improve oral delivery by tuning a wide range of parameters, as described below.

2.1

size

Considerable effort has been expended over the past two decades to develop methods that enable the fabrication of drug particles with precisely controlled properties at the micro- and nanoscale. There are two main reasons for the interest in nanoparticles for oral dosage forms. First, a large number of newly developed drug candidates and marketed substances belong to BCS classes II and IV, which are not bioavailable due to low solubility as a common feature. Nanoscaling of drugs has become a key approach to address this issue since it can significantly improve the dissolution profile of compounds (Kesisoglou et al., 2007). Second, seminal reports of oral absorbability of nanoparticles have stimulated efforts aimed at developing drug delivery systems that efficiently enter the systemic circulation. Admittedly, this route of administration is extremely complex. However, since oral formulations require a high degree of patient compliance, several strategies for administering highly potent yet sensitive biotechnological drugs have been explored (Sood and Panchagnula 2001).

2.1.1

nanoscale impact

Due to their small size, nanoscale drug particles distribute faster and more uniformly in the gastrointestinal tract compared to conventional unit dosage forms (Asghar and Chandran, 2006). Due to the high surface

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The use of nanoscale drugs should also lead to longer residence times and more consistent bioavailability due to increased interaction of particles with the epithelial lining and mucus (Keck and Müller 2006). However, perhaps the most pronounced effect of the nanoscale on the bioavailability of a substance is based on changes in the dissolution behavior of the compound. In this regard, the effect of shrinking drug crystals to the nanoscale on dissolution rate is mainly attributed to the increase in particle specific surface area. This can be illustrated by considering a microparticle (radius ¼ 10 mm) "broken" into 106 nanoparticles (radius ¼ 100 nm). Despite having the same total mass, the nanoparticles have a hundred times larger dissolved surface area. The transfer rate of solvated molecules from this surface region into the bulk solution is determined by the hydrodynamic boundary layer thickness (Borm et al., 2006). For particles >50 mm in size, the layer thickness is fairly constant, typically around 30 mm. In contrast, for smaller micro- and nanoparticles, the predicted thickness is comparable to the diameter or radius of the particle (Galli 2006). Due to the reduced boundary layer associated with this large surface area, dissolved molecules can be transferred more quickly into the total solution. The link between these effects explains the higher solubility of nanoformulated substances and is illustrated in the extended Noyes-Whitney equation (Equation 1) of Nernst and Brunner: dCt D A ðCs Ct Þ ¼ h dt

(1)

where CS is the equilibrium solubility of the substance, Ct is the concentration at time t, D is the diffusion coefficient, A is the particle surface area, and h is the hydrodynamic boundary layer thickness (Noyes and Whitney 1897; Brunner 1904; Nernst 1904). In addition to changing dissolution rates, nanoscale active substance crystals are also characterized by increased saturation solubility due to greater surface curvature of the particles (Borm et al., 2006; Wu and Nancollas, 1998). This relationship is given in the Ostwald-Freundlich equation (2) for solid-liquid dispersions: log

Cr 2sV ¼ C1 2:303 R T r r

(2)

Cr represents the solubility of a particle of radius r, C1 the solubility of an infinite particle, s the interfacial tension of the substance, V the molar volume of the particle material, R the absolute gas constant, T the absolute temperature and r the density of the solid. Significant surface curvature effects are expected at particle sizes below about 1 micron, which lead to increased dissolution pressure and thus saturation solubility (Müller and Peters 1998). However, since these effects are based on the assumption of spherical particles, it should be emphasized that this is actually the case

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The observed solubility of broken crystallites may differ significantly from that predicted (Tang et al. 2004; Borm et al. 2006). Typically, when the dissolution rate of a poorly water soluble substance is optimized, a faster rate of absorption is monitored in vivo. Therefore, if insufficient dissolution rate is the factor limiting the bioavailability of a substance, reducing the size of drug particles to the nanoscale has been reported to be an effective way to improve drug performance (Liversidge and Cundy 1995) and accelerate its action. action (Wu et al., 2004). ; Hanafy et al., 2007). The latter is of interest, especially in relation to the administration of analgesic substances, in which case there should be no significant delay in therapeutic plasma concentrations. In a study comparing the bioavailability of nanocrystalline naproxen with two commercially available products, Naprosyn1 (suspension) and Anaprox1 (tablets), a 50% reduction in the time to peak plasma drug concentration was found for the nanoscale material. (Merisko-Liversidge et al., 2003). In addition to a faster onset of action, it has been reported that the administration of drug nanocrystals may benefit compounds whose bioavailability often differs between fasted and fasted states. In particular, poorly water-soluble compounds often exhibit increased dissolution rates mediated by food components and gastrointestinal secretions (i.e., bile). By enhancing the solubility properties of these substances independent of external factors, the nanoscale can normalize bioavailability and thus therapeutic response (Wu et al., 2004). This is also an issue for BCS class IV medicinal products, where additional incorporation of penetration enhancers into the formulation may prove to be a valuable concept to achieve higher plasma concentrations (Merisko-Liversidge et al., 2003; Wu et al., 2004 Year). If sustained release of not only rapidly dissolving but poorly bioavailable drugs is also an issue, coating nanoscale drug crystals with mucoadhesive polymers or incorporating drug molecules into polymer particles with bioadhesive properties can improve drug delivery performance of the system (Takeuchi et al. 2001; Bernkop-Schnürch 2005).

2.1.2

Nanoscale Technology for Drug Particles

Most techniques for nanoscaling are based on treating aqueous suspensions of drug particles containing surfactants to stabilize the resulting ultrafine colloids. As a rule of thumb, particle regrowth due to Ostwald ripening should not be a problem if the water solubility of the processing compound is less than 1 mg ml-1 (Merisko-Liversidge et al., 2003). The water content must be reduced if the substance is more soluble or if the nanocrystals are to be converted into a solid dosage form. This is achieved by lyophilization or spray drying after the size reduction process, which improves the long-term stability and redispersibility of the particles. The most established methods for nanoscale solid drug particles are described below. As shown in Table 2, several companies offer formulation development services based on these processes. to clear

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Table 2 Overview of commercially applied micro-nanoscale solid drug technologies and marketed products (if applicable) technology (company) process marketed product (active ingredient; company; year of FDA approval; dosage form) Rapamune1 (sirolimus) wet milling. ; Wyeth; NanoCrystal1 Technology 2000; Tablets) (Elan Nanosystems) TriCor1 (Fenofibrate; Abbott; 2001; Tablets) Emend1 (Aprepitant; Merck; 2003; Capsules) Megace1 ES (Megegestrol; Bristol-Myers Squibb ; 2005; Liquid Oral Suspension) Insoluble Drug Delivery Ultrasound, Homogenization, Triglide1 ( Fenofibrate; Sciele Pharma; 2005; Tablet) Platform Technology Milling, Microfluidization (IDD1-P) (SkyePharma) Precipitation followed by High Pressure Homogenization - NANOEDGE1 ( Baxter) Dissolution - Precipitation - Nanomorph1 (Soliqs/Abbott) High Pressure Homogenization - DissoCubes1 (SkyePharma) High Pressure Homogenization - Nanopure1 (PharmaSol) (in Water/PEG/Glycerol/Oil) BioAqueousSM - Precipitation - Solubilization - Template - Emulsion Service (Anton et al. 2008) (The Dow Chemical - Spray Freezing into Liquid (SFL) -Company) (Hu et al. 2003 ) – Evaporation Precipitation in Aqueous Solution (EPAS) (Chen et al. 2002)

Information on shrinkage methods is not always provided (such as in the case of IDD1-P), but most techniques show similarities. Alternative concepts contained in The BioAqueousSM Solubilization Services (The Dow Chemical Company) are not described in detail but are provided as a reference for the interested reader. Overall, given the relatively recent FDA approval dates for marketed products, an increasing number of formulations containing nanoscale drug particles are expected to emerge in the future.

Bottom-up process In order to produce nanoparticles by precipitation, the substances to be formulated must be soluble in water-soluble solvents that are as non-toxic as possible. In practice, this is a limiting factor for some new chemical entities that are neither

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Soluble in aqueous media, insoluble in most organic solvents. After mixing an organic solution of a compound with an aqueous solution of a stabilizer, environmental conditions must be controlled to induce the precipitation of the desired crystalline or amorphous drug particles in the nanometer or micrometer size range (400–2,000 nm) (Moschwitzer and Müller 2007). At some stage in the process, particle growth must be stopped, which is achieved by adding selected polymers (gelatin, chitosan, poloxamer, etc.). These substances act as growth inhibitors by adsorbing on the particle surface and increasing the viscosity of the suspension. Finally, the resulting drug nanoparticles are stabilized, for example by spray drying with excipients (lactose, mannitol, etc.). This step is necessary to obtain particle size and amorphous state, which determine the drug's dissolution rate and thus its bioavailability in vivo (Rasenack et al., 2003). Top-down process A common feature of top-down drug nanosizing approaches is the application of external forces to induce fragmentation of drug particles at crystal defects. Depending on the technique used, the crushing forces are generated by collisions with grinding media, cavitation, high shear forces, collisions between particles or turbulence (Moschwitzer and Müller 2007). Common to all techniques is that a certain particle size threshold must be accepted in practice, depending on the drug properties and the technique-specific practical energy input. This lower limit in practically relevant process times is a consequence of the reduction in available lattice defects per grain of fractures associated with miniaturization. Therefore, further particle size reductions can only be achieved with an exponential increase in energy input after several treatment cycles. Since reaching this point is determined by the substance-specific hardness of the active ingredient particles, the optimized process parameters leading to the desired particle size cannot be directly transferred to other substances. However, it is often reported that most particles become visibly nanoscale within a few process cycles. Continued reduction does not necessarily lead to a further reduction in the average particle size, but to a narrower particle size distribution. Wet milling Wet milling has been used for many years to produce nanoscale drug particles and is the fundamental size reduction technique of the patented NanoCrystal1 technology (Elan). Basically, the milling chamber containing the milling media (beads), dispersion medium (usually water), drug granules and stabilizer as a whole undergoes a rotating motion, or its contents are agitated by an agitator. The rotating beads create high shear forces and collisions between the particles inside the mill, which together result in a size reduction of the drug crystals to around 80-400 nm. It should be noted that high inner surface

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Grinders also increase product loss due to sticking of ground drug particles. This can lead to cross-contamination and is especially undesirable when dealing with expensive materials, as new chemicals are often only available in small quantities. Another problem is the contamination of grinding media wear, which consists of stainless steel, glass and ceramics based on yttria-stabilized zirconia (Moschwitzer and Müller 2007). This problem can be solved by using durable and high-quality grinding media, such as beads made of highly cross-linked polystyrene resin, with an abrasion range of only 0.005% (w/w) depending on the active ingredient concentration of the dispersion (Merisko-Liversidge et al., 2003). When the success of a formulation technology is measured by the number of products in clinical trials and commercialization, the NanoCrystal1 technology has thus far outperformed its competitors in terms of commercial success (Table 2). High Pressure Homogenization Piston Gap Homogenizer High pressure homogenization has been standard technology in the food and dairy industry for decades. In pharmaceutical formulations, its use has flourished as the method is not limited to emulsions or suspensions but can be performed in the laboratory (0.5 mL) and in large quantities (2,000 Lh–1) (Keck and Müller 2006; Möschwitzer and Müller 2007; Müller and Peters 1998). With high-pressure homogenization, the particle suspension is pushed through the narrow gap of the homogenization valve. Since the volume of the suspension must remain constant over time in front of the valve and within the much narrower gap (~20 mm), the velocity of the liquid suddenly increases as it enters the channel. This causes the static pressure of the liquid to drop below its vaporization point, even at room temperature. This leads to the formation of air bubbles, which is often referred to as "cavitation". After passing through the gap, the fluid velocity is reduced again, causing the bubbles to implode, creating high-energy microjets. Although cavitation is the main mitigating principle for high-pressure homogenization, cavitation is likely to be complemented by other factors such as interparticle collisions and high shear forces (Moschwitzer and Müller 2007). For solid drug granules, no general rules have been established for the addition of surfactants during homogenization. The amount and type of surfactant added was reported to be associated with a reduction only during emulsion processing (Keck and Müller 2006). This can be explained by a decrease in surface tension, which reduces the energy input required to break up the droplet. Compared to wet grinding, product loss is very low due to particle abrasion of the inner surfaces of the equipment as well as machine parts. When using "hard" homogenization conditions of 1,500 bar pressure and 20 repeated cycles, atomic absorption spectroscopy detected only about 0.7 ppm iron in the final suspension (Krause et al., 2000).

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Although high-pressure homogenization of aqueous suspensions is considered highly beneficial for production, the use of water as a dispersant for subsequent conversion of nanocrystalline drug particles into conventional dosage forms such as tablets or capsules is not advisable. In most cases, the water content of the suspension must be reduced before further processing, which requires time- and energy-intensive steps such as freeze-drying, spray-drying or fluid-bed drying (Moschwitzer and Müller 2007). In this case, a more suitable technique was developed and licensed under the name Nanopure1 (PharmaSol GmbH, Berlin), in which the nanocrystals are suspended in reduced water (water-glycerol/PEG mixture) or anhydrous dispersion ( oil, liquid PEG, heat-melt polyethylene glycol). Interestingly, drug particles can also be reduced by high-pressure homogenization in these media, although the possibility of cavitation is significantly reduced in systems with low vapor pressure. In this case, it is assumed that shear forces, interparticle collisions, and turbulent flow mainly lead to size reduction of dispersed drug crystals (Keck and Müller 2006). The main advantage of this method is the ease of handling the suspension after the high pressure homogenization step. When using hot-melt PEG as a dispersant, the nanosuspension can be filled directly into soft or hard gelatin capsules (solidification in the capsules) or cooled, resulting in crystallization of the PEG matrix and particle entrapment. The resulting solid is ground into a powder that can be compressed into tablets or filled into hard gelatin capsules. This provides an added advantage, as the cured PEG matrix stabilizes the nanoparticles, thereby extending the shelf life of the formulation. In addition, PEG melts at physiological temperature, allowing the release of monodisperse nanoscale drug crystals in the gastrointestinal tract. Microfluidization The technology is based on the reduction of suspended drug crystals in the Microfluidizer1 processor. Basically, a booster pump drives a particle suspension through a fixed-geometry microchannel until stream "Y" splits and gradually narrows. As with piston gap homogenizers, the reduction in channel diameter results in a gradual increase in flow velocity, which leads to cavitation and high shear rates. Eventually, the two streams collide head-on, causing further impact and reduction. Although very high homogenization pressures can be achieved using this technique, it has been reported that certain substances require a large number of processing cycles to achieve sufficient particle size reduction (Keck and Müller 2006; Moschwitzer and Müller 2007).

2.1.3

Nanoparticle Absorption

The uptake of nanoparticles at the time of oral administration has been extensively studied for the delivery of highly potent protein and peptide drugs since standard

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Especially in the context of chronic therapy, parenteral use is associated with low patient acceptance (des Rieux et al., 2006). The methods discussed below are mainly based on drug-loaded polymer nanoparticles, for which various fabrication processes have been developed. The first reports of particle uptake from the gastrointestinal tract appeared decades ago. Surprisingly, micron-sized (150 mm) particles such as starch granules, pollen, silicate crystals, and powdered rabbit hair were found in the subepithelial layer of the mucosa after feeding to animals (Volkheimer 1974). This observation has been attributed to a rather passive mechanical process called "absorption," which occurs at low frequencies. Since then, more efficient routes have been identified for the oral administration of particles that eventually enter the systemic circulation. These pathways primarily target normal intestinal epithelial cells and M cells, which reside in the follicle-associated epithelium (FAE), the single or associated lymphoid follicles that cover the gut-associated lymphoid tissue. Sparse M cells exhibit several properties that may prove to be favorable for particle uptake. Compared to normal enterocytes, they are said to be covered with a thinner layer of glycocalyx, which is said to facilitate granule attachment. In addition, M cells have adapted to take up a variety of substances and are located in close proximity to lymphoid tissue (Shakweh et al., 2004). Evidently, this latter fact has sparked great interest in the development of microparticle systems that can induce mucosal immunity (O'Hagan et al., 2006). However, extrapolation of animal data to humans is difficult due to the different FAE anatomy in rodents (10-50% M cells) compared to humans (~5% M cells) (Florence 2005; des Rieux et al. 2006 ). Thus, to date, it has been unclear whether individual M cells could serve as resorption sites through which nanoparticles are taken up in sufficient quantities to produce systemic effects. In this regard, normal intestinal epithelial cells may be preferred given their abundance in the gut. Using polymer nanospheres as a model for oral drug delivery systems, several studies have found the optimal size for effective absorption to be in the 50-100 nm range (Jani et al., 1990). These results are consistent with other findings from more detailed studies focusing on enterocytes (Jani et al., 1992) and inflammatory bowel disease treatments (Lamprecht et al., 2001), which reported the highest uptake of 100 nm particles. In the latter case, high uptake by local diseased tissue rather than systemic uptake is ideal, possibly due to increased uptake of particles by macrophages and entrapment due to increased mucus secretion in inflamed tissues. In general, the role of mucus in the uptake of nanoparticulate materials is controversial. On the one hand, it has been suggested that entanglement of particles between glycoproteins inhibits uptake, and on the other hand, entrapment in the mucus layer may increase residence time in the intestine and contact of particles with the intestinal wall. Clearly, some important questions remain to be resolved if systemic responses to oral nanoparticles are to be successfully achieved.

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For example, in the case of insulin, several studies have shown conflicting therapeutic effects in animals (Delie and Blanco-Prieto 2005). A consequence of these observations is that, especially for drugs with narrow therapeutic windows, when formulated as nanoparticles, normalization of uptake efficiency must be a major research topic in controlling oral bioavailability (Dearn AR 1997). This must include a systematic study of the effect of the type of polymer making up the particle matrix and the type and concentration of surfactant used in the manufacturing process (Delie and Blanco-Prieto 2005). In addition, the drug loading efficiency of polymeric carriers is also an issue, especially in the case of water-soluble compounds such as proteins. Finally, if one considers the complex process of particle uptake into the epithelial cell layer, it is conceivable that the basolateral export and subsequent translocation into the bloodstream represents an underestimated hurdle still to be successfully overcome (Florence 2006).

2.2

pH dependent drug delivery

Due to the significant variation in the pH of the environment in different parts of the gastrointestinal tract (see point 1.2 "Physiological conditions of the gastrointestinal tract" of this chapter), systems administered by the oral route are exposed to various external environmental conditions during passage. Since the uptake of many active substances is known to vary significantly in successive regions of the intestinal mucosa, providing high concentrations of active substances in the fraction with optimal uptake is essential to achieve adequate bioavailability. To this end, physiological changes in pH can be used as chemical triggers to release drugs specifically at the target area, thereby protecting the compound from areas of intended uptake and bypassing deleterious effects in these areas. Most pH-dependent delivery systems involve prevented release in the acidic gastric environment, followed by unhindered or broad release in the small or large intestine with a higher pH range. It is important to note that natural differences in gastrointestinal pH between individuals, as well as transient fluctuations caused by dietary factors, disease, or co-administration of other drugs, can affect the performance of pH control systems and should be considered when designing for safety Up to this point the system must be delivery strategy. A general advantage of enteric-coated tablets or multiparticulate systems is ease of manufacture and relatively low cost compared to other, more complex drug delivery systems (Friend 2005). Current approaches to the development of pH-dependent dosage forms are based on the application of polymer coatings to preformed tablets, capsules, pellets or granules, or the embedding of the drug in pH-sensitive matrices or hydrogels. In addition to traditional forms of enteric coatings, the field of colon targeting is an emerging area that has experienced a particular boost through pH-triggered systems.

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pH Sensitive Coatings

Today, a variety of polymers with different backbone modifications and side chain conjugates are available, enabling precise control of release at a given pH threshold. Among them, methacrylic acid copolymer or Eudragit1 (trademark of Evonik Röhm, Darmstadt, Germany) is one of the best known and most widely used copolymers for the production of pH-sensitive coatings or hydrogels. Various types of Eudragit1 powders and ready-to-use suspensions are available on the market, which are copolymers of methacrylic acid and methyl methacrylate with different functional groups. The ratio of free acid to ester groups determines the degree of swelling and/or dissolution rate of the polymer, and the number and type of surface charged groups determine the predominant anionic or cationic character (Moustafine et al., 2005). By introducing positively charged units, such as dimethylamino or quaternary amino groups, polymers can be regarded as polycations (Eudragit1 E, RL, RS types). These polycations are insoluble at neutral saliva pH and, depending on the nature of the charged group, dissolve completely in the stomach (dimethylamino, Eudragit1 E) or swell without complete decomposition in the case of quaternary amino groups (Eudragit1 RL and RS ). For negatively charged polymers (Eudragit1 L, S, and FS), hydration of the ionized carboxylate groups at pH 5.5 or higher results in insolubility in acidic gastric fluid, followed by alkaline released in the environment. This prevents exposure of the encapsulated compound to the harsh conditions of the stomach and allows unimpeded delivery to the preferred site of absorption. Fine-tuning the polymer composition and the type and degree of derivatization can control the pH threshold with relative precision. This offers the possibility of more accurately determining the site of dissolution in the small or large intestine, but is limited by the fact that the pH of the lower GI tract varies much less than the transition from gastric juice to the duodenum and may vary with nutritional status. changes with changes. The application of different Eudragit1 coatings in pharmaceutical dosage forms is widely used in numerous products marketed for topical or systemic therapy (eg Asacol1, Salofac1, Claversal1, EntocortTM EC, Budenofalk1) (Friend 2005; Fedorak and Bistritz 2005). However, since the pH profile of the gastrointestinal tract is variable, precise control of the release site is not possible, so today's trend is to use combination delivery systems with more than one coating or targeting principles. Various systems have been published covering combinations of pH and time-dependent release, mainly consisting of an outer shell of Eudragit1 type L, S or FS and an inner core of filled polymers such as Eudragit1 types RL and RS, ethylcellulose or hydroxypropylmethylcellulose (Vandelli et al., 1996; Gupta et al., 2001; Edsba¨cker et al., 2003). Moustafine and colleagues investigated the formation of interpolyelectrolyte complexes between different types of Eudragit1 polymers, which showed different release profiles compared to conventional physical mixtures of individual polymers (Kumar 2000; Moustafine et al. 2005). Similar ion complexes can be observed with other polymers used for controlled purposes

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releases, for example, chitosan or carboxymethylcellulose, and can be made pH-sensitive by incorporating one or more pH-sensitive compounds (Lorenzo-Lamosa et al., 1998). As mentioned earlier, more sophisticated delivery systems are gaining in popularity, as the combination of different targeting strategies allows further fine-tuning of release locations. With regard to pH-dependent drug delivery systems, these new advances relate in particular to two areas, the field of colon targeting on the one hand and the development of novel pH-sensitive hydrogels on the other.

2.2.2

pH and Colon Targeting

In recent years, targeted therapy in the distal gastrointestinal tract has gained increasing attention not only as a way to improve the treatment of localized diseases such as infection, cancer or Crohn's disease, but also offers interesting possibilities for systemic absorption (Yeh et al., 1995). Fragile compounds such as proteins or peptides are exposed to far fewer digestive proteolytic enzymes in the colon than in the rest of the GI tract and spend a long time in the colon, accounting for almost 80% of the total GI transit time, resulting in The main challenge for drug uptake is the absorption window. The pH-dependent degradation of polymeric drug carriers or coatings has been cited as an important principle for the successful implementation of such colon-targeting strategies, although selectivity of release within the target range remains problematic. Due to the lack of pronounced pH changes in the lower GI tract, there are no effective stimuli that can reliably trigger breakdown of target systems via pH alone, especially in the upper colonic region. Furthermore, as is now known, the pH in the colon can vary and be even more acidic than in the small intestine (Lorenzo-Lamosa et al., 1998; Friend 2005). Therefore, many recently developed colon-targeting strategies are based on the combined application of pH-triggered and second-principles based on different modes of operation. Notably, these alternative principles include enzymatic systems that exploit the specific microbial communities present in the colon and use different types of bacterially cleavable polymers containing azo groups or glycosidic linkages (Brondsted and Kopecek 1992; see "Using Nanoparticle Technology for Cancer Therapy", section 6). .The main disadvantage of these synthetic or natural polysaccharides (such as chitosan, pectin, dextran or chondroitin sulfate) is their high solubility in gastric juice (Chen et al., 2004) A combination of strategies to avoid premature decomposition seems obvious. To bypass the upper GI tract, acid-insoluble pH-dependent or time-delayed time-dependent coatings can be used. However, the timing system should be used with caution, as the usual transit time to the colon in the fasted state is known to be 4 to 6 hours, but may be significantly increased due to delayed gastric emptying in the fasted state. Therefore, a pH-controlled delivery system that is less susceptible to fluctuations in nutrient status may be preferred.

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Lorenzo-Ramosa et al. A colonic multiparticulate drug delivery system was developed consisting of chitosan micronuclei encapsulated in Eudragit1 L-100 or S-100 microspheres (Lorenzo-Lamosa et al., 1998). Combination of a biodegradable chitosan matrix with an acid-resistant Eudragit1 coating in the colon resulted in uncompromised integrity, no release during passage through the gastric lumen, and subsequent near zero-order kinetic drug release in the small intestine and colon. A coherent interplay between pH-dependent dissolution, swelling and bacterial degradation, made possible by the formation of ionic crosslinks between chitosan and acrylic polymers, is responsible for the observed release profile. In addition to the combination with enzyme targeting, the pH-controlled principle with slow-release matrices has also been successfully used for colonic delivery. EntocortTM EC capsules (AstraZeneca, Lund, Sweden) for the treatment of Crohn's disease consist of a Eudragit1 L100-coated shell containing drug-carrying ethylcellulose particles (Edsba¨cker et al., 2003) . Budenofalk1 capsules (Falk Pharma, Freiburg, Germany) (Fedorak and Bistritz 2005) function on the basis of microparticles coated with Eudragit1, which dissolve at pH >6.4. Krishnamachari et al. The use of poly(D,L-lactide-glycolide) (PLGA) microparticles coated with Eudragit1 S-100 was recently proposed to deliver budesonide to disease sites (Krishnamachari et al., 2007).

2.2.3

pH-responsive hydrogel

Swellable and biodegradable hydrogels have proven to be important tools for enhancing drug delivery to specific parts of the small intestine or colon, and can be specifically designed to use changes in pH as stimuli to induce swelling or dissolution. Likewise, different targeting strategies can be employed depending on the chemistry of the hydrogel. The most common approach, however, is to design the gel to prevent swelling in the acidic gastric juices and not release the incorporated drug until it reaches the neutral or basic regions of the lower GI tract. This selective swelling behavior can be achieved by using matrices containing, for example, carboxyl groups, which maintain a tight network at low pH and increase swelling upon ionization at higher pH. Various polymers or copolymers have been investigated for this purpose, most of which are prepared by copolymerization or crosslinking of polymer precursors. It should be noted that hydrogels with similar composition may exhibit different degradation modes depending on how they are fabricated (Yeh et al., 1995). This can be attributed to the different network structures resulting from different synthesis methods. Crosslinking between polymers can occur through covalent bonding, hydrogen bonding, or physical entanglement, and can strongly affect drug release kinetics by favoring surface erosion or bulk erosion as the degradation mode. In general, the higher the crosslink density, the lower the swelling of the gel in aqueous media and the slower the drug release (Yeh et al., 1995). Once complete swelling is achieved, drug release can be further enhanced by chemical hydrolysis or by disrupting the matrix by

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Combined with biodegradable cleavage sites, these sites are now available to degradative enzymes (Akala et al., 1998). All pH-responsive hydrogels are polyelectrolytes containing acidic or basic groups that accept or donate protons according to changing environmental conditions (Qiu and Park 2001). It is important to note that due to the proximity of adjacent charged groups in the polyelectrolyte matrix, the overall degree of ionization may be less than a similar amount of the corresponding monoacid or monobase (Mayo-Pedrosa et al., 2008). The degree of swelling of hydrogels is mainly determined by the electrostatic repulsion between trapped ionized groups, which is much higher than that of uncharged polymers. The use of comonomers such as hydroxyethyl methacrylate and methyl methacrylate leads to different hydrophobicities of the gels and allows further tuning of the pH response. Cationic polymers swell at the acidic pH of the stomach and can consist of, for example, chitosan or N,NO-dimethylaminoethyl methacrylate (Siegel et al., 1988; Patel and Amiji, 1996) . Hydrogels with anionic groups begin to swell or dissolve in the neutral or alkaline environment of the small or large intestine, and can be prepared from acrylic acid, poly(acrylamide), poly(methacrylic acid), poly(vinyl acetal diethyl Aminoacetate) of various derivatives. or dextran (Akala et al., 1998; Aikawa et al., 1998; Chiu et al., 1999; Qiu and Park, 2001). Otherwise, inert hydrogel matrices, which do not change with pH, can be crosslinked by using a suitable crosslinker, such as crosslinker. For example, poly(vinyl alcohol) hydrogels cross-linked with maleic acid (Peppas and Peppas 1990; Gohil et al. 2006; Basak and Adhikari 2008). Chitosan, a copolymer of D-glucosamine and N-acetylglucosamine, is widely used for drug delivery due to its good biocompatibility, but it is often found in the intestinal tract due to its rapid dissolution in the acidic environment of the stomach drug delivery is limited. Chen et al., 2004). However, complexes of chitosan derivatives and polyanionic alginates can be used as pH-sensitive hydrogels that can withstand the acidic conditions in gastric juice if the chitosan is additionally immobilized with a suitable cross-linking agent. Due to toxicological reasons, the use of glutaraldehyde is restricted, Chen et al. The use of genipin as a low toxicity cross-linking agent of natural origin with alginate-N,O-carboxymethyl chitosan hydrogel system has been proposed (Chen et al. 2004). Sufficient retention of encapsulated model drugs and subsequent release at neutral pH can be demonstrated. Recently, in situ coating of cellulose/PVP pellets with pH-sensitive hydrogels via photopolymerization has been reported, allowing specific tuning of the pH response and incorporation of drugs into the pellet matrix and coating gel (Mayo-Pedrosa et al. ,Year 2008).

