Membrane Technology: Volume 1: Membranes for Life Sciences

Membranes for the Life Sciences Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Membranes for the Life Scie...

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Membranes for the Life Sciences

Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes

Membranes for the Life Sciences Volume 1

Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes

The Editors Dr. Klaus-Viktor Peinemann GKSS Research Center Institute of Polymer Research Max-Planck-Strasse 1 21502 Geesthacht Germany Dr. Suzana Pereira Nunez GKSS Research Center Institute of Polymer Research Max-Planck-Strasse 1 21502 Geesthacht Germany

All books published by Wiley-VCH are carefully produced. Nevertheless, authors, editors, and publisher do not warrant the information contained in these books, including this book, to be free of errors. Readers are advised to keep in mind that statements, data, illustrations, procedural details or other items may inadvertently be inaccurate. Library of Congress Card No.: applied for British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library. Bibliographic information published by the Deutsche Nationalbibliothek The Deutsche Nationalbibliothek lists this publication in the Deutsche Nationalbibliograﬁe; detailed bibliographic data are available in the Internet at . # 2008 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim All rights reserved (including those of translation into other languages). No part of this book may be reproduced in any form – by photoprinting, microﬁlm, or any other means – nor transmitted or translated into a machine language without written permission from the publishers. Registered names, trademarks, etc. used in this book, even when not speciﬁcally marked as such, are not to be considered unprotected by law. Typesetting Thomson Digital, India Printing Strauss GmbH, Mo¨rlenbach Binding Litges & Dopf GmbH, Heppenheim Cover Design Adam-Design Weinheim Wiley Bicentennial Logo Richard J. Paciﬁco Printed in the Federal Republic of Germany Printed on acid-free paper ISBN: 978-3-527-31480-5

V

Contents

Preface XI Contributors ((Peinemann Vol. 1))

1 1.1 1.2 1.3 1.3.1 1.3.1.1 1.3.1.2 1.3.1.3 1.3.1.4 1.3.1.5 1.3.1.6 1.3.2 1.3.2.1 1.3.2.2 1.3.2.3 1.3.2.4 1.4 1.4.1 1.4.2 1.4.3 1.4.4 1.5 1.6 1.6.1 1.6.2

XIII

Membranes in Hemodialysis 1 Jo¨rg Vienken Introduction 1 Historical Achievements 2 Membranes for Hemodialysis: Polymers and Nomenclature 7 Membranes from Regenerated Cellulose 8 Modiﬁed Cellulosic Membranes 9 Cellulose Acetates 9 DEAE-Modiﬁed Cellulose, Hemophan 10 Benzyl-Modiﬁed Cellulose (Synthetically Modiﬁed Cellulose, SMC) PEG-Grafted Cellulose 11 Vitamin E-Modiﬁed Cellulosic Membranes 12 Synthetic Membranes 12 Polyacrylonitrile (PAN) 13 Polymethylmethacrylate (PMMA) 14 Polysulfone (PSu) 15 Polyamide (PA) 17 Dialyzer Constructions 17 Hollow Fiber Dialyzers 17 Housing 18 Potting Material 18 Fiber Bundle 19 Dialysis Membranes and Performance: Principles of Membrane Transport 20 Dialysis Membranes and Biocompatibility 24 Some Basic Information on Membranes and Biocompatibility Parameters 24 Thrombogenicity of Different Types of Dialyzers and Filters 26

Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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Contents

1.6.3 1.6.4 1.6.4.1 1.6.5 1.6.5.1 1.6.6 1.6.6.1 1.6.7

1.6.8 1.6.9 1.7 2 2.1 2.2 2.2.1 2.3 2.4 2.4.1 2.4.2 2.5 2.6 2.6.1 2.6.2 2.7 2.7.1 2.7.2 3 3.1 3.2 3.3 3.4 3.4.1 3.4.2 3.5 3.6 3.7

Complement Activation by Different Dialyzers and Filters 29 Cell Activation by Different Types of Dialyzers and Hemoﬁlters 31 Apoptosis 31 Oxygen Species Production – Induction of Oxidative Stress 32 Degranulation of Neutrophils 34 Stimulation of Cytokine Generation by Different Types of Dialyzers and Hemoﬁlters 35 The Impact of Membrane Types on LPS-Stimulated IL-1b Secretion 36 The Impact of Large-Pore Dialysis Membranes on the Inﬂammatory Response in HD Patients by Cytokine Elimination 36 The Effect of Different Dialyzers on the Acute Phase Reaction 37 Activation of the Kinin System by Different Types of Dialyzers and Hemoﬁlters 37 Conclusion 39 Membranes for Artiﬁcial Lungs 49 Frank Wiese Introduction 49 History of Blood Oxygenation 49 Membrane Oxygenators 50 Principle of Gas Transfer 53 Membranes and Membrane Properties 55 Microporous Membranes 55 Dense Membranes/‘‘Diffusion Membranes’’ 57 Membrane Production 59 Operational Modes and Membrane Makeup in Oxygenators Microporous Capillary Membranes, Blood Inside 62 Microporous Capillary Membranes, Blood Outside 62 Extracorporeal Circulation 65 Cardiodiapulmonary Bypass (CPB) 65 Lung Support Systems 65 Membranes for Blood Fractionation/Apheresis 69 Frank Wiese Introduction 69 History of Plasmapheresis 70 Principles of Plasmapheresis 73 Membranes and Membrane Properties 76 Plasma Separation Membranes 76 Plasma Fractionation Membranes 79 Membrane Production 81 Operational Modes in Plasmapheresis Procedures 83 Medical Indications for Blood Plasma Treatment 88

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4

4.1 4.2 4.2.1 4.2.2 4.2.3 4.2.4 4.2.5 4.2.6 4.2.6.1 4.2.6.2 4.2.6.3 4.3 4.3.1 4.3.1.1 4.3.1.2 4.4 4.4.1 4.4.1.1 4.4.1.2 4.4.1.3 4.4.1.4 4.4.1.5 4.4.2 4.4.2.1 4.4.2.2 4.4.2.3 4.4.2.4 4.4.2.5 4.5 4.5.1 4.5.2 4.5.2.1 4.5.3 4.5.3.1 4.5.3.2 4.5.4 4.5.4.1 4.5.4.2

Membranes in the Biopharmaceutical Industry 91 Anthony Allegrezza, Todd Ireland, Willem Kools, Michael Phillips, Bala Raghunath, Randy Wilkins, Alex Xenopoulos Introduction 91 Microﬁltration Membranes Used in the Biotech Industry 93 Introduction 93 Microﬁltration Membranes: Development of Industrial Membranes 94 Effect of Membrane Structure on Properties 95 Aspects of Cartridge Design 97 Membrane Surface Modiﬁcation 98 Sterilizing Filters 99 Retention 99 Permeability 101 Capacity 102 Practical Membrane Considerations for Sterile Filtration by Microporous Membranes 102 Sterile Filtration Process Considerations 103 Filter Selection 104 Device Selection 110 Ultraﬁltration and Virus Filtration Membranes for Biopharmaceutical Applications 113 Ultraﬁltration Membranes 113 Membrane Suppliers 114 Membrane Selection 114 Membrane Structures 115 Characterization 115 Devices 116 Virus Filtration Membranes 117 Membrane Suppliers 117 Membrane Structures 118 Devices 118 Membrane Selection 119 Characterization 120 Applications of Ultraﬁltration Membranes in Biopharmaceutical Manufacturing 121 Ultraﬁltration Theory 122 Typical Ultraﬁltration Process 123 Process Development and Optimization 124 Processing Plan Optimization 128 Mode of Operation 128 Diaﬁltration Mode/Strategy 129 Scale-up Considerations 131 Process Implementation Considerations 132 Process Robustness 133

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VIII

Contents

4.5.5 4.5.5.1 4.5.5.2 4.6 4.6.1 4.6.2 4.6.3 4.6.4 4.6.5 4.6.6 4.6.7 4.7 4.7.1 4.7.2 4.7.2.1 4.7.2.2 4.7.2.3 4.7.3 5 5.1 5.2 5.2.1 5.2.2 5.2.3 5.2.4 5.3 5.3.1 5.3.1.1 5.3.1.2 5.3.2 5.3.2.1 5.3.2.2 5.3.3 5.3.4 5.3.5 5.3.6 5.3.7 5.3.8 5.4 5.5

System Considerations 134 Equipment Options 134 Process Control Options 134 Practical Aspects of Virus Filtration Process Design and Implementation 135 Membrane Selection 135 Process Development and Optimization 138 Capacity 140 Small-Scale Simulation 141 Pilot-Scale Studies 141 Virus Validation Studies 142 Implementation 143 Membrane Adsorbers 145 Membrane Chemistries 146 Current Applications 147 Flow-Through Polishing 147 Flow-Through Precapture 148 Large Molecule Bind–Elute Puriﬁcation 149 Future Trends 149 Membrane Applications in Red and White Biotechnology 155 Stephan Lu¨tz, Nagaraj Rao Introduction 155 Types of Membrane Processes in Red and White Biotechnology 156 Bubble-Free Aeration 156 Filtration Processes 157 Dialysis and Electrodialysis 157 Adsorption of Microorganisms 158 Examples of Membrane Processes in Biotechnology 158 Bubble-Free Gassing 158 Hydrogen 158 Oxygen 159 Membranes for Cell Retention 161 Higher Cells/Red Biotechnology 161 Whole-Cell Biotransformation 162 Membranes for Enzyme Retention 162 Membranes for Cofactor Retention 166 Application of Dialysis and Electrodialysis in Biotransformations Application of Pervaporation and Stripping in Biotransformations 168 Nanoﬁltration and Ultraﬁltration in Biotechnology 170 Bioelectrochemical Applications 170 Summary 172 Acknowledgment 172

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Contents

6 6.1 6.2 6.2.1 6.2.2 6.3 6.3.1 6.3.2 6.3.2.1 6.3.2.2 6.4 7

7.1 7.2 7.2.1 7.2.2 7.2.3 7.2.3.1 7.2.3.2 7.2.3.3 7.2.3.4 7.3 7.3.1 7.3.1.1 7.3.1.2 7.3.1.3 7.3.1.4 7.3.2 7.3.2.1 7.3.2.2 7.3.2.3 7.4 8