2.3

swelling

Swellable hydrophilic polymers are widely used components in controlled release systems. They are mainly used in the production of

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For paint. Drug release from a swellable matrix is mainly controlled by the polymer swelling rate and the diffusion of the API through the gel layer. A variety of designs are available, allowing different timed-release regimes, from pulsed release to long-term sustained release. Cellulose ethers such as hydroxypropylmethylcellulose (HPMC), methylcellulose (MC) and hydroxyethylcellulose (HEC) are the most popular polymers in this class due to their good swelling and compressibility properties. important representation. They have good properties, most have generally recognized as safe (GRAS) status, and are available in a variety of viscosities and alternative grades. In addition, copolymers of methacrylic acid, galactomannans, alginates, and many other natural and synthetic polymers can be used. Upon contact with water, these polymers swell and change from a glassy state to a more permeable rubbery state, forming a gel layer. At first, the penetration of water dominates, the source front representing the inner boundary of the gel layer continues to move, and the thickness of the gel layer increases. At the same time, part of the polymer dissolves or erodes at the outer boundary of the gel layer, the so-called erosion front. If water absorption and polymer erosion are in balance, layer thickness will stagnate. Finally, when all the polymers are hydrated, the thickness of the gel layer will decrease. Furthermore, another boundary within the gel layer, the so-called diffusion front, has been described, which represents the boundary between dissolved and still undissolved active substances (Colombo et al., 2000; Gazzaniga et al., 2008) . As mentioned above, the thickness of the gel layer is time-dependent, which is important for drug release kinetics. Depending on the rate-determining step, drug release can be diffusion-controlled, swelling-controlled or chemically controlled. Diffusion-controlled delivery can be modeled by Fick's law with constant or variable diffusion coefficients. When drug diffusion is faster than polymer swelling, drug release depends on the swelling rate of the polymer network. A prominent example of this swelling-controlled drug release is HPMC. In chemically controlled release, a chemical reaction within the delivery matrix is the rate-limiting principle, such as hydrolysis or enzymatic degradation of the polymer or cleavage of the conjugated drug from the network (Lin and Metters 2006). Another key parameter for active substance release is the mesh size of the hydrogel, which is typically between 5 and 100 nm in the swollen state. If the API is smaller than the grid size (which is the case for most APIs), its diffusion is not hindered. However, for nonporous hydrogels and porous hydrogels with pore sizes similar to drug sizes, diffusion is sterically restricted and the diffusion coefficient decreases. By choosing appropriate parameters, the structure and mesh size of swollen hydrogels and the degree of gel swelling and degradation can be tuned (Lin and Metters 2006). This enables the development of specially tailored control delivery devices. Sustained release independent of pH, ionic strength, or temperature is of particular interest for chronotherapy of chronic diseases that exhibit circadian rhythm fluctuations, such as rheumatoid arthritis or bronchial asthma, which appear especially in the early morning. Delayed release allows for bedtime

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Dosing without exposing the patient to the API overnight (Gazzaniga et al., 1994). The most common extended-release systems use swellable hydrophilic polymers and are either coated depot systems or capsule systems with hydrophilic polymer stoppers (Gazzaniga et al., 1994). Reservoir system coatings with hydrophilic polymers are typically applied by pressure coating techniques, spraying, dipping or powder coating. An example is the ChronotopicTM delivery system, which consists of a drug core with one or more units or a gelatin capsule surrounded by a layer of HPMC responsible for the lag phase (Gazzaniga et al., 1994). It can be used for sustained release after a programmed lag phase, or for time-dependent, colon-specific delivery when an additional external enteric coating film is applied (Sangalli et al., 2001). Bimodal plasma levels are achieved through a reservoir system comprising a fraction of the API in the outer layer and a fraction of the API in the core, resulting in an immediate initial dose followed by sustained release. However, even more general platforms can be used, such as multi-unit devices with various inner cores with different release properties in hard gelatin capsules. Combinations of different coated cores even enable multi-pulse delivery. Another recently introduced controlled-release technology is SyncroDoseTM (Penwest Pharmaceuticals, USA), which consists of a medicated core coated with an erodible layer of a mixture of xanthan gum and locust bean gum. The delay time of these systems depends on the composition of the coating. In general, the coating step can be tricky. For example, obtaining a uniform coating of HPMC by double densification, which is critical for release performance, is a major challenge. Consequently, alternative methods of compression coating were evaluated, resulting in devices such as the EncoreTM system and the one-step dry-coated tablet (OSDRC1) (Ozeki et al., 2004; Gazzaniga et al., 1994). The polymer-plug capsule system consists of an insoluble shell surrounding an API core, a soluble cap, and a stopper composed of a hydrophilic polymer sheet that seals the open end. After the plug expands or corrodes, the active substance is released. Commercial examples include the Pulsincap™ system and the Egalet1 system, which consists of a cylinder containing an API core and erodible plugs at both ends (Bar-Shalom et al., 1991; Gazzaniga et al., 1994). In addition to sustained release systems, swellable polymers have many other applications for controlled release technology. For example, they are used in buoyant gastric retention delivery systems to retain low-density components, and in expandable systems, where they remain in the stomach longer due to their size. Scalable systems must be small enough to swallow. In the stomach, they expand rapidly by swelling, or unfolding, making them too large to pass through the pyloric sphincter. After they have released a certain amount of active substance, they must be emptied from the stomach, preferably completely digested. Expandable systems are easier to manufacture than deployable systems and should generally have a better chance of clinical implementation (Klausner et al., 2003).

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Most currently used hydrogel delivery systems contain polymers that do not degrade in the gastrointestinal tract. Nonetheless, enzymatically degradable hydrogels are becoming increasingly important, especially for site-specific delivery in the colon (Friend 2005; Lin and Metters 2006). Furthermore, the field of stimuli-sensitive hydrogels is also evolving. Their swelling or deswelling behavior and drug release depend on external factors such as pH, ionic strength or temperature (Qiu and Park 2001). In conclusion, hydrophilic polymers are used in a variety of controlled-release formulations ranging from site-specific devices for gastric retention or colonic delivery to time-dependent devices such as osmotic pumps. They are expected to be used in newly developed delivery systems as well as many new hydrophilic polymers with tailored properties in the future.

2.4

Osmotic pressure

Oral osmotic pumps have received increasing attention over the past three decades because they are largely unaffected by pH, the presence of food, and other physiological parameters. Their function is based on the osmotic pressure difference between the interior of the delivery system and the environment. Osmotic pumps have well-known properties that can be tailored to achieve zero-order release or structured release with a high degree of in vitro-in vivo correlation. They have been used in many therapeutic areas such as cardiovascular medicine, endocrinology, urology and central nervous system diseases. Their advantages are stable drug concentrations in plasma and less frequent dosing (Conley et al., 2006). Osmotically controlled delivery systems consist of an osmotically active drug-containing core surrounded by a semipermeable membrane. In an aqueous environment, water permeates the membrane and forms a saturated solution of drug and/or other osmotic agents with a high osmotic pressure. The solution can only be dispensed through predetermined dispensing openings in the membrane. In general, drug release depends on drug solubility, osmotic pressure gradient, release pore size, and membrane properties. For osmotic administration, ideally the drug should have moderate solubility (50–300 mg ml–1). Otherwise, solubility modifying excipients can be added to the active ingredient core. For example, highly water soluble drugs can be formulated with sodium chloride to reduce their solubility. Drugs with pH-dependent solubility can be co-formulated with acidic or basic ingredients. Interestingly, poorly water-soluble drugs can even be released as a suspension. In this case, dispersed drug is extruded due to swelling of co-formulated hydrophilic polymers or due to effervescent excipients such as sodium bicarbonate and citric acid. Another approach is to use poorly soluble drugs in the form of cyclodextrin complexes (Verma et al. 2000, 2002). However, if the osmotic pressure of the saturated liquid is insufficient, sugar, water-soluble salts or other osmotically active excipients can be added

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core. Alternatively, hydrophilic polymers may be incorporated into the core material. When water flows in, they expand, increasing hydrostatic pressure (Thombre et al., 2004). Another important consideration is the size of the dispensing port. In many cases, these openings are formed by mechanically or laser drilling holes in the membrane. In addition, osmotic pumps with controlled porosity have also been developed. In this case, the openings are due to the incorporation of leachable materials that dissolve when exposed to an aqueous environment (Zentner et al., 1985). For a membrane, two basic requirements must be met: first, it must be able to admit water, but must prevent the penetration of dissolved substances; second, the membrane must withstand the internal pressure. Commonly used materials include cellulose esters such as cellulose acetate, various types of Eudragit1, and ethyl cellulose. As the membrane thickness is typically 200-300mm, adequate water inflow may be problematic. This can be alleviated by adding hydrophilic flow enhancers such as polyethylene glycol (Verma et al. 2000, 2002). Alternatively, asymmetric membranes have been developed, consisting of a porous layer for increased mechanical strength and a thin semipermeable layer for diffusion control. They can be used to increase water influx, which may facilitate higher drug release rates for low solubility drugs (Herbig et al., 1995; Thombre et al., 1999). Drug release from osmotic pumps usually follows zero-order kinetics; H. Over time, a constant amount of active ingredient is released. Once the osmolarity gradient is established, this delivery regimen begins over time and continues until all osmotically active solids are dissolved (Verma et al., 2000). In addition, more complex systems have been developed allowing rise, delay or pulse release. Among the currently available osmotic delivery system designs, the most prominent example is the elementary osmotic pump (EOP), a single-chamber device primarily used for moderately water-soluble drugs. The osmotic technology OROS1 (Alza Corporation, USA) has been used in many commercially available products such as the phenylpropanolamine formulation Acutrim1 (Ciba-Geigy) (Shokri et al., 2008). To apply EOP technology to poorly water-soluble drugs, swellable EOPs containing drugs dispersed in hydrophilic polymers and osmotic agents were developed. When water flows in, internal hydrostatic pressure increases due to polymer expansion and osmosis. This releases the active ingredient in dispersed form in the gel (Shokri et al., 2008). In addition to single-chamber devices, many multi-chamber systems have also been developed. Osmotic push-pull pumps consist of a double-layer core surrounded by a semipermeable membrane with small openings. The top layer near the opening contains the drug and the bottom layer contains the osmotic polymer. Water can cause swelling in the lower body. The increased internal pressure ("push") leads to extrusion ("pull") of the active substance. Similarly, the Sandwich Osmotic Tablet System (SOTS) consists of a three-layer core with a push layer in the middle and two API layers attached (Liu et al. 2000). In addition to the push-pull system

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Various osmotic pumps have been created with non-expandable second chambers as the second chamber expands. These include means for diluting the API solution through a second chamber. This may be beneficial for some drugs when concentrated solutions can irritate the gastrointestinal tract (Verma et al. 2000, 2002). Recent developments include extended-release multiparticulate systems consisting of various API-containing pellets coated with a semipermeable membrane, and L-OROS1 (Alza Corporation) and OROS-CT1 ( Alza). Corporation) for site-specific delivery to the colon (Verma et al. 2000, 2002). Different release profiles can be achieved by combining osmotic systems with other principles of controlled drug release. For example, a drug-containing coating may additionally be applied to provide an initial dose. Another approach is based on sustained burst release of the drug from capsules coated with a semipermeable layer filled with an API and a swellable polymer. After a delay time which can be set via the coating, the internal pressure increases sharply and the capsule "explodes". Among the many other pulsatile release approaches, an example is the multiparticulate system, which combines osmotically active particles with different coatings that release their drug load sequentially (Anal 2007). Examples of marketed products based on osmotic delivery are the nifedipine formulation Procardia1 XL (Pfizer), a calcium channel blocker used in the treatment of angina, and Covera-HS1, a verapamil formulation. Glucotrol1 XL (Pfizer) is a formulation of glipizide used in the treatment of type 2 diabetes. Also, doxazosin preparation Cardura1 XL (Pfizer) for benign prostatic hyperplasia (BPH), oxybutynin preparation Lyrinel1 XL (Janssen-Cilag) for overactive bladder, methylphenidate preparation Concerta1 (Janssen-Cilag) for attention deficit hyperactivity disorder (ADHD), Invega1 is a formulation of paliperidone for schizophrenia, and Jurnista1 (Janssen-Cilag) is a formulation of hydromorphone for pain ( Conley et al., 2006). In conclusion, osmotic pumps are an extremely versatile platform for controlled release profiles. Over the years, the range of drug candidates has expanded from osmotically active drugs with moderate water solubility to liquid drugs with good and poor water solubility, and many new applications can be expected in the future.

2.5

density

The density of the dosage form, among other physical properties, can be used for site-specific drug release. It is used in gastric retention delivery systems that remain in the stomach for a longer period of time. Thus, bioavailability and therapeutic effect of multiple substances with local action or absorption are achieved

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If they are sufficiently stable at acidic pH, they can improve metabolism in the stomach and proximal small intestine. These include drugs with a small absorption window, such as levodopa or furosemide, but also drugs with limited stability in the distal parts of the gut. Delayed release in the stomach results in less fluctuating plasma levels, for example in the treatment of Parkinson's disease with levodopa (Crevoisier et al., 1987). Another area of application is topical administration to treat gastric ulcers, such as misoprostol or antibiotics against Helicobacter pylori (Singh and Kim 2000). The residence time of a pharmaceutical dosage form in the stomach depends on a number of physiological factors such as B. feeding status, quality and quantity of food consumed, age and health status. As it also depends largely on the size and density of the preparation, buoyancy and expansion systems and mucoadhesive systems are the most promising approaches to increase gastric residence time. Floating devices have a lower overall density compared to gastric contents (1.004-1.010 g cm-3) and thus are buoyant in the stomach, while inflating systems expand or unfold after swallowing, making them too large to pass through the pylorus Sphincter (Klausner et al. 2003). On the other hand, the mucoadhesive system adheres to the gastric mucosa. While longer gastric residence times can be achieved with such systems, there is a risk of esophageal adhesions, which can lead to drug-related injuries (Talukder and Fassihi 2004; Streubel et al 2006). Another method of gastric retention is a high-density system that owes its weight to heavy inert excipients such as barium sulfate, zinc oxide, titanium dioxide, or iron powder (Singh and Kim 2000). They are thought to sink into the fundus of the stomach and become trapped in the folds of the antrum. However, high-density systems have not been shown to be effective in vivo (Bardonnet et al., 2006). Additionally, magnetic systems can be used. An external magnet on the stomach holds the dosage form in the desired position, which also contains a small magnet. The main disadvantage of this method is the difficulty in accurately positioning the external magnets (Bardonnet et al. 2006). Floating drug delivery systems (FDDS) and low density systems can be single or multiple dosage forms. Based on their mechanism of action, they are divided into effervescent and non-effervescent agents. Effervescent FDDS become buoyant after a certain lag time after ingestion due to gas release. In most cases, carbonate or bicarbonate is used. They release carbon dioxide when they come into contact with stomach contents. Various designs have been developed. Swellable hydrophilic polymers such as methylcellulose or chitosan are used in combination with carbonate/bicarbonate and tartaric/citric acid to produce matrix systems. When in contact with water, the polymer swells, producing and retaining carbon dioxide. After a delay, the device begins to suspend and release its drug load (Arora et al. 2005). Another approach is to coat the active ingredient core with an inner layer of effervescent dal. 1991). A similar principle applies to floating capsules containing a large number of particles with varying residence times in the stomach

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according to their coating. Increasing the thickness of the swellable membrane layer leads to better swimming performance on the one hand and longer lag times and reduced release rates on the other (Arora et al., 2005). In addition, various multilayer tablets have been developed, such as an expandable asymmetric three-layer tablet containing a separate gas-producing layer (Arora et al., 2005; Yang et al., 1999). For controlled zero-order release in the stomach, multilayer tablets consist of two barrier layers and a drug layer. Thus, barrier erosion is associated with increased API exposure. Another interesting approach is to combine flotation and expansion by incorporating rapidly expandable polymers into effervescent materials (Arora et al., 2005). A special case is that of raft-forming systems that include hydrophilic polymers that, due to trapped CO2 bubbles, create a viscous gel layer that floats on gastric fluid (Bardonnet et al., 2006). As indicated above, most designs are based on combinations of effervescent substances with swellable polymers. However, ion exchange resins loaded with bicarbonate and surrounded by a semipermeable membrane can also be used (Arora et al., 2005). Another method of gas generation was described in the 1970s. Volatile liquids such as cyclopentene or diethyl ether are sealed in specific chambers where they vaporize at body temperature (Arora et al. 2005; Singh and Kim 2000). However, most effervescent systems are based on carbonates or bicarbonates. Non-foaming FDDS are buoyant due to entrapped air or fatty excipients, and the hydrodynamic balance system (HBS) is one of the most important methods. They contain swellable hydrophilic polymers that trap air bubbles, resulting in an overall density that is less than that of the stomach contents. The flotation of these devices depends on the type of polymer and the presence of air pockets in the matrix. The mixture of active ingredient and swellable polymer can be encapsulated or compressed into tablets. Another prominent approach is the production of hollow particles with active substance-containing shells, so-called microbubbles. They are prepared by solvent evaporation techniques, for example using B. polycarbonate, various Eudragit1 types, polypropylene or poly(methyl methacrylate). Other types of buoyancy beads include alginate beads made by precipitating alginate in a calcium chloride solution or beads made by spray drying. In addition, polypropylene foam powders have been proposed as low-density additives for single-component and multiparticulate systems (Streubel et al., 2006). Other inventive designs include sheet-like multilayer devices with drug-containing layers and barrier membranes sealed together to trap small air bubbles, and combinations of technologies such as flotation devices coated with mucoadhesive polymers (Arora et al., 2005 ; Singh and Kim 2000). Commercially available products based on FDDS include floating capsules such as Madopar1 HBS (Roche), an extended-release formulation of levodopa and benserazide for the treatment of Parkinson's disease, and the diazepam formulation Valrelease1 (Roche). Another example is Cytotec1 (Pfizer), a bilayer floating capsule for the topical delivery of misoprostol for the treatment of gastric ulcers. In addition, raft formation systems are commercially available, such as the antacid Gaviscon1

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(GlaxoSmithKline) and Topalkan1 (Pierre Fabre), for the treatment of gastroesophageal reflux (Arora et al., 2005; Singh and Kim, 2000). However, some limitations of currently available technologies still exist. One disadvantage of single unit systems is that they leave the stomach prematurely. Their "all or nothing" urination process can lead to unreliable and non-reproducible therapeutic effects. Therefore, multi-unit systems are preferable (Singh and Kim 2000; Bardonnet et al. 2006). Due to its mechanism, non-effervescent systems with immediate buoyancy are considered safer than effervescent systems, which only buoy after a certain delay and involve an increased risk of premature gastric emptying (Streubel et al., 2006). In general, effective flotation depends on having enough fluid in the stomach. Therefore, patients need to drink water frequently, which may affect patient compliance (Hwang et al. 1998). Gastric retention technology has yet to prove its added value when one also considers that traditional non-disintegrating controlled-release tablets may have significant gastric residence time (Waterman 2007). Nevertheless, promising results were obtained both in vitro and in vivo. Extensive research efforts are still underway as many drug treatments would benefit from reliably prolonging gastric retention. Therefore, more successful floating drug delivery systems can be expected in the future.

2.6

Enzyme-mediated release in the colon

Enzyme-mediated release of drugs from their carriers is a widely studied research topic in the field of drug delivery. Much interest in this oral administration-triggered release mechanism stems from the different enzyme patterns and activities that occur during gastrointestinal transit (Table 1). The colon is a particularly interesting site because of its high bacterial population and associated enzyme-rich environment dominated by (oxido)reductase (nitroreductase, hydrogenase, azoreductase, nitric oxide reductase, sulfoxide reductase, etc.) and hydrolases (b-glucuronidase, glycosidase, esterase, amidase, sulfatase, etc.) (Bourgeois et al. 2005; Scheline 1973). Since the presence and high activity of these enzymes is characteristic of the large intestine, these facts can be used for local release of active ingredients with simultaneous low absorption in the upper gastrointestinal tract. In practice, this is of interest for the treatment of chronic disease states such as Crohn's disease involving the colon, ulcerative colitis (Friend 2005) or colon cancer, but also for oral vaccinations and administration of proteins and peptides significance. Although it is not entirely clear to date whether the colon can serve as a site for adequate protein absorption to produce systemic effects, it has a relative lack of protein-degrading enzymes, a relatively long residence time of digesta, and a highly reactive enhancer of absorption in this area. There are continuing efforts in the field (Rubinstein et al. 1997; Mackay et al. 1997). In general, enzymatically degradable drug delivery systems for colon-specific drug delivery are based on (polymer) prodrugs or polymers used as drug core coating materials, as well as drugs embedded in matrices or hydrogels. Embedding media in gel systems (Sinha and Kumria 2003). ).

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Prodrug

Primarily in the context of topical treatment of inflammatory bowel disease (IBD), efforts have been made to use prodrugs to deliver therapeutic agents to the colon. Conjugation of active substances with hydrophilic or high molecular weight molecules is said to reduce premature absorption and increase the availability of active substances in the large intestine (Friend 2005). To be able to formulate prodrugs, drug molecules must have chemically active hydroxyl, amino, nitro or carboxyl groups. Using these functionalities, prodrugs mainly with azo and glycosidic linkages were synthesized for colon-specific Sexual delivery (Minami et al., 1998). Glucuronic acid (Simpkins et al., 1988), dextran (Larsen et al., 1991), as well as sulfonamides and p-aminobenzoic acid (Bourgeois et al., 2005). It must be emphasized that careful selection of carrier molecules is an essential development step. This is clearly demonstrated in the case of sulfasalazine, used to treat Crohn's disease and ulcerative colitis, which consists of sulfapyridine (carrier) azo-linked to mesalazine (active ingredient). It was found that during the enzyme-mediated cleavage of the azo bond, high concentrations of sulfapyridine were absorbed, subsequently causing side effects (Peppercorn 1984). This led to the development of mesalamine prodrugs containing safer carrier molecules, such as ipsalazine and balsalazide or olsalazine consisting of two mesalazine molecules, as the most elegant examples (Chan et al., 1983 Year). Other approaches include linking mesalamine to higher molecular weight carriers such as polysulfonylvinyl (polyasa1) or polyamidoamine dendrimers (Wiwattanapatapee et al., 2003), so that the drug releases in the colon and the carrier follows the Fecal excretion. In general, when the prodrug is formed with a high molecular weight polymer, the potential for absorption of the conjugate in the upper GI tract decreases dramatically, resulting in more efficient delivery in the colon. However, care must be taken when conjugating lipophilic APIs to hydrophilic polymer chains, as sufficient water solubility of prodrugs is a key prerequisite for their enzymatic degradability. Another approach to the treatment of IBD relies on the β-glycosidic derivatization of nonpolar anti-inflammatory corticosteroids such as dexamethasone, prednisolone, or hydrocortisone with carbohydrates, resulting in hydrophilic prodrugs (Friend and Chang 1984). In addition to delivery to specific sites of inflamed tissue, the resulting change in the polarity of the compound reduces absorption in the small intestine, thereby reducing systemic side effects and therapeutic dose.

2.6.2

Polymer Coatings and Substrates

Encapsulation of solid dosage forms with environmentally responsive coatings is an essential technique for the development of highly effective oral dosage forms. Enzymatic polymers represent promising coating and matrix materials when preferential drug release in the colon is required.

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The suitability of several classes of polymers and polysaccharides for this purpose has been extensively studied (Friend 2005). For successful introduction into the large intestine, the coating material should have good film-forming properties, low permeability, and poor water solubility. However, for effective enzymatic degradation, some degree of swelling must be present. The first biodegradable polymers for colonic drug delivery were azopolymers, usually composed of hydrophilic, hydrophobic and azo moieties (Bourgeois et al., 2005). By varying the ratios of the three components, polymers with improved degrading (more hydrophilic) and protective (more hydrophobic) properties can be developed (Van den Mooter et al., 1992, 1993). However, the disadvantage of this approach is that the reduced azoaromatic linker can be toxic and the coating material degrades rather slowly in the colon. Alternatively, the use of (naturally occurring) polysaccharides may be preferable due to their biodegradability, availability and cheapness, as well as colon-resident bacteria (Havgaard and Bronsted 1996). To compensate for the poor film-forming properties and often high water solubility of these substances, modifications by changing the hydrophilicity of the main chain or blending with synthetic polymers have been investigated (Chourasia and Jain 2004). The most commonly used polysaccharides include amylose (Milojevic et al., 1996), guar gum (Tugcu-Demiroz et al., 2004), inulin (Vervoort and Kinget, 1996), and chondroitin sulfate. For the latter mucopolysaccharide, the degree of crosslinking with dicyclohexylcarbodiimide can control the drug release profile of the formulation (Rubinstein et al., 1992). The polycationic polysaccharide chitosan has also been widely used to develop a colonic release form, but exhibits rather high water solubility at gastric pH (Hejazi and Amiji 2003). To overcome this problem and improve the integrity of chitosan-coated nuclei in the upper gastrointestinal tract, acid-resistant coatings using Eudragit L-100R1 or L-100S1 (see the Haag chapter of this book) have been proposed (Lorenzo-Lamosa et al. , 1998). Chitosan has also been used to formulate beads for multiparticulate hydrogel systems via polyelectrolyte complexation with tripolyphosphate. It was found that the degradation of the hydrogel matrix depends on the molecular weight and degree of deacetylation of chitosan (Zhang et al., 2002). For another commonly used polysaccharide gum, its high solubility in the upper gastrointestinal tract also limits its use as a coating material, however, thick compressed layers or combinations with ethylcellulose have been shown to significantly reduce the use of Premature drug release from pectin-coated delivery systems (Ahmed 2005). Interested readers are referred to the comprehensive collection of studies involving pectin-containing formulations triggered by microorganisms (Bourgeois et al., 2005). Another potentially more advanced colon-targeted delivery system (CODES™) contains the synthetic disaccharide lactulose. The system consists of a core containing lactulose and active substance, an inner acid soluble coating (Eudragit1 E) and an outer enteric coating (Eudragit1 L). The latter dissolves in the small intestine, causing the intestinal fluid to diffuse into the nucleus, where lactulose dissolves. Lactulose diffuses from the nucleus to the lumen of the colon where it resides

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It is then broken down by bacterial enzymes. This results in a local acidic environment, which in turn leads to dissolution of the acid-soluble layer and release of the drug (Katsuma et al., 2002). Given the success of delivering locally acting therapeutic agents to the colon via enzymatic release formulations, exploiting this site-specific mechanism of oral delivery of biotechnological drugs remains an interesting prospect for future research.

2.7

biological detection

A rather sophisticated approach to improving oral drug delivery is the use of biometric mechanisms. They can be used for bioadhesion to achieve longer residence times in the gastrointestinal tract, or for highly specific delivery to specific tissues. Among bioadhesives, a widely studied method is the mucoadhesion of polymers such as chitosan, polyacrylates, and polythiolated polymers, which is caused by the nonspecific interpenetration of polymer chains and mucus. Sincerely. Unfortunately, these classical mucoadhesives have not proven successful for oral administration in humans (Gabor et al. 2004; Varum et al. 2008). A more site-specific approach is based on cell adhesion. To this end, drug delivery systems are functionalized with ligands that bind specific structures on the surface of specific cell types. Typically, this is accomplished by covalently coupling the targeting device to the surface of the delivery vehicle. Once bound to the target site, the drug is released. If the target is a receptor that can mediate endocytosis, binding can even lead to cell invasion—provided the delivery system is the right size. However, a prerequisite is the identification of suitable targets and then suitable ligands that can be surface-coupled to the delivery system and resist premature degradation in the gastrointestinal tract. Among the many different approaches to targeting the small intestine, sugar targeting using lectins is one of the most popular. Lectins are proteins that specifically bind specific carbohydrate moieties (Sharon and Lis 2004). In the gut, sugars are present in the mucus and in the glycocalyx of enterocytes and are part of the "sugar code", an important biorecognition mechanism (Gabius 2000). It is known from the literature that the sugar composition of the glycocalyx is a tissue-specific feature that can be used for targeting (Gabor et al., 2004). Targeting enterocytes has attracted widespread interest in recent years, especially for enhancing the absorption of nanoparticles for the purpose of delivering drugs that cannot be orally administered due to poor solubility, absorption, or stability in the gastrointestinal tract. In addition, targeting M cells to improve oral vaccine delivery is being studied in detail. Their main function is to deliver antigens from the lumen to the GALT. In addition to targeted attacks on healthy intestinal epithelial cells, there are methods such as targeting Helicobacter pylori with antibiotics and administering anticancer drugs to colon cancer cells.

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381

targeting intestinal cells

Glycosylation patterns in enterocytes were assessed based on studies of the binding properties of various fluorescein-labeled lectins with different carbohydrate specificities. N-acetylglucosamine-binding lectins, such as wheat germ agglutinin (WGA) from wheat and tomato agglutinin (LEA) from tomato, bind extensively to intestinal epithelial cells. WGA interacts with the glycosylated extracellular domain of the epidermal growth factor (EGF) receptor, leading to receptor-mediated endocytosis (Lochner et al., 2003). The intracellular localization of internalized WGA was studied using Caco-2 cells, a model of human intestinal epithelial cells. Most internalized lectins are found in lysosomes (Wirth et al. 2002). Caco-2 monolayer experiments also showed some degree of transcytosis of WGA. For WGA functionalized nanoparticles, cell adhesion and cell invasion were also observed using Caco-2 cells (Weissenboeck et al., 2004; Fillafer et al., 2008). Therefore, WGA may be a promising candidate for enhanced oral drug delivery. It not only mediates mucoadhesion and cell adhesion of enterocytes, but also mediates cell invasion and even partial transcytosis. Also, it is not broken down by gastrointestinal enzymes. While some plant lectins are highly toxic, WGA is a dietary lectin that is also found in wheat products. No toxic effects are expected at the concentrations required for sugar targeting, but this remains to be confirmed by in vivo studies. Another issue that requires further elucidation is the risk of eliciting an immune response to WGA during its use for drug delivery (Gabor et al. 2004). In addition to lectin-mediated targeting, other approaches include targeting vitamin B12 receptors for endocytosis (Russell-Jones et al., 1999), peptide transporters (Walter et al., 1996), and bile acid-mediated translocation (Swaan et al., 1996). Targeting of the transferrin receptor has also been investigated, but a major drawback of this strategy is that this receptor is mainly expressed on the basolateral side of enterocytes.

2.7.2

Targets M cells

M cells are specialized epithelial cells that line lymph nodes and mesenteric lymph nodes (Miller et al. 2007). They are able to take up antigens from the intestinal lumen and deliver them to the part of the immune system connected to the gut. Therefore, they are prime targets for improving the effectiveness of oral vaccinations. Oral vaccines have better adherence than parenteral vaccines, are less expensive, have no risk of contracting blood-borne diseases, and can generate both mucosal and systemic immunity (Aziz et al., 2007). Although the overall experience with marketed products has been positive, few oral vaccines are available (Foxwell et al., 2007). To date, most oral vaccines are based on live attenuated organisms (O'Hagan et al., 2006). most famous

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One example is the oral poliovirus vaccine (OPV), which was introduced in the 1960s and successfully helped reduce polio cases worldwide. Nevertheless, live attenuated poliovirus still carries the risk of regaining the neurovirulence and transmissibility properties of wild poliovirus and should therefore - indeed in Europe - be replaced by inactivated poliovirus vaccine Complete eradication of polio (Bonnet and Dutta). 2008). Other examples of available oral vaccines are live attenuated cholera, typhoid and rotavirus vaccines. In addition to antimicrobial vaccination, oral immunotherapy against type I allergy is an important goal. About a quarter of the population in industrialized countries suffers from allergies. The only disease-causing treatment in this area is allergen-specific immunotherapy, which aims to shift the antibody response from a Th2 to a Th1 immune response, thereby alleviating allergic symptoms. To date, it is only available by subcutaneous and sublingual routes (Roth-Walter and Jensen-Jarolim 2007). For intestinal mucosal immunity, antigens must pass through the intestinal epithelium and enter the gut-associated lymphoid tissue (GALT). This transcytosis occurs primarily on M cells. Their main function is to uptake and transport antigens from the intestinal lumen to GALT, thereby enabling the immune system to monitor the gut contents. M cells are well suited for this purpose: their apical membranes have reduced glycocalyx and reduced membrane hydrolase activity. In addition, the number of lysosomes is drastically reduced, allowing intact antigen trafficking to GALT. In lymphoid follicles, antigens are captured by antigen-presenting cells, which triggers an immune response that ultimately leads to antibody production. Mucosal immunity is primarily dependent on secretory immunoglobulin A (IgA), which captures antigens at the mucosa and neutralizes viruses and endotoxins without causing tissue damage (Brandtzaeg 2007). The intestinal mucosa contains at least 80% of the body's activated B cells, approximately 90% of which produce IgA (Brandtzaeg et al., 1989). Therefore, mucosal vaccines are of particular interest against pathogens that cause mucosal infection or enter the mucosa (Levine 2003). In addition to mucosal immunity, systemic immunity can also be induced by persistent serum IgG and IgA responses. Protection ranges from symptom relief to complete suppression of infection. More recently, antimicrobial vaccine research has focused on subunit and DNA vaccines (Peek et al., 2008). Subunit vaccines use only part of the pathogen, such as a specific protein, as an antigen. They have the advantage of not being able to revert to toxic forms and being free of pathogen-derived contaminants. However, oral DNA and subunit vaccines and allergens must be protected from low gastric pH and digestive enzymes. A number of vaccine delivery systems have been developed in recent years. These include live bacterial and viral vectors, virus-like particles, non-living delivery systems and plant-based genetically modified "edible vaccines". In nonliving systems, liposomes, proteosomes, and a range of polymeric particles have been studied (O'Hagan et al. 2006). made of particles

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The biocompatible and biodegradable polymer poly(D,L-lactide-co-glycolide) is the most widely studied nonviral delivery platform (Brayden and Baird 2001; Perrie et al. 2007). Particle uptake in M cells is influenced by their size and surface properties. Although it is generally accepted that particles smaller than 1 mm are efficiently transencapsulated by M cells, there is some controversy regarding the maximum particle size that can be accommodated. Some authors have suggested that the optimal size for transcytosis may be up to 2, 3 or even 5 mm (Roth-Walter and JensenJarolim 2007; O'Hagan et al. 2006; Brayden 2001). Furthermore, hydrophobic particles are more readily taken up by M cells than hydrophilic particles, and negatively or neutrally charged particles are better than positively charged particles (des Rieux et al. 2006). Uptake of antigen from non-targeting systems is often insufficient to elicit mucosal immunity, whereas systems attached to M cells can efficiently transcyst. A major challenge is identifying M cell-specific targets. One of the major drawbacks is the high M cell variability between species and individuals, and between different physiological states and ages (des Rieux et al. 2006; Brayden and Baird 2001). Another reason is the lack of suitable in vivo and in vitro test models (Foxwell et al., 2007). Various approaches have been investigated so far, with more or less promising results. The two main strategies are the use of lectins for sugar targeting and mimicking known pathogen-cell interactions (Brayden et al., 2005). The lectin most commonly used for M cell targeting is Ulex europaeus lectin 1 (UEA-1), which specifically binds a-L-fucose (Foster and Hirst 2005). Aleuria aurantia lectin (AAL) is an alternative (Roth-Walter et al., 2004, 2005). UEA-1 and AAL binding were examined in mice. In humans, however, α-L-fucose is absent on the apical surface of M cells. Another major strategy follows the example of pathogens using M cells as entry portals. Bacteria such as Yersinia, Salmonella and Shigella interact with M cells through adhesins. A Yersinia adhesin called invasin has been shown to interact with the b1 integrin on the cell surface (Clark et al., 1998). Unlike other polarized epithelial cells, b1 integrin in M cells is not restricted to the basolateral side but is also expressed apically. This has also been used to target nanoparticles with the RGD peptide motif (des Rieux et al., 2006). Another adhesin that has been studied for M cell targeting is reovirus adhesin s1 (Miller et al., 2007). In addition, the ganglioside GM1 (the B subunit of the cholera toxin receptor) is also thought to enhance vaccine delivery. Although it is also present on enterocytes, the anatomical features of the M cell make the receptor more accessible and thus may appear more attractive. Furthermore, IgA-mediated interactions with M cells may be another approach (des Rieux et al. 2006). Although fundamental challenges remain in successfully targeting M cells, some encouraging results hold promise for future improvements in the induction of immunity against various pathogens and oral immunotherapy against type I allergy.

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other

Targeted therapy has also been proposed to eradicate H. pylori from the stomach. This pathogen, which lives deep in the mucus of the stomach, can cause ailments such as indigestion, gastritis and ulcers. It was also associated with a significantly increased risk of gastric cancer and MALT lymphoma. Targeted systems can be used to deliver antibiotics more efficiently, in addition to other gastric retention dosage forms. One approach uses sugar targeting. Interestingly, in this case sugar was the target and lectins were the target. Bacterial adhesin BabA2 binds to fucosylated tissue blood group antigen Lewis b on the surface of gastric cells. Thus, fucose-functionalized nanoparticles preferentially attach to bacteria. Unfortunately, not all H. pylori strains express BabA2 (Bardonnet et al., 2006). In addition, cancer-specific cell surface structures can be used for targeted delivery. For example, the β-galactoside-specific lectins galectin-1 and -3 are upregulated in many epithelial tumors such as colon cancer and can be used as targets (Minko 2004).

2.8

absorption enhancer

2.8.1

general considerations

Today, it is well known that several factors can negatively affect the bioavailability of drugs, including insufficient dissolution, rapid degradation of compounds in gastric or intestinal fluids, poor membrane permeability, and extensive first-pass metabolism. Uptake by intestinal epithelial cells has been of particular interest and represents a critical step in the chain of events involved in drug absorption and its transport to the site of action. When examining the physicochemical properties of specific compounds, various properties are associated with reduced membrane permeability, such as low octanol/water distribution, presence of charged functional groups, large polar surface area, or high molecular weight (see Section 1). 1.1). Much work has been done to improve mucosal transport, which can be broadly divided into strategies that tend to alter the physicochemical properties of the drug itself and those that tend to alter membrane permeability. Since the principle of chemical modification of the active substance itself—that is, the development of derivatives or prodrugs with the same effect—is often difficult or impossible and must be done individually for each specific active substance, membrane thinning strategies will is a great advantage and can be done without applying more changes to different fabrics. When developing formulations based on the simultaneous administration of one or more absorption-enhancing excipients, several key aspects must be considered. First, clarify whether the improvement in oral bioavailability is achieved by

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Enhancers are possible, so it is important to accurately identify the main limiting factors in the pharmacokinetic profile of a drug. Particular attention should be paid to a comprehensive understanding of the uptake mechanism, as the principles of uptake enhancement are not limited to affecting membrane permeability but can include other strategies such as affecting secretory trafficking. After a thorough evaluation of the drug's kinetic behavior in humans, the choice of absorption enhancer should be based on a detailed knowledge of its mechanism of action (Yeh et al. 1995). Unfortunately, penetration-enhancing agents have not yet been able to reach their full potential due to common problems with most substances regardless of their chemical origin. Several difficulties to be overcome are inextricably linked to the principle of co-management. To effectively enhance absorption, drugs and excipients must be present in sufficient concentrations at the site of uptake, while simultaneously being delivered to the intended target area. Therefore, it is clearly important to detect differences in gut distribution. If necessary, release can be improved through the use of additional delivery strategies, such as enteric coating (Aungst 2000). The biggest problem common to all types of membrane penetration enhancers is the ever-present threat of toxic side effects. The risk of severely damaging the mucosa by attempting to manipulate its permeability is inherent in this approach and has sparked debate about the usefulness of absorption enhancers. It's difficult to determine the extent of the damage done to the membrane by treatment with permeation enhancers—it may be able to recover—but from today's perspective, many of the compounds studied so far appear to be at least as harmful as they are helpful. However, some promising results have been published on different chemical classes of excipients that appear to enhance drug absorption without affecting cell viability, even when used at concentrations effective to enhance uptake (Hastewell et al. 1994 2008; Whitehead et al. 2008). Steep concentration-response curves and relatively low safety margins have been demonstrated in vitro for most enhancers studied to date. In many cases, there was no distinction between effective and membrane-damaging concentrations (Whitehead et al., 2008). It is worth mentioning that most enhancers have a higher cytotoxic potential in cell culture models than in intact intestinal membranes, attributing to the presence of repair mechanisms that allow some degree of recovery after trauma (Aungst 2000 ). However, the curative potential is limited and often insufficient to sufficiently extend the therapeutic concentration window to safe levels.

2.8.2

Strategies to Improve Uptake

In order to have a positive impact on bioavailability through the use of absorption enhancers, a thorough understanding of the pharmacokinetics of the drug and an understanding of the mechanism of action of the penetration enhancers are required. Unfortunately until now

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Not all principles governing enhancer behavior are fully understood. In addition to the concept of directly affecting mucosal physiology, other aspects related to drug absorption should also be considered, such as the possibility of improving intestinal dissolution of poorly soluble drugs by solubilizing excipients. This feature is known to be the main reason or at least a supporting factor for the absorption-enhancing potential of many absorption enhancers (Stegemann et al., 2007). Looking closely at intestinal epithelial cells, enhancers can primarily act as promoters of transcellular or paracellular pathways, or interfere with active secretory trafficking. In general, with regard to membrane-specific effects, one should keep in mind that, similar to how APIs may have preferential uptake sites throughout the GI pathway, responses to specific enhancers may vary. Therefore, the effect of enhancing absorption can also be regioselective and must match the distribution curve of the active substance (Aungst 2000).

Solubilizing excipients Today's trend towards higher molecular weight or increased lipophilicity for new chemicals often presents problems with limited solubility in the gastrointestinal tract and thus may be related to limited oral bioavailability (Stegemann et al., 2007). Especially for BCS class II substances (poorly soluble but permeable, see section 2), an increase in dissolution rate is expected to lead directly to increased absorption after oral administration. There are many ways to influence the solubility of active substances, including modification of the molecule itself, for example by conversion into various salts (see section 1.1) and simultaneous administration of solubilizing excipients. The field of solution enhancement, in addition to utilizing various traditional chemical principles (Strrickley 2004), has had a major impact in recent years through the application of newly developed technologies such as cyclodextrins or self-emulsifying drug delivery systems, and is far from being fully explored. Cyclodextrins are cyclic oligomers of glucose that can be divided into different groups according to their molecular weight and chemical properties. In oral formulations, β-cyclodextrin and various newly developed derivatives are of great importance (Stegemann et al., 2007). By forming clathrates with lipophilic molecules, cyclodextrins are able to temporarily alter the physicochemical properties of substances and significantly increase the solubility and stability of compounds (Davis and Brewster 2004). Another advantage of cyclodextrins is that they are inert to biofilms in living tissues, ie they increase the solubility and absorption of the active substances they contain, but they are not absorbed themselves (Gould and Scott 2005). Lipid-based drug delivery systems have received increasing attention in recent years not only because of the postulated ability to utilize the lymphatic absorption pathway, which involves direct uptake in the lymphatic system, but also because

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Ability to formulate self-emulsifying drug delivery systems. Such systems are based on mixtures of several lipophilic, amphiphilic or hydrophilic excipients that together form thermodynamically stable microemulsions of nanoscale droplets when placed in an aqueous environment (Porter et al., 2008) . Medium-chain glycerides have been shown to be suitable candidates for the development of such self-emulsifying systems (Constantinides et al., 1996) and, in addition to acting as solubilizers, are said to facilitate transcellular uptake (Muranushi et al., 1981) .Furthermore, some excipients included in self-emulsifying systems may inhibit cytochrome P enzyme and P-glycoprotein mediated transport from the blood to the lumen, such as Tween 80 or Cremophor RH40 (Mountfield et al. 2000; Hugger et al. 2002). So far, several formulations utilizing the principle of self-emulsification have been successfully marketed (eg Sandimmun Neoral1, Norvir1).

Transcellular pathways Most compounds known to enhance absorption are said to have a direct effect on cell membranes, thereby facilitating transport via transcellular pathways. This group includes surfactants, medium chain triglycerides, fatty acids, steroidal detergents, acylcarnitines or alkanoylcholines, alpha-amino acids, etc. (Aungst 2000). The exact mechanism that makes the cell membrane more permeable remains unknown but may involve several different factors such as increased bilayer fluidity, dissolution of membrane components, and pore-forming effects. To date, most analytical studies have been limited to detecting the intracellular uptake of labeled molecules or the release of cellular or membrane components into the surrounding medium, both indicators of increased membrane permeability, without providing further insight into the rationale. information. Of course, although the effects on membrane integrity using several enhancers appear to be at least partially reversible, this approach is strongly associated with potential toxic side effects.

Paracellular pathways Compared to transcellular pathways, enhanced transport via the paracellular pathway requires at least temporary opening of interconnected tight junctions without compromising overall membrane integrity and thus may be less detrimental to cell viability. The maximum dimension of the paracellular space is 3–5 nm, which may limit such channels to solutes with molecular radii no larger than 1.5 nm (~3.5 kDa) (Salama et al. 2006). For penetration enhancers, the triggers leading to the opening of tight junctions have not been clearly identified, but based on the fact that the zonules of closure are tightly bound to the underlying cytoskeleton, Ca2+-dependent contraction of actin fibers surrounding the junctions was discussed as a possible mechanism (Fasano and Uzzau 1997). Additional breakdown of extracellular Ca2+ by chelating agents or other lysing agents

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The effect can make the closure band more permeable. In this regard, an interesting approach describes the use of substances that mimic endogenous tight junction modulators and proposes these molecules as promising candidates for potentially less harmful reversible enhancement of uptake. Zonula occludens toxins and their biologically active fragments have been successfully used for this purpose (Fasano and Uzzau 1997).

Secretory transport inhibitors Although certain drugs are lipophilic and thus not inherently limited in their ability to transmembrane, their oral bioavailability may be reduced due to active transport from enterocytes to the lumen of the intestine. Any inhibition of these transport systems, including, for example, P-glycoprotein and other members of the multidrug resistance-associated protein family, would lead to an increase in net adsorption permeation (Aungst 2000). However, care should be taken to avoid systemic side effects when applying this principle, since these transporters are not exclusively present in the intestinal mucosa. In addition to the proven inhibitors of these transporters such as cyclosporine (Terwogt et al., 1998), the activity of several long-used pharmaceutical excipients such as Tween 80 and Cremophor EL has now also been demonstrated. Secretory transport is affected (Nerurkar et al., 1996; Porter et al., 2008; Hugger et al., 2002). 2.8.3

appearance

For the critical evaluation of absorption enhancers, it is important to mention that the development and testing of penetration enhancers is usually performed using in vitro models (e.g. Caco-2 cells), although the in vivo performance of the same enhancers has been shown to be important in many situations Down is different. Often, results obtained in cell culture or chamber assays cannot be successfully replicated in vivo, and it is unclear whether in vitro cytotoxicity assessments are predictive of in vivo situations (Aungst 2000). The comparative evaluation of absorption enhancers is also made more difficult up to now by the large number of different test methods which have yielded very different results. For solubilizing excipients in particular, novel test models that take into account, for example, the role of endogenous emulsifiers in digestive fluids are urgently needed, which could further our understanding of increased absorption (Porter et al., 2008). Whitehead et al. A recently published study of Caco-2 cell monolayers, including 153 known penetration enhancers, aimed to assess their "therapeutic concentration window" by comparing their enhancing effects and toxicity potential. Considering that the obtained results still need to be confirmed by in vivo experiments, the study still clearly shows that effective penetration enhancement is possible without compromising safety (Whitehead et al., 2008).

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3 Future Outlook The large and growing market share of oral formulations in the future will include developments that can be divided into two categories: personalized medicine and formulation of malabsorbed chemicals, including biotech drugs and established drugs. While each individual's individual formula is still fiction, the fact is that age has a great influence on the required dose and frequency of dosing. Children are often "therapeutic orphans" because drug development and clinical trials are designed for adults. Medicinal drops or lollipops are not the ultimate solution to pediatric formulation needs, at least in Europe. Physiological changes in gastrointestinal pH, residence time, or blood flow, as well as possible morphological changes such as mucosal atrophy or decreased liver function, and sometimes motor function must be considered when designing dosage forms for the growing elderly population compromised account. Oral formulations that address these age-related challenges promise to significantly improve treatment outcomes. Another challenge is to consider circadian rhythms such as heart rate, blood pressure, plasma hormone levels, gastric pH, renal function and disease states such as B. asthma or allergies. Addressing these parameters requires ingenious modified-release formulations that reduce side effects and improve efficacy. Today, the genetic mechanisms behind more and more diseases such as cancer or various metabolic disorders have been elucidated. Needless to say, there is a great deal of interest in using this knowledge for new treatment options. However, the main limitation is the lack of suitable delivery equipment. Therapeutic genes as well as proteins and peptides must be protected from premature degradation and targeted to their molecular targets. While fundamental hurdles still stand in the way of successful gene delivery, some researchers have begun experimenting with the hope that one day even oral gene therapy drugs will be available. Despite enormous efforts, oral delivery of therapeutic proteins and peptides in sufficient quantities remains a challenge. Nature has equipped the GI tract with many protective mechanisms to limit the access of bioactive proteins or genetic compounds. Therefore, the success of oral bioengineered therapeutics will depend on our ability to cross the GI barrier without compromising gut physiology.

References Ahmed IS (2005) Effect of simulated gastrointestinal tract conditions on drug release from pectin/ethylcellulose as a film coating to the colon. Drug Dev Ind Pharm 31:465-470 Aikawa K, Matsumoto K, Uda H, Tanaka S, Shimamura H, Aramaki Y, Tsuchiya S (1998) pH-responsive polymer polyvinyl acetal-diethylaminoacetate ( AEA) hydrogelation. International Journal of Pharmacy 167:97-104

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Wirth M, Kneuer C, Lehr CM, Gabor F (2002) Lectin-mediated drug delivery: evidence for distinguishing cell adhesion from cell invasion and lysosomal accumulation of wheat germ lectin in the Caco-2 model. J Drug Target 10:439–448 Wiwattanapatapee R, Lomlim L, Saramunee K (2003) Dendrimer conjugates for colonic delivery of 5-aminosalicylic acid. J Control Release 88:1–9 Wu W, Nancollas GH (1998) New insights into the relationship between solubility and particle size. J Solution Chem 27:521-531 Wu Y, Loper A, Landis E, Hettrick L, Novak L, Lynn K, Chen C, Thompson K, Higgins R, Batra U, Shelukar S, Kwei G, Storey D (2004) The The Role of Biopharmaceuticals in the Development of MK-0869 Clinical Nanoparticle Formulations: The Beagle Model Predicts Improved Bioavailability and Reduced Effect of Food on Human Absorption. Int J Pharm 285:135–146 Yang L, Eshraghi J, Fassihi R (1999) A novel intragastric drug delivery system for the treatment of Helicobacter pylori-associated gastric ulcers: in vitro evaluation. J Control Release 57:215–222 Yeh PY, Berenson MM, Samowitz WS, Kopecˇkova´ P, Kopecek J (1995) Site-specific drug delivery and penetration enhancement in the gastrointestinal tract. J Control Release 36:109-124 Yoshioka M, Hancock BC, Zografi G (1994) Indomethacin crystallizes from an amorphous state below and above its glass transition temperature. J Pharm Sci 83:1700-1705 Yoshioka M, Hancock BC, Zografi G (1995) Inhibition of indomethacin crystallization in poly(vinylpyrrolidone) co-precipitates. J Pharm Sci 84:983-986 Zentner GM, Rork GS, Himmelstein KJ (1985) Controlled porosity osmotic pumps. J Control Release 1:269–282 Zhang H, Alsarra IA, Neuau SH (2002) In vitro evaluation of a chitosan-containing multiparticulate system for macromolecule delivery to the colon. International Journal of Pharmacy 239:197-205

Transdermal Drug Delivery Richard H. Guy

content 1 2

introduce . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Current status of transdermal drug delivery. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Barrier function of the skin. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Drugs administered transdermally. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3 Improve transdermal drug delivery. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4 Topical and "subcutaneous" administration. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Conclusion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .refer to. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract: Transdermal drug delivery is a proven technology that contributes significantly to the global supply of medicines. Since the 1980s, the field has seen impressive growth and many commercial successes; importantly, a new chemical has recently been developed and approved for transdermal delivery without the need for injectables previously used as or oral dosage forms. The progress made is based on a clearer understanding of skin barrier function and the physicochemical, pharmacokinetic, and physiological factors underpinning the feasibility of transdermal drug delivery. Novel non-invasive methods of enhancing and controlling drug transport across the skin are being extensively investigated, and some technologies, such as B. iontophoresis, have reached real maturity. "Local" subcutaneous delivery of drugs (eg, to underlying muscles and other sites).

R.H. Guy Department of Pharmacy and Pharmacology, University of Bath, Claverton Down, Bath, BA2 7AY, UK Email:[email protected]This chapter is a revised version of the article "Transdermal Science and Technology - An Update" published by R.H. Guy in Drug Delivery System, 22:442-449 (2007). M. Schäfer-Korting (Ed.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_13, # Springer-Verlag Berlin Heidelberg 2010

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organizations) are gaining recognition and new opportunities can be envisaged in this underappreciated field. Key words transdermal drug delivery skin improvement iontophoresis minimally invasive technique

1 Introduction The ultimate goal of drug therapy is a non-invasive device that senses a patient's drug needs and then delivers the correct dose at the correct rate. Currently, this is achieved by taking a blood sample and analyzing the drug itself or for biomarkers, and then administering it, usually by intravenous injection or infusion. Passive transdermal delivery meets non-invasive criteria and enables drug delivery at a substantially sustained and controlled rate over extended periods of time (Figure 1). However, most patches only provide zero-order deployment, with no monitoring capabilities. In contrast, iontophoresis and a new generation of so-called "minimally invasive" technologies, such as microneedling and thermal poration methods, differ in the way they treat the skin as a platform for information exchange. This chapter examines the latest technology in transdermal drug delivery and examines the potential for future developments. Although the discussion focuses on drugs with systemic effects, the concepts discussed are equally applicable to the "topical" subcutaneous administration of anti-inflammatory drugs.

Figure 1 Steady State Plasma Concentrations of Oxybutynin Following Repeated Transdermal Dosing Over Several Days (http://www.oxytrol.com/index.asp)

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2 Current status of transdermal drug delivery 2.1

Barrier function of the skin

Over the past 30 years, significant progress has been made in the understanding of skin barrier function. The structure of the stratum corneum (SC) is a major resistance to drug transport and it has been established that the organization of the stratum corneum intercellular lipids is clearly related to its biophysical properties and its role in barrier function (Bouwstra and Ponec 2006. ). From this In this study, a simple algorithm for predicting skin permeability (Fig. 2) was derived (Potts and Guy 1992), so that the physicochemical feasibility of transdermal drug delivery could be tested in silico before conducting experiments. However, while the potential "availability" of a given drug can be assessed, the pool of potential drug candidates for transdermal delivery has not expanded significantly. The criteria to be met are the same as those established before the launch of scopolamine, the first approved transdermal drug, in the early 1980s. Although the physical properties of the skin barrier are now well understood, its biology remains a real challenge since many chemicals, including some transdermal candidates, can cause skin irritation or sensitization. Currently, there is no solution to this problem, and drugs that cause irritation are not an option for transdermal delivery. Therefore, any drug/agent combination evaluating transdermal feasibility must be tested for irritation and sensitization early in the development process.

2.2

transdermal drugs

Transdermal drug delivery has several well-known advantages (Delgado-Charro and Guy 2001). Systemic prometabolism is avoided, which means that lower daily doses can be administered. Blood or plasma levels of the drug can be maintained longer within the therapeutic window, prolonging the effect of the drug and reducing the frequency of dosing required. This reduces inter- and intra-patient variability and increases patient compliance and acceptance. Finally, drug delivery can be stopped by removing the patch. Figure 2 Empirical algorithm for predicting the permeability coefficient (kp) of a drug from an aqueous carrier through the skin (Potts and Guy 1992)

log kp = -2.7 + 0.71 * log Ko/w - 0.0061 * MW Ko/w = octanol-water partition coefficient. MW = molecular weight. The unit of kp is cm/h.

402 2.0

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Figure 3 Linear relationship between systemic steady-state plasma concentrations (Cp) and transdermal drug delivery system area for clonidine, nitroglycerin, and fentanyl (Delgado-Charro and Guy 2001)

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Transdermal drug delivery, on the other hand, is limited by several factors (DelgadoCharro and Guy 2001). It should only be used on potent drugs with daily doses on the order of 10 mg or less. Higher doses can only be achieved by increasing patch size (Figure 3). From a practical and aesthetic point of view, 100 cm2 is generally considered an acceptable maximum. Transdermal drugs are "small" molecules with a molecular weight (at least currently) of 500 Da or less; lipophilicity is essential, but some water solubility is also required since the drug must leave the SC to be absorbed systemically. Of course, transdermal delivery is suitable for drugs with poor oral bioavailability (if any) and short biological half-lives (implying significant first-pass effects and unfavorable dosing regimens). Finally, as previously stated, the medicinal product must not be locally irritating or sensitizing. Transdermal administration is approved for the following drugs: scopolamine (motion sickness), nitroglycerin and isosorbide dinitrate (angina), clonidine (hypertension), estradiol (hormone replacement therapy), fentanyl (sedation pain), nicotine (smoking cessation), testosterone (hypogonadism), norelgestromin + ethinyl estradiol (contraception), oxybutynin (urinary incontinence), selegiline (depression), methylphenidate (caution motor deficit/hyperactivity disorder), buprenorphine (pain relief), rivastigmine (dementia), rotigotine (anti-Parkinson's drug), and granisetron (anti-Parkinson's drug). emetic). The range of diseases treated by transdermal drug delivery is very wide - the only commonality between these drugs is their effectiveness. Rotigotine is a particularly noteworthy example because it is the first novel chemical agent designed from scratch for transdermal delivery. All other drugs were approved for administration by other, more conventional routes of administration, usually oral, until transdermal routes were developed. Oxybutynin (molecular weight 357.5 Da, log(octanol-water partition coefficient (P)) ~4.7) can be used to characterize the drug for transdermal delivery (Dmochowski et al., 2003). The Oxybutynin Patch is a thin, flexible, clear acrylic adhesive system that is applied every 3-4 days (e.g. on the abdomen or upper arms) to significantly reduce incontinence episodes. Observed skin irritation was limited. Delivers approximately 4 mg of drug per day

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Figure 4 Oxybutynin and its major metabolite N-desethyloxybutynin ( mean steady-state plasma concentration of N-DEO). ASP)

From a 39 cm2 patch, with a total oxybutynin loading of 36 mg, the system maintained relatively constant plasma concentrations (Figure 1). When administered orally, the drug exhibits extensive first-pass effects due to biotransformation by CYP3A4. Transdermal administration completely eliminated this problem, as evidenced by the significantly lower systemic concentrations of the drug's major metabolite (Figure 4). In addition, the dose of oxybutynin required for transdermal administration was significantly reduced.

2.3

Improve transdermal drug delivery

The SC barrier completely eliminates the passive transport of charged drugs or compounds with a molecular weight greater than 1,000 Da. Getting around this deadlock requires enhanced technology. The options at hand are divided into three approaches: l l l

Formulations containing chemical penetration enhancers; physical effects on the drug itself, such as iontophoresis; and physical and/or mechanical energy applied to the barrier, such as ultrasound, microneedling, thermoporation.