8.1 8.1.1 8.1.2 8.2

Membranes in Controlled Release 175 Nicholas A. Peppas, Kristy M. Wood, J. Brock Thomas Introduction 175 Controlled Release Kinetics 176 Diffusion in Membrane-Controlled Release 176 Physical Parameters of Controlling Release 180 Membranes and Solute Transport 180 Characterization of Membranes 180 Solute Transport in Network Membranes 182 Structural Parameters of Membranes 182 Determination of Molecular Pore Sizes 183 Applications in Drug Delivery 184 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane 191 Dimitrios F. Stamatialis Introduction 191 Human Skin – Fundamentals of Skin Permeation 192 Human Skin Structure 192 Stratum Corneum – Main Drug Barrier 193 Drug Transport Through the Skin 195 Passive Diffusion 195 Iontophoresis 198 Electroporation 204 Other Methods 207 Transdermal Drug Delivery System – Structure/Design 212 Passive TDD Systems 212 Types 212 Materials 213 Skin or Device-Controlled Delivery 214 Commercialization – Patents 215 Active TDD Systems 216 Types 216 Materials – Devices 216 Commercialization – Patents 218 Conclusions and Outlook 221 Application of Membranes in Tissue Engineering and Biohybrid Organ Technology 227 Thomas Groth, Zhen-Mei Liu Introduction 227 Application of Membranes in Blood Detoxiﬁcation 227 Requirements to Support Adhesion and Function of Cells Application of Membranes in Tissue Engineering 231

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Contents

8.2.1 8.2.2 8.2.3 8.2.4 8.3 8.3.1 8.3.2 8.3.3 8.4 8.5 9

9.1 9.1.1 9.1.2 9.2 9.2.1 9.2.2 9.2.2.1 9.2.3 9.2.3.1 9.2.3.2 9.2.3.3 9.2.4 9.2.5 9.2.5.1 9.3 9.4

Introduction to Tissue Engineering and Membrane Applications 231 Tissue Engineering of Skin 232 Tissue Engineering of Bone 236 Further Tissue Engineering Applications of Membranes Membranes in Biohybrid Organ Technology 242 Organ Failure and Biohybrid Organ Technology 242 Biohybrid Liver 245 Biohybrid Kidney 249 Summary and Conclusions 253 Acknowledgments 254

240

Membranes in Bioartiﬁcial Pancreas – An Overview of the Development of a Bioartiﬁcial Pancreas, as a Treatment of Insulin-Dependent Diabetes Mellitus 263 Ana Isabel Silva, Anto´nio Norton de Matos, I. Gabrielle M. Brons, Marı´lia Clemente Velez Mateus Introduction 263 Diabetes and Its Treatment 263 The Bioartiﬁcial Organ Concept 265 Bioartiﬁcial Pancreas 266 Immunoprotection and Biocompatibility of Implanted Devices Vascular Devices 268 Biocompatible Materials in Vascular Devices 269 Extravascular Devices 272 Implantation Sites 272 Macrocapsules 273 Microcapsular Devices 280 Inﬂuence of Recipients Sensitization to Donor Antigens in Graft Survival 308 Islet Oxygenation Studies 308 Oxygenation of Ba-Alginate Microencapsulated Islets 311 Final Comments 312 Acknowledgment 313 Index

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XI

Preface Life without membranes has never been possible. They are essential components of our body. But in the past 50 years the membrane scientists managed to incorporate artiﬁcial membranes in medical and engineering processes, which improved the life quality of a great part of the world population. Just to mention two examples, there are currently around 1.5 million hemodialysis patients worldwide, who decades ago would have been condemned to die at an early age. Two billion people are affected by water shortages in over 40 countries. Reverse osmosis desalination plants assure water supply to a large area of the planet, which barely has access to alternative sources. In the past decades the membrane technology established its place in the chemical industry. Rapidly growing applications are the nanoﬁltration of ﬁne chemicals and the recovery of monomers in polymerization plants. The electronic industry depends on clean water obtained by membrane systems. The globalization of the food market requires active packaging ‘‘membrane’’ systems to allow the transport of fruits and vegetables for long distances and storage times. Membrane processes have been for a long time the integrated steps for the cheese and wine production, caring for the improvement of the product quality and for the introduction of new products. But this is just the start. Recently, the concern about global warming and the contribution of the greenhouse effect has increased dramatically. As a consequence, there is a growing demand for renewable energy carriers with low emission of CO2. Membranes are the core components of fuel cells and electrolysers. Apart from this, the implementation of the membrane processes in a modern society that is dependent on hydrogen technology and renewable energy is a long-term process, and in this transition period membranes are expected to play an important role. Fossil fuels are still the predominant source of hydrogen. Membranes can make the hydrogen production and processing in reﬁneries much more effective. Processes such as steam reforming and water shift reaction can be signiﬁcantly improved with the use of membranes. Membrane technology can also be involved in the CO2 emission reduction in reﬁneries and coal plants. After publishing the book ‘‘Membrane Technology in the Chemical Industry,’’ we clearly saw a demand for more information on membranes and their Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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Preface

applications in different ﬁelds. A series of six books with the following topics was then initiated: Membranes in Life Science, Membranes for Energy, Membranes in the Food Industry, Membranes for Water Treatment, Membranes for Chemical Technology, and Membrane Preparation. Suzana Nunes, Klaus-V. Peinemann

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Contributors ((Peinemann Vol. 1)) Anthony Allegrezza Millipore Corporation Bioprocess R & D 80 Ashby Road Bedford, MA 01730 USA

Willem Kools Millipore Corporation Bioprocess Applications Engineering 900 Middlesex Turnpike Billerica, MA 01821 USA

I. Gabrielle Brons University of Cambridge Department of Surgery Cambridge Institute of Medical Research CIMR Box 139 Addenbroke’s Sie Cambridge CB2 2XY UK

Zhen-Mei Liu Martin-Luther-Universita¨t HalleWittenberg Institute of Pharmacy Kurt-Mothes-Strasse 1 06120 Halle (Saale) Germany

Thomas Groth Martin-Luther-Universita¨t HalleWittenberg Institute of Pharmacy Kurt-Mothes-Strasse 1 06120 Halle (Saale) Germany Todd Ireland Millipore Corporation Bioprocess Sales Management 900 Middlesex Turnpike Billerica, MA 01821 USA

¨tz Stephan Lu Forschungszentrum Ju¨lich GmbH Institut fu¨r Biotechnologie 2 Leo-Brandt-Str. 52425 Ju¨lich Germany Marilia Mateus IBB-Institute for Biotechnology and Bioengineering Centre for Biological and Chemical Engineering Instituto Superior Te´cnico Av. Rovisco Pais 1049-001 Lisboa Portugal

Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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Contributors ((Peinemann Vol. 1))

Anto´nio Norton de Matos University of Porto ICBAS and HGSA-Santo Anto´nio General hospital Division of Angiology and Vascular Surgery Largo Professor Abel Salazar 4099-001 Porto Portugal

Ana Isabel Silva IBB-Institute for Biotechnology and Bioengineering Centre for Biological and Chemical Engineering Institute Superior Te´cnico Av. Rovisco Pais 1049-001 Lisboa Portugal

Klaus-Viktor Peinemann GKSS Research Center Institute of Polymer Research Max-Planck-Strasse 1 21502 Geesthacht Germany

Dimitrios F. Stamatialis University of Twente Faculty of Science and Technology Biomedical Technology Institute Membrane Technology Group PO Box 217 7500 AE Enschede The Netherlands

Nicholas Peppas The University of Texas Departments of Chemical and Biological Engineering University Code C 0400 Austin TX, 78712 USA Michael Phillips Millipore Corporation Bioprocess R & D 80 Ashby Road Bedford, MA 01730 USA Bala Raghunath Millipore Corporation Bioprocess Applications Engineering 900 Middlesex Turnpike Billerica, MA 01821 USA Nagaraj Rao Rane Rao Reshamia Laboratories Pvt. Ltd. Plot 80, Sector 23, CIDCO Industrial Area 400705-Navi Mumbai India

J. Brock Thomas The University of Texas Departments of Chemical and Biological Engineering University Code C 0400 Austin TX, 78712 USA Jo¨rg Vienken BioSciences Department Fresenius Medical Care Else Kroenerstrasse 1 61346 Bad Homburg Germany Frank Wiese Consulting Membrane Technology/ Application Starenstrasse 100 42389 Wuppertal Germany

Contributors ((Peinemann Vol. 1))

Randy Wilkins Millipore Corporation Bioprocess Field Marketing 900 Middlesex Turnpike Billerica, MA 01821 USA

Kristy M. Wood The University of Texas Departments of Chemical and Biological Engineering University Code C 0400 Austin TX, 78712 USA

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1

1 Membranes in Hemodialysis Jo¨rg Vienken 1.1 Introduction

The treatment of kidney patients by hemodialysis with tubular sheet or capillary membranes represents a convincing success story in medical therapy. When John Abel in the United States and Georg Hass in Germany started to investigate the application of membranes at the onset of the twentieth century, they would never have dreamt that nearly a hundred years later, in 2005, about 1.9 million kidney patients would undergo treatment for end-stage renal disease and nearly 1.5 million patients would owe their lives to hemodialysis and related devices such as membranes, tubing systems, and so on [1] (Figure 1.1). One reason for this overwhelming success may be the successful miniaturization of dialyzers that facilitates the mass production of these devices. Further, and as of today, the exclusive use of capillary membranes considerably contributes to better performance properties of dialysis membranes. Therefore, the following ﬁgures may not be surprising: In total, in 2005, dialysis patients all over the world are currently in need of about 150 million dialyzers annually. Compared to 2002, this ﬁgure represents an increase of 15%. Taking into account that one dialyzer contains a length of about 3 km of capillary membranes, nearly 450 million km of these materials have to be manufactured annually. The ﬁgure reﬂects the unimaginable dimension of an annual capillary membrane production being equivalent to nearly three times the distance between the earth and the sun. The technical skill of manufacturing capillary membranes in such large quantities represents half a century of an effective research and development. A reproducible high capillary membrane quality in terms of constant geometry, adequate performance, and high-level blood compatibility guarantees the safety of patients. Further, membrane production at a large scale allows for a reasonable cost of production and consequently for a facilitated access to these devices all over the world. Membranes are the centerpiece of dialyzers and more than 650 different types of dialyzers are currently available on the market. How they are characterized and how they differ, will be the scope of this chapter.

Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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1 Membranes in Hemodialysis

Fig. 1.1 The ESRD patient population is continuously increasing and adds up to 1 900 000 patients worldwide in 2005. A majority of 1 297 000 patients is treated by hemodialysis.Source: FMC database 2005.

1.2 Historical Achievements

Following Gekas [2], a membrane is both an intervening phase to separate two phases and a barrier to the transport of matter between phases adjacent to it. The separation of water and toxins from blood, as observed in the glomeruli of the kidney, is also a membrane process. Therefore, it is of no surprise that physicians searched for ways to treat renal failure through the application of artiﬁcial membranes [3] (Table 1.1). The ﬁrst artiﬁcial membranes were handmade from collodium, a cellulose-nitrate derivative. These collodium membranes are considered to be the forerunners of today’s capillary hollow ﬁber membranes. They were used in the early experiments by Abel [4] on rabbits and dogs. Abel termed for the ﬁrst time the devices containing collodium membranes as ‘‘artiﬁcial kidneys.’’ The membrane transport process applied in these experiments was dialysis, that is, transport of low-molecular-weight compounds such as urea and creatinine through a dense membrane driven by a concentration gradient (Figure 1.2)

Tab. 1.1

1748 1861 1913 1923 1932 1938 1943 1968 1969 1974 1983 1987 1997 2000 2001 2005

Chronicle of capillary membranes and persons/companies behind. Pig bladder as a membrane Deﬁnition of term ‘‘dialyzer’’ Collodium (cellulose nitrate) tubes in dog dialysis Collodium tubes for hemodialysis in humans Continuous production of tubular membranes Cellophane, Heparin Rotating drum with cellophane Capillary dialyzer with cellulose acetate Polyacrylonitrile (PAN, AN69) Cuprophan hollow ﬁber Polysulfone (PSu), high ﬂux, biocompatible DEAE-modiﬁed cellulose Hemophan Vitamin E-bonded membrane Polyamide–PES membrane Helixone, PSu, nanocontrolled spinning Vitabran E, vitamin E-bonded PSu

Jean Antoine Nollet Thomas Graham John Abel George Haas Richard Weingand William Thalhimer Willem Kolff Richard Stewart Hospal Company Werner Bandel Ernst Streicher Enka Company Terumo Company Gambro Company Fresenius Medical Care ASAHI Medical Ltd

1.2 Historical Achievements

Fig. 1.2 The famous Chemist and Scotsman Thomas Graham (1805–1869) was the ﬁrst to deﬁne the term ‘‘dialyzer’’ in 1861.

The credit for pioneering the ﬁrst hemodialysis in humans, however, goes to the German physician Hans Georg Haas (1886–1971), who developed a system for hemodialysis in 1923 independent of Abel in Baltimore, USA. Like Abel, Georg Haas used collodium tubes [5–7]. Haas commented on his experimental efforts in membranes by saying, ‘‘I’d like to say, it was a way of the Cross, because once, one obstacle had been removed another followed.’’ and ‘‘Above all, it was necessary to ﬁnd a suitable dialysis membrane. I have tried a series of different dialyzers from a variety of materials, animal, vegetable membranes and paper dialyzers. The best performances were obtained by dialyzers from Collodium with respect to fabrication, dialysis effects, safe sterilization and, because they can be obtained in any geometric shape’’ [7]. Haas could proﬁt from the results published earlier in 1921 by Arnold Eggerth from the Hoagland Laboratories in New York. Eggerth had shown that it was possible to achieve a tailor-made membrane clearance by just choosing the appropriate dilution of the alcoholic solvent [8]. In 1932, Weingand [9] proposed an apparatus for a continuous manufacturing of membrane tubes from cellulose solutions. The cellulose solution issues from the annular nozzle and ﬂows – guided by gravity – along a guide member. ‘‘The solidiﬁed tube is treated in a series of baths: wasting, bleaching, desulfurizing, and plasticizing baths.’’ The application of a desulfurizing bath in Weingand’s experiments leads us to the conclusion that the cellulose solution was a viscose solution (viscose – cellulose xanthate in NaOH). Progress in the development of dialysis membranes and dialyzers was not made until 1938, when William Thalhimer [10] discovered that a membrane used in the sausage industry could be employed in removing solutes from the blood. This manmade cellulose hydrate material called cellophane was manufactured either by the Visking Company in Chicago, USA, or the Kalle Company in Germany. Cellophane membranes turned out to be uniform in thickness and strength, and one that could be produced in large quantities. Thalhimer experimented with dog blood and used for the ﬁrst time heparin as an anticoagulant in hemodialysis [11].

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1 Membranes in Hemodialysis

Fig. 1.3 Dialysis became a reality when Kolff developed his ‘‘Rotating Drum.’’ Kolff’s concept of a rotating drum with wound tubular membranes from Cellophane/Cuprophan was the basis of dialysis therapy until the early sixties of the last century. The right-hand photo

shows Kolff (right, and 89 years old in 1999) ¨rg Vienken explaining his ‘‘Rotating Drum’’ to Jo (left) and Horst Klinkmann (center) during a visit in Bad Homburg in 1999. A modiﬁed Kolffdialysis machine in use in Glasgow until 1965 is shown on the left.

This was the ﬁrst contribution to the long lasting success story of cellulosic membranes in hemodialysis that were later considered to be the golden standard. The next milestone in membrane application for hemodialysis was the development of the ‘‘Rotating Drum’’ by Willem Kolff. Supported by the local industry, Kolff [12–14] treated his ﬁrst patient in a hospital in Kampen, Netherlands. His dialyzer contained a tubular membrane made from cellophane. By this means, Kolff was able to separate 40 g urea from blood in 6 h (Figure 1.3). The ﬁrst dialyzer of Kolff [12] contained membrane tubes (Cellophane, later Cuprophan) with a sufﬁciently large membrane surface area and made an efﬁcient clinical dialysis possible. Kolff commented at that time: ‘‘It had an additional advantage: one could repair a leak in the cellophane tubing.’’ Kolff reported later that ‘‘Once a nurse handed me a pair of scissors over the artiﬁcial kidney. Between of us, we dropped it and there were light holes in the cellophane. We then stopped the rotating drum, cut out the damaged spots of cellophane tubing, spliced the two ends over a glass tube coated with soft rubber, and went on with the dialysis’’ [14] (Figure 1.4). According to Klinkmann, Falkenhagen, and Courney [3], the Swede Niels Alwall can be considered to be the inventor of ultraﬁltration. Ultraﬁltration needs a pressure difference as a driving force and can separate molecules with molecular weights up to 1 million by convectional forces. Alwall [15] used a device that allowed to exert a negative pressure on the dialysis membrane by inserting the membrane into a pressure-stable housing. By this means, he could control ultraﬁltration and consequently overhydration in uremic patients.

1.2 Historical Achievements

Fig. 1.4 The chemical formula of the major membrane polymers. Cellulosic membranes are mainly hydrophilic by means of their hydroxyl groups, whereas most synthetic membranes are basically hydrophobic and have to be rendered partially hydrophilic by adding hydrophilizing agents, such as polyvinylpyrrolidone.

The ﬁrst artiﬁcial kidney with tiny capillary membranes was described by Richard Stewart in 1964 [16]. The membrane material used for the production of membranes at that time was cellulose acetate. Later, cellulose acetate hollow ﬁber membranes were accompanied by membranes from unmodiﬁed regenerated cellulose. Based on the traditional German Bemberg experience with the Cuprophan/Cuoxam process [17], the ENKA Company in Wuppertal, Germany (later Membrana GmbH, Wuppertal, Germany) developed the ﬁrst cellulosic tubular membranes in cooperation with the American nephrologist L. Bluemle, followed by the development of capillary membranes. For both the membrane types, cuoxam was used as a solvent for the membrane polymer. Cuprophan hollow ﬁber membranes represent the classical dialysis membranes with a dense morphology. Its ultrastructure consists of cellulosic microﬁbrils, which

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1 Membranes in Hemodialysis

leads to a high mechanical stability of the tiny capillaries, while maintaining its high ﬂexibility. As a consequence, it was possible to produce capillary membranes with very small wall dimensions: thickness between 5 and 11 mm is the standard till date. These small dimensions allowed the manufacturing of compact and handy dialyzers with a small extracorporeal blood volume. In 1974, the ENKA Company started the routine production of Cuprophan hollow ﬁber membranes [18] in Europe followed by the ﬁrst clinical studies in 1977. Since that time and as a standard, dialyzers contain about 11 000 capillaries representing a membrane surface area of >1 m2 [17]. In contrast to ENKA, CD Medical Inc. (USA), the inventor of capillary membranes for dialyzers, still continued to use cellulose acetate as a raw material for the following reason: by saponiﬁcation of cellulose acetate the morphology of the cellulose membrane could be rendered more spongeous. As a result, membrane transport properties could be improved and resulted in the formation of a membrane with an increased ultraﬁltration proﬁle. This ﬁnally leads to the development of the so-called high-ﬂux dialysis [19]. Today, standardly applied high-ﬂux membranes are typically made by new synthetic materials such as polysulfone (PSu), polyethersulfone (PES), polyacrylonitrile (PAN), polyamide (PA), or polymethylmethacrylate (PMMA). In 1983, Fresenius [20] started to produce hollow ﬁbers made of polysulfone for high-ﬂux dialysis in parallel to the development of dialysis machines with ultraﬁltration control. Ultraﬁltration control is the prerequisite for high-ﬂux dialysis; it allows to run membranes under high convective forces. In parallel to Fresenius, the Gambro Company [21] with its subsidiary in Germany started to manufacture polyamide capillary membranes for hemoﬁltration. In 1985, Henderson and Chenoweth published an article entitled ‘‘Cellulose membranes, time for a change?’’ [22]. Based on investigations on blood/membrane interaction, they challenged cellulosic membranes for their alleged lack of blood compatibility in terms of cell and complement activation. Previous experiments had shown that blood compatibility depends on the chemical composition of the membrane surface. An optimal surface modiﬁcation would then improve blood membrane interaction. Investigations lead to the development of membranes from both modiﬁed cellulose and synthetic polymers, which showed domain-like surface structures with lipophilic/hydrophilic areas [23]. Membranes with this composition proved to show reduced direct interaction with blood and its components and thus a better biocompatibility. Modiﬁed cellulosic membranes such as cellulose acetate and cellulose-ether Hemophan, as well as the synthetic counterparts of cellulosic materials, polysulfone, polyamide, or polymethylmethacrylate are examples for such biocompatible membranes. During recent years, the performance of dialysis membranes has been further improved through the introduction of nanotechnology into the production process. With the nanocontrolled spinning technology [24] deﬁned porosities in membranes could be achieved as found in the new Helixone membranes in the FX Series of dialyzers from Fresenius Medical Care. A further improvement for the Helixone membrane bases on a special Moire´-like ondulation structure of the ﬁber, facilitating