A literature review on penetration enhancers is extensive (Smith and Maibach 2005) and provides a consistent message: that is, although several patches on the market contain excipients that allow the drug to pass through the skin more quickly, the Incorporation of enhancers into skin where transdermal delivery systems are not possible can lead to increased incidence of irritation. In fact, there are direct

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An enhancer is the correlation between the effectiveness of enhancing drug penetration and its inherent ability to cause irritation; thus, the better the enhancer, the more severe the irritation observed. Thus, enhancer-drug combinations that significantly enhance transport while causing minimal or no stimulation have not yet been defined (Karande et al., 2005). Regulatory hurdles to incorporating declared dermal penetration enhancers into new formulations are also significant. It is not unreasonable for authorities such as the US Food and Drug Administration to require evidence that penetration enhancers actually work; moreover, if the excipient is not a Generally Recognized as Safe Substance (GRAS), information is required to demonstrate its safety. Alternative strategies involving the manipulation of topical formulations include the incorporation of liposomes or other "vehicles" and the use of supersaturation. Despite significant efforts in this area, no transdermal products utilizing this technology have been commercialized to date. Drug flux can be improved with supersaturated carriers without irritation (Figure 5). However, there is currently a major stability issue that makes this method difficult in practice (Moser et al. 2001). However, spray formulations are under development (Thomas and Finnin 2004), and these may eventually show that the technology is feasible. Iontophoresis improves drug delivery across the SC without significantly disrupting the barrier itself. In iontophoresis, a potential gradient applied to the skin drives the movement of (primarily) ionized species through the SC, which is normally impermeable for transport of charged entities (Mudry et al., 2006). Ion electromigration is determined by the current flowing in the iontophoresis circuit. Therefore, the transdermal drug delivery is proportional to and controlled by the charge flowing between the electrodes

Placebo Gel 0.02% Supersat. Gel Untreated Control 0.1% Cream Placebo Cream 0

Steroid Effects (Vosoconstrictor Units) Figure 5 Hydrocortisone-induced vasoconstriction (measured in arbitrary units) in a commercial cream formulation (0.1%, w/v) and a supersaturated gel containing only 0.02% After dosing (w/v). Three different controls were also evaluated as previously described (A.F. Davis, personal communication)

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Figure 6 Iontophoresis of ropinol through the skin as a function of (a) current density and (b) concentration of drug hydrochloride in the region of the driving (positive) electrode (LuzardoAlvarez et al., 2001)

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Contact with the skin (Figure 6) (Luzardo-Alvarez et al., 2001). However, the ionized drug is only one of several available charge carriers in the system: for example, the background electrolyte may contain ions of the same charge that can compete with the drug for charge transport into the skin; moreover, oppositely charged ions in the body can Charge is transported in the opposite direction. Therefore, the efficiency of iontophoresis is always less than 100%, and often very low. Furthermore, electromigration is not the only mechanism involved in iontophoresis. Under normal physiological conditions, the skin has a net negative charge and is therefore preferentially permeable to cations. Because the membrane is charged, applying an electric field to the skin induces convective electroosmosis of the solvent in the anode-to-cathode direction, increasing cation transport, which is delayed by anions, and providing a means of enhancing transdermal delivery is more neutral, water-soluble sexual compounds. The contribution of electromigration and electroosmosis to overall iontophoretic transport depends on the size and mobility of the drug; the latter becomes increasingly important as the size of the molecule increases (Guy et al. 2000). Although the principles of iontophoresis have been known for more than a century, transdermal devices based on this technology have only recently entered the market (though, at least so far, without some degree of commercial success). Integrated delivery systems for lidocaine (for local anesthesia) (Vyteris 2008) and fentanyl (for pain relief) (Janssen-Cilag 2009) have received regulatory approval and iontophoresis devices have been introduced for diabetes Glucose monitoring in patients (Sieg et al., 2005). In the latter example, an electrical current "pulls" glucose out of the body's extracellular fluid, through the skin to the surface, where it is analyzed on-site and the results are transformed by a pre-calibrated A measure of blood sugar levels. Therefore, iontophoresis is a well-established technique for drug delivery and non-invasive monitoring. Other systems are in development with potential applications in fertility treatment and Parkinson's disease treatment. The development of technologies capable of "shorting-circuiting" the skin's barrier function has been the subject of much activity in recent years. These methods are characterized by creating new paths through SC

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48.3 microns wide

41.7 microns wide

100 microns

Figure 7 Confocal microscopic image of human skin after perforation with a microneedle array (MicroCorTM, Corium International, Inc., Redwood City, CA) (G.W. Cleary, personal communication)

Into the skin (Guy 2003). The dimensions of these tubes are measured in microns (Figure 7), so they are large enough to transport large molecules such as vaccines, proteins, and DNA. Methods of skin perforation include mechanical means, through microneedle arrays or high-velocity particles, or the application of physical energy, such as the use of ultrasound (phonophoresis) or heat or lasers (thermoporation). The term "minimally invasive" is used to describe these new technologies, reflecting the claim that they produce little or no physical sensations in the body. Among these methods, microneedling, thermoporation, and sonophoresis are the most advanced. Tremendous advances in micromachining and microfabrication technologies have enabled the fabrication of various microneedle arrays (Sivamani et al., 2007). They can be very small, only penetrating about 100 mm through the skin (and thus causing hardly any sensation). Various materials are used to make microneedles, which can be solid or hollow. Their application strategy could be insertion followed by immediate removal and application of a patch, or the use of biodegradable microneedle arrays to allow slow release of the drug of interest in situ. Alternatively, microprojections can be coated with an inoculum that desorbs upon insertion and then stimulates the skin's immune system to enhance the desired pharmacological effect. Thermal energy can be used to create micron-sized portals in the skin using heated microwire arrays or using radiofrequency microablation (Altea 2008; Levin et al. 2005). Pore depth induction

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patch ab

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Figure 8. Insulin serum concentration versus time profiles (mean SEM; n=8) in normal healthy subjects 12 hours after transdermal administration through heated porous skin (Altea Therapeutics 2008). The minimum effective value is 9-10 mU/ml

407

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Ensure access to viable tissue and ensure that the tubing is then quickly filled with extracellular fluid. In this way, therapeutic doses of drugs impermeable to normal skin (for example, salts of active substances are much more stable than corresponding non-ionized substances) can be delivered through relatively small skin areas. Thermal ablation of SCs significantly reduces membrane impedance, allowing feedback control of energy pulses delivered to the skin. Like microneedles, the size of the micropores allows for the transdermal delivery of proteins such as insulin (Figure 8), human growth hormone, and interferon-alpha. For over a decade, attempts have been made to increase drug delivery through the skin by using low-frequency (20–100 KHz) ultrasound (Mitragotri and Kost 2004)—so-called sonophoresis. Although approved products are not yet on the market, phonophoresis has been shown to significantly increase the delivery of small and large molecules, and its use in noninvasive blood glucose monitoring is under active development. Cavitation has been found to be the mechanism by which ultrasound creates a low resistance path through the skin. The rate of formation of these networks can be significantly accelerated by pretreating SCs with surfactants such as sodium lauryl sulfate. The increased permeability lasted for approximately 12 hours, and barrier restoration was reported to occur within one day of treatment. Overall, the described minimally invasive technique offers several interesting applications, especially in the delivery of vaccines and other macromolecules, and noninvasive glucose monitoring. Despite the high intensity of activity currently underway, the field is still in a so-called "early" stage of development, with proofs of principle clearly demonstrated but long-term validation of the resulting successful products yet to be validated. The main open questions concern newly created The patency of the route, the overall safety associated with skin pore opening, and the economics—that is, will the cost of the technology be competitive with simple needles and syringes and will medical systems and insurance companies reimburse such advanced therapies nationwide?

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Topical and "subcutaneous" administration

Barriers to drug delivery to the skin and directly to underlying tissue to treat local inflammation in dermatological conditions appear to be similar to those encountered with transdermal drug delivery. Achievement of both goals has been disappointing so far, as clearly inefficient formulations have dominated the market. Therefore, there is a real opportunity to significantly improve medical treatment in these areas. Skin problems are increasingly part of the GP's workload. This is explained by the fact that the incidence of atopic eczema in children increased from 5% in 1950 to over 25% in 2000 and is still increasing (Cork et al., 2006). However, for most topical drug formulations, only a few percent of the administered dose is actually absorbed; that is, the "bioavailability" of the topical drug is low. Therefore, there is clearly a need to improve the situation by better matching the drug to its carrier and ensuring improved availability of the active substance (as this will undoubtedly lead to more reproducible treatments with fewer or no side effects). Influence). As mentioned earlier, supersaturation is a means by which more of the administered dose can be mobilized into the skin, and the creative use of volatile and non-volatile excipients is an approach being explored to take advantage of this sub- Steady state without having to be so difficult to solve stability problems. The use of NSAIDs to treat localized pain and inflammation of subcutaneous targets (muscles, joints, tendons, etc.) (Lee and Maibach 2006) has been successfully and widely used in Asia. While there has been skepticism about the effectiveness of this approach in the United States, and relatively little activity in Europe, interest in the opportunities offered by the strategy is growing. The key, of course, is to provide convincing evidence that topical administration to the subcutaneous tissue is at least as good as oral administration in terms of how quickly and to what extent the drug reaches the target site. An important benefit in achieving this goal is that equivalent efficacy is achieved with very low or negligible patient systemic exposure to the drug, while significantly reducing adverse side effects. The segment is expected to witness significant growth.

3 Conclusions Several outstanding contributions of transdermal drug delivery can be identified: 1. Drug delivery that differs significantly from traditional methods such as oral delivery. 2. Elimination of the "peaks and valleys" in drug plasma levels that occur with more conventional routes of administration.

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3. Develop different patch technologies and delivery mechanisms to achieve the desired drug delivery profile. 4. Linear operation of the drug dose by changing the patch area. 5. Avoid the first-pass effect and change the ratio of drug to metabolite (thereby reducing some important side effects). 6. Applications of the technology in various therapeutic areas. 7. Provides continuous and controlled drug delivery from 12 hours to 7 days. 8. Drugs that are difficult to formulate have been successfully administered transdermally. 9. Improve patient compliance. 10. Improve medication use. The success of transdermal delivery is expected to continue, although the rate at which new conventional passive delivery candidates are considered is unlikely to change significantly. Also, predictable is an area of significant growth. For example, while iontophoresis has yet to hit a commercial home run, the maturity of the technology means it will be available once the right opportunity is found. The range of minimally invasive approaches may somehow affect macromolecule drug delivery, and many interesting drug-technical 'fits' have been explored in depth. Finally, there is no doubt that the development of the transdermal space will also facilitate new methods of delivering drugs into the skin and subcutaneously, and important potential applications in these areas are evident.

References Altea Therapeutics (2008) http://www.alteatherapeutics.com Bouwstra JA, Ponec M (2006) The skin barrier in health and disease. Biochim Biophys Acta 1758:2080–2095 Cork MJ, Robinson DA, Vasilopoulos Y, Ferguson A, Moustafa M, MacGowan A, Duff GW, Ward SJ, Tazi-Ahnini R (2006) New perspectives on epidermal barrier dysfunction in atopic dermatitis : Gene-Environment Interactions. J Allergy Clin Immunol 118:3-21 Delgado-Charro MB, Guy RH (2001) Transdermal delivery. In: Hillery AM, Lloyd AW, Swarbrick J (eds) Drug delivery and targeting for pharmacists and pharmaceutical scientists. Harwood Academic Publishers, London, pp. 207–236 Dmochowski RR, Sand PK, Zinner NR, Gittelman MC, Davila GW, Sanders SW, Transdermal Oxybutynin Study Group (2003) Treated patients, urgent and mixed Patients with urinary incontinence. Urology 62:237-242 Guy RH (2003) "Minimally invasive" techniques for transdermal drug delivery and clinical chemistry. Drug Delivery Challenges of the New Millennium. Bulletin Technique Gattefosse' 96:47-61 Guy RH, Kalia YN, Delgado-Charro MB, Merino V, Lopez A, Marro D (2000) Iontophoresis: Electrorepulsion and Electroosmosis. J Control Release 64:129-132 Janssen-Cilag (2009) http://www.janssen-cilag.co.uk/product/detail.jhtml?itemname=ionsys_info

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Karande P, Jain A, Ergun K, Kispersky V, Mitragotri S (2005) Design principles of chemical penetration enhancers for transdermal drug delivery. Proc Natl Acad Sci USA 102:4688–4693 Lee CM, Maibach HI (2006) Deep percutaneous penetration of muscles and joints. J Pharm Sci 95:1405–1413 Levin G, Gershonowitz A, Sacks H, Stern M, Sherman A, Rudaev S, Zivin I, Phillip M (2005) Transdermal delivery of human growth hormone via radiofrequency microchannels. Pharm Res 22:550-555 Luzardo-Alvarez A, Delgado-Charro MB, Blanco-Méndez J (2001) Iontophoretic delivery of ropinirole hydrochloride: effects of current density and carrier formulation. Pharm Res 18:1714–1720 Mitragotri S, Kost J (2004) Low frequency sonography: a review. Adv Drug Deliv Rev 56:589–601 Moser K, Kriwet K, Naik A, Kalia YN, Guy RH (2001) Passive enhancement of skin penetration and its in vitro quantification. Eur J Pharm Biopharm 52:103-112 Mudry B, Guy RH, Delgado-Charro MB (2006) Iontophoresis in transdermal drug delivery. In: Touitou E, Barry BW (eds.) Enhancement of drug delivery. Taylor and Francis, New York, NY, pp. 279–302 Potts RO, Guy RH (1992) Predicting skin permeability. Pharm Res 9:663–669 Sieg A, Guy RH, Delgado-Charro MB (2005) Noninvasive and minimally invasive methods for transcutaneous glucose monitoring. Diabetes Technol Ther 7:174–197 Sivamani RK, Liepmann D, Maibach HI (2007) Microneedling and transdermal applications. Expert Opinion Drug Deliv 4:19-25 Smith EW, Maibach HI (2005) Percutaneous Penetration Enhancers 2nd Ed. CRC Press, Boca Raton, FL Thomas BJ, Finnin BC (2004) The Transdermal Revolution. Drug Discov Today 9:697-703 Vyteris (2008) http://www.vyteris.com/home/Our_Products/Lidosite.php

In the Brain's Perspective - Overcoming or Circumventing the Blood-Brain Barrier Heidrun Potschka

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introduce . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 412 The structure and function of the blood-brain barrier. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 412 Limitations of the blood-brain barrier as a treatment for central nervous system disorders factor. . . . . . . 414 Regulation of blood-brain barrier function. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 416 4.1 Opening of the BBB. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 417 4.2 Inhibition of efflux transport. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 418 4.3 Prevention of disease-related or treatment-related blood-brain barrier changes. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 420 5 turns Cross the blood-brain barrier. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 423 5.1 Nanocarrier systems and drug conjugates. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 423 5.2 Intranasal Administration. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 424 5.3 Intracerebral Administration. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 424 6 Conclusion. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 426 References. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 426

Conclusions: Components of the blood-brain barrier, including its efflux transport system, can effectively limit the entry of potential CNS therapeutics into the brain. Efficient excretion of transporters from the brain is a common reason why the pharmaceutical industry excludes new compounds from further development of CNS therapeutics. Furthermore, high levels of transporter expression present in individual patients or often associated with pathophysiology appear to be a major cause of treatment failure in various CNS disorders, including brain tumors, epilepsy, brain HIV infection, and psychiatric disorders. Increased knowledge of the structure and function of the blood-brain barrier provides the basis for developing strategies to increase brain uptake. 16, 80539, Munich, Germany Email: p[email protected]M. Schäfer-Korting (Hrsg.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_14, # Springer-Verlag Berlin Heidelberg 2010

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Useful pharmaceutical compounds. The various strategies discussed in this review aim to modulate the function of the blood-brain barrier or to bypass components of the blood-brain barrier. Key words blood-brain barrier, controlled release system, efflux transporter nanoparticles, P-glycoprotein, pharmacodynamics

1 Introduction The limited membrane permeability of cells that are part of the blood-tissue barrier plays a key role in protecting tissues from putatively harmful xenobiotics. On the other hand, effective barriers may limit the brain penetration of central nervous system (CNS) therapies. Consequently, many promising compounds in CNS drug development fail due to limited access to brain targets. In addition, the efficacy of marketed drugs can be reduced by restricting brain pathways, leading to ineffective or even outright drug resistance. Penetration may be limited by underlying barrier function. In certain diseases, there may be pathophysiologically induced changes in the blood-brain barrier (BBB), which further restrict access to the brain. In addition, the drug can also affect the blood-brain barrier, further strengthening the barrier. In individual patients, inheritance of physiological factors that contribute to BBB structure and function may additionally influence brain pathways and the efficacy of CNS treatments. It is of particular interest to develop strategies to overcome, modulate, or bypass the BBB with the goal of optimizing brain pharmacokinetics.

2 Structure and function of the blood-brain barrier The BBB plays a crucial role in controlling the passage of compounds from the blood into the central nervous system. The main component of the BBB is the single layer of capillary endothelial cells of the brain. The limitations of brain penetration are due to the presence of tight junctions between adjacent endothelial cells and the relative absence of fenestrae and pinocytic vesicles. Basement membranes, pericytes, and astrocyte foot processes surround the capillary endothelium of the brain. Tight junctions between capillary brain endothelial cells and surrounding astrocytes appear to be critical for the induction of endothelial layer barrier function, including the formation of interendothelial tight junction complexes. Because of the blood-brain barrier, circulating compounds can only enter the brain by lipid-mediated passive diffusion of small nonpolar molecules or, less commonly, by catalytic transport (Pardridge 1999). as a

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Thus, for most CNS drugs, there is a strong positive correlation between lipophilicity and brain pathways. However, uptake may be lower than expected for compounds that are excreted from the brain by active BBB efflux transporters. A number of membrane transporters have been described in BBB endothelial cells involved in the influx or efflux of various essential substrates, including electrolytes, nucleosides, amino acids, and glucose (Lee et al., 2001), as well as xenobiotics. Efflux transporters of the BBB protect CNS tissue from environmental changes by limiting penetration into brain tissue and promoting extrusion of brain tissue (Leslie et al. 2005). Brain efflux transporters also act by restricting the entry of drugs used to treat CNS disorders into the brain, as transporters do not differentiate between harmful xenobiotics and active pharmaceutical ingredients (APIs) of drugs used to treat CNS disorders with adverse effects (Loscher and Potschka 2005a). Several brain efflux transporters are associated with restricted brain penetration of CNS-active drugs, which can limit drug efficacy and even lead to complete drug resistance. P-glycoprotein (Pgp, ABCB1) was the first drug efflux transporter discovered in the BBB more than a decade ago (Cordon-Cardo et al., 1989; Thiebaut et al., 1987). Since then, accumulating data have shown that various BBB efflux transporters play key roles in limiting brain uptake of various therapeutic agents (Loscher and Potschka 2005b). The most important efflux transporters identified at the BBB to date belong to the ABC transporter class. The ABC transporter consists of two transmembrane domains and two nucleotide-binding domains (Rosenberg et al. 2003), which may also be encoded by two separate polypeptides. For several ABC transporters, such as Pgp, there is more than one substrate-binding site per transporter, allowing a wide range of substrates. The transport of compounds is associated with large conformational changes in the transporter molecule (Martin et al., 2001). ATP binding appears to trigger conformational changes associated with changes in the affinity and orientation of the substrate-binding site, resulting in release of the substrate at the extracellular surface of the membrane (Martin et al., 2001). Subsequently, hydrolysis of ATP resets the transporter for the next cycle (Senior et al., 1995). ABC genes are divided into families (ABCA, ABCB, ABCC, ABCD, ABCE, ABCF, and ABCG) based on the structural features of the encoded transporters. Since the old term is often used in the publications cited in this review, it is used throughout the chapter. The new nomenclature will be shown at least once when the corresponding operator is mentioned for the first time. Efflux transporters expressed at the BBB include members of the ABCB, ABCC, and ABCG families: Pgp (ABCB1), members of the multidrug resistance-associated protein family (MRP/ABCC family), and breast cancer-associated protein (BCRP/ABCG2).Development The question of whether an agent is a transport substrate is of particular interest to the pharmaceutical industry. Low affinity for BBB efflux

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Transporters facilitate the development of CNS therapeutics that require high concentrations in the brain. Conversely, high affinity for BBB efflux transporters facilitates the development of drugs that should act in the periphery to avoid CNS side effects.

3 Blood-brain barrier as a limiting factor in the treatment of central nervous system diseases As mentioned earlier, the function of the blood-brain barrier can decisively affect the efficacy and tolerance of drugs. Compounds may be too hydrophilic for efficient entry by diffusion, and transport by brain efflux transporters may be the limiting factor for brain uptake of lipophilic compounds. Treatment success for many central nervous system disorders, including brain cancer, epilepsy, depression, schizophrenia, and HIV-associated encephalopathy, is limited by poor response or outright resistance to drug therapy. Among other mechanisms, changes in drug uptake by target cells in the brain or brain parenchyma are thought to be an important cause of treatment failure (Loscher and Potschka 2005a; Thuerauf and Fromm 2006). It has been hypothesized that disease-associated or therapy-associated changes in the expression of efflux transporters decisively affect the pharmacokinetics of a large number of important CNS drugs in the brain. Many brain tumors are highly resistant to drug therapy, and systemic chemotherapy often does not improve outcome. A key factor in the failure of systemic chemotherapy is the limited penetration of effective chemotherapeutic agents into the BBB (Nies 2007). Anticancer drugs were among the first to be identified as substrates of BBB efflux transporters; H. From Pgp as well as MRP and BCRP. Several studies have shown that the poor efficacy of systemically administered anticancer drugs is at least in part due to the activity of BBB efflux transporters (Kemper et al., 2004). In 30-40% of people with epilepsy, antiepileptic drug therapy does not adequately control seizure activity. Transporter inhibitor microdialysis experiments, knockout mouse experiments, and in vitro studies have shown that several antiepileptic drugs are transported by BBB-Pgp and some by MRP (Cucullo et al. 2007; Loscher and Potschka 2005b; Marchi et al. 2005 2002; Rizzi et al. 2002; Sills et al. 2002). Since antiepileptic drugs generally penetrate the brain well, they can only be low to moderate affinity substrates. Penetration of antiepileptic drugs into the brain is limited only when seizure activity leads to overexpression of BBB efflux transporters. Therefore, tests with sufficient sensitivity to determine whether antiepileptic drugs are substrates of BBB transporters are needed. Ignoring this fact has led to some inconsistent data in recent years. Therefore, further studies are needed to determine whether all clinically relevant antiepileptic drugs are substrates of BBB efflux transporters, especially the human subtype.

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In general, seizure-induced overexpression of BBB efflux transporters in epileptic brains may explain multidrug resistance in epilepsy due to restricted access to antiepileptic drugs to their target sites. Experiments with two different models of drug-resistant epilepsy provide important support for this concept. Pgp expression in drug-resistant rats significantly exceeds that in drug-responsive rats (Potschka et al. 2004; Volk and Loscher 2005). Recent experimental studies have shown that transporter inhibitors can overcome drug resistance in seizures, providing further evidence for the multidrug transporter hypothesis in drug-resistant epilepsy (Brandt et al., 2006; Clinckers et al., 2005) . Further development or validation of new strategies to overcome drug-resistant epilepsy must take into account that the problem is multifactorial, and therefore the relative importance of overexpression of efflux transporters must be elucidated. Genetic defects in Pgp in mice lead to improved brain pathways for several antidepressants, suggesting that these drugs are secreted into the blood through Pgp (Grauer and Uhr 2004; Uhr and Grauer 2003; Uhr et al. 2003, 2000). Whether this active efflux transport contributes to treatment failure in depression remains to be clarified. The lack of models for treatment-resistant psychiatric illness makes it difficult to test the validity of this hypothesis, so the hypothesis remains speculative. The first indirect support for the impact of BBB efflux transporters on treatment success in psychiatric disorders came from a genetic analysis of schizophrenic patients treated with broperidol. MDR1 genotype is associated with treatment response to bromperidol (Yasui-Furukori et al., 2006). More recently, Uhr et al. (2008) reported that MDR1 genotype in patients with depression is a strong predictor of success in multiple antidepressant treatments. The development of HIV protease inhibitors has led to major advances in the treatment of HIV infection. A major limitation of their effectiveness, however, is restricting access to the brain, leaving the brain's viral reservoir intact. Pgp-mediated efflux has been hypothesized to result in limited brain penetration of HIV protease inhibitors such as saquinavir, amprenavir, nelfinavir, and indinavir (Banks et al., 2006; Edwards et al. et al., 2002; Kim et al., 1998; Washington et al., et al., 2000). Upregulation of Pgp at the BBB by HIV Tat protein (Hayashi et al., 2005) may further reduce the penetration and efficacy of HIV protease inhibitors in long-term AIDS survivors. In addition to Pgp, MRP1, MRP2 and MRP4 also accept HIV protease inhibitors as substrates and thus may be involved in limiting their entry into the brain. Riluzole is the only drug recognized to prolong survival in patients with ALS. Studies in Pgp knockout mice have shown that both riluzole and minocycline, a compound that can delay disease onset, are substrates of Pgp (Milane et al., 2007), and thus this BBB efflux transport Protein may affect the therapeutic effect. Pgp-mediated brain extrusion has a dramatic impact on the efficacy of opioids and opioid analgesics. Modulation of Pgp function significantly affects the analgesic effects of morphine (King et al. 2001; Letrent et al. 1999; Thompson).

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Table 1 CNS Therapeutic Drugs as Substrates of BBB Efflux Transporters: Example Pharmacology Examples of Transporters Involved Cancer Drugs Doxorubicin, Daunorubicin, Vinblastine, ABCB1/Pgp, ABCC-Vincristine, Paclitaxel , Etoposide, Topotecan Transporter/MRP, ABCG2/BCRP Analgesics Morphine, Methadone, Fentanyl ABCB1/Pgp HIV Protease Amprenavir, Indinavir, Saquinavir ABCB1/Pgp, ABCC Inhibitor Transporter/MRP Antipsychotics Olanzapine, Amisulpride ABCB1/Pgp Drugs Antiepileptics Phenytoin, Carbamazepine, Oxcarbazepine, ABCB1/Pgp, ABCC Drugs Lamotrigine, Phenobarbid Topiramate, Felbamate, Transporters/MRPs Valproic Acid, Topiramate Antidepressants Amitriptyline, Nortryptiline, Venlafaxine, ABCB1/Pgp-Paroxetine

wait. 2000). Thus, Pgp appears to be an important factor in pain control with opioid analgesics, which can affect the onset, intensity, and duration of the analgesic response (Dagenais et al. 2004). More recently, this hypothesis was further confirmed when Hamabe et al. (2007) reported an inverse correlation between the analgesic effect of morphine and individual Pgp expression levels in the cortex in a mouse model. In addition, Pgp appears to limit the brain distribution of certain antimicrobials, including fluoroquinolones and erythromycin (Sasabe et al 2004; Schinkel 1999). Excretion of these antibiotics from the brain may contribute to their limited or ineffective efficacy against microbial infections of the CNS. In conclusion, BBB efflux transporters have been shown to significantly affect the CNS effects of many drugs (Table 1), and this effect is clinically relevant for many of these drugs.

4 Modulation of blood-brain barrier function Increasing knowledge of the impact of the blood-brain barrier and its efflux transport systems on response to CNS therapy has prompted efforts to develop strategies to optimally target drugs to brain tissue (Fig. 1; Table 2) ).. The development of positron emission tomography-based imaging methods offered the possibility of non-invasively studying Pgp-mediated transport and its regulation in vivo in individual patients (Elsinga et al. 2004; Hendrikse and Vaalburg 2002; Langer et al. 2007; Lee et al., 2006). Advances in these diagnostic techniques will open up the possibility of selecting patients who can benefit from new strategies designed to evade or bypass the blood-brain barrier.

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brain

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Astrozytäre Endfüße basement membrane

Basolateral

1 tight knot

brain capillary endothelial cells

PGP

Apical/luminal MRP

BCRP

Blood Figure 1 Strategies to increase brain penetration of CNS therapeutics. (1) The opening of the blood-brain barrier can be achieved by dissolving tight junction complexes. (2) Modulation of BBB efflux transporter function or inhibition of efflux transporter induction resulted in a more specific increase in brain penetration. (3) The use of nanoscale delivery systems and drug conjugates can bypass BBB efflux transport molecules. (4) Direct intracerebral administration (e.g. via an implanted delivery system) bypasses the barrier and results in higher local drug concentrations

4.1

BBB opening

Disruption of the blood-brain barrier must be transient and reversible in order to play a role in the delivery of drugs, such as anticancer drugs, to the brain. Various hypertonic solutions have been used to disrupt the blood-brain barrier (Kroll et al. 1998). Due to its approval for use in patients, mannitol is the most commonly used drug in preclinical and clinical studies. Mannitol-mediated opening of the BBB has been used in combination with anticancer drugs to treat patients with metastatic or primary brain tumors. Some studies have shown this strategy with some success and the lowest morbidity. Briefly, it was hypothesized that mannitol causes osmotic contraction of endothelial cells, which exerts tension on tight junctions and then disintegrates. However, the events leading to increased permeability are likely to be much more complex due to the multiple structural and functional changes that occur in endothelial cells in response to mannitol. More recently, Farkas et al. (2005) reported that hypertonic mannitol induced β-catenin phosphorylation. Since β-catenin is a key component of the adapter complex, its phosphorylation may be important for mannitol-induced reversible opening of the BBB. In general, osmotic disruption of the BBB is insufficient to rule out the possibility of toxic xenobiotics entering the CNS. In addition, blood-brain barrier opening and albumin extravasation

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Table 2 Strategies to increase brain penetration B. Brain tumors Inhibition of efflux transport by mannitol bradykinin analogues Alkylglycerols Brain tumors Epilepsy Focal cerebral ischemia Altered Transporter Expression Epilepsy Focal Cerebral Ischemia BBB Nanocarrier System Bypass Brain Tumors and Drug Conjugates Epilepsy Focal Cerebral Ischemia HIV Mental Illness Intracerebral Administration Brain Tumor Epilepsy

experimental evidence

clinically proven

+ + +

+ +

Can promote and even cause epileptogenesis (Ivens et al 2007; Tomkins et al 2007; van Vliet et al 2007). Therefore, more specific strategies to target APIs to the brain without compromising the integrity of the blood-brain barrier would be beneficial. Administration of a bradykinin analogue (RMP-7) has been proposed as an alternative. RMP-7 opens tight junctions through a receptor-mediated mechanism, thereby facilitating the delivery of the cytostatic drug carboplatin to glioma implanted in rat brains (Matsukado et al., 1996). However, bradykinin analogues failed to improve the efficacy of carboplatin in phase II and III clinical trials (Prados et al., 2003). Other options include intracarotid administration of alkylglycerols, which affect the BBB in more subtle ways. Increased drug transport via the paracellular pathway has been described in rodents (Erdlenbruch et al. 2003a,b). As far as we know, there are no human data so far.