1.3 Membranes for Hemodialysis: Polymers and Nomenclature

the easy entry of dialysis ﬂuid into the center of the membrane bundle and thus optimizes the concentration gradient of diffusible uremic toxins.

1.3 Membranes for Hemodialysis: Polymers and Nomenclature

A wide spectrum of hemodialyzers and ﬁlters combined with a multitude of different membranes are currently offered in the market. In 2004, more than 650 different dialyzer types with membranes made of 22 different polymers were commercially available [25]. Chemical and physical behaviors of a dialysis membrane are primarily determined by its polymer composition. The sterilization mode of the ﬁnal dialyzer, resistivity against sterilizing agents possibly used for reprocessing of the dialyzer, and biocompatibility are inﬂuenced by the type of polymer. Polymers for dialysis membrane manufacturing typically originate from the textile industry due to its spinning expertise. The ideal polymer suitable for dialysis should easily be manufactured to a biocompatible membrane family whose members exhibit different hydraulic permeabilities and display a considerable physical strength and excellent diffusive properties, as well as the resistance to all chemicals and sterilizing agents used in hemodialysis procedures, including steam (temperature to stand: >121 8C). Additionally, some appreciate the ability of a hemodialysis membrane to adsorb endotoxins at the outer surface because it is of substantial beneﬁt against microbial contamination of the dialysis ﬂuid. However, only the polysulfone and the polyamide membrane families fulﬁll these demands completely at the moment. As can be concluded from historical details, hemodialysis membranes of today are high-tech products that are tailored to the scientiﬁc-based demands of hemodialysis therapy. Even ‘‘early’’ materials such as regenerated cellulose are subject to permanent improvement by their manufacturers, and actual types vary in several aspects from their counterparts made 10 years ago, although they may not reach the maximal demands formulated above. Two classes of materials are currently used for the production of dialysis membranes: cellulose and synthetics that can be divided into two and three subclasses, respectively. Some structural differences exist between these two classes beside their different polymers: cellulosic membranes have relatively thin walls (in the range of 6.5– 15 mm) and a uniform (symmetric) composition across their entire capillary wall in order to achieve a high diffusive solute transport. This thin membrane does not provide enough strength to withstand high ultraﬁltration rates as are employed in convective therapies. Therefore, most cellulosic membranes are not suitable for convective dialysis treatments like hemodiaﬁltration or hemoﬁltration with the only exception of cellulose triacetate (CTA) membranes. In contrast to the classical cellulose membranes, synthetic membranes exhibit a membrane wall thickness of 20 mm and more, which may be symmetric (e.g., PMMA) or asymmetric (e.g., Fresenius Polysulfone) (Figures 1.5 and 1.6).

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Fig. 1.5 Cross section of a capillary membrane made from the regenerated cellulose (Cuprophan). The membrane wall exhibits a homogeneous structure with a tiny wall thickness of 8 mm and an inner diameter of 200 mm.

1.3.1 Membranes from Regenerated Cellulose

For the production of cellulosic membranes, puriﬁed cellulose (e.g., linters) is solved in an ammonia solution of cupric oxide (origin of the brand name Cuprophan). The

Fig. 1.6 Cross section of a capillary membrane made from Fresenius Polysulfone. The polsulfone membrane is represented by a small layer of about 1 mm thickness at the luminal side of the capillary. A heterogeneous spongeous wall of about 39 mm thickness guarantees mechanical stability.

1.3 Membranes for Hemodialysis: Polymers and Nomenclature

cellulosic polymer chains are newly arranged in the spinning process, therefore, the term ‘‘regenerated cellulose’’ is used. The result of both the spinning processes is a macroscopically homogenous structure that is extremely hydrophilic, sorbs water, and thus forms a hydrogel. Actually, diffusion of solutes takes place through water swollen amorphous regions. Regions of crystalline cellulose give the necessary mechanical strength. Therefore, cellulosic membranes can be made very thin and down to a thickness of 6.5 mm. It is considered as an advantage that the hydrophilic or water attracting properties of the membrane allow only little protein adsorption and thus avoid secondary layer formation. The Japanese manufacturer Asahi Medical terms its cellulosic membrane ‘‘cuprammonium rayon’’ and the German manufacturer Membrana GmbH sells its product under the brand name Cuprophan. Originally, cellulosic membranes have all been of low hydraulic permeability. Today, dialyzers with regenerated cellulose membranes and higher ﬂuxes with ultraﬁltration coefﬁcients up to 50 mL/h mmHg (Bioﬂux, H 2000, IDEMSA, Spain) are available on the market. In RC-HP 400 the average diameter of pores is increased from 2.76 nm for the standard regenerated cellulose to 7.23 nm that was achieved in the production process by the following ways: (i) an alteration of the concentration in the polymer solution, (ii) the solvent content in the coagulation bath, and (iii) the velocity of membrane formation through the coagulation bath [26]. In order to maintain the physical strength, wall thickness also had to be elevated to 18.5 mm (RC-HP 400) and 20 mm (Xanthogenate), respectively. The sieving coefﬁcient (SC) for b2-m, however, is only 0.3 for RC-HP 400, and elimination takes place solely by membrane transfer. However, these more permeable membranes are not suitable for convective treatments, like hemoﬁltration or even hemodiaﬁltration, because their mechanical strengths are not appropriate and their sieving coefﬁcients for b2-m are too low. Membranes made of regenerated cellulose have – due to their low wall thickness – a good low-molecular-weight clearance that is comparable and in a good dialyzer even superior to low-ﬂux synthetic membranes. The other advantage of unmodiﬁed regenerated cellulosic membranes is their resistance to all sterilization modes. A disadvantage, however, is the poor biocompatibility, and bacterial products are not retained by some synthetic low-ﬂux membranes. 1.3.1.1 Modiﬁed Cellulosic Membranes As it became obvious that cellulosic materials and especially, the nucleophilic hydroxyl groups of the cellulose polymer interact with humoral systems of blood, for example, with complement proteins, modiﬁcations of the polymer were produced mostly in the direction of substitution of the hydroxyl groups [23] in the molecular backbone of the cellulose molecule. 1.3.1.2 Cellulose Acetates In cellulose acetate (CA) membranes, at least two of the three hydroxyl groups of the glucose monomer have been replaced by acetyl groups. In the presence of sulfuric acid, cellulose is treated with pure acetic acid and acetic anhydride to form a cellulose

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acetate ester. This ester bond is responsible for the thermolability – CA cannot be sterilized by steam – and the chemolability of the CA polymer at pH >7. Some reports exist about polymer degradation resulting in polymer fragments such as acetylated carbohydrate derivatives that may lead to adverse reactions in the patients within 24 h after cellulose acetate treatment, for example, anaphylactoid reactions, scleritis, iritis, tinnitus, conjunctivitis, red eye syndrome, visual loss, or hearing loss [27,28]. The acetate substitution makes the hydrophilic cellulosic backbone polymer more hydrophobic and allows some protein adsorption on the inner surface. Several types of CA membranes are currently on the market, differing in their degree of hydroxyl-group substitution, their hydraulic permeability, and their manufacturing process (meltspun or coextrusion). All these differences may lead to some small advantages of one CA over the other [29]. Altogether, CA membranes exhibit the good performance characteristics of their unmodiﬁed cellulosic counterparts but are improved in their biocompatibility, although not reaching the excellent proﬁle of some synthetic membranes. Althin Medical AB, Sweden (acquisition by Baxter Healthcare Corporation, USA, in the late 1990s) took over the Cordis Dow meltspun process from 1963 and produced a symmetric cellulose diacetate membrane under the brand name Althane that used to be available in a high-ﬂux as well as in a low-ﬂux version. Dialyzers with this kind of membrane are held responsible for a number of deaths among dialysis patients in Spain, Croatia, and Texas, USA, in 2001. Intensive investigations [30–32] revealed that residual perﬂuorocarbon (PF5070), a processing ﬂuid used for the repair of leaky dialyzers, is the reason for these fatal reactions. The production of this kind of Althane dialyzers had therefore been ceased in 2001, according to a press release of Baxter Healthcare in 2001. Cellulose triacetate with a thin uniform skin structure was ﬁrst used only as highly porous membrane for high-ﬂux dialysis or hemodiaﬁltration (Nissho/Nipro) and was the ﬁrst and only cellulosic membrane capable for more convective therapies than hemodialysis. Today low-ﬂux types are also available. The highly permeable CTA membrane is also suitable for continuous renal replacement therapy [33]. 1.3.1.3 DEAE-Modiﬁed Cellulose, Hemophan Assuming that cellulose acetate was the ‘‘oldest’’ polymer used for hemodialysis, N,N-dimethyl-aminoethyl (DEAE) modiﬁed cellulose was the ﬁrst cellulosic membrane that was modiﬁed on purpose in order to increase biocompatibility (brand name Hemophan, Membrana Germany, was available during the years in dialyzers from Baxter, Bellco/Sorin, B. Braun, Gambro AB, Haidylena, Cobe, IDEMSA, JMS, Kawasumi, NephroSystems/Meditech, and Nikkiso). Only 1.5 % of all hydroxyl groups of the cellulose molecule is replaced in this polymer by tertiary amino groups through stable ether bonds [34]. These positively charged bulky groups ﬁrst set hydrophobic spots on a hydrophilic surface and then led to a steric hindrance of the interaction between complement factors and the membrane. A marked improvement of the biocompatibility proﬁle in comparison to unsubstituted regenerated cellulose with equal performance characteristics is observed. In respect to C5ageneration and activation of leukocytes no statistically signiﬁcant difference could be

1.3 Membranes for Hemodialysis: Polymers and Nomenclature

observed during dialysis with DEAE-cellulose in comparison to ETO-sterilized polysulfone low-ﬂux dialyzers. But the positively charged DEAE groups are under suspicion to lead to an increased platelet activation as well as to heparin consumption during treatment. Heparin adsorption could be demonstrated with DEAE cellulose from saline as well as from blood [35,36]. Advantageous effects are found in that the positively charged DEAE groups attract the negatively charged phosphate molecules from plasma, improving phosphate clearance along with that [37,38]. DEAE-cellulose exhibits the hydraulic properties of regenerated cellulose (Cuprophan) and is available as low-ﬂux and high-ﬂux types (Hemophan HP, UFcoeff 24 and 32 mL/h mmHg, respectively, g-wet Nikkiso [39]). Due to its stable ether modiﬁcation of cellulose, the membrane is sterilizable by all current methods. In early 2006, Membrana GmbH, Wuppertal, Germany, announced the termination of the production of cellulosic membranes by the end of 2006. Thus, membrane polymers, which have inﬂuenced the development of kidney therapy to a high degree over decades, will disappear from the market and leave the forum to their synthetic counterparts. 1.3.1.4 Benzyl-Modiﬁed Cellulose (Synthetically Modiﬁed Cellulose, SMC) Another example for the creation of hydrophobic domains on a hydrophilic surface is synthetically modiﬁed cellulose) produced by Membrana GmbH Wuppertal [40]. It is available in different housings from different manufacturers (under the brand name Polysynthane from Baxter, as SMC from Bellco/Sorin, B. Braun and Kawasumi). In this polymer, less than 1 % of the hydrophilic hydroxyl groups (OH) of the cellulosic backbone have been replaced by hydrophobic benzyl groups through ether bonds. This modiﬁcation leads to an improved biocompatibility in comparison to unmodiﬁed regenerated cellulose, although it does not reach that of low-ﬂux polysulfone [41]. Furthermore, biocompatibility parameters such as elastase release, complement activation, and leukopenia are timely delayed during dialysis with benzylcellulose: the nadir of leukopenia is at 30 min rather than around 15 min as it is with other modiﬁed or unmodiﬁed cellulosic membranes. Maximum complement activation and elastase release are also observed after 30 min [42]. However, all the cellulosic properties are sustained in this low-ﬂux membrane type, which is available as for sterilizability with all current methods, mechanical strength, and good low-molecular-weight clearances [43]. 1.3.1.5 PEG-Grafted Cellulose Improvement of biocompatibility with cellulosic membranes was also achieved by grafting the cellulosic backbone of cuprammonium rayon with a polyethyleneglycol (PEG) layer (AM-BIO membrane, Asahi Medical Co. Ltd, Tokyo, Japan). Alkylethercarboxylic acid (PEG acid) is esteriﬁed with its terminal carboxyl group to the hydroxyl group of the cellulose [44]. These PEG chains form a so-called hydrogel layer on the cellulosic surface (thickness 2.4 nm) that may act as a buffer zone between the cellulosic backbone and blood hindering the direct contact of plasma proteins with the membrane surface. This may lead to a reduction in platelet adhesion and

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complement activation. Lower platelet adhesion could be observed with this kind of membrane in vitro in comparison to unmodiﬁed and PEG-grafted cellulosic membranes [45]. 1.3.1.6 Vitamin E-Modiﬁed Cellulosic Membranes The ﬁrst attempt to create a ‘‘bioreactive’’ dialysis membrane is the development of a vitamin E-coated (D-a-tocopherol) cellulosic membrane with the brand name ‘‘Excerbane’’ from the Terumo Company in Japan (today in 2006: ASAHI Medical [46]). Here, performance capabilities of a high porosic cellulosic membrane are combined with the biocompatibility features of a synthetic copolymer together with the therapeutic approach to reduce the oxidative stress during treatment by supporting the body’s own antioxidant defense mechanisms with the supplementation of vitamin E [47]. The surface modiﬁcation is carried out during the ﬁber spinning process: the modifying solution consists of a hydrophilic acrylic polymer with reactive epoxy groups. This polymer, based on ﬂuorescein, possesses an inhibitory effect on complement activation. A second polymer, an oleyl alcohol chain, shows inhibitory properties with platelet aggregation. Both are dissolved into the core solution. The original solved regenerated cellulose and the core solution, containing the modiﬁer, are both coextruded through the spinerette into the coagulation bath, where two phases are formed. The outer circumference of the hollow ﬁber is composed of the cellulosic membrane, and the primary hydrophilic inside is covered by a hydrophobic layer of the modiﬁer. This takes place by covalent bonding of the reactive epoxy group of the modiﬁer with the hydroxyl groups of the cellulosic membrane. The amount of vitamin E immobilized via hydrophobic bonding to oleyl alcohol is around 150 mg/m2 [48]. Some clinical reports exist already with this kind of membrane showing an improved biocompatibility in comparison to regenerated cellulose. The long-term therapeutical effect of this kind of vitamin E supplementation has to be further elucidated. 1.3.2 Synthetic Membranes

The majority of synthetic membrane polymers currently on the market are basically hydrophobic and have to be made more hydrophilic by additives or copolymers during the production process. The polymer mixture is extruded through a spinerette followed by phase inversion and immersion. Partial evaporation of the solvent is responsible for skin formation. Phase separation determines the structure of the membrane as well as its crystallinity. The main purpose to develop synthetic membranes was to create membranes with higher porosity in order to mimic more the natural kidney ﬁltration process [20] and remove middle molecules and higher molecular weight uremic toxins like b2-m. As an indication for a possible pore size dimension of high-ﬂux dialysis membranes, the

1.3 Membranes for Hemodialysis: Polymers and Nomenclature

molecular radius (Stokes radius) of b2-m is described as being 1.6 or 2.2 nm. It was calculated that the pore radius of a b2-m removing membrane should be greater than the double radius of the molecule, which is approximately 5 nm for b2-m removal and lower than 8 nm to avoid albumin loss [49]. 1.3.2.1 Polyacrylonitrile (PAN) The company Rhone Poulenc in France has been the ﬁrst manufacturer who brought a highly permeable, symmetric, synthetic membrane on the market, which is a blend of a copolymer of the hydrophobic polyacrylonitrile with the hydrophilic methallylNa-sulfonate. The medium-sized pores are distributed in high density over the homogenous polymer, and glycerol is used as a pore ﬁller. This membrane produced and marketed under the brand name AN69 by Hospal (Gambro) till date has been of great success over the years. ASAHI Medical is the other PAN producer (PAN, PAN DX, ASAHI Medical, Japan). Their membrane consists of the hydrophobic monomers, acrylonitrile and methacrylate and is made hydrophilic by the addition of acrylic acid. Due to the special process, the membrane is asymmetric and exhibits a skin layer with pores in a wide range determining the sieving properties of the membrane. Until the early 1990s AN69 was held as one of the most biocompatible dialysis membranes, although it has one drawback: like all PAN membranes it is not sterilizable by heat. Moreover, since 1990, reports were published reporting anaphylactoid reactions with AN69 in combination with the consumption of angiotensinconverting enzyme (ACE) inhibitors [50–53]. Contact phase activation of the kallikrein–kinin system on the negatively charged surface of the AN69 membrane resulting in bradykinin formation was identiﬁed as the underlying mechanism. Surface electronegativity (zeta potential, see also Table 1.4) of the membrane, the pH of the rinsing solution [54,55], as well as the dilution factor of the plasma were found to be the inﬂuencing factors for the extent of reaction. These observations have also been found true for the second polyacrylonitrile membrane currently on the market, which also leads to bradykinin generation but to a lower extent: its zeta potential was found to be lower, which was 60 mV in comparison to 70 mV of AN69. Due to its microstructure and its surface electronegativity, AN69 exhibits through negative charges of the sulfonate groups, a high adsorption capacity for proteins, especially in positively charged proteins [58]. This was proven for complement factor D, b2-microglobulin, and low-molecular-weight proteins [56,57]. This feature is of advantage with respect to complement activation and b2-m elimination, but negative with respect to high-molecular-weight kininogen adsorption resulting in contact activation and higher residual blood volume in comparison to polysulfone membranes. Furthermore, therapeutic proteins like erythropoietin (EPO) are also adsorbed, probably diminishing their effectiveness. In cases where EPO is administered subcutaneously after hemodialysis treatment, this is of no clinical importance. But protein adsorption in substantial amounts may be a disadvantage, if dialyzers should be reused, because they have to be cleaned with aggressive chemicals like sodium hypochlorite that may damage their ﬁbers in order to rebuild their performance.