4.2

Inhibition of efflux transport

Growing awareness of the impact of efflux transporters on successful drug treatment of CNS disorders has prompted efforts to develop regulatory strategies

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Transporter function (Loscher and Potschka 2005a; Thuerauf and Fromm 2006). Since Pgp is known to transport a large number of commonly prescribed drugs, efforts to date have focused specifically on this transporter. Mechanisms that can regulate Pgp activity in the BBB include direct inhibition by specific inhibitors, functional modulation, and transcriptional regulation (Bauer et al. 2005). First, drugs used for other indications and known to inhibit Pgp in cell culture, such as verapamil, cyclosporine A, and quinidine, were tested as Pgp modulators (Fox and Bates 2007). Due to the low binding affinity for Pgp, high doses of these early inhibitors were required and excessive toxicity was observed in patients. Second-generation inhibitors are developed as analogs of the original drug. Valspodar (PSC-833), a non-immunosuppressive derivative of cyclosporin D, is an example of these drugs being developed (Fox and Bates 2007). The compound proved to be better tolerated than first-generation inhibitors. However, Valspodar inhibits CYP enzymes, resulting in decreased clearance and increased systemic exposure of co-administered compounds. While first- and second-generation Pgp inhibitors were hindered by additional pharmacodynamic effects or additional effects on drug metabolism (Thomas and Coley 2003), the development of third-generation Pgp inhibitors brought selectivity and more Effective modulators such as tariquidar, laniquidar, zosuquidar, etc. produced by elacridar (Bates et al., 2002; Thomas and Coley, 2003). Three generations of Pgp modulators include competitive inhibitors that are themselves substrates and noncompetitive inhibitors that induce conformational changes that affect transport efficiency. Given the complexity of efflux transport, the goal is to develop dual or pluripotent inhibitors. Jekerle et al. (2006) recently reported the development of a new inhibitor, WK-X-34, which modulates Pgp and BCRP in experimental models. In clinical settings, co-administration of Pgp inhibitors with anticancer drugs in oncology has shown some efficacy (Breedveld et al., 2006), although not all studies provided promising data. Therefore, further development of these agents must await to determine the true potential of Pgp-mediated reversal of multidrug resistance in the treatment of brain tumors and other CNS disorders. In this context, a recent study finding differences in the sensitivity of Pgp across different cellular and blood-tissue barriers is of particular interest (Choo et al., 2006). Pgp located in the BBB was found to be more resistant to inhibition than Pgp in other tissues (Choo et al. 2006). This resistance can be overcome with sufficiently high doses of inhibitors. However, whether this can be safely achieved in clinical settings remains to be clarified. Experimental studies using a rodent glioblastoma model and a rodent melanoma brain metastasis model demonstrated the effectiveness of this strategy (Fellner et al., 2002; Joo et al., 2008). The brain penetration and effectiveness of systemically administered paclitaxel can be significantly increased by coadministering the Pgp inhibitors Valspodar or HM30181A (Fellner et al., 2002). Co-administration of Pgp inhibitors also enhanced antiepileptic drug responses and even helped to overcome isolated antiepileptic drug resistance in several animal models (Brandt

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wait. 2006; Clinker et al. 2005). In these studies, the combination was shown to be well tolerated. Coadministration of third-generation Pgp inhibitors has also been shown to be a promising neuroprotective strategy in rodent models of focal cerebral ischemia (Spudich et al., 2006). Tariquidar increases the accumulation and neuroprotective efficacy of the neuroprotectants FK506 and rifampicin. Given the experimental success, it is important to consider that any modulation of transporter function is associated with specific hazards. First, complications of combining Pgp inhibitors or modulators with CNS-active drugs may be related to the intended target. Affecting the pharmacokinetics of therapeutic drugs does not only affect target tissues or brain regions. Elevated drug concentrations in other brain regions and peripheral tissues contribute to increased side effects. Consequently, several studies combining anticancer drugs and Pgp inhibitors had to be terminated prematurely due to increased chemotherapy-related toxicity (Fox and Bates 2007). Second, multidrug transporters such as Pgp have multiple physiological functions, including protection against xenobiotics. Other xenobiotics absorbed by the body may be more harmful due to the affected distribution of efflux transporter inhibitors. Furthermore, Pgp and MRP protect brain parenchymal cells from apoptosis (Gennuso et al. 2004; Pallis et al. 2002), thus transporter inhibition may promote cell death. Nonetheless, temporary inhibition of efflux transporters by short-term use of inhibitors may be a tolerable strategy to reverse or prevent drug resistance. In terms of specific brain targeting, there is evidence that modulation of efflux transporter function can indeed enhance brain penetration of CNS treatments; however, transporter activity is also impaired in other blood-tissue barriers, hematopoietic cells, and excretory organs.

4.3

Prevention of disease-associated or treatment-associated changes in the blood-brain barrier

The expression of efflux transporters is highly dynamically regulated. This regulatory process can be viewed as a mechanism capable of adapting to changing detoxification and tissue protection requirements. The expression regulation of Pgp has been most extensively studied. Understanding the regulation of BBB efflux transporter activity is of particular interest as it may provide a molecular basis for developing strategies for targeted manipulation of BBB function to improve drug therapy in CNS diseases. This underscores the particular importance of further research focusing on different regulatory mechanisms and their interactions.

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A variety of xenobiotics, including several APIs, have been shown to induce the expression of multidrug transporters. In the treatment of brain tumors, chemotherapeutic drug-induced expression of efflux transporters in tumor cells and BBB endothelial cells is a well-established mechanism that limits drug concentrations in target tumor cells and leads to treatment failure (Lee and Bendayan 2004; Loscher and Potschka ). 2005a). The strong induction of anticancer drugs may be due to their pronounced cytotoxic effect on cells and induction of cellular stress responses. Orphan nuclear receptors are recognized as master regulators of drug-induced changes in the expression of metabolic enzymes and members of the multidrug transporter family (Masuyama et al., 2005). Orphan nuclear receptors PXR/SXR (known as pregnane X receptor [PXR] in rodents and steroid and exogenous receptors [SXR] in humans) have been shown to be expressed in rat brain capillaries ( Bauer et al., 2004). The observation that the PXR ligand dexamethasone increases Pgp expression and Pgp-specific transport demonstrates its functional relevance to regulation of efflux transporters in the BBB (Bauer et al., 2004). Therefore, PXR/SXR may be an important exogenous sensor in brain capillary endothelial cells mediating the induction of Pgp. Targeting these allosensors by administering antagonists has been proposed as a means to overcome therapy-induced resistance mechanisms (Ekins et al., 2007). Ongoing molecular characterization of receptor binding sites has important implications for the future discovery of molecules that are more selective and potent than currently available antagonists. Using colon and lung cancer cell lines, it could be demonstrated that the induction of efflux drug transporters by xenobiotics, especially chemotherapeutic agents, is not necessarily dependent on PXR (Huang et al., 2006). Since the group also showed that the mode of regulation may be cell-specific, it is unclear whether these data can be extrapolated to brain capillary endothelial cells. Baker et al. (2005) reported that epigenetic changes in the MDR1 promoter in response to chemotherapy drugs would then amplify the MDR phenotype. Significant changes in the temporal and spatial patterns of histone modifications occurred in the 5'-hypomethylated region of MDR1, which was directly associated with the upregulation of MDR1 (Baker et al., 2005). Further studies could provide the basis for identifying additional targets to prevent therapy-induced overexpression of transporters. Several CNS pathologies are associated with altered expression or function of efflux transporters. Epilepsy, characterized by recurrent spontaneous seizures, is one of the most common neurological disorders. In animal models of epilepsy, transient increases in Pgp and MRP2 expression in brain capillary endothelial cells, astrocytes, and neurons were observed after seizures, suggesting that seizures themselves can induce overexpression of drug transporters (Loscher and Potschka 2005a; Sisodiya 2003). This seizure-related overexpression was found to be restricted to brain regions involved in seizure onset and spread. These data are consistent with studies of human epileptogenic tissue dissected from drug-resistant patients undergoing epilepsy surgery

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Indicates high expression rates of efflux transporters (Loscher and Potschka 2005a; Sisodiya 2003). However, the lack of sufficient control tissues makes it difficult to draw definitive conclusions from these studies, as successfully treated patients often do not undergo surgical resection of epileptogenic foci. The cellular mechanisms involved in seizure-induced overexpression of efflux transporters remain to be elucidated. Of particular interest regarding the release of excess glutamate associated with seizures, glutamate has been shown to upregulate Pgp expression through an NMDA receptor mechanism (Zhu and Liu 2004). Recently, we were able to demonstrate that extracellular glutamate signaling through NMDA receptors and COX-2 in brain capillaries increases BBB Pgp expression after seizures (Bauer et al., 2008). Consistent with our hypothesis, exposing isolated rodent brain capillaries to glutamate increases Pgp expression and trafficking activity. These increases were blocked by the NMDA receptor antagonist MK-801 and the selective COX-2 inhibitor celecoxib. In rats, intracerebral microinjection of glutamate leads to a local increase in Pgp expression in brain capillaries. Furthermore, using the pilocarpine rat model of status epilepticus, we achieved attenuation of the seizure-induced increase in capillary Pgp expression by administering the nonselective COX inhibitor indomethacin. These data suggest that inhibition of COX could increase brain uptake of antiepileptic drugs and overcome transporter-mediated drug resistance (Bauer et al., 2008). Consistent with seizure-related molecular changes in the BBB, upregulation of Pgp after focal cerebral ischemia has also been described (Spudich et al., 2006). Since enhanced glutamate release is also a hallmark of ischemic brain injury, this induction may also be related to glutamate release and subsequent activation of inflammatory events and prevent the use of the same strategy previously demonstrated in epilepsy models. Further elucidation of the mechanisms involved in transporter regulation in CNS diseases could open up the possibility of new strategies to improve brain penetration of CNS therapeutics. In addition to changes in the receptor or promoter regions involved, multiple mechanisms contribute to the cellular stress response, including phospholipase C, protein kinase C, mitogen-activated protein kinase cascade, mobilization of intracellular Ca2+ , cytokines, nuclear factor kappa B, and heat shock factor 1 regulate multidrug transporter genes such as MDR1 (Ho and Piquette-Miller 2006; McRae et al 2003; Shtil and Azare 2005; Tchenio et al 2006). Felix and Barrand (2002) used primary cultured rat brain endothelial cells to study the effect of oxidative stress on transporter expression and found a stress-related increase in Pgp expression and function, whereas no such changes were observed for MRP1. Hartz et al. (2004, 2006) defined the signaling pathway as part of the innate immune response that rapidly regulates Pgp activity. Their results suggest that the inflammatory cytokine tumor necrosis factor (TNF)-α reduces Pgp activity through activation of the TNF-R1 receptor, endothelin-1 release, and endothelin-B receptor signaling. All of these findings refine our view of regulatory mechanisms and stimulate research efforts to prevent transporter-mediated drug resistance in CNS diseases.

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5 Bypassing the blood-brain barrier 5.1

Nanocarrier Systems and Drug Conjugates

An alternative approach that does not compromise the protective function of efflux transporters is to bypass the transporter molecule. Different strategies have been used here, including encapsulation of nanoparticles (Huwyler et al. 1996; Kreuter 2001) or conjugation of transport substrates (Mazel et al. 2001). Nanoscale carrier systems, including polymers, emulsions, micelles, liposomes, and nanoparticles, can deliver their contents to the brain through passive targeting (de Boer and Gaillard 2007). In these systems, the rate of distribution in the brain is often limited. Permeability generally depends on physicochemical properties and physiological conditions. A variety of compounds, including several anticancer drugs, have been formulated into nanoscale delivery systems. Efficient dosing has been described in rodents. While free doxorubicin had no relevant effects in a rat brain tumor model, liposomal doxorubicin prolonged survival (Sharma et al., 1997). There is evidence that the first dose of doxorubicin promotes subsequent uptake of higher doses in the brain due to toxic effects on proliferating endothelial cells and reduction of the angiogenic factor VEGF (Zhou et al., 2002). Polymer-based particles have been less studied. In most cases, they are formulated by adsorbing the drug onto the particle surface (Kreuter 1995). The particles are then phagocytosed into cells that release the drug. Increased uptake of doxorubicin into the brain was demonstrated when the drug was applied to the surface of solid poly(butyl cyanoacrylate) particles (Gulyaev et al., 1999). In a rat glioblastoma model, 20% of animals achieved long-term remission with this formulation (Steiniger et al., 2004). Following administration of liposomal daunorubicin to patients, the central tumor mass achieved therapeutic drug concentrations (Zucchetti et al., 1999). In clinical studies, the authors considered the response rates to liposomal doxorubicin or daunorubicin to be promising (Hau et al 2004; Koukourakis et al 2000a,b; Lippens 1999). However, the conclusions were complicated by the fact that the patients also received radiation or other chemotherapeutic agents. Therefore, further studies are needed to definitively determine the therapeutic potential. Active drug targeting strategies involve the use of techniques that exploit endogenous transport mechanisms for site-specific delivery (de Boer and Gaillard 2007). Ligands for targeting can be bound to the drug itself or attached to the surface of drug-loaded particles. With respect to the BBB, ligand-mediated site-specific delivery involves receptor-mediated transcellular systems at the BBB to reach extracellular or intracellular targets in the brain. Interestingly, targeting strategies can benefit from pathophysiological mechanisms when the target is induced during a disease process. Gaillard et al. (2005) have identified a novel carrier protein for targeted delivery that utilizes the diphtheria toxin receptor, which is expressed under neuroinflammatory conditions such that B.

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These occur in many brain diseases including Alzheimer's, Parkinson's, multiple sclerosis, ischemia, encephalitis, epilepsy, tumors, etc. mediated transcytosis and brain uptake (Gaillard et al. People, 2005). This delivery strategy can be applied using CRM197-drug conjugates or CRM197-coated drug-loaded liposomes. The most commonly characterized receptor-mediated transcytosis system for CNS targeting is the transferrin receptor (Pardridge 2002). Targeting the drug to this receptor can be achieved by using endogenous ligands or by using antibodies against this receptor (OX-26). The insulin receptor is another receptor-mediated transcytosis system that has been used to target drugs to the central nervous system (Pardridge 2005). Another approach is based on the LRP1 and LRP2 receptors, which are known multiligand scavengers. Several ligands for these receptors, including melanotransferrin, apolipoprotein, and aprotinin, have been used to facilitate brain targeting (de Boer and Gaillard 2007).

5.2

intranasal administration

Administration through the nose has been considered as a possible way to target drugs to the central nervous system (Graff and Pollack 2005). Three routes are generally assumed for drug entry into the nasal cavity. These pathways include direct transport to the brain, eg along the nerve sheath, transport along the axons of neurons and into the blood through the nasal mucosa (Graff and Pollack 2005). The extent to which a molecule traverses these pathways and thus actually reaches the CNS depends critically on its chemical nature and its formulation (Ugwoke et al., 2001). A recent meta-analysis of all published studies claiming evidence of direct nose-to-brain transmission found that only two rat studies provided results that could be considered indicators of direct nose-to-central nervous system transmission (Merkus and van den). Berger 2007). The same analysis found no pharmacokinetic evidence to support the claim that nasal administration in humans enhances drug delivery to brain targets compared with intravenous administration under similar dosage conditions. Therefore, it is currently rather doubtful whether intranasal administration can be considered an effective means of brain targeting.

5.3

intracerebral administration

Typically, significantly higher drug concentrations in the brain can be achieved by direct administration into the CSF or brain parenchyma

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(Huynh et al., 2006). Intraventricular or intrathecal drug infusion delivers drugs into the cerebrospinal fluid (CSF), bypassing the blood-brain barrier and hepatic metabolism. After the drug enters the cerebrospinal fluid, it still needs to pass through the ependymal brain-cerebrospinal fluid barrier. This is possible for many small molecules and lipophilic compounds. However, penetration into the parenchyma is limited due to tortuosity, transcapillary loss, cellular uptake, and binding (de Boer and Gaillard 2007). Therefore, drugs injected into the CSF are less likely to diffuse into the parenchyma. Therefore, intraventricular administration of CNS drugs is thought to be particularly useful in the treatment of meningiomas, since tumor metastasis can be prevented through their effects on CSF-floating cells, but not in gliomas (Huynh et al., 2006). The intrathecal route of administration does not result in drug accumulation in parenchymal structures deep in the brain and is therefore more suitable for the treatment of disseminated meningeal or spinal disorders (Groothuis et al., 2000). Implantable controlled-release systems can be designed from degradable and non-degradable polymers (Sawyer et al., 2006). When properly designed, polymers can provide reliable, sustained release over a period of days to years. Biodegradable polymers are more common than nondegradable systems because the permanence of the delivery system limits clinical use. Controlled-release systems have been used clinically to treat brain tumors (Sawyer et al., 2006). Intracranial implantation of carmustine (BCNU)-loaded wafers in patients with malignant glioma following surgical resection of the tumor was well tolerated and resulted in a modest improvement in patient survival (Brem et al 1995; Engelhard 2000) . A general disadvantage of controlled release systems for different indications is that the local penetration of the drug is limited due to diffusion limitations through the brain parenchyma. Convection-enhanced delivery was developed to transport compounds in the brain, thereby overcoming the diffusion barrier encountered in polymer-controlled release (Bobo et al., 1994). The method is based on continuous infusion of the API through a larger area of tissue using convective propulsion (Huynh et al., 2006). Benefits come from greater distribution and sustained exposure due to longer infusion times compared to boluses in the brain parenchyma (Sawyer et al., 2006). The technology has been used in chemotherapy, gene therapy and immunotherapy. Clinical utility has been demonstrated in glioma patients with tumor recurrence. Patients receive an infusion locally or directly into the tumor after surgical resection. Clinical studies have shown that convection-enhanced delivery is suitable for delivering drugs to large tumor volumes. For example, clinical trials using immunotoxins in glioblastoma tumors have been able to achieve complete regression with minimal systemic toxicity in some patients (Husain et al., 2003). Alternatives to direct delivery include gene therapy using viral, lipid, polymer-based, and cell-based delivery strategies. For details on these methods, the reader is referred to review articles focusing on these techniques (de Boer and Gaillard 2007; Huynh et al. 2006).

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6 Conclusions In recent years, understanding of the impact of brain efflux transporters on the treatment of CNS disorders has increased. Accumulated knowledge of the structure, function, localization, and substrate specificity of brain efflux transporters facilitates the development and validation of strategies to manage the activity of these transporters in the clinical setting. Several strategies for targeting drugs to the brain are already in clinical use, especially for the treatment of brain tumors. Recently, particular attention has been paid to the regulation of transporter expression or function under pathophysiological conditions that contribute to disease development or progression but may also affect the outcome of drug therapy. Further research could provide new ways to prevent the enhancement of BBB function in CNS diseases and thus prevent the development of drug resistance.

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Carriers for Topical Treatments in Dermatology Hans Christian Korting and Monika Schäfer-Korting

content 1 2

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introduce . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Absorption (through) the skin. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Skin morphology and barrier function. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Route of skin penetration. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3 Assessment of (through) skin absorption. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4 Improved (trans)dermal absorption. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Drug carrier - technical aspects. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1 Liposomes. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 Liposomes. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3 Solid lipid nanoparticles and nanostructured lipid carriers. . . . . . . . . . . . . . . . . . . . 3.4 Microemulsions and nanoemulsions. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.5 Polymeric particles, including dendrimers. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.6 Different nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Current therapies for key target indications of API-loaded nanoparticle delivery systems. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Carrier-Loaded API Clinical and Preclinical Data. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.1 Liposomes and liposomes. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.2 Lipid nanoparticles. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.3 Particulates. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.4 Microemulsions. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.5 Polymeric particles, including microparticles and dendrimers. . . . . . . . . . . . . . . . . . . . . .

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H.C. Korting (*) Clinic and Polyclinic for Dermatology and Allergology, Ludwig-Maximilians-University, Frauenlobstraße 9-11, 80337 Munich, Germany Email:[email protected]M. Schafer-Korting (*) Institute for Pharmacy (Pharmacology and Toxicology), Free University of Berlin, Koenigin-Luise-Strasse 2-4, 14195 Berlin, Germany 电子邮件：[email protected]

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6 different methods. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 461 7 Conclusions. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 461 References. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 461

Summary Topical medications are less likely to have serious systemic side effects than systemic medications. Starting from liposomes, various types of nano- and micro-sized drug carriers have been developed to increase the well-known low permeability of drugs to the skin, which limits not only topical but also transdermal treatments of dermatological diseases. Today, formulations based on liposomes and microsponges are approved for the treatment of dermatophytosis, acne, and actinic keratoses. Drug carriers such as lipid nanoparticles, polymer particles, dendrimers, and multishelled nanocarriers with dendritic cores are being investigated. Based on the rapidly increasing research in this field in both academia and industry, a breakthrough seems likely once stability issues (nanoparticles) and safety concerns (dendrimers) are overcome. The technical methods and results of in vitro, ex vivo and in vivo tests are described, taking pharmacokinetic, efficacy and safety aspects into consideration. Key words skin penetration, skin absorption, drug carrier, liposome, solid lipid nanoparticles, microemulsion, polymer particle

缩写词 5-FU API CAT-1 CMS CNT CPA FITC GRAS logP MCT NLC OECD PLA PLGA PASI QSPR SCCP SLN TAC

5-Fluorouracil Active Pharmaceutical Ingredient 4-Trimethyl-ammonium-2,2,6,6-tetramethylpiperidine-1-oxyiodide Multishell core (nanocarrier) (Carbon-) Nanotube Cyproterone acetate Ketofluorescein 5-isothiocyanate generally recognized as safe octanol:water partition coefficient medium-chain triglyceride nanostructured lipid carrier ) Psoriasis Area and Severity Index Quantitative Structure-Permeability Relationship Consumer Products Scientific Council Solid Lipid Nanoparticle Triamcinolone Acetonide

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1 Introduction In the topical treatment of dermatological diseases, the active pharmaceutical ingredient (API) comes into contact with the target site before entering the systemic circulation, and it is usually mandatory for non-target site contact. Therefore, there are fewer systemic side effects compared with parenteral or oral administration. The application of drugs (API preparations) to the skin surface is not only used in the treatment of dermatological diseases, but also in antirheumatic therapy to control the gastrointestinal side effects of NSAIDs. In addition, applying the drug to the skin surface avoids the first-pass effect of hepatic metabolism and the large fluctuations in plasma levels that are characteristic of rapidly clearing drugs after repeated oral administration. Thus, transdermal delivery is becoming increasingly important for systemic treatments, such as the use of drugs with extensive first-pass elimination, such as glyceryl trinitroglycerin and estrogens, and the use of opioids for chronic pain suppression. It was not until the 1960s and 1970s that pioneering research opened up the field of skin pharmacology, especially skin pharmacokinetics, which provides insight into drug effects on skin tissue, target API levels, and systemic absorption of active ingredients and toxicants. Quantitative Insights. Only then will receptor affinity and drug carrier influence on therapeutic efficacy or suitable surrogate parameters [eg skin bleaching for anti-inflammatory effects of glucocorticoids; (McKenzie and Stoughton 1962)] become clear. Since overcoming the stratum corneum (stratum corneum) of the epidermis proved to be the greatest challenge, its structure and methods of overcoming or bypassing the stratum corneum were studied in detail. In fact, early pioneering studies showed that renal hydrocortisone excretion was related to stratum corneum thickness at the target site (Feldmann and Maibach 1969). Later, a shunt pathway for drug penetration through hair follicles was demonstrated (Hueber et al., 1994)—the latter being most relevant for studying percutaneous absorption in fur animals.

2 (Trans)dermal absorption When overcoming the stratum corneum, active substances and toxic substances first reach the viable epidermis (Fig. 1). A concentration gradient from the epidermis to the vascularized dermis then allows the active ingredient to enter the blood vessels.

2.1

Skin Morphology and Barrier Function

The skin consists of the epidermis and the underlying dermis (figure 1). Further down is the subcutaneous tissue, which is rich in triglycerides. Micron-sized hair follicles protruding from the dermis and even subcutaneous tissue

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Figure 1 Human skin morphology (with permission from Mutschler et al. (2008))

The most common skin appendages are those that extend to the surface of the skin and thus penetrate the epidermis. Compared to fur animals, humans have very few hair follicles, with openings covering only about 0.1 percent of the skin's surface. In men, sex differences are small (Jacobi et al. 2005a). The vascularized dermis (1-2 mm thick) is rich in fibers (collagen, elastin) embedded in the matrix surrounding fibroblasts, the most numerous cells in the dermis. The most abundant cells in the nonvascularized epidermis are keratinocytes; cell contacts are formed by desmosomes. Cellular differentiation recognizes four epidermal layers: the basal layer (separated from the dermis by a basement membrane to which keratinocytes contact via hemidesmosomes), the spinous layer, the granular layer, and the uppermost stratum corneum (composed of approximately layer composition) a flat layer of apoptotic keratinocytes, the so-called keratinocytes. The thickness of the stratum corneum is about 10-15 mm, and the thickness of the active epidermis is 50-100 mm. Other cell types are embedded between basal keratinocytes, such as melanin-producing melanocytes and antigen-presenting Langerhans cells. In inflammatory skin conditions such as dermatophyte dermatitis, psoriasis vulgaris, and fungal skin infections, white blood cells invade the skin. Furthermore, in psoriasis, the thickness increases due to increased proliferation and impaired differentiation of keratinocytes. The primary permeability barrier protecting humans from excessive water loss and environmental toxins (and microbes) is the stratum corneum. During differentiation and reaching the surface, keratinocytes break down phospholipids and synthesize ceramides, which are packaged into granules and secreted into the intercellular space, where ceramides combine with cholesterol, long-chain (mainly C22 and C24) free fatty acids Combining sulfate with cholesterol forms a highly ordered lipid layer that acts as a covering. In healthy skin, these lipids are predominantly solid due to orthogonal packing, whereas hexagonal packing and a gel-like structure are characteristic of most lipids

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The superficial layer of the stratum corneum (Bouwstra and Honeywell-Nguyen 2002). At the same stage of differentiation, the nuclei, especially the DNA of keratinocytes, are degraded and the water content of the superficial skin drops below 20%. In the human body, dozens of cells overlap laterally by about 15% and come together to form clusters, the basic unit of the permeability barrier. The connections between the clusters protrude into the underlying skin in small grooves up to a few micrometers in size. Hydrophilic channel-like structures (nanopores) up to 20 nm in diameter have been reported to funnel into the space between keratinocytes (for review see Cevc 2004). The permeability barrier is less effective in neurodermatitis, psoriasis, and other skin conditions. The significantly reduced barrier function in patients with neurodermatitis is due to reduced amounts of ceramides (especially ceramide 1) and a higher proportion of hexagonal lateral packing of epidermal lipids (Choi and Maibach 2005; Bouwstra and Ponec 2006). In addition, there is an irregular pattern of lipid organization and an irregular structure of desmosomal protein particles. Changes in ceramide composition (again notably decreased ceramide-1 levels) as well as increased proliferation and incomplete differentiation of keratinocytes can also be found in psoriatic plaques, leading to the formation of highly irregular stratum corneum (Choi and Maibach 2005). .In fact, the similarities between these disorders are becoming increasingly apparent (Wilsmann-Theis et al., 2008), which dovetail well with their therapeutic overlap. The lipid layer structure is more altered and protein function is disturbed in lamellar ichthyosis, which can be attributed to genetic defects in protein or lipid metabolism (Akiyama and Shimizu 2008).

2.2

skin penetration pathway

In general, drugs penetrate and penetrate the skin mainly through the tortuous intercellular pathways between the corneocytes of the stratum corneum. Therefore, both the physicochemical properties and the anatomical range of the active ingredient are important. Absorption is fairly high on the forehead, lower in the space behind the ears, lower in the abdomen and arms, and lowest in the palms and soles of the feet (Feldmann and Maibach 1969; Rougier et al 1987; Tsai et al 2003). Uptake of fentanyl and sufentanil from the feet, thighs, chest, and abdomen was very similar, and no effects of age or sex were observed (Roy and Flynn 1990). Furthermore, highly lipophilic active ingredients can penetrate the skin through the hair follicle (Ogiso et al., 1996; Munster et al., 2005), and unlike the stratum corneum, the micron size does not preclude penetration into the follicle opening (Hoffman, 1998); Lademann et al. 1999, Otberg et al. 2004, Toll et al. 2004). Indeed, follicular penetration has recently gained increasing attention, especially with the development of novel drug delivery systems. In addition to its function as a barrier to penetration, the stratum corneum is also a reservoir for topically applied substances. API also binds to keratin in keratinocytes

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Because miscibility with lipid domains contributes to reservoir function (Heard et al., 2003). Four days after application of glucocorticoids under occlusion, reocclusion of the treated area releases glucocorticoids from its reservoirs, again inducing skin bleaching (Vickers and Fritsch, 1963). From the reservoir, the active ingredient can diffuse into the viable skin and enter the dermal vessels; alternatively, elimination from the stratum corneum can be achieved by desquamation of the corneocytes without contact with the viable skin.