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Both polyacrylonitrile membranes are available only with high hydraulic permeability suitable for high-ﬂux dialysis, hemodiaﬁltration, and hemoﬁltration. AN69 is additionally available in ﬂat sheet format incorporated into plate dialyzers for acute hemoﬁltration. In order to overcome problems with anaphylactic reactions, Hospal developed a new AN69 type by coating the polyacrylonitrile ﬂat sheet membrane with polyethylenimine (PEI), a polycationic polymer [59]. During the manufacturing process PEI is sprayed onto the membrane surface until an optimized concentration of 9 mg/m2 is reached. In the resulting AN69ST (ST stands for surface treated [59]) the zeta potential is reduced to around 0 mV. Negatively charged sulfone groups of the AN69 polymer should be masked by the polycationic polymer and further provide a steric barrier via its thickness of the layer. First reports exist about clinical experiences with such a modiﬁed plate dialyzer (Crystal ST): No hypersensitivity reaction could be observed during 3 years and 10 630 dialysis sessions, among them 3400 with ACE inhibitor therapy. These data have to be conﬁrmed by other centers to make sure that this new membrane is safe for the patients even in combination with ACE inhibitor therapy [60]. Coating the membrane with a positively charged polymer furthermore leads to another effect, the binding of negatively charged heparin. First clinical trials show that if heparin was brought onto the membrane during the priming procedure no further systemic heparinization was needed during treatment. Heparin release from the membrane was undetectable. This anticoagulation regimen was possible only with unfractionated heparin, whereas low-molecular-weight heparin desorbs quickly. It will be of great interest to know how the introduction of positive charges on the surface will alter biocompatibility in comparison to the ‘‘old’’ AN69 type. That biocompatibility will be different is quite sure because changes in electronegativity of the surface have a great impact on protein adsorption. 1.3.2.2 Polymethylmethacrylate (PMMA) PMMA membranes were introduced into the market in 1977 in Japan by Toray Industries and have been the ﬁrst g-ray sterilized synthetic membranes [49]. They consist of a hydrophobic, nonpolar polymer produced from methylmethacrylate monomers, homo-PMMA, or are in some type also copolymerized with the addition of small amounts of p-styrene sodium sulfonate. The polymer forms a membrane that is symmetric, almost homogeneous, and isotropic. In current PMMA membranes the pore radius varies between 2 and 10 nm, having a volume fraction of pores (porosity ranging from 50 to 70 % [49]). A membrane family consisting of nine different members was created with lowﬂux types (B2, B3 series), high-ﬂux types (B1 series), which are mostly used for HD today, and a version with increased b2-m removal (BK series U/P/F). The last was achieved by increasing the pore size, resulting also in increased adsorption properties for b2-m of the membrane [61]. The BK-F model is the most strongly b2-m adsorbing HD-membrane currently on the market, where convective b2-m removal is negligible [49,62]. The larger pores (10 nm) also allow the removal of

1.3 Membranes for Hemodialysis: Polymers and Nomenclature

larger uremic toxins. An erythopoiesis-inhibiting fraction, KR4-0 and its subfraction YS-1 (MW 40 000), has been isolated from dialyzate of a PMMA BK-F HD treatment [63]. A clinical beneﬁt could be demonstrated through the fact that EPO doses could be diminished to half in patients receiving BK-F treatment after 2 years in comparison to HD with standard cellulosic membranes. Whether this is a unique property of the BK-F membrane or it can also be achieved with other highly permeable membranes has to be further elucidated. A multicenter, randomized, controlled trial with 84 patients for 12 weeks could not conﬁrm any effect on anemia with highly porous membranes (BK-F PMMA) in comparison to low-ﬂux cellulose hemodialysis [64]. The strong adsorbing property of PMMA has also some disadvantages: it also results in an undesired adsorption of platelets to the membrane with all its effects on ﬁbrin formation. A higher residual blood volume was also reported for PMMA membranes. Due to the negative surface, anaphylactoid reactions occurred under ACE inhibitor therapy. Therefore, with respect to biocompatibility, the membrane is placed only at an upper level among the synthetic membranes. All dialyzers containing PMMA are sterilized by g-irradiation. The material is not steam sterilizable. 1.3.2.3 Polysulfone (PSu) Not less than 18 suppliers of polysulfone dialyzers are currently on the market underlining the great success of this membrane polymer. Polysulfone ﬁts all demands of a modern polymer: It is sterilizable with all methods (g-ray, b-ray, ethylene-oxide, steam), biocompatible, has physical strength, and chemical resistance. The membrane material exhibits its good performance characteristics both in its low-ﬂux as well as in its high-ﬂux versions that remove considerable amounts of b2-m by ﬁltration. Moreover, polysulfone is suitable as an endotoxin adsorber and thus an active protection system for contaminated dialysis ﬂuids (Table 1.2). Because of all these advantages more and more membrane producers have developed their own polysulfone, although it is sometimes hidden among difﬁcult nomenclature. Fresenius introduced the ﬁrst high-ﬂux polysulfone in 1983 [20], followed by the low-ﬂux version in 1989. Table 1.2 provides an overview about the different manufacturers of polysulfones and some special features of the particular polymer. Due to patent protection, all polysulfones developed till date have to be different from the original Fresenius Polysulfone. They differ in their basic copolymer/polymer alloy, the addition of polyvinylpyrrolidone (PVP) (polysulfone alone is hydrophobic and has to be made more hydrophilic, which happens in most cases by blending the polymer with the hydrophilic PVP or not), and in their entire production processes, resulting in different morphologies. In chemistry, the terminus ‘‘polysulfone’’ comprises simply a group of polymers containing sulfone groups and alkyl or aryl (e.g., arylether) groups. However, according to chemical convention, all such polymers that additionally contain isopropyliden groups are termed as polysulfones (Fresenius Polysulfone, Asahi Polysulfone, Toraysulfone). Those dialysis membrane polysulfones that do not contain isopropyliden groups are termed as polyarylethersulfones or shortly as

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1 Membranes in Hemodialysis Membranes of the polysulfone (PSu) and polyarylethersulfone (PES) family on the market and their characteristics.

Tab. 1.2

Manufacturer

Brand name

Sterilization

Flux

Fresenius Medical Care

Fresenius Polysulfone

InLine steam ETO e-beam

Low

Helixone ASAHI Medical Gambro AB Hospal/Cobe Kimal Membrana GmbH (Allmed, Baxter, Bellco, Helbio, Haidylena, IDEMSA, Kawasumi, Saxonia) Minntech Nikkiso Saxonia (B. Braun) Toray Industries

High Low High High

APS Vitabran E Polyamix (PES) PolyamideSa Arylane (PES) Polyethersulfone DIAPES

g-wet g-wet Heat g-dry g-dry b-dry

Low High High High Low

Purema (PES) Minntech PS Polyphen PEPA a-Polysulfone Toraysulfone

g-dry, steam ETO ETO g-wet g-dry g-wet

High High High High High High

a According to the manufacturer, PolyamideS is a copolymer of polyarylethersulfone and only a small amount of polyamide. Therefore, it is mentioned under polysulfones [65].

Source: data from Ref. [1].

polyethersulfones (DIAPES, Arylane). This is a little bit confusing because, as mentioned above, all dialysis membrane polysulfones include an arylether. PEPA and PolyamideS contain another polymer in addition to polyarylethersulfone: PEPA polyarylate and PolyamideS polyamide and PVP [66]. The polyamide in PolyamideS is a matter of debate at the moment because some investigators could not ﬁnd any polyamide in PolyamideS [65]. PEPA is the only polysulfone membrane that does not contain PVP and therefore exhibits some special characteristics: it adsorbs larger quantities of b2-m, which is uncommon for all other membranes, but despite this adsorption removal rates for b2-m are even lower than that of the other polysulfone membranes. Recently an improved version of the original Fresenius Polysulfone, the Helixone membrane, was developed [24,67]. The polymer is unchanged, but the wall thickness (from 40 to 35 mm) and inner diameter of the ﬁber (from 200 to 185 mm) are reduced. By applying nanotechnology-based fabrication procedures [24,67] for the ﬁrst time in hemodialysis, the nominal average pore size has been increased in Fresenius Polysulfone from 3.10 to 3.30 nm. With the advanced production process, it was possible to create an almost uniform pore distribution at the dense innermost layer and a homology in pore size that results in a sharper

1.4 Dialyzer Constructions

molecular weight cut-off. The sieving coefﬁcient for b2-m was extended to 0.8, whereas the sieving coefﬁcient for albumin was preserved in the range between 0.001 and 0.01. First clinical studies revealed no difference in biocompatibility in comparison to that of Fresenius Polysulfone, due to no changes in polymer composition [68]. A considerably better performance was found with helixone due to higher mass transfer coefﬁcients, when compared with that of urea, creatinine, phosphate, and b2-m clearances, which is explained by a new geometric dialyzer, ﬁber design and pore size dimensions [69]. 1.3.2.4 Polyamide (PA) Polyamide membranes (in Polyﬂux dialyzers, Hemoﬂux hemoﬁlters, and FH hemoﬁlters, Gambro Hechingen, Germany) consist of a hydrophobic aromatic-aliphatic copolyamide that is blended with hydrophilic polyvinylpyrrolidone [70]. This mixture leads to a microdomain structure with alternating positively and negatively charged regions on the surface. This is held responsible for the good biocompatibility of the polymer. The membrane is asymmetric with three distinguishable regions: a thin skin of 0.1–0.5 mm on the blood side followed by a sponge structure of 5 mm that is supported by a ﬁnger structure of about 45 mm. Pore size increased dramatically from the blood side to the dialyzate side, being smallest at the skin layer with around 5 nm. Polyamide dialyzers and ﬁlters exhibit good b2-m removal with a SC of 0.6. This is due to their performance characteristics and not due to adsorption because protein adsorption is very low with this membrane. However, bacterial products are successfully rejected at the dialyzate side. Polyamide dialyzers and ﬁlters are available only in ETO-sterilized form. In order to overcome this disadvantage, polyarylethersulfone was added to the blend. The new membrane PolyamideS is heat sterilizable and described under polysulfone membranes.

1.4 Dialyzer Constructions

Biocompatibility, the size of removed particles, and the possibility of sterilization mode are mainly determined by the dialysis membrane, and all other aspects of effective dialysis are related to the design of a dialyzer. Current dialyzers are available only as hollow ﬁber devices. 1.4.1 Hollow Fiber Dialyzers

A current hollow ﬁber dialyzer consists of a housing that contains one membrane ﬁber bundle. This is ﬁxed at its both ends in the housing by polyurethane (PUR). An advanced cutting process is necessary to form a smooth surface after embedding in order to minimize activation of humoral or cellular systems in the blood. This surface

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is covered by caps that contain the ports for the inlet and outlet of blood respectively, and in recent developments also for dialysis ﬂuid. Normally the dialyzer housing contains two ports for the inlet and outlet of the blood and the dialysis ﬂuid, respectively. Behind this relatively simple construction stand the 60 years of knowledge and development. The trend to minimize devices is evident, if Kolff’s artiﬁcial kidney from 1943 [12] is compared to the latest dialyzer of today. Current dialyzers are optimized with respect to nearly each component. 1.4.2 Housing

The size of each dialyzer housing is tailored to the performance characteristics of the included ﬁber bundle. The blood compartment volume and resistance to blood ﬂow have to be as low as possible, and each ﬁber has to be equally washed round with dialysis ﬂuid. The number of ﬁbers and the ﬁlling degree of the dialyzer housing increase with surface area until an increase in surface is inefﬁcient for a rise in performance. Then the housing has to be increased. The comparison of dialyzers from various manufacturers that contain the same membrane clearly demonstrates the differences and what an advanced technology stands behind the creation of a good housing design. Current trends in device technology are the introduction of materials that can be disposed in an environmentally friendly manner. Polypropylene is such a material that is used in the newest housings instead of polycarbonate [69]. Furthermore comparing dialyzer performances, housings are getting smaller and smaller since the development of the ﬁrst disposable hollow-ﬁber dialyzer. This is possible because membranes with higher performance characteristics are available due to reduction of wall thickness, reduction of inner surfaces of ﬁbers, and a special bundle design. Smaller dialyzers avoid waste, have a sparing effect on transport costs, and are beneﬁcial for the patient because blood contact surface area as well as blood volume in the extracorporeal circuit is reduced. An advantage for the clinic is the saving of storage space. 1.4.3 Potting Material

Composition of the potting compound changed over years in order to minimize risks of toxic substances that may evolve after sterilization of the polyurethane. Especially, the irradiation with b- or g-beams may lead to the ﬁssion product 4,40 methylene dianiline, a proven carcinogenic substance [71]. The amount of PUR was considerably reduced over time, as it was obvious that it functions as a reservoir for ETO leading to allergic reactions in the patients [72]. Furthermore, effective surface area is increased with less potting area. Some manufacturers treat the reduced surface of the potting area in a special manner in order to avoid blood activation (e.g., Hospal Arylane series, FMC FX class series). Also, polycarbonate and/or

1.4 Dialyzer Constructions

silicon rings were introduced that allowed the PUR content of the hollow ﬁber dialyzers to be reduced. 1.4.4 Fiber Bundle

Of considerable importance for the performance of a dialyzer is the ﬁber bundle construction that has been dramatically improved till date. In early dialyzers, kinked ﬁbers resulting from insufﬁcient potting technology were a big problem leading to considerable thrombus formation. In modern high quality dialyzers, this problem has been overcome due to sophisticated embedding and cutting procedures. Even ﬁber distribution is nearly uniform in high quality products. In order to improve dialysis ﬂuid ﬂow around the ﬁbers, several bundle conﬁgurations have been developed and tested. The dialysis ﬂuid ﬂowed across the ﬁbers rather than along them. Multiple bundles, with the use of a solid central core with ﬁbers wound in a spiral manner, or warp knitted hollow-ﬁber mats have been introduced into dialyzer technology. Today, the newest ﬁber developments use an undulation of the ﬁber itself in order to provide space for a continuous uniform ﬂow of the dialysis ﬂuid. Besides this, ﬁbers with ﬁns and bundles with spacer yarns are also on the market [73–76] (Figures 1.7 and 1.8). Fiber bundle size and swelling of the membrane determine the priming blood volume, which is an important parameter in the choice of the dialyzer for patients with low blood volume, especially children. Today, the blood volume in most dialyzers is smaller than that of the blood tubing sets.

Fig. 1.7 Dialyzer performance considerably depends on the arrangement of capillaries in the membrane bundle. In order to allow for an easy access and ﬂow of dialysis ﬂuid to the center of the bundle, capillary membranes have to be either separated through the insertion of spacer yarns (b in left-hand panel) or have to be

ondulated (a in left-hand panel). An ondulation ´ structure, as following the geometry of a Moire realized in FX-dialyzers with helixone membranes (c) guarantees optimal clearance characteristics through a homogeneous ﬂow of dialysis ﬂuid. Source: modiﬁed from Ref. [73].

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Fig. 1.8 The new FX-class of dialyzers exhibits improved clearance characteristics due to a modiﬁed PSu membrane and a special ondulation conﬁguration of the capillaries. As a consequence, urea clearance is increased as compared to traditional dialyzer conﬁgurations. This allows for reducing dialysis ﬂuid consumption at comparable blood ﬂows.

1.5 Dialysis Membranes and Performance: Principles of Membrane Transport

Dialysis is a membrane separation process in which one or more dissolved species ﬂow across a selective barrier in response to a difference in concentration and a difference in pressure. Concentration differences refer to the mode of transport of diffusion, where separation occurs because small molecules diffuse more rapidly than larger ones. In the absence of difference of pressure and temperature across the membrane, A. Fick’s phenomenological description of diffusion, published in 1855 [77], states that solutes will move from regions of greater to regions of lesser concentration [DC] and at a rate proportional to the differences (Figure 1.9). J ¼ DADC=Dd;

ð1Þ

with J representing solute ﬂow, D the diffusion coefﬁcient, A the area, and d the distance between the two separated compartments, that is, the membrane thickness in the dialysis situation. The minus sign for the diffusion coefﬁcient accounts for the

1.5 Dialysis Membranes and Performance: Principles of Membrane Transport

Fig. 1.9 Annual growth rates for dialyzer sales show that high-ﬂux and synthetic membranes have become the dominating products in dialysis therapy. Source: re. FMC market survey 2003.

convention that ﬂux is considered positive in the direction of decreasing concentration. According to Albert Einstein, diffusion concentration decreases roughly in proportion to the square root of molecular weight. In order to increase solute ﬂow, membrane manufacturers have tried for years to have a membrane thickness as small as possible. Complications may arise from the observation that blood in contact with artiﬁcial surfaces, like a hemodialysis membrane, leads to the formation of a secondary layer of proteins. This boundary layer affects diffusional transport across a membrane. A further effect for reduced diffusional clearance can be attributed to drug administration to the patient, such as human recombinant erythropoietin (rhEPO). Because the removal of solutes by the process of hemodialysis is dependent on the ﬂow of the solute and water from the blood compartment to the dialyzate compartment, it is comprehensible that the fraction of water in the blood is reduced by raising the hematocrit (Hct). Indeed, increasing Hct from 20 to 40 %, leads to a reduction of creatinine and phosphate clearances of 8 and 13 %, respectively. Urea removal, however, is less affected [78]. Based on these observations, it appears prudent to increase hemodialysis prescription, that is, treatment time, by 10–15 %, when Hct is raised to near 40. Dialysis membranes are primarily categorized according to their permeability for water and solutes. The ultraﬁltration factor (UF) and sieving coefﬁcients are the appropriate parameters. Ultraﬁltration factors refer to the amount of ﬁltered water in mL/h in relation to the applied transmembrane pressure (TMP) in mmHg, which is exerted through the blood pump of the dialysis machine. The sieving coefﬁcient is equivalent to the amount of a given solute removed (Figure 1.10): SC ¼

2CF þ CBo ; CBi

ð2Þ

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Fig. 1.10 The weekly performance of the human kidney overrides by far the performance of classical dialysis therapies, such as low- and high-ﬂux dialysis. A closer bridging the gap between the performance of conventional therapy and the human kidney will certainly be based on convective therapies such as hemodiaﬁltration.

whereby CF refers to the concentration of solute in ﬁltrate and CBi and CBo to the solute concentrations in blood inlet and blood outlet, respectively. Consequently, the maximum value for a sieving coefﬁcient is ‘‘1’’ representing a 100 % sieving. Following Table 1.3 and a general agreement, low-ﬂux membranes are characterized by an UF factor of less than 10 mL/h mmHg, whereas high-ﬂux membranes have an UF factor greater than 10 mL/h mmHg. In addition to their UF factors, hemoﬁlters and high-ﬂux dialysis membranes are characterized by their sieving coefﬁcient for b2-microglobulin (b2-m). b2-m is a small protein of a molecular weight of 11.818. High serum levels of b2-m are considered to be causative for dialysis-related amyloidosis. The molecular weight cutoff of high-ﬂux membranes, which refers to the molecular weight of molecules that cannot pass through the dialysis membrane, is typically around 60 000. It implies that albumin is retained in the patient’s blood.

Tab. 1.3

Categories for hemodialysis membranes.

Category Low ﬂux High efﬁciency High ﬂux Hemoﬁlter

UF factor (mL/h mmHg)

Cut-off molecular weight

Sieving coefﬁcient (for b2-m)

Surface area (m2)

<10 >10 >10 >20

–— –— <60 000 >60 000

–— –— >0.6 >0.6

<1.5 >1.5 –— –—

Low- and high-ﬂux refer only to the hydraulic permeability.