2.3

Evaluation of (per)dermal absorption

Quantitative structure-permeability relationship (QSPR) methods based on molecular weight and octanol:water distribution have been developed to assess the intrinsic permeability of drugs (for reviews see Moss et al., 2002; Geinoz et al., 2004). These methods have been improved recently (Katritzky et al., 2006; Majumdar et al., 2007; Ottaviani et al., 2007). However, the results may only be valid when applied to drugs whose physicochemical properties are close to those of the training set. In fact, transdermal absorption quantified ex vivo using human and porcine skin and reconstructed epidermis is not relevant to QSPR data when studying drugs with widely varying molecular weights and lipophilicity (Schäfer-Korting et al., 2016). 2008a).. In any case, experimental studies are essential to elucidate the influence of carriers on skin absorption (Schäfer-Korting et al. 2008b). Skin penetration and penetration can be quantified ex vivo on human or animal skin, in vitro on reconstituted human skin, and in vivo (e.g., humans and rats). The experimental approach described here enables an understanding of the predictability of preclinical and clinical data generated when testing topical drug delivery systems. Various methods for assessing (per)dermal absorption have been established. For regulatory toxicology purposes, the OECD has issued in vivo Guidance 427 (OECD 2004a) and in vitro (ex vivo) Guidance 428 (OECD 2004b) based on experiments in rats. The ex vivo technique was tested using a Franz diffusion cell (Bronaugh and Stewart 1985). Based on a strictly controlled protocol, this method was recently validated using human epidermis, pig skin, and reconstructed human epidermis as test matrices (Schäfer-Korting et al., 2006, 2008a). The level of the agent of interest is monitored in the recipient fluid which is separated from the donor vehicle through the skin and applied to the outer surface of the skin. To follow the penetration of active substances into the skin, the active substances were quantified after the skin was removed from the Franz cell (Wagner et al. 2004; Luengo et al. 2006; Stecova et al. 2007; Lombardi Borgia et al. 2008; Schäfer et al. -Korting et al. 2008a). Reasonable agreement between ex vivo and in vivo data generated in humans has been reported (Bronaugh and Franz 1986; Wagner et al. 2002b). Drug release from topical formulations, such as the use of synthetic membranes instead of skin in drug development (Jenning et al. 2000b; Schmook et al.

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2001). While release may herald the vehicle's enhanced skin penetration, it is the skin that reveals the drug's intrinsic penetration capability. In fact, the better skin penetration of ibuprofen than ketoprofen at pH 7.4 was only evident when in vitro (rat) skin was used as the test matrix (Takahashi et al., 2002). However, recent efforts have been made to develop matrix systems (such as PAMPAskin) to quantify transdermal absorption without the need for excised skin or reconstructed epidermis (Ottaviani et al., 2007). For parenterally or orally administered APIs, once a (pseudo)steady-state drug distribution has been achieved, plasma levels reflect the drug concentration at the target site. However, this does not apply to topical treatments. There, the drug passes through the skin of the target organ before entering the bloodstream. Therefore, in clinical trials of topical dermal drug absorption, plasma levels should not be monitored except to investigate systemic exposure and the possibility of systemic side effects. A non-invasive alternative recommended for use in humans is the tape stripping method for quantification of drug levels in the stratum corneum (Weigmann et al. 1999, 2005; Shah 2001; Wagner et al. 2002a; Löffler et al. 2004; Jacobi et al. et al. 2005b) – must be relayed by any compound applied to the skin in order to have a dermal or systemic effect. A transparent and highly flexible adhesive film is applied to stretched skin with standardized pressure and peeled off, removing superficial keratinocytes (Bashir et al 2001; Lademann et al 2005). Since the carrier affects the amount of cuticle removed per strip, the amount of API must correlate with the amount of cuticle recovered (Turner et al. 2005; Jacobi et al. 2005b). In studies of topical glucocorticoids (Pershing et al., 1992; Pelchrzim et al., 2004), drug levels in the stratum corneum have been reported to correlate with dermal availability (Shah et al., 1998) and skin relaxation responses. However, for drugs that target specific living skin layers, peel-off tape may not provide meaningful results. Differential peeling is a procedure in which the first step removes the cuticle by tape peeling and the second step removes the contents of the follicle by cyanoacrylate peeling (Schaefer and Lademann 2001; Teichmann et al. 2005). Thus, differential exfoliation allows selective access to material that has penetrated the follicular opening.

2.4

Improve (per)dermal absorption

In many skin diseases such as atopic eczema and psoriasis vulgaris, a partially disturbed barrier facilitates the passage of xenobiotics through the stratum corneum (Anigbogu et al., 1996). Although this is a major concern when considering the safety of nanotechnology (Stern and McNeil 2008), adequate skin penetration remains an important challenge in the development of topical dermatology. Traditional active ingredient carrier systems, such as creams and ointments, result in an active ingredient absorption rate of only a few percent, which is often associated with high fluctuations in the absorption rate. Thus, API levels in diseased skin may be subtherapeutic in some patients, while causing adverse local or systemic side effects in others.

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To be well absorbed, a substance should have all l l l

The molecular weight is less than 500 g mol 1 , the melting point is low, and it has sufficient solubility in oil and water. The octanol:water partition coefficient (logP) is about 1-3.

Saturated or supersaturated systems promote skin penetration (Moser et al 2001; Barry 2004). In addition, penetration enhancers can improve penetration into the skin. These active ingredients can be classified into the following categories (Thong et al., 2007): l

Lift

Solvents and hydrogen bond acceptors (e.g. dimethylsulfoxide, dimethylformamide), fatty acids and alcohols (lauric acid, oleic acid, ethanol, n-octanol, etc.) and weak surfactants (e.g. azone 1, terpene ene).

Penetration enhancers are present in very high concentrations in microemulsions. Particular attention should be paid to propylene glycol, which is not uncommon in vehicles used as a reference when testing microparticle formulations. The mechanisms of performance enhancement are still only partially understood. Based on chemical structure, interact with the lipids of the stratum corneum, reducing the rigidity of the barrier and favoring the use of nonpolar pathways, or altering protein conformation (protein denaturation), allowing the use of polar signaling pathways. Some enhancers can also work bidirectionally (Thong et al. 2007). Azone causes the formation of an additional liquid phase in the epidermal lipid mixture (Pilgram et al. 2001). The channel-like structures between keratinocyte clusters can be modified by detergents and fatty acids to facilitate drug penetration (Barry 2004). A mixture of penetration enhancers such as sodium laureth sulfate and phenylpiperazine was more effective than the drug alone and allowed even macromolecular drugs such as heparin, luteinizing hormone, and oligonucleotides to penetrate rat skin (Karande et al. people, 2004). In general, detergent-induced changes are especially associated with general skin irritation, leading to irritant dermatitis in severe cases (Fang et al. 2003; Gloor 2004). Risk increases with the effectiveness of penetration enhancers and the use of skin closure to further increase penetration (for review, see Thong et al., 2007).

3 Drug Carriers - Technical Aspects New drug carriers for dermatological diseases are often developed based on nanotechnology principles. However, to increase the loading capacity of APIs, nanocarriers for biomedical purposes have a size of 100–1000 nm (Anton et al.

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2008), exceeding the size of engineered nanomaterials and exotic air pollution components, which by definition are at least an order of magnitude smaller than 100 nm (Stern and McNeil 2008). According to the structure, nanoparticles can be divided into nanocapsules with a typical core-shell structure and uniform nanospheres. Thus, microcapsules and microspheres range in size from >0.1 to 1 mm. Biocompatibility is an important requirement for pharmaceutical use, and it is a challenge to design the formulation of the carrier to match the physicochemical properties of the active substance (e.g. avoid chemical instability) and therapeutic requirements (e.g. avoid chemical instability). B. Improve skin penetration and prolong absorption). The design and production of the various types described below are discussed in detail elsewhere (Anton et al., 2008). If drug approval is sought, the safety of all ingredients must of course be demonstrated for regulatory purposes.

3.1

Liposomes

Depending on their size, liposomes are either nanospheres or microspheres. Small unilamellar vesicles (SUVs) range in size from 20 to approximately 100 nm, large unilamellar vesicles (LUVs) are greater than 100 nm, and multilamellar vesicles (MLVs) exceed 500 nm in size. Vesicles are composed of amphiphilic molecules such as phospholipids that form bilayers in their cavities that contain water and are dispersed in an aqueous medium. The polar headgroups of phospholipids such as lecithin form the interface with aqueous media. Lipophilic drugs are incorporated into the bilayer, and charged hydrophilic drugs reside in the aqueous phase within the vesicle. However, hydrophilic active ingredients are also often present in large amounts outside the liposomes. However, liposomes are metastable systems and their pharmaceutical use is often limited by instability. Instability may be due to vesicle size changes and formation of ester hydrolysis and oxidation products due to vesicle leakage, aggregation or fusion. With liposomes composed of unsaturated phospholipids (liposomes in liquid state). The first liposomal product marketed in 1988 was a formulation of econazole for the topical treatment of dermatophytosis (Fig. 2; Pevaryl1 Lipogel). Recently, liposomes loaded with UV-A and UV-B absorbers (Daylong1 Actinica) have been approved in the EU as a medical device for the prevention of ) of actinic keratoses. For a long time (Ulrich et al 2008). In addition, liposomal diclofenac (Diclac1 Lipogel) is approved for topical treatment of osteoarthritis (Table 1).

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Figure 2 Econazole-loaded liposomes. (a) Freeze-fracture electron micrograph of a commercial-scale batch of econazole liposome dispersion (ELD) (Pevaryl1 Lipogel) six months after production (with permission of Kriftner, 1992). (b) Interaction of ELD with germinal spores of Candida albicans (CA) and reconstructed epidermis (keratinocytes, C). The cell membrane of a CA blastospore is partially detached from the CA cell wall (arrow). The cell membrane of a keratinocyte (C) is impregnated with an electron-dense material (double arrow) representing the applied drug (courtesy Korting et al., 1998)

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Table 1 For topical treatment of skin diseases and transdermal application API vehicle commercial company Indicated product econazole liposome Pevaryl cilag, Dermatomykosen Lipogel schweiz schweiz schweiz spirig, von methoxygene methoxycarbamic acid candy candy schey daylong scheyacticlib katimenticlia kytimectimentiliclia keyatiCeliben keyatiCeliben, German Liposome diclac Hexal, Osteoarthritis Lipogel Tre tinoin microsponges Retin-A OrthoAcne vulgaris Micro Neutrogena, American Fluorouracil Microsponges Carac Sanofi Aventis, Actinic Keratosis USA

3.2

Neosomen

Niosomes are derived from liposomes, which are nonionic surfactant vesicles composed of polyoxyethylene alkyl ethers, polyoxyethylene alkyl esters, or sucrose diesters. Thus, niosomes are also related to micelles, but in a larger size. For example, lipoids capture retinoic acid with high efficiency (>91%) and protect labile drugs from photodegradation (for a review, see Date et al., 2006).

3.3

Solid Lipid Nanoparticles and Nanostructured Lipid Carriers

Solid lipid nanoparticles (SLN) are nanospheres or nanosheets composed of lipids that are solid at room temperature and fatty acid glycerides. Nanodispersions obtained by hot or cold homogenization under high pressure are stabilized by adding surfactants such as poloxamers, polysorbates or sucrose esters. SLN nanodispersions are stable for many years. The generally rather poor loading capacity of SLNs due to limited space for APIs within the lattice is ameliorated by the use of lipid mixtures (e.g., medium-chain triglycerides, MCTs) that are solid and liquid at room temperature; the resulting nanoparticles were Known as nanostructured lipid carriers (NLCs), it is actually an intermediate between SLNs and nanoemulsions (for a review, see Muèller et al., 2007). A detailed physical study of NLC by Compritol/MCT revealed that liquid lipids associate with a solid lipid matrix to form nanospoon-like structures (Jores et al., 2005). Despite the limited loading capacity, which limits the amount of active ingredient relative to the lipid, the encapsulation efficiency can be very high, often exceeding 90%.

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2007); clobetasol propionate logP 3.98 (Hu et al. 2002); prednicarbate logP 4.02 (Santos Maia et al. 2002); econazole logP 5.32 (Sanna et al. 2007)). Recently, a versatile method, paraelectric spectroscopy, was developed to distinguish between APIs attached to particle surfaces and those embedded in the particle matrix (for a review, see Blaschke et al., 2007). The reliability of the results was validated by comparison with independent methods, such as measurements of fluorescence anisotropy (solvochromism) and electron paramagnetic resonance using the lipophilic benzoxazone dye Nile Red (Lombardi Borgia et al. 2005) research (Braem et al. 2007). Detailed insights into carrier and guest interactions can be gained by altering the signals emitted by guest molecules, but this limits the applicability to model agents. Usually, it doesn't apply to APIs. Due to the polymorphic transition of the lipid lattice, drug release can be very rapid (burst release), but sustained drug release is also possible, and the initial burst release can be followed by a delayed release (for a review, see Schäfer-Korting et al. 2007).

3.4

Microemulsions and Nanoemulsions

Both microemulsions and nanoemulsions have droplet diameters less than 100 nm. Microemulsions are clear dispersions of oil and water stabilized by interfacial films of surfactant (and co-surfactant) molecules. Microemulsions form spontaneously without high shear equipment. Active ingredients are dissolved so they penetrate quickly into the skin. Increased thermodynamic activity, the presence of (co)surfactants as penetration enhancers, and occlusiveness improve skin penetration to varying degrees (Santos et al., 2008). In contrast to microemulsions, nanoemulsions are metastable systems whose structure depends on the manufacturing process: spontaneous emulsification or using a high-shear device. Nanoemulsions appear as watery liquids, emulsions, or crystal clear gels.

3.5

Polymer particles including dendrimers

In contrast to the aforementioned carriers (nanoscale or microscale physical aggregates), polymer particles are nanoscale or microscale macromolecules, but they can also form aggregates (Haag and Kratz 2006). Porous polymer systems have an open structure; in principle, capsules and spheres are possible. The large inner surface of nanospheres and microspheres allows high elasticity; both hydrophilic and lipophilic molecules can be charged. Bullets can deliver charged material over extended periods of time with or without burst discharge. Nanocapsules have a core-shell structure: the core acts as a liquid reservoir for active ingredients (mainly oils) and the shell consists of biodegradable substances

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polymer. In addition to high drug encapsulation efficiency and protection of unstable active ingredients from degradation, another advantage is the low polymer content. The particles are composed of biodegradable synthetic materials such as polylactic acid (PLA), poly(D,L-lactic-co-glycolic acid) (PLGA), or polycyanoacrylate (Anton et al., 2008); in addition, Modified natural substances such as chitosan and albumin can also be used. Dendrimers are macromolecules with randomly branched but well-defined dendritic structures that can reversibly bind guest molecules; alternatively, guest molecules can be covalently attached to reactive sites on the dendrimers. These nanoparticles can be engineered with specific core-shell structures for active release of charged drugs via pH-triggered shell cleavage (Haag and Kratz 2006). Guest molecules can be loaded into a hydrophilic core (such as polyglycerol or polyethyleneimine) or into a lipophilic shell formed by covalent bonding of an alkyl ketone. Good encapsulation efficiency requires a minimum core size (about 3,000 g mol 1 ) and a highly branched architecture (Kramer et al., 2002). For details, see R. Haag's chapter in this volume. Microsponges1 are spherical delivery systems made of non-collapsible polymer structures (mainly substituted acrylates or styrene-divinylbenzenes) with large porous surfaces and 5-300 mm in diameter (microspheres: Dimensions below 50 mm, large balls: approx. 100-200). mm), thus absorbing high API loads up to their own weight. Loaded into microsponges, unstable APIs can be protected from environmental influences (Date et al., 2006). Formulations containing fluorouracil (CaracTM) for actinic keratoses and retinoic acid (Retin-A Micro1) for acne are in clinical use in the United States (Table 1).

3.6

various nanoparticles

Titanium dioxide and zinc oxide nanoparticles reflect and scatter UV light; the most effective are titanium dioxide particles with a size of 60–120 nm. Titanium dioxide particles are considered safe (SCCP 2007). Additionally, UV-B and UV-A absorbers, active ingredients in sunscreens, were tested when loaded onto nanoparticles. (Carbon) nanotubes (CNTs) are lightweight, high-strength construction materials; quantum dots consist of a colloidal core surrounded by one or more surface coatings (shells) that reduce metal leaching from the core. So far, these nanoparticles have not been used for biomedical purposes. However, due to their strong photostable fluorescence, quantum dots have potential for use in diagnostics. For reasons of occupational safety, the skin permeability of spherical and ellipsoidal QDs was evaluated; these particles could penetrate the stratum corneum of porcine skin ex vivo but were not detected in receptor media (Ryman-Rasmussen et al., 2006 Year). A similar behavior was also observed for spiked quantum dots. However, contact with a viable epidermis, such as diseased skin, may elicit an inflammatory response in humans due to contact

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Keratinocytes (Zhang et al. 2008). Since nanotubes and quantum dots are not currently used as drug carriers, they will not be considered further here.

4 Current therapies for major target indications of API-loaded nanoparticle delivery systems In dermatology, research into new drug substances and drug delivery systems has focused on common, often difficult-to-treat diseases, especially acne and psoriasis. In severe cases, it is not uncommon to prescribe highly potent active ingredients for systemic use, which can also lead to serious adverse reactions. Acne vulgaris. Almost 85% of young people aged 12-25, even 8% of adults aged 25-34 and 3% of people aged 35-44 suffer from acne. The disorder manifests as comedones when androgens induce sebum production during puberty, and hyperkeratosis of the hair follicles impedes sebum delivery to the skin surface. Propionibacterium and micrococci colonizing the sebaceous ducts break down sebum and induce the formation of inflammatory mediators, leading to the formation of papules and pustules. Treatment options (Table 2) are based on the pathology of the predominant symptom, taking into account its severity - mild, moderate, or severe. Topical treatments (retinoids, benzoyl peroxide, azelaic acid, antibiotics) are indicated for mild and moderate acne, while more severe acne may require a combination of topical and systemic treatments. Effectiveness requires frequent use of anti-acne medications, but side effects often reduce patient compliance. Oral isotretinoin Table 2 API efficacy of anti-acne active ingredients Retinoic acid: tretinoin, isotretinoin, adapalene

Rich in Mild to Moderate Acne: Isotretinoin, Oral for Even Severe Acne

Benzoyl Peroxide for Mild Acne

Azelaic acid is sufficient to treat mild acne. Antibacterial drugs: good to high erythromycin, clindamycin, tetracycline. Antiandrogen: Cyproterone Homoacetate

side effects stimulation

Irritation, burning Well tolerated Induces resistance Antiandrogenic effects Exclude use in male patients

Promising drug delivery systems Liposomes, microsponges, microemulsions, SLN: improve retinoic acid skin penetration and reduce side effects Liposomes, microsponges: improve efficacy and tolerability Liposomes (enhance clindamycin efficacy)

Liposomes, SLN: Improve skin penetration; Liposomes: Low serum levels, effectiveness equivalent to oral administration

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This application is the first choice in the treatment of severe acne, especially in male patients. Long-term remission is possible. However, increased liver enzyme values in blood and plasma lipids and severe skin irritation were dose-limiting side effects. Teratogenic effects caused by retinol are a serious problem for women. For severe acne in women, the antiandrogen/progestin cyproterone acetate (CPA) is an option. This medicine is used by mouth. To avoid the risk of exposing the unborn child to antiandrogens, CPA is given in fixed combination with ethinyl estradiol to inhibit conception. Industrial development of equivalent topical forms of isotretinoin and CPA has failed. Atopic dermatitis and psoriasis vulgaris. Potent anti-inflammatory agents such as glucocorticoids remain a major cause of atopic and contact allergic dermatitis. The main risk is skin atrophy. After an initial reversible thinning that can be detected by skin ultrasound, the procedure can lead to irreversible striations within weeks with potent corticosteroids (Schäfer-Korting et al., 1996). Recovery of skin thickness after effective conventional glucocorticoids takes time (Korting et al., 2002) and may not fully recover before the next exacerbation of dermatitis occurs, requiring further glucocorticoid therapy. Inhibitors of nuclear factor of activated T cells (NFAT), such as tacrolimus and pimecrolimus, are moderately effective alternatives in the treatment of dermatitis, while cyclosporine is used in severe psoriasis and dermatitis. However, cyclosporine is only effective when used systemically. In fact, the high molecular weight of these active ingredients exceeding 800 g mol 1 affects skin penetration. Oral methotrexate is another option for treating refractory psoriasis. However, methotrexate can cause hepatotoxicity and cyclosporine immunosuppression.

5 Clinical and preclinical data on drug loading Advances in drug delivery systems may make the topical use of these drugs safer. Although only limited in vivo data have been reported, a wealth of information is available from ex vivo studies. Due to the importance of intercellular pathways for overcoming the human cuticle, the results of studies in human and porcine skin are superior to studies in fur, both ex vivo and in vivo.

5.1

liposomes and liposomes

To improve skin penetration and targeting, and transdermal therapy, liposomes have been studied since the early 1980s and are the subject of extensive reviews (eg Bouwstra and Honeywell-Nguyen, 2002; Elsayed et al., 2007) . Therefore, we focus on the most important and up-to-date data. basic research. Loaded liposomes can increase skin penetration, while pretreatment of skin with empty liposomes is generally less efficient. while in

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Previously, some authors suggested skin penetration of intact vesicles, but today this is no longer a problem, except for very fluid vesicles. In fact, liposomes tend to fuse on the surface of the skin. When loaded with more flexible liquid liposomes, only band 9 vesicles were detected by freeze-fracture electron microscopy. Thus, rigid gel-state liposomes, which are not detectable even in a large portion of the superficial stratum corneum, result in slower skin penetration compared to liquid liposomes. However, depending on the phospholipids used to make liposomes, significant changes in the structure of the stratum corneum, such as intercellular deposition of liposome contents, could be induced—possibly due to the known lysophospholipid content that disrupts lipid membranes. In general, skin penetration is reduced when applied under occlusion, which promotes vesicle fusion at the skin surface (for review, see Bouwstra and Honeywell-Nguyen 2002). Instead, vesicle size and lamellar structure appear to be secondary. Liposomes not only improve penetration of the stratum corneum, but also target the sebaceous unit (for review, see Elsayed et al. (2007)). Liposomes have also recently gained interest in immunization, since antigen-presenting Langerhans cells of the skin are abundantly located near hair follicles. Highly flexible liposomes (surfactants added to reduce interfacial tension), so-called transfersomes, not only increase skin penetration and penetration of small molecules (glucocorticoids, ketoprofen), but also albumin, etc. Skin penetration and penetration of higher molecular weight active substances and insulin-thus reducing blood glucose levels in mice and humans (Cevc 2004). Although not generally accepted (Bouwstra and Honeywell-Nguyen 2002; Elsayed et al. 2007), small channel-like structures in the skin have been hypothesized to compress intact transferosome 1 (Cevc 2004). Another way to improve skin penetration is to increase the flexibility of the vesicles by adding ethanol; these liposomes are known as ethosomes (Dayan and Touitou 2000; Touitou et al. 2000). The increased impact of API penetration into ethosomes compared to ethanol solutions or liposomes can be explained by the known synergistic effect of increased ethanol penetration and high fluidity of the vesicle membrane (Dubey et al., 2007; Elsayed et al. al, 2007). . ). Most interestingly, ethanol bodies transported the marker rhodamine red into the deeper layers of the skin; maximal rhodamine red staining was obtained at a depth of approximately 130 mm, which corresponds to the superficial dermis. For conventional liposomes, staining is most evident at a depth of 20-60 mm (Dubey et al., 2007). This suggests that the delivery system can also regulate the location of the API in skin tissue. Liposomes can also act as drug depots. In gel-state multilamellar liposomes (composed of hydrogenated lecithin/cholesterol) loaded with the hydrophilic spin probe CAT-1 and injected subcutaneously, 60% of the spin probe remained in the in intact vesicles. In contrast, the signal associated with the spin probe decreased rapidly after injection of CAT-1 solution due to partitioning and bioreduction (Moll et al., 2004). Medicines to treat acne. With regard to drug delivery systems for the treatment of acne, the gold standard retinoic acid is the most studied (Table 2). different

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Animal and human dermal absorption tests have been performed on liposome-encapsulated tretinoin formulations. It is important to also conduct clinical studies of clinical effectiveness and tolerability in acne patients. The resulting parameters were compared with those obtained using conventional formulations (for an overview see Date et al., 2006; Schäfer-Korting et al., 2007). Liposomal formulations have higher topical retinoid concentrations when applied to animal and human skin compared with gels or solutions. Accumulation in the skin was most pronounced with negatively charged liposomes or with the addition of hyaluronic acid as a drug localizer, whereas positively charged vesicles appeared to favor skin penetration. The clinical effect on acne vulgaris was examined in two double-blind studies. Liposomal retinoids with the same efficacy (same active ingredient concentration and reduced active ingredient concentration) were better tolerated than approved gels. In addition, if the same concentration of retinoids is used, the effectiveness can also be improved. Liposomal encapsulation can also increase the effectiveness of benzoyl peroxide for mild acne while reducing irritation that may be caused by hair follicles. Most interestingly, in an open-label clinical study in women with moderate to severe acne, topical liposomal CPA was found to be as effective as oral CPA plus ethinyl estradiol. Both formulations were superior to placebo. When applied topically, plasma concentrations of CPA are only about 10% of those after oral application (for review, see Date et al., 2006). Thus, liposomes have the potential to overcome the well-known poor skin penetration of CPAs, which was actually the reason for the early shelving of topical CPA formulation development. Improvements in CPA penetration after alcohol solutions (Iraji et al 2006) and after exposure to SLN (Stecova et al 2007) have not been described until recently. antifungal medicines. Ringworm is a common fungal infection; dermatophytes, the related pathogen, occur on and in the stratum corneum of bald skin. Although in most cases, incorporation of azole antifungals into traditional semisolid vehicles such as creams is curative, there is room for improvement. Stabilized liposomal formulations of econazole embedded in gel matrices (Table 1; Figure 2A) have been launched in several European countries. In an in vitro model of skin infection caused by C. albicans, econazole liposome gels appeared to be superior to creams of equal strength (Fig. 2B; Korting et al. 1998; Schaller et al. 1999). Furthermore, mycological cure rates for tinea pedis were 80% (liposomal gel) and 73% (econazole cream), but superiority could not be demonstrated statistically (Korting et al., 1997). Glucocorticoids. Transfersomes enhanced the skin penetration and activity of triamcinolone acetonide in standard human volunteer tests, which were skin bleaching and UV-induced erythema suppression tests. Interestingly, transfersome-loaded triamcinolone acetone made skin thinning less severe when glucocorticoid concentrations were reduced (90%) to equivalent concentrations in commercially available creams and ointments (Fig. 3; Fesq et al. ,Year 2003). Reduction in skin thickness persists after liposomal triamcinolone acetonide

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Figure 3 Transfersomes loaded with triamcinolone acetonide (TCA). (a) Activity in UVB erythema test and skin bleaching test. (b) Skin thinning effect (courtesy of Fesq et al., 2003)

The anti-inflammatory effect, if this finding can be confirmed in future studies, opens the door to safer use of effective conventional glucocorticoids in severe chronic dermatitis in the future. Antipsoriatic drugs. Methotrexate has very poor penetration into the skin from aqueous solutions (Weinstein et al., 1989). Liposomes and niosomes may have the potential to enable topical methotrexate therapy, replacing current oral treatments for severe and refractory psoriasis. Methotrexate seeps into the skin

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Improved after application of a 50% propylene glycol solution (to improve drug penetration) and elastic liposomes (Trotta et al., 2004). A recent study of methotrexate-loaded ethanol bodies (69% encapsulation efficiency, stable for at least 120 days) demonstrated increased penetration of methotrexate through human skin (Dubey et al., 2007). Niosomal methotrexate incorporated into chitosan gel was tested on volunteers and psoriasis patients. Efficacy from Psoriasis Area and Severity Index (PASI) reduction was superior to methotrexate gel; no significant irritation and sensitization effects were observed (Lakshmi et al., 2007). antineoplastic drugs. Furthermore, loading niosomes with 5-fluorouracil (5-FU) (consisting of Bola surfactant, Span 801, and cholesterol, with a loading capacity of approximately 40%) resulted in increased penetration of human stratum corneum membranes compared to aqueous 5-fluorouracil solutions. octafold (5-fluorouracil). FU solution and quadruple, when empty niosomes are added to the aqueous solution. Furthermore, 5-FU exhibited increased cytotoxicity against the human melanoma cell line SKMEL-28 as well as the spontaneously transformed keratinocyte line HaCaT (Paolino et al., 2008). This is consistent with significantly increased 5-FU potency and reduced irritative effects when loaded with microsponges (Menter et al., 2008). Transdermal application. As mentioned above, liposomes have been and are often tested for transdermal applications. In a multicenter, randomized, double-blind study of 397 patients with knee osteoarthritis, transfersome-loaded ketoprofen was used daily for six weeks. Placebo and celecoxib 100 mg (recommended daily dose 200 mg) served as reference. Mean reductions in pain scores after ketoprofen (18.2) and celecoxib (20.3) were similar to and better than placebo (9.9). The gastrointestinal effects of the transfer body ketoprofen and placebo were comparable (Rother et al., 2007).

5.2

Lipid-Nanopartikel

basic research. Studies with lipophilic drugs (e.g., retinol, prednicate, CPA) and the lipophilic labeling dye Nile Red loaded with Compritol-based SLNs have demonstrated faster, faster Greater ex vivo penetration of human and porcine skin. In fact, drug levels in the skin may be approximately four times higher than in cream and gel formulations after a few hours (Jenning et al., 2000a; Santos Maia et al., 2002; Lombardi Borgia et al., 2005; Stecova et al. et al., 2007). Later, the difference in skin drug concentrations after SLN and conventional formulations was reduced. The initial increased absorption appears to be due to the sudden release of solid particles following evaporation of water from the skin surface and changes in lipid modification leading to tighter packing of the lipid lattice (Jenning et al., 2000b). In fact, the addition of empty SLN to prednisate cream was not more effective than unmodified cream in permeating prednisate in isolated human skin (Santos Maia et al., 2002).

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The efficiency of SLN and NLC (average diameter: 140–180 nm) in enhancing skin penetration was compared with that of the dye Nile Red embedded in a lipid matrix, where NLC was preferentially present in the liquid lipid phase (Jores et al., 2005; Lombardi Borgia and others, 2005). When loaded with SLN from Compritol or Precirol, Nile Red uptake increased approximately four-fold compared to the uptake after the reference cream. NLC was less efficient when using MCTs and using the penetration enhancer oleic acid as the liquid lipid (Lombardi Borgia et al., 2005). Scanning electron microscopy has shown that SLN applied to the skin surface degrades fairly quickly, which seems to facilitate the contact of the charged drug with the skin surface (Küchler et al., 2008). This action helps to improve skin penetration. Another proposed mechanism is skin closure, which is inversely proportional to particle size and thus more pronounced in nanoparticles than microparticles (Müller et al., 2002). On the other hand, the irritating effect often observed with the use of penetration enhancers can be ruled out, since neither Poloxamer 188 nor the nanoparticle dispersion used for nanoparticle stabilization disturbed phospholipid membranes, as examined using atomic force microscopy (Blaschke et al. et al., 2016). in the news). However, lipid films formed by degraded lipid particles can also delay skin penetration, as observed in porcine epidermis penetration studies with econazole. Load antifungal agents onto Precirol/Tween 80-based SLNs. As the lipid content of the particles embedded in the gel matrix increased, the lag time increased and the cumulative amount of permeated econazole decreased (Sanna et al., 2007). Most importantly, SLN loading not only increases skin penetration, but also induces epidermal drug release, as described for prednicarbate (Santos Maia et al., 2002) and podophyllotoxin (Chen et al., 2006) , opening up the field of vision for more selective drug action and reducing adverse side effects related to the function of the skin and extracutaneous organs. When studying the interaction between a guest molecule and its carrier, epidermal targeting appears to be related to the binding of the guest molecule to the carrier surface (Sivaramakrishnan et al., 2004; Stecova et al., 2007). For guest molecules incorporated into lipid matrices, no targeting was observed (Lombardi Borgia et al., 2005). As described for liposomes, SLNs also have the potential to deliver loaded active substances to hair follicles and sebaceous glands (Münster et al. 2005). Medicines to treat acne. As with liposome loading, retinoic acid-loaded SLNs may be better locally tolerated than retinoic acid gels, as demonstrated in the rabbit Draize test (for review, see Date et al., 2006). In addition, isotretinoin, the retinoic acid isomer, was loaded onto 30 and 50 nm diameter SLNs, composed of Precirol and stabilized by Tween 80 and lecithin. Despite significant accumulation in tissues, no infiltration of isotretinoin into rat skin was observed in SLN-loaded conditions—in contrast to retinoids dissolved in 95% ethanol (Liu et al., 2007 ). This could also offer new ways to treat severe acne. Currently, only oral isotretinoin can permanently suppress severe acne. However, due to the high risk of teratogenic side effects, this should only be used in women who use strict contraception during and for 2 months after using isotretinoin.

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To avoid side effects, topical antiandrogen therapy using lipid nanoparticles may be a future option, provided skin penetration is sufficiently high and systemic absorption is low. When loaded with SLN, NLC, nanoemulsions and microparticles, CPA penetration in ex vivo human skin closely correlated with the effect observed when loading the lipophilic dye Nile Red: SLN quadrupled CPA uptake, NLC and microparticles were more efficient low, and CPA levels in the skin after application of the nanoemulsion did not exceed levels after application of the CPA-loaded cream. Penetration of human skin is related to skin penetration (Stecova et al., 2007). In addition to steroid antiandrogens such as CPA, there are nonsteroidal drugs with high androgen receptor selectivity, such as RU 58841. While RU58841 was unable to load SLNs, it could load the lipophilic myristate prodrug without issue. Interestingly, despite rapid release of active androgens by keratinocytes, fibroblasts, dermal papilla cells, and sebocytes, no prodrug was detected in receptor media following ex vivo application of RU58841 myristic acid to porcine skin or drug (SLN loaded or unloaded). More importantly, the reconstituted human epidermis lacks permeability (Münster et al. 2005), since this in vitro test matrix is hyperactive compared to human skin (Schreiber et al. 2005). Since SLN-loaded Nile Red is able to detect high concentrations of the dye at a depth of approximately 650 mm in hair follicles and sebaceous glands (Münster et al., 2005), SLN has potential not only against acne API but also against androtype at disease sites Active ingredient for hair loss - liposome-encapsulated RU 58841 is also claimed to have this active ingredient (Bernard et al., 1997). Glucocorticoids. Loading of lipophilic prednisolone prednicate diester with SLN (Compritol/Poloxamer 188, mean particle size 144 nm) increased the in vitro absorption through human skin by approximately four-fold after 6 hours, compared to commercial milk with steroids. cream, because both SLN dispersions had been applied or creams containing prednicarbate SLN. After 24 hours, the difference in epidermal prednicarate concentrations decreased after SLN and conventional formulations. Interestingly, prednicarbate targets the epidermis (Figure 4; Santos Maia et al., 2002). This was not observed for example with betamethasone 17-valerate, which binds poorly and causes systemic overload at relevant concentrations (Sivaramakrishnan et al., 2004). Although prednicarbate-SLN has only been tested in vitro to date, a clinical study using clobetasol propionate-loaded SLN in patients with dermatitis showed that the corresponding glucocorticoid The expected efficacy of treatment has improved (Kalariya et al., 2005). Whether the skin atrophy potential of glucocorticoid-loaded SLNs is reduced, as observed with the delivery agent triamcinolone acetonide (Fesq et al., 2003), remains to be demonstrated. Despite an improved risk-benefit ratio compared to equivalent glucocorticoids, prednisate in traditional formulations is not completely free from skin-thinning effects (Schäfer-Korting et al 1993; Korting et al 2002). Antivirus system. Podophyllotoxins are the standard drug for the treatment of genital warts; their development is caused by human papillomavirus infection of epithelial cells. Since podophyllotoxin can cause severe skin irritation, a well-tolerated alternative was sought by including SLN in the active ingredient. Podophyllotoxin was loaded onto tripalmitin-based SLNs stabilized by lecithin and poloxamer 188 (36-208 nm in diameter).

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Figure 4 (a) Absorption and epidermal targeting (0-100 mm skin depth) of prednicate loaded in SLNs (black bars) was improved compared to the corresponding cream (grey bars). The formulation was applied ex vivo to human skin for 6 hours. (b) Addition of empty SLN to prednisylate cream (white column) had no effect on skin penetration and distribution of prednisylate compared to PC cream (with permission from Santos Maia et al., 2002).

Tween 80 (bimodal particle size distribution with peaks at 44 and 194 nm). No podophyllotoxin penetration was observed in isolated porcine skin using either of the SLN dispersions compared to the ethanol solution used as a reference. However, in skin tissue, podophyllotoxin concentrations were similar after ethanol solutions and Tween-stabilized SLN dispersions; the concentrations were even nearly four-fold higher after lecithin/poloxamer SLN dispersions. Due to the abundant detection of podophyllotoxin-associated fluorescence near the epidermis and hair follicles, the authors concluded that both transepidermal and follicular uptake are relevant (Chen et al., 2006). These results may facilitate systematic studies of the effect of particle structure and the development of delivery systems for topical treatments such as herpes virus infections. Especially topical treatment of cold sores still needs to improve the cure rate, while topical treatment of genital herpes is not yet possible. Anti-UV. The potential protective effects of unloaded and loaded SLNs against sunburn and skin cancer were evaluated. Interestingly, cetyl palmitate nanodispersion acts both as a particulate UV blocker itself and as a vehicle for the sustained release of UV absorbing 2-hydroxy-4-methoxybenzophenone. This triples the UV protection compared to two separate components (Müller et al. 2002). UV protection was recently investigated by testing 3,4,5-trimethoxybenzoylchitin-loaded SLNs; efficacy increased with the addition of tocopherol (Song and Liu 2005). Transdermal application system. As described for liposomes, SLN loading can also delay API release. This has been described with flurbiprofen-loaded SLN composed of stearic acid/cholesterol stabilized with different concentrations of lecithin/poloxamer 188

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relationship; hernias averaged 72–93%. While pulsed release was observed with the SLN dispersion, the release was sustained when the particles were embedded in the polyacrylamide gel. In an acute carrageenan-induced paw edema model in rats, the strength of flurbiprofen-induced edema inhibition was in the order: SLN embedded in gel matrix > SLN dispersion > flurbiprofen dissolved in pH 7.4 saline ( Jain et al., 2005). Rat paw edema - acute carrageenan-induced and persistent Freund's adjuvant-induced - was also tested with tristearin stabilized with lecithin/polyethylene glycol (400) monostearate (mean diameter 123 nm) SLN of glycerides, in which triptolide is a diterpenoid triepoxide. Compared with the solution (propylene glycol/water), the penetration of triptolide into rat skin was increased 3.4-fold with the optimized SLN dispersion; no lag phase was observed. The microemulsion tested in parallel released triptolide more efficiently (seven times than the solution) - albeit after a delay period of about 5 hours. The rapid penetration of SLN after dispersion is very consistent with the most effective inhibition of acute edema and the highest absorption of microemulsions, with a very strong effect on persistent edema (Mei et al., 2003). safety aspect. The lipids used to produce SLNs generally have a GRAS status (generally recognized as safe) (Müller et al., 2007), but this describes a safety profile for oral exposure, with less relevance for dermal application. SLN appears to be well tolerated in various cell lines (e.g. HL60) and primary cells (mouse peritoneal macrophages, human keratinocytes, and fibroblasts) (Müller et al. 1997; Scholer et al. 2001, 2002; Santos Maia et al. 2002; Weyenberg et al. 2007). SLN has been reported to be less cytotoxic than polycyanoacrylate and PLGA nanoparticles (Müller et al., 1997). However, addition of cationic surfactants such as stearylamine or dimethyldicoctadecylammonium bromide for stabilization reduces the viability of macrophages exposed to SLNs (Weyenberg et al., 2007). Application of SLN dispersions to reconstructed human epidermis has been shown to be well tolerated (Santos Maia et al., 2002; Küchler et al., 2008), which can be well explained by the barrier function of the construct, This is absent in monolayer cultures. Since reconstituted human epidermis has a lower barrier function to the test formulation than normal human skin (Schäfer-Korting et al. 2008a), it should be well tolerated in humans. However, it must be remembered that topical skin care is developed for skin conditions and then the function of the barrier may be greatly reduced.

5.3

microparticles

In addition to liposomes, microsponges containing anti-acne drugs have also been investigated. Loading microsponges with retinoic acid or benzoyl peroxide increases the local tolerance of these actives while positively affecting efficacy, also against P. acnes (for an overview see Date et al., 2006). Recently introduced microsponge delivery system

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Sustained release of benzoyl peroxide has also been described (Jelvehgari et al., 2006), but the results of in vitro and in vivo tests are still pending. In an open-label study of 28 patients, microsponges loaded with 4% hydroquinone and 0.15% retinol for sustained drug release resulted in a well-tolerated formulation that improved post-inflammatory pigmentation in melasma Hyperpigmentation and hyperpigmentation (Grimes 2004). In addition, microparticle formulations that form a UV protective film on the skin appear to be useful in sunscreens (Lademann et al., 2004). A growing global health burden is actinic keratosis, which can progress to squamous cell carcinoma. Topical 5-FU 5% cream is currently one of the first-line treatment options, but the local tolerance is poor. 0.5% 5-FU (Table 1) loaded onto microsponges was found to be almost as effective. According to a recent study of 356 patients with actinic keratoses, a formulation based on 5-FU microsponges proved to be more effective than the corresponding vehicle after only one week of use. The therapy was well tolerated, with no patients discontinuing treatment due to side effects (Menter et al., 2008). The skin can also be affected by the side effects of systemically applied drugs. A dose-limiting effect, such as in doxorubicin chemotherapy, is palmoplantar erythema, which is caused by the active ingredient being excreted in sweat and then diffusing to the skin surface and penetrating the stratum corneum. Since there are obvious reservoirs in the thick cuticles of the palms and soles of the feet, large amounts of free radicals are produced that can seriously damage the skin. To avoid this side effect by using an effective protective cream, microparticles 10-100 mm in size consisting of a mixture of Hippospongia communis and silicic acid are loaded with an antioxidant mixture and incorporated into the cream carrier. The formula forms an even film on the skin's surface for 6 hours, even in deep grooves and wrinkles. The openings of hair follicles and sweat glands are also covered. This barrier impairs the penetration of doxorubicin, while the antioxidants loaded on the particles impair free radical production. In contrast, the reference microparticle-free cream did not protect the follicle opening (Lademann et al., 2008). This study points to a new protective strategy against skin side effects of systemically applied anticancer drugs. Tristearin lipid microparticles stabilized with hydrogenated phosphatidylcholine (particle size 10-40 mm) and incorporated into oil-in-water emulsions enhanced UV-A absorber butylmethoxyl compared to oil-in-water. The penetration of dibenzoylmethane in the stratum corneum is only slightly emulsion. Water emulsion when applied to human volunteers (Iannuccelli et al., 2008). This is consistent with previous findings (Stecova et al., 2007; Küchler et al., 2008).

5.4

micro emulsion

The solubility potential of the microemulsion facilitates skin penetration by increasing the concentration of lipophilic active ingredients, thus increasing the concentration

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gradient against tissue. The main site of solubilization is the lipophilic part of the micellar surfactant membrane. Furthermore, the generally very low interfacial tension enables good contact to the entire field of application, including e.g. B. wrinkles and microcracks. In addition, isopropyl palmitate and surfactants can enhance permeability by disrupting stratum corneum lipid organization or increasing the solubility of the API in the skin after entering the dermis (for review, see Kreilgaard 2002; Santos et al. 2008). basic research. Methyl nicotinate applied in an isopropyl palmitate microemulsion gel rapidly caused intense erythema. However, this effect wears off quickly, whereas liposomes, which do not damage the cuticle and form a reservoir in the cuticle, give a longer-lasting effect. Water-in-oil (w/o) microemulsions and liposomes (from hydrogenated lecithin/cholesterol) accelerated the penetration of benzyl nicotinate into mouse skin compared to hydrogels by providing better skin contact. However, the effect of skin oxygenation due to vasodilation was not affected (for review see Kreilgaard 2002). In contrast, microemulsion and liposome formulations accelerated butyl nicotinate-induced vasodilation in mouse skin over reference hydrogels. A further improvement was achieved by the synergistic combination of the penetration enhancer N-lauroyl sarcosine/sorbitan monolaurate added to the microemulsion, as increased stratum corneum lipid fluidity facilitated drug penetration ( Abramovic et al., 2008). Combined penetration enhancers have been reported to be generally well tolerated (Karande et al., 2004). However, they can cause some skin irritation (Abramovic et al., 2008). Medicines for skin diseases. In addition to liposomes and microsponges, microemulsions containing anti-acne drugs have also been studied (for an overview see Date et al., 2006). Incorporation of retinoic acid into microemulsions improves skin levels while reducing penetration from isolated porcine skin. Furthermore, erythema formation was less severe in the Draize test in rabbits. The penetration of nonanediol through mouse skin was increased, further enhanced by the addition of the penetration enhancer dimethyl sulfoxide. However, skin irritation from microemulsions has been observed. Although skin penetration of hydrocortisone was improved compared to amphiphilic creams, skin discoloration and increased blood flow due to skin irritation were counteracted (Lehmann et al., 2001). Uptake of methoxsalen increased eightfold compared to water; efficacy was related to the level of penetration enhancer. Furthermore, skin penetration of methotrexate was increased tenfold in lecithin-water-propylene glycol-decanol-benzyl alcohol microemulsions compared to water-propylene glycol solutions (Kreilgaard 2002). Transdermal application. To improve osteoarthritis treatment, microemulsions were developed that increased the penetration of diclofenac, indomethacin, ketoprofen, celecoxib, and rofecoxib into human skin three to eightfold. The anti-inflammatory effects of celecoxib and rofecoxib are faster than those of conventional active substances (for an overview see Kreilgaard 2002; Santos et al 2008). Microemulsion loading also improved the penetration of triptolide into rat skin and efficacy on long-term rat paw edema (see above). This is consistent with the high levels of surfactants and co-surfactants used to prepare the microemulsions. No erythema was noted at the treated site (Mei et al., 2003).

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Polymeric particles, including microparticles and dendrimers

Compared to the studies of liposomes and lipid nanoparticles, polymer-based particles have been considerably studied. However, these studies, especially when using biostable particles, can provide insight into skin penetration processes and thus have important implications for understanding the general function of nanoparticle drug delivery systems. Biodegradable poly(ε-caprolactone) nanoparticles increase skin absorption of lipophilic agents such as the UV protector methoxycinnamate (Alvarez-Roman et al., 2001) and the dye Nile Red (Alvarez- Roman et al., 2004b), while flufenamic acid reduced uptake when loaded onto PLGA nanoparticles (Luengo et al., 2006). In fact, flufenamic acid often penetrates human skin effectively, even when applied in aqueous solution (Schäfer-Korting et al., 2008a). Loading biostable poly-n-butylcyanoacrylate nanocapsules (188 nm in diameter) with indomethacin (77% entrapment efficiency) enhanced drug release. The nanodispersion enhanced the penetration of indomethacin into the skin by almost 10 times that achieved with the indomethacin-Pluronic1 gel, and when the nanocapsules were incorporated into the Pluronic gel, the absorption was increased by almost 4 times . Nanoparticles were detected within the stratum corneum by parallel dye labeling (Miyazaki et al. 2003). Interestingly, for the first time the penetration of Nile red bark was compared with core multishell nanocarriers (CMS) synthesized using dendrimer technology (Kra¨mer et al., 2002), which formed aggregates with a size of approximately 40 nm , and SLNs (with an average size of 180 nm). Clear advantages of CMS nanotransporters (Küchler et al. 2008). Whether this is due to differences in particle size or due to the material making up the particles remains to be clarified. The mechanism of penetration enhancement is best assessed by covalent attachment of dyes to biostable particles. Nanoparticles labeled with fluorescein 5-isothiocyanate (FITC) localize preferentially at the follicle opening. Particle accumulation increases over time and is most pronounced for smaller particles (Alvarez-Roman et al. 2004a; Luengo et al. 2006). In addition, a series of studies by Lademann's group showed that carriers as large as 3 mm can deliver charged drugs into open hair follicles. The optimal size is around 300-750 nm (Lademann et al., 1999; Schaefer and Lademann, 2001; Toll et al., 2004; Teichmann et al., 2005), as the particles use a pumping mechanism, depending on the hair The movement and jagged structure of the hair surface is caused by the desquamation of keratinocytes. While spontaneous hair movement transports particles deep into the follicle in vivo, in ex vivo studies the hair must be massaged to move (Teichmann et al. 2006). Due to low sebum flow, particles are stored in the follicle for about 10 days, whereas free actives are stored for only 4 days (Lademann et al., 2007). Perhaps this will enable targeted use in specific areas of skin treatment in the future. Fluorescein-labeled nanoparticles also accumulated in skin ridges, but penetration of intact particles was not observed (Alvarez-Roman et al. 2004b; Luengo et al.

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2006). Thus, unlike transfersomes, stable particles with a diameter of 20 nm cannot squeeze through the tubular structure of the stratum corneum and must release a charged API in order to penetrate the stratum corneum. safety aspect. While dendrimers, especially those with an amine surface, can be quite toxic, a polyethylene shell surrounding the dendritic core reduces toxicity, which is said to also reduce the tendency of nanoparticles to aggregate (Stern and McNeil 2008 ). This approach was taken in the development of CMS nanocarriers, which actually show very low toxicity to keratinocytes (Küchler et al., 2008).

6 different methods Iontophoresis and microneedling improve penetration of ionized active ingredients into intact skin. These procedures are described elsewhere in this guide. Stimulating results have been reported in relation to iontophoresis for the treatment of eg. B. Acne scars (tretinoin gel) and hypertrophic scars have been reported (for review see Patravale and Mandawgade 2008).

7 Conclusions: Active ingredient loading of nanoparticles, microparticles and microemulsions for dermal and transdermal applications can alter skin penetration. Depending on the type of active substance and formulation, the permeability can be increased or decreased. Increasing absorption often goes hand-in-hand with increasing efficiency. If the drug concentration is adjusted, local tolerance can be improved. Currently, only a few drugs based on nanoscale or microscale application systems are approved for topical use and brought to the market. However, it should be kept in mind that progress in the field of transdermal therapeutic systems has also been rather slow due to technical challenges and high standards of drug development to ensure efficacy and safety.

References Abramovic Z, Sustarsic U, Teskac K, Sentjurc M, Kristl J (2008) Effect of a nanoscale delivery system containing benzyl nicotinate and a penetration enhancer on skin oxygenation. Int J Pharm 359:220–227 Akiyama M, Shimizu H (2008) Recent advances in the molecular aspects of nonsyndromic ichthyosis. Exp Dermatol 17:373-382

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Yaskawa Tsutomu and Ogura Yuichiro's Treatment of Blindness

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introduce . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 470 Target vitreoretinal disease. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 473 2.1 Cytomegalovirus (CMV) retinitis. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 474 2.2 Non-infectious uveitis. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 475 2.3 Macular edema. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 475 2.4 Retinitis pigmentosa. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 476 2.5 Age-related macular degeneration (AMD). . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 477 2.6 Proliferative vitreoretinopathy (PVR). . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 478 3 Non-biodegradable device of. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 479 3.1 Vitrasert1: Non-biodegradable implant containing ganciclovir. . . . . . . . . . . . . . . . . . . . . . . . . 479 3.2 Retisert1: Non-biodegradable implant containing fluocinolone. . . . . . . . . . . . . . 481 3.3 I-vationTM: Non-biodegradable implant containing triamcinolone acetonide. . . . . . . . . . . . 482 3.4 Medidur1: Non-biodegradable, used with fluocinolone. . . . . . . . . . . . . . . . 482 4 Biodegradable equipment. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 483 4.1 Posurdex1: Dexamethasone-containing biodegradable insert. . . . . . . . . . . . . . . . . . . . . . . . . . . 483 4.2 Injectable microspheres. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 484 5 Triamcinolone acetonide crystal suspension. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 484 6 NT-501: encapsulated cells Technology (ECT). . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 485 7 Conclusions. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 485 References. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 486

Abstract: The development of intraocular drug delivery systems (DDS) is urgently needed for the treatment of intraocular diseases, especially in the posterior segment of the eye (vitreous cavity, retina, and choroid), most of which are unresponsive to conventional pharmacological approaches. T. Yasukawa (*) Department of Ophthalmology and Visual Science Nagoya City University Graduate School of Medical Sciences, 1 Kawasumi, Mizuho-cho, Mizuho-ku, Nagoya Aichi 467 -8601, Japan Email:[email protected]

M. Schäfer-Korting (Hrsg.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_16, # Springer-Verlag Berlin Heidelberg 2010

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Eye. Repeated intravitreal injections of anti-angiogenic drugs are effective in the treatment of age-related macular degeneration, but repeated injections still have the risk of serious side effects such as endophthalmitis. Intraocular DDS can solve these problems. Intraocular sustained drug release from implantable or injectable devices has been investigated for the treatment of vitreoretinal diseases. A nonbiodegradable reservoir implant was first marketed in 1996 for the treatment of cytomegalovirus retinitis secondary to acquired immunodeficiency syndrome. Clinical trials of various effective devices followed to treat other difficult eye diseases. An injectable stick insert that releases steroids is currently in Phase III study for the treatment of macular edema secondary to diabetic retinopathy or retinal vein occlusion. Therefore, in the near future, various types of intraocular DDS will be available for the treatment of visually impaired intraocular diseases. Key words age-related macular degeneration, biodegradable polymer, controlled release intraocular drug delivery system, microsphere

Abbreviations AIDS AMD API CMV CNV DDS ECT EVA HAART PLGA PVA PVR RPE VEGF

Acquired Immunodeficiency Syndrome Age-Related Macular Degeneration Agents Cytomegalovirus Choroidal Neovascularization Drug Delivery System Encapsulation Cell Technology Ethylene Vinyl Acetate Highly Active Antiretroviral Therapy Poly(lactic-glycolic acid) Poly(vinyl alcohol) Proliferative Vitreoretinal retinal pigment epithelium vascular endothelial growth factor

1 Introduction The eye is a photosensitive organ with a specific structure similar to that of a camera. The eye should not only maintain the transparency of the ocular medium, but also its configuration to stabilize refractory structures (e.g. cornea, etc.).

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lens). To achieve this, the most efficient shape is a balloon-like structure, which has few blood vessels, cellular components, and clear liquid inside. Aqueous humor, the circulating fluid of the eye, is in good condition and contains sufficient non-vascular structural nutrients and strict limits of macromolecules, such as proteins and lipids, for homeostasis and clarity. This forms the blood-water barrier, which consists of the unpigmented ciliary epithelium and the iris vascular endothelium with tight junctions. The outer and inner blood-retinal barriers formed by the retinal pigment epithelium (RPE) and retinal vascular endothelium regulate retinal homeostasis (Maurice and Mishima 1984; Geroski and Edelhauser 2001; Ambati and Adamis 2002; Yasukawa et al. 2004, 2006, 2007). In this way, the intraocular space is isolated from the systemic circulation (Fig. 1), so systemically administered drugs cannot easily enter the intraocular space (Macular Degeneration Research Group Pharmacological Therapeutics 1997; Ip and Gorin 1996). Therefore, to be effective, high doses of the drug must be used, often causing side effects in other healthy tissues. On the other hand, the eye is covered with collagen layers (such as cornea and sclera) and epithelial and endothelial barriers (such as cornea and RPE). These barriers, continuous tear production, forward flow of aqueous humor, and choroidal circulation limit the penetration of topically applied drugs (eg, eye drops and ointments) (Lang 1995; Kim et al 2007) (Figure 1). Therefore, eye drops must be instilled frequently or in high concentrations to achieve therapeutic concentrations also in the anterior segment. Because of the longer time, it is much more difficult to deliver the drug to the posterior segment

Figure 1 Barriers limiting drug delivery to the posterior segment of the eye. Cornea, tear drainage, episcleral blood flow, and counterrotating intraocular convection limit delivery of topically administered drugs to the posterior segment of the eye. Drugs administered systemically do not readily enter the vitreous cavity and retina due to the blood-retinal barrier

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Diffusion pathways from the ciliary body to the Schlemm's canal and counter-rotating intraocular convection. In light of these concerns, recent treatments for vitreoretinal diseases such as exudative (wet) age-related macular degeneration (AMD) and macular edema have been administered with periocular or intraocular injections (Ip et al., 2004; Gillies et al., 2006); Rosenfeld et al. 2006; Brown et al. 2006; Augustin and Schmidt-Erfurth 2006). However, in the vitreous cavity, most low-molecular-weight drugs have a half-life of only a few hours, and even large molecules such as antibodies have a half-life of several days (Bakri et al. 2007a,b). Therefore, repeated intravitreal injections may be required, which may be associated with side effects such as cataract formation, vitreous hemorrhage, endophthalmitis, and retinal detachment (Cantrill et al., 1989; Cocherau-Massin et al., 1991; Heinemann, 1989; Ussery et al., 1988) . ). The drug delivery system (DDS) can overcome the above-mentioned problems of pharmaceutical methods. With regard to DDS in the anterior segment, several inserts intended to replace eye drops are already on the market: Ocusert1 Pilo (controlled-release pilocarpine, Alza Co., Palo Alto, CA), Mydriasert1 (IOLTech, La Rochelle, France), and Lacrisert1 (hydroxypropyl pilocarpine). Cellulose-based Eye Insert, Merck & CO., Inc., Whitehouse Station, NJ) (Fig. 2). The Ocusert1 pilo insert, introduced in 1974, contains a core reservoir composed of pilocarpine and alginic acid. the core is surrounded

Figure 2 Controlled release system. Ocusert1, Mydriasert1, Vitrasert1 and Retisert1 are clinically available, non-biodegradable (reservoir) devices. Lacrisert1 is a biodegradable (monolithic) insert in the conjunctival sac. *Commercial medical equipment. **Products in clinical trials

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A hydrophobic ethylene vinyl acetate (EVA) copolymer film that controls the diffusion of pilocarpine, an active pharmaceutical ingredient (API), from the insert into the eye. Pilocarpine can reduce intraocular pressure in glaucoma patients. While commonly used eye drops are supposed to be administered four times a day, the insert releases pilocarpine for a week after it comes into contact with the conjunctival surface. Mydriasert1 is an insoluble matrix posterior ocular insert containing phenylephrine and tropicamide as active ingredients to achieve sustained mydriasis during surgery or fundus examination. Lacrisert1 is biodegradable and is inserted daily into the conjunctival sac in place of eye drops to treat dry eye. These types of inserts can be easily removed when side effects occur. However, these inserts do not offer any significant advantages over the use of traditional eye drops. On the other hand, developing DDS in the posterior segment of the eye is certainly beneficial, but is more complicated and difficult. Implants should be placed under the tendon of the eye, within the sclera, or within the vitreous. In this case, any mechanical effects of the remaining matrix and any pharmacological effects of the drug used should be considered. In this section, we discuss medical devices that are clinically available, clinically trialed, or preclinically tested for the treatment of posterior segment disorders.

2 Targeted vitreoretinal disease Vision loss slightly affects quality of life. AMD, diabetic retinopathy and retinitis pigmentosa, and glaucoma are the leading causes of blindness in developed countries, while cataracts and corneal disease continue to pose a threat to vision in developing countries. Retinal diseases are among the most challenging in the body and include AMD, diabetic retinopathy, retinitis pigmentosa and macular edema secondary to diabetic retinopathy, uveitis, pseudophakia and retinal vein occlusion. Some eyes with glaucoma do not respond to medicines and surgery. In this condition, retinal ganglion cells are damaged, leading to progressive visual impairment. Consequently, most difficult-to-treat eye diseases are associated with altered retinal function. Since the retina is the neurosensory organ responsible for binocular coordination, the retina, especially its central macula, requires homeostasis in terms of (1) structure and position, (2) transparency, and (3) neurophysiology. AMD and macular dystrophies often result in irreversible damage to the neurophysiological function of the macula. Macular edema affects not only the structure of the retina, but also its transparency and physiological function. So far, advances in surgical techniques have even allowed special instruments to peel off the innermost lining of the retina, which has recently greatly improved vision in macular holes. However, most other macular diseases remain difficult to treat. Due to the difficulties of pharmacological approaches described above, many vitreoretinal diseases are resistant to the effects of current drugs. for this reason,

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DDS may be required to treat various intraocular disorders. Potentially, controlled (delayed) release intraocular systems could be used in diseases where repeated topical administration may be effective, such as: B. wet AMD, macular edema, cytomegalovirus (CMV) retinitis and uveitis. In addition, additional implantation of a controlled-release intraocular device during surgery may improve recovery for conditions that initially required vitreoretinal surgery for which complications or recurrence may occur (eg, proliferative vitreoretinopathy (PVR), choroidal neovascularization formation (CNV) and diabetic retinopathy). .In addition, controlled release systems may be required for the treatment of chronic diseases for which there is no satisfactory therapy, such as: B. Geographic atrophy (dry AMD), macular edema, and retinitis pigmentosa.

2.1

Cytomegalovirus (CMV) retinitis

CMV retinitis was once a matter of public concern because it occurs in approximately 25% of patients with acquired immunodeficiency syndrome (AIDS) and is the leading cause of blindness in terminally ill patients (Gross et al. 1990). In this context, there is an urgent need to develop a new treatment modality. Preserving vision requires early diagnosis and effective treatment. Systemic administration of ganciclovir or foscarnet was initially shown to be effective in slowing the progression of CMV retinitis. However, these drugs often have serious side effects: ganciclovir caused myelosuppression and foscarnet renal dysfunction, which may require discontinuation of treatment (Henderly et al 1987; Holland et al 1986; Jabs et al 1987). Therefore, intravitreal administration of ganciclovir was followed but had to be repeated to maintain intraocular drug concentrations within the therapeutic range, which may be associated with inherent risks of retinal detachment, cataracts, vitreous hemorrhage, and endophthalmitis (Cantrill et al. 1989; Cocherau-Massin et al. 1991; Heinemann 1989; Ussery et al. 1988). The urgent need to preserve the quality of life of AIDS patients focused on the development of novel intraocular DDSs, leading to the first commercial application of a controlled-release system of ganciclovir with a non-biodegradable polymer device (Vitrasert1) (Sanborn et al., 2008). 1992). . This has been clinically proven to be effective and biocompatible, although surgery is required to remove depleted units or to repeat implantation (Morley et al. 1995; Sanborn et al. 1992; Smith et al. 1992). Recently, highly active antiretroviral therapy (HAART), a combination of reverse transcriptase-inhibiting nucleosides and human immunodeficiency virus type 1-specific protease inhibitors, has been identified for normalizing immunity and promoting several opportunistic infections including CMV retinitis (Autran et al 1997; Mitchell et al 1999; Vrabec et al 1998). Thus, HAART significantly reduces the need for intraocular devices in AIDS patients. Nonetheless, many scientists have made significant discoveries in the field of intraocular DDS and now continue to develop new devices to treat a variety of other difficult vitreoretinal diseases.

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noninfectious uveitis

The uvea is the intraocular tissue containing melanin granules involving the iris, ciliary body, and choroid. Uveitis is an autoimmune or inflammatory eye disease that occurs in the uvea and adjacent tissues such as the sclera and retina. Uveitis has acute or chronic symptoms affecting localized or diffuse areas of the eye, with the potential for recurrence. The disease often requires long-term medical treatment with steroids, immunosuppressants, antibiotics, or combinations thereof, to suppress inflammation or, in individual cases, to prevent recurrence. Persistent anterior segment inflammation can sometimes lead to posterior iris synechiae that limit iris dilatation, secondary glaucoma with or without peripheral anterior iris synechia and cataract formation, and chronic posterior segment inflammation can sometimes lead to blurred vitreous humor leads, macular edema, effusion Ischemic or ischemic retinal vascular disease and other retinal dysfunction. The need for long-term medication led to the development of a controlled-release system for fluocinolone using a nonbiodegradable polymer device (Retisert1), which was approved by the US FDA in 2005 (Jaffe et al., 2000).

2.3

macular edema

When macular edema persists, it often results in irreversible vision loss and is associated with a variety of conditions, including diabetic retinopathy, uveitis, and retinal vein occlusion (Figure 3). The disease is characterized by leakage of serum from retinal capillaries followed by swelling of the macula in the center of the retina. Diabetic macular edema is one of the leading causes of blindness in diabetic patients. Approximately 500,000 cases in the United States require treatment each year. The permeability of the retinal vasculature is increased by vascular endothelial growth factor (VEGF), which is released from the cellular components of the retina upon microcapillary occlusion and subsequent ischemia. Macular edema is also affected by several systemic diseases: the degree of diabetic nephropathy as well as retinopathy, hypertension, anemia, and hyperlipidemia. Diabetic macular edema is divided into two types: focal and diffuse. Focal macular edema is often caused by microaneurysms with impaired blood-retinal barrier integrity and can be treated by laser photocoagulation of microaneurysms. Diffuse macular edema, on the other hand, is caused by dilated retinal capillaries around the macular area. Diffuse macular edema is currently treated with grid laser therapy, intravitreal or subfascial steroid injections, triamcinolone acetonide (Ip et al., 2004; Gillies et al., 2006), or vitrectomy. Recently, several non-biodegradable or biodegradable devices are being clinically tested with various corticosteroids for intraocular use (Medidur1, I-vation1, and Posurdex1) and for intraocular injection of anti-VEGF drugs (Lucentis1 and Avastin1) to Treatment of diffuse diabetic macular edema.

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Figure 3 Macular edema due to central retinal vein occlusion. (a) Fluorescein angiography showing cystoid macular edema. (b) Optical coherence tomography showing retinal segments with cystic spaces at the macula. Vision is 20/100. (c) Intravitreal injection of crystalline triamcinolone acetonide (4 mg) reduced macular edema and improved visual acuity to 20/50. However, macular edema often recurs, requiring repeated dosing, laser photocoagulation, or vitrectomy

2.4

retinitis pigmentosa

Retinitis pigmentosa is a group of clinically and genetically heterogeneous eye diseases in which primary retinal and RPE degeneration results from various genetic alterations in the physiological function of photoreceptors and underlying RPE cells that act as sensors of visible light. Gene mutations include autosomal dominant and recessive, X-linked, and mitochondrial patterns. Most of the causative genes may be related to the retinoid cycle, including eg rpe65, abca4, vmd2. Usually one

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The electroretinogram was negative or unrecordable before the visual field defect occurred. In many cases, there is progressive visual field damage, in some cases accompanied by vision loss. So far, there have been no successful treatments to ameliorate the genetic alterations or even delay the progression of symptoms. However, a recent report on gene therapy offers some hope for a cure. Intravitreal injection of an rpe65-encoding adenoviral vector delayed the deterioration of visual function in patients with congenital amaurosis (Maguire et al., 2008). This finding suggests that a long-term supply of a defective or altered protein derived from a genetic mutation could save eyes with retinitis pigmentosa. Intracellular delivery of proteins and gene therapy using viral or non-viral vectors, siRNA, etc. may have great potential for the treatment of retinitis pigmentosa.

2.5

Age-Related Macular Degeneration (AMD)

AMD is the leading cause of legal blindness in persons over the age of 50 in most developed countries (Klein et al., 1992; Ryan et al., 1994). AMD affects 15 million Americans and many more in other Western countries. About 14-24% of U.S. citizens ages 65-74 and 35% of people over age 75 have the disorder. There are approximately 200,000 new cases of AMD each year in the United States, and the annual incidence is expected to increase as the population ages. AMD is divided into two types: choroidal neovascularization (CNV; exudative or wet AMD) and geographic atrophy (dry AMD). Geographic atrophy represents the death of photoreceptors and underlying RPE cells, as well as atrophy of the underlying choroidal capillaries, leading to progressive impairment of central vision without successful treatment. In contrast, CNV usually results in acute or subacute irreversible vision loss (Ryan et al., 1994) (Figure 4). CNV has been treated with surgical resection, macular displacement using scleral retraction sutures or 360-degree retinotomy, laser photocoagulation, photodynamic therapy, transpupilular thermotherapy, or radiation therapy (American Academy of Ophthalmology 2000; Bressler 1997; Chakravarthy et al. 1993; Kaplan 1996; Macular Photocoagulation). Research Group 1986, 1990, 1994a,b).

Figure 4 Exudative age-related macular degeneration. Choroidal neovascularization (asterisk) invades subretinal space with fibrin and hemorrhage (arrow). Even with the latest treatments (anti-VEGF therapy), there is limited improvement in already weakened vision.

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Despite multiple treatment options, vision recovery in CNV eyes remains very limited. Furthermore, recurrence of CNVs is a serious problem. Macular translocations with 360-degree retinotomy often result in significant improvements in visual acuity (Eckardt et al 1999; Pertile and Glaes 2002). However, serious complications such as postoperative PVR and diplopia have not been minimized. Recently, the development of pharmacological therapies for AMD has been of great interest to many clinicians, researchers, and pharmaceutical companies, as photodynamic therapy using photosensitizers, topical steroid therapy, and more recently, the use of anti-VEGF drugs have proven effective. Has been shown to be relatively effective in exudative AMD (Blumenkranz et al 2002; Rosenfeld et al 2006; Brown et al 2006; Augustin and Schmidt-Erfurth 2006; Adamis and Shima 2005). Many clinical trials are being completed, ongoing or in preparation. Recent successes have been with intravitreal or subfascial injections, as past pharmacological approaches to the treatment of exudative AMD with systemic interferon alfa and thalidomide have failed (Ip and Gorin 1996; Maguire et al 2001; Pharmakologische Macular Degeneration Research Group 1997 Treatment). However, newer drugs have been or will be tested, not only by intravitreal or subfascial injection, but also by systemic administration and even eye drops. Thus, a major recent trend in the development of new compounds that has often blinded many subsequent innovators is that limiting access to pharmacological agents is a critical issue. There is a need to develop intraocular controlled-release or drug-targeting systems, as well as to develop new potent anti-angiogenic drugs.

2.6

Proliferative Vitreoretinopathy (PVR)

PVR is a major cause of failure in retinal detachment surgery and is associated with pathological wound healing in the eye: formation of a fibrous membrane of RPE cells, glial cells, macrophages, and fibroblasts on or under the retina (Jerdan et al. 1989; Machemer and Racca). 1975; Retina Society Terminology Committee 1983). RPE cells that spread into the vitreous or subretinal space after retinal detachment produce extracellular matrix proteins, proinflammatory cytokines, and chemokines. This leads to disruption of the blood-retinal barrier, recruitment of other inflammatory cells, and myofibroblast transdifferentiation of RPE cells (Scheiffarth et al., 1988). The resulting contractile forces within the fibrous tissue that forms can eventually lead to retinal detachment and/or macular folds, jeopardizing vision. RPE cells and fibroblasts are usually the predominant cell types in the PVR epiretinal membrane. Several cytokines may play a role in the pathogenesis of PVR, including transforming growth factor beta (Gonzalez-Avila et al., 1995), platelet-derived growth factor (Milenkovic et al., 2003), and basic fibroblast growth factor (Hueber et al. al al 1996-1997; La Heij et al 2002), tumor necrosis factor-alpha (Armstrong et al. 1998), epidermal growth factor (Milenkovic et al. 2003), interleukin-1beta (El-Ghrably et al. al. 2001), interleukin 6

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(El-Ghrably et al. 2001; La Heij et al. 2002), interleukin 8 (El-Ghrably et al. 2001), interferon gamma (El-Ghrably et al. 2001), hepatocyte growth factor (Hinton et al. 2002) ).).). ), connective tissue growth factor (Hinton et al. 2002) and VEGF (Armstrong et al. 1998). Therefore, many substances can be candidates for the treatment of PVR. On the other hand, a single therapy targeting one of these substances may not be sufficient to treat PVR. Since many cases require surgical intervention to mechanically reduce retinal stretch, some intraocular controlled-release drug devices can be used as an adjunct to reduce the occurrence of PVR.

3 Non-biodegradable devices Since the late 1980s, the development of intraocular controlled-release systems has been vigorously explored due to the urgent need to establish a new therapeutic modality to replace repeated intravitreal injections of ganciclovir for the treatment of CMV retinitis in AIDS patients . These efforts resulted in the first commercial product, Vitrasert1, a nonbiodegradable implant containing ganciclovir (Figure 2) (Sanborn et al., 1992). This reservoir implant is made of non-biodegradable polymers such as polyvinyl alcohol (PVA), EVA, and silicone laminate with an API inside. This type has the most stable and long-lasting drug release profile compared to other implant types because it can store a large amount of drug and modulate drug release only by the total surface area and thickness of the permeable PVA polymer (Fig. 5a) (Okabe et al. 2003; Yasukawa et al. 2004, 2006). EVA or silicone laminate is an impermeable polymer used to limit the actual surface area for drug penetration, while PVA is used to build the frame of the device and regulate drug permeability. On the other hand, disadvantages of this type include the relatively large size of the device, which requires a large incision for implantation, which increases the risk of vitreous hemorrhage, subsequent epiretinal membrane and retinal detachment, and may require removal surgery for replacement Potential complications of implant or treatment, such as retinal detachment and API-induced side effects. In fact, the same implant type, Retisert1, releases fluocinolone acetone for the treatment of chronic noninfectious uveitis, steroid-induced cataracts and glaucoma with very high rates. When developing controlled-release systems, researchers must consider not only sustainable effects, but also possible adverse effects of active ingredients and non-biodegradable devices.

3.1

Vitrasert1: non-biodegradable implant containing ganciclovir

Vitrasert1 (Bausch & Lomb, Rochester, NY), a reservoir implant that delivered the antiviral drug ganciclovir intraocularly in patients with AIDS-associated CMV retinitis, was approved by the FDA in 1996. Ganciclovir is a synthetic nucleoside

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Figure 5. API release profiles from non-biodegradable implants (a) and biodegradable implants (b). Non-biodegradable implants facilitate stable and sustained release, while biodegradable implants do not require removal or replacement surgery. While biodegradable implants (solid squares in b) generally exhibit a triphasic release pattern of API, hybrid biodegradable implants (solid circles in b) can improve the stability and periodicity of API release. Replication of Okabe et al. (2003) and Yasukawa et al. (2000)

A 20-deoxyguanosine analog that inhibits the replication of viruses such as CMV, herpes simplex virus types 1 and -2, Epstein-Barr virus, and varicella-zoster virus. This intravitreal implant contains ganciclovir tablets consisting of 4.5 mg ganciclovir and 0.25% magnesium stearate as inactive ingredients and coated with PVA and EVA. PVA and EVA are non-degradable polymers used to control drug release rates. Hydrophobic EVA is used to limit the surface area available to release hydrophilic drugs. After fluid has entered the device, ganciclovir is slowly released through a water-permeable PVA membrane over 5-8 months. Ganciclovir partially dissolves in the ingested water to saturating concentrations and continues to diffuse across the PVA membrane. Depot implantation resulted in zero-order release kinetics, while the internal solution remained saturated with ganciclovir. Vitrasert1 requires surgical implantation. A 5.5 mm sclerotomy was performed 4 mm posterior to the limbus. After trimming the vitreous prolapse, a ganciclovir implant was inserted into the vitreous cavity via sclerotomy and secured with anchor sutures. The sclerotomy and conjunctiva are then sutured. Care should be taken when handling the device so as not to damage the PVA membrane. usually after implantation

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Over the next 2-4 weeks, patients will notice an immediate decrease in vision. This temporary visual impairment may be caused by the surgery itself. If reimplantation is required or complications such as retinal detachment occur, the implanted device must be removed.

3.2

Retisert1: non-biodegradable implant containing fluocinolone

Retisert1 (Bausch & Lomb, Rochester, NY) is a reservoir implant of the same shape as Vitrasert1 for the treatment of chronic noninfectious uveitis (Jaffe et al. 2000) (Fig. 6a). The implant consists of a tablet containing 0.59 mg of the active substance fluocinolone and inactive ingredients such as microcrystalline cellulose, PVA and magnesium stearate. The implant was coated with a PVA and silicone laminate to allow sustained release of the corticosteroid, fluocinolone acetone, at a constant rate of 0.3-0.6 mg/day to the posterior segment of the eye for 30 months. The device is 5mm long, 2mm wide and 1.5mm thick, making it smaller than the Vitrasert1. Fluocinolone acetonide reduces inflammation and reduces intravitreal VEGF levels in the eye. At 34 weeks post-implantation, uveitis recurrence rates were reduced to 7-14%, compared to 40-54% in the control group. The implant also reduces the need for systemic and topical steroids. The implant procedure is the same as Vitrasert1. Care should be taken when handling the implants to avoid damaging the PVA membrane, which may lead to an unexpected increase in the drug release rate from the implants. Ironically, however, the sustained release of steroids into the eye leads to a high incidence of steroid-induced intraocular complications. Within 34 weeks of implantation, approximately 60% of patients required medical treatment for steroid-induced glaucoma. Additionally, approximately 32% of patients required glaucoma (filter) surgery within two years of implantation (Callanan 2007). Even within two years of implantation, most phakic eyes develop cataracts and require cataract surgery. despite this,

Figure 6 Various medical devices used for DDS in the posterior segment of the eye. (a) Retisert1 (Bausch & Lomb) contains a tablet of fluocinolone encapsulated in a non-biodegradable polymer. (b) I-vationTM (SurModics) is a helical, non-biodegradable implant containing triamcinolone acetonide that is inserted through a small incision. (c) Posurdex1 (Allergan) is an injectable dexamethasone stick with a special applicator

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Since long-term persistent intraocular inflammation can itself be associated with the development of secondary glaucoma and cataracts, the sustained and superior efficacy of implants in the treatment of chronic uveitis is most notable.

3.3

I-vationTM: a non-biodegradable implant containing triamcinolone acetonide

A unique non-biodegradable implant, I-vationTM (SurModics, Irvine, CA), is currently in Phase I clinical trials for the treatment of diabetic macular edema. The device contains 0.925 mg of the corticosteroid, triamcinolone acetone, and delivers it slowly to the back of the eye. The device is shaped as a unique stent designed for minimally invasive implantation, unlike Vitrasert1 and Retisert1, which require a more invasive surgical procedure to implant (Fig. 6b). The small diameter of the implant allows for implantation through a sclerotomy using a 25-gauge needle. The unique helical design maximizes the surface area available for drug delivery and allows for fixation of the implant on the sclera. An attached thin cap is placed below the subconjunctival space. Triamcinolone acetonide is coated with a mixture of polybutylmethacrylate and polyEVA. The safety and biocompatibility of I-vationTM are derived from preclinical studies up to 9 months after implantation. SurModics is sponsoring a Phase I safety study called STRIDE (Sustained Triamcinolone Release for Inhibition of Diabetic Macular Edema) for the treatment of diabetic macular edema. Preliminary clinical results indicate that I-vationTM is safe and well tolerated. The average macular thickness in eyes with diabetic macular edema decreased to 230 mm at 6 months, compared to a preoperative thickness of 376 mm. Subjects are supervised for a period of three years. Drug release profiles can be tailored by varying the ratio, thickness, and total surface area of the polymer components in the coating. It remains to be seen whether the sclera will tolerate anchoring of the device over longer periods of time and after device removal.

3.4

Medidur1: non-biodegradable insert with fluoride

Medidur1 (Alimera Sciences Inc., Alpharetta, GA; pSivida Inc., Watertown, MA) is an injectable, nonbiodegradable intravitreal implant for the treatment of diabetic macular edema. The Medidur1 Insert is only 3.5 mm long and 0.37 mm in diameter and is designed to deliver a constant amount of fluocinolone to the posterior segment of the eye over a period of 18 to 36 months. Inserts, like other non-biodegradable implants, are reservoir implants. Unlike these, however, it is designed for seamless adoption in office environments

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The 25 gauge transconjunctival syringe system eliminates the need for surgical incisions and subsequent sutures in the conjunctiva and sclera. The insert delivers 0.2 mg or 0.5 mg of fluocinolone daily into the vitreous cavity. A phase III study for the treatment of diabetic macular edema is currently underway. It should be clarified whether an insert suspended without fixation in the vitreous cavity would have deleterious effects and could easily be surgically removed if necessary.

4 Biodegradable devices Compared with non-biodegradable implants, biodegradable implants have the following advantages: no need for surgical removal and flexible shape. They can be made in various configurations such as microparticles, rods, discs, tablets, and implants (Figure 2) (Yasukawa et al., 2004, 2006, 2007). Posurdex1, a rod that can be injected with a special syringe using a 22-gauge needle, was recently in phase III trials for the treatment of macular edema secondary to retinal vein occlusion or diabetic macular edema. The implant consists of a homogeneous mixture of dexamethasone as the API and poly(lactic-co-glycolic acid) (PLGA), a biodegradable polymer classified as a monolithic type. In general, there are three phases to this type of drug release: (1) the first burst, caused by drug deposition on the implant surface; (2) the diffusion phase, driven by osmotic pressure and polymer biodegradation; (3) Final eruption due to sudden disintegration of the implant matrix (Fig. 5b) (Yasukawa et al., 2004). Therefore, researchers should consider the first and last bursts if the active substance has toxic effects at high concentrations. Compared to a depot implant, the API release profile of a monolithic device should be influenced by several factors, including the type and molecular weight of the polymer, the mixing ratio of the polymer and API, and the total surface area of the device. Yasukawa and Kunou et al. showed that a mixture of two polymers of different molecular weight resulted in a reduced final eruption and a more stable and longer release of the API (Fig. 5b) (Yasukawa et al. 2000; Kunou et al. 2000). Thus, the release profile of the API in biodegradable implants can become as stable as that of non-biodegradable implants, whereas the duration of API release can be shorter due to the limited amount of API contained.

4.1

Posurdex1: biodegradable insert containing dexamethasone

Posurdex1 (Allergan Inc., Irvine, CA) is a PLGA-fabricated biodegradable polymer matrix that releases dexamethasone over approximately five weeks. Benefit in persistent macular edema associated with diabetic retinopathy, retinal vein occlusion, uveitis and cataract surgery in Phase III clinical trials. A prototype of the insert was implanted through a No. 20 scleral incision in the vitreous cavity at the surgical site. Phase II clinical trial

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It was shown that patients who received an insert containing 0.7 mg of dexamethasone had the greatest improvement in vision, and most of these patients had a three-line rise on an eye chart, which measures vision. No side effects such as increased intraocular pressure or cataract formation were observed in the treatment group. A novel delivery system was then developed for Posurdex1 using a disposable applicator with a 22-gauge needle (Figure 6c). With this applicator, the insert is injected in practice. The clinical trial is currently in phase III to study the efficacy and safety in more patients with diabetic macular edema. Surodex1 (Oculex Pharmaceuticals, Sunnyvale, CA), a PLGA granule containing 0.06 mg dexamethasone, ensures sustained release of dexamethasone for 7–10 days after insertion into the anterior chamber. Surodex1 achieves higher intraocular drug levels than conventional dexamethasone eye drops and is effective in reducing inflammation after cataract surgery (Tan et al., 1999). The insert is approved for cataract surgery in Singapore.

4.2

injection of microspheres

Although the injectability of microspheres was an advantage, in the 1990s it was thought that intravitreally injected microspheres would impair the clarity of the intraocular medium. Since then, intravitreal injection of crystalline triamcinolone acetonide has been used to treat macular edema and exudative AMD (Ip et al 2004; Augustin and Schmidt-Erfurth 2006; Gillies et al 2006). Triamcinolone acetonide is hydrophobic and is mostly suspended in the vehicle. Nevertheless, crystalline triamcinolone acetonide sinks mainly into the lower part of the vitreous cavity, where it does not affect ocular media and visual function. The successful use of this drug suspension will provide a great opportunity for intraocular DDS with microparticles. Ongoing Posurdex1 clinical studies clearly demonstrate that the biodegradable polymer is biocompatible. Many types of biodegradable implants and microparticles will enter clinical trials in the near future. In Japan, Sub-Tenon's betamethasone microbead injection (DE-102, Santen Pharmaceuticals, Ikoma, Japan) is currently in Phase II/III clinical trials for the treatment of diabetic macular edema. Microspheres containing pegaptanib, an aptamer with affinity for VEGF (Macugen; Pfizer Inc., New York, NY), were also examined in the laboratory. When these microspheres were injected into the vitreous, the rabbits slowly released pegaptanib into the vitreous cavity over four months.

5 Triamcinolone acetonide crystal suspension Triamcinolone acetonide crystal suspension is of great significance in the treatment of vitreoretinal diseases; triamcinolone acetonide intravitreal and subfascial administration is widely used in the treatment of macular diseases

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Edema, exudative AMD, and uveitis (Ip et al., 2004; Gillies et al., 2006; Augustin and Schmidt-Erfurth, 2006) (Fig. 3c). Due to its hydrophobic nature, the infused crystals gradually dissolve, resulting in a sustained release of triamcinolone acetonide. Intravitreal injections of 4 mg or subfascial injections of 20 mg triamcinolone acetonide provided 3-month efficacy in chorioretinal tissue. Intravitreal injections, although more effective than sub-tenon injections, carry a higher risk of steroid-induced glaucoma and cataract progression. In most cases, however, both complications are treatable. Therefore, the crystal suspension was considered as a DDS without a polymer matrix as a base. Future steroid delivery device candidates may need to exhibit a release profile superior to intravitreal injection of crystalline triamcinolone acetonide.

6 NT-501: Encapsulated Cell Technology (ECT) ECT is a new technology developed by Neurotech Pharmaceuticals, Inc. (Lincoln, RI, USA) that enables the sustained delivery of cell-based factors to the posterior segment of the eye. ECT implants consist of cells genetically engineered to secrete therapeutic factors and an envelope of semipermeable hollow fiber membranes (Figure 2). The permeability of the hollow fiber membranes allows the flow of oxygen and nutrients and prevents direct contact of the encapsulated cells with the cellular and molecular components of the immune system, allowing long-term cell survival. The encapsulated cells continuously produce the therapeutic protein, which diffuses out of the implant at the target site. The protein was delivered to the vitreous cavity for up to 18 months via the ECT device NT-501, which contains human cells genetically engineered to secrete ciliary neurotrophic factor. Like Vitrasert1 and Retisert1, NT-501 is secured to the ciliary body with sutures. In early 2006, the National Eye Institute completed a phase I clinical trial of 10 patients diagnosed with retinitis pigmentosa. ECT-mediated ciliary neurotrophic factor administration has been shown to be safe over a six-month treatment period. Also, unexpectedly, visual acuity tended to improve to varying degrees from baseline. NT-501 is currently in a Phase II/III clinical trial for the treatment of retinitis pigmentosa and a Phase II clinical trial for the treatment of dry AMD. The ECT machine may also release antibodies and cytokines.

7 Conclusions In the treatment of exudative AMD, repeated intraocular injections of anti-VEGF drugs have recently been shown to be significantly suppressive compared with other existing treatment modalities such as photodynamic therapy, laser photocoagulation, and surgery. Many drug candidates are currently undergoing clinical or preclinical studies.

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However, some studies refer to systemic use or even eye drops without considering the specificity of ocular pharmacokinetics. Some scientists have forgotten or failed to recognize the history of previous failures of interferon and thalidomide. Other single-compound therapies may not be as effective as anti-VEGF therapies because VEGF must be a key regulator in many steps of inflammation and angiogenesis, so its blockade can directly and rapidly inhibit the inflammatory response and angiogenesis. Compared with repeated intravitreal injections of anti-VEGF drugs, controlled release systems may be required to further improve efficacy and reduce the incidence of side effects. In the 1980s, AIDS-associated cytomegalovirus retinitis accelerated the development of intraocular drug delivery systems following the recognition of clinical use of nonbiodegradable implants. Various biodegradable implants and microparticles are now available for clinical use. The most difficult vitreoretinal diseases are now being targeted.

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right of withdrawal

Novel biosensing and miniaturized drug delivery devices for controlled and responsive drug delivery Andrea A. Robitzki and Randy Kurz

M. Schäfer-Korting (Hrsg.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_3, # Springer-Verlag Berlin Heidelberg 2010

The book chapter "Biosensing and Drug Delivery at the Microscale: Novel Devices for Controlled and Responsive Drug Delivery" is published in Handbook of Experimental Pharmacology, vol. 197: Drug Delivery retracted by the authors because the current version is based on a preliminary manuscript that contains incorrect citations, references, and wording that is partially identical to other original manuscripts. The authors state that they have done everything possible to correct this error and demonstrate that their intent was to keep everyone involved happy. The author deeply regrets the error.

A.A. Robitzki (*) Center for Biotechnology and Biomedicine, University of Leipzig, Deutscher Platz 5, 04103 Leipzig, Germany Email:[email protected]

M. Schäfer-Korting (Hrsg.), Drug Delivery, Handbook of Experimental Pharmacology 197, DOI 10.1007/978-3-642-00477-3_17, # Springer-Verlag Berlin Heidelberg 2010

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