1.5 Dialysis Membranes and Performance: Principles of Membrane Transport

By increasing the transmembrane pressure (e.g., by means of the peristaltic pump), ﬁltration ﬂow will increase. This increase is almost linear as long as aqueous solutions are used. With whole blood, the increase of ﬁltration ﬂow tends to become nonlinear because of the formation of a secondary layer protein coat by protein adsorption, which is a fast process. Albumin, one of the proteins with the highest blood concentration, is already deposited in 50 ms after the contact with the dialysis membrane [79]. These observations have to be taken into account when ﬁxing ultraﬁltration rate and dialysis time. The amount of ultraﬁltered water, for example, in the treatment of acute renal failure, may determine the mortality of patients as shown by Storck et al. [80], who compared pump-driven and spontaneous continuous hemoﬁltration in postoperative renal failure. Increasing the ultraﬁltration volume per day improved mortality by about 20%. Recent investigations in the treatment of kidney patients have focused on the removal of large molecular weight solutes such as cytokines, complement proteins, and proteins modiﬁed either by glucose (advanced glycation end products [AGEs]) or by oxidative stress pathways. Transport of large solutes across a dialysis membrane is improved if convectional transport mechanisms are applied. Convective clearance is deﬁned by CConvection ¼ SC QF;

ð3Þ

and thus determined by the membrane’s sieving coefﬁcient (SC) and the ﬁltrate ﬂow (QF). It is possible to increase the sieving coefﬁcient for a given solute by switching from a low-ﬂux to a high-ﬂux membrane. Filtrate ﬂow depends on internal ﬁltration, which is deﬁned as the total water ﬂux across the membrane within the closed blood and dialyzate compartment of a dialysis ﬁlter. Internal ﬁltration depends on the pressure gradient characteristics (Dp) along the length of dialyzer and may be given by a simpliﬁed approach through the application of Hagen–Poiseulle’s law: Dp ¼

8hLQB ; NpR4

ð4Þ

where h is the blood viscosity, L is the length of dialyzer membrane, QB is the blood ﬂow, N is the number of capillaries in the dialyzer, and R is the internal radius of a capillary membrane. Consequently, a better convective clearance (higher Dp) directly depends on blood viscosity to be adapted through the administration of erythropoietin, the length of the dialyzer, and the blood ﬂow. Applying higher blood ﬂows, that is, >300 mL/min is therefore mandatory in high-ﬂux dialysis or in hemodiaﬁltration. Convective clearance depends indirectly on the number of capillary membranes in a dialyzer and on the membrane’s luminal radius to the power of four. Thus, reducing the inner diameter of a capillary membrane would considerably increase the convective clearance [73,81].

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Fig. 1.11 Data sheets of dialysis membranes refer to ultraﬁltration in relation to the applied transmembrane pressure. However, the amount of ultraﬁltrate (QF) depends on the exclusive presence of water/saline or blood. Due to a secondary layer formation by blood proteins, QF is leveling off at higher TMPs.

Modern capillary membranes, for example, Helixone polysulfone membranes (Fresenius Medical Care, Germany, inner diameter: 185 mm) have a reduced inner diameter as compared to the classical diameter dimension of 200 mm for typical dialysis membranes. Decreasing the inner diameter from 250 mm over 200 mm down to 175 mm has no impact on small solute clearances such as urea, but shows considerable consequences for the removal of larger solutes such as b2-microglobulin, for which convective clearance increases threefold [81] (Figures 1.11–1.13).

1.6 Dialysis Membranes and Biocompatibility 1.6.1 Some Basic Information on Membranes and Biocompatibility Parameters

The biocompatibility of a dialyzer, particularly that of its membrane, is one of the main criteria that inﬂuences dialyzer choice. During the ‘‘Consensus Conference on Biocompatibility held in 1993, biocompatibility was deﬁned as ‘‘the ability of a material, or device, or system to perform without a signiﬁcant host response in a speciﬁc application’’ [82]. The deﬁnition was weakened by adding ‘‘signiﬁcant’’ in comparison to ﬁrst explanations because it was realized over the years that any foreign material would lead to some kind of reaction in the host. In hemodialysis treatment, blood/artiﬁcial surface interactions (as special aspect of biocompatibility also termed as hemocompatibility) are of special importance because of its chronic treatment character.

1.6 Dialysis Membranes and Biocompatibility

Fig. 1.12 Reducing the inner diameter of a capillary membrane leads to a high pressure drop along the length of the dialyzer and thus favors convective forces for the removal of larger solutes. Here, a reduction

of the inner diameter of a polysulfone capillary membrane from 200 to 175 mm leads to a nearly 80 % increase of inulin- and vitamin B12-clearance. Source: modiﬁed from Ref. [73].

Fig. 1.13 Improved clearances can be obtained for small proteins, such as b2-microglobulin (b2-m), if the inner diameter of a capillary membrane is reduced. b2-m clearance is found to be up to threefold higher, when the inner

diameter is reduced from 250 mm over 200 mm down to 175 mm. With this simple change in ﬁber geometry convective therapies, such as hemoﬁltration or hemodiaﬁltration, gain an even better performance.

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Parameters of biocompatibility are manifold and their number increases continuously with scientiﬁc knowledge. In the beginning of chronic hemodialysis (CRF) treatment, thrombogenicity and hemolysis were the main matters of concern for physicians [83]. Manufacturers focused on the development of less thrombogenic surfaces and on the toxicity of the material that should not liberate plasticizers and chemical additives. In the late 1970s, hypersensitivity reactions in response to ethyleneoxide (ETO) sterilization of the dialyzer were reported [84]. At the same time, complement gained attention and was linked to the transient leukopenia and sequestration of neutrophils in the lung during dialysis with cellulosic membranes [85,86]. The inﬂuence of chronic hemodialysis treatment on the immune system was recognized in more detail over the years with the formulation of the ‘‘Interleukin Hypothesis’’ in 1983 being a milestone in this kind of research [87]. Stimulation of immune cells during dialysis, the release of mediators such as cytokines, and the pathological consequences thereof have from then on been topics of intensive research. In the early 1990s, the possible interaction between extracorporeal treatment and pharmacological action of some drugs became clinically evident: Anaphylactoid reactions had been observed in ACE inhibitor patients in combination with the use of certain negatively charged dialysis membrane, which until then had been considered ‘‘biocompatible’’ [88]. Investigations revealed that the kallikrein–kinin system was involved in this sort of side effects leading the attention to the new biocompatibility parameter bradykinin (Figure 1.14). What are the biocompatibility parameters of today? Some parameters can be measured relatively easily in the clinical setting, such as the differential cell count, whereas factors of the complement and clotting cascade need specialized laboratories. On a more in-depth scientiﬁc level, an assessment of gene expression in relation to cytokine or b2-microglobulin generation, and the stimulation of immune cells and their reactions (receptor expression, generation of reactive oxygen species [ROS], etc.) have become state of the art.

1.6.2 Thrombogenicity of Different Types of Dialyzers and Filters

The thrombogenic potential of a dialyzer or ﬁlter is strongly dependent on the type of membrane polymer, its permeability, the design of the device inﬂuencing ﬂow conditions, and its manufacturing quality. Surface free energy, charge, roughness, and chemical composition of a dialyzer membrane have been identiﬁed as being responsible for the variable thrombogenic potential of dialysis membranes [89]. As the extrinsic system of the clotting cascade is activated by negatively charged surfaces, membrane polymers that exhibit an anionic surface, such as polyacrylonitrile membranes (PAN DX, PAN; AN69) could be considered more thrombogenic than more neutral polymers such as polysulfone.

1.6 Dialysis Membranes and Biocompatibility

Fig. 1.14 Contact of blood with negatively charged biomaterial surfaces leads to the activation of the contact phase and subsequently to the formation of the nonapeptide bradykinin via the kallikrein pathway. If degradation of bradykinin by angiotensin-converting enzyme is blocked

through the administration of ACE-inhibitors, blood pressure is downregulated by prostaglandin generation. Generally speaking, synergistic effects between biomaterial or membrane properties and the administration of pharmaceutical drugs may not be neglected in the future.

In fact, the highest factor XII adsorption and autoactivation to factor XIIa can be observed with AN69 membranes in vitro [90]. However, the activation of the coagulation cascade also depends on other plasma constituents, for example, highmolecular-weight kininogen, prekallikrein, plasma inhibitors of XIIa, and the action of heparin [90]. High-molecular-weight kininogen (HMWK), another constituent of the trimolecular complex initiating contact activation, could be eluated in its intact form from the used cellulose acetate (Cordis Dow) and regenerated cellulose (Cuprophan), whereas it was cleaved at PMMA and polyacrylonitrile (PAN; AN69) surfaces suggesting activation [91]. Moreover, low-molecular-weight fragments of plasminogen were detected in eluates from polyacrylonitrile (AN69) dialyzers indicating that the ﬁbrinolytic system was activated in response to clotting activation. Intact plasminogen was adsorbed also with all other materials [91] (Figure 1.15). Clinical studies investigating coagulation activation with different types of membranes using the ex vivo formation of the thrombin–antithrombin III complex (TAT) as clotting parameter showed a higher TAT formation with polyacrylonitrile membranes (Asahi PAN and Hospal AN69) than with polysulfone (Fresenius Polysulfone) and regenerated cellulose (Cuprophan) [92]. In another study higher levels with PMMA than with EVAL or polysulfone [93] were shown. These ﬁndings were conﬁrmed in controlled ex vivo studies, which found higher TAT generation, platelet factor 4 release, and platelet consumption with AN69 (Nephral 300,

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Fig. 1.15 Bradykinin formation after in vivo contact of patient blood with the surfaces of different dialysis membranes. Bradykinin levels in the venous line of negatively charged PAN (AN69) dialyzers is far above normal and

explains adverse events in these patients. Note that the concentration of bradykinin in these cases is found to be in the femtomolar (fmol) range. Source: adapted from Ref. [157].

Hospal) in comparison to regenerated cellulose (Cuprophan, Renak RA-15U, Kawasumi) and polysulfone (Fresenius Polysulfone, Hemoﬂow F6HPS) [94]. A low TAT generation with regenerated cellulose is in agreement with the ﬁndings that hydroxyl-bearing surfaces are relatively inert with respect to activation of the intrinsic coagulation pathway; such membranes inhibit the development of the trimolecular complex (HMWK, PK Factor XII) that is necessary for contact activation [95]. DEAE-modiﬁed cellulose (Hemophan), a membrane that is a low level platelet activator, led to more TAT generations than high-ﬂux polyamide [96]. This is not the result of a higher contact system activation by DEAE cellulose, but it is rather due to the positive DEAE groups adsorbing negatively charged heparin from whole blood. This adsorption inactivates heparin and, consequently, sufﬁcient amounts of heparin are not available for anticoagulation. This is normally reﬂected in an increased heparin consumption for Hemophan during treatment [97,98]. However, another in vivo investigation failed to ﬁnd such an increased heparin requirement and also did not observe any increase in TAT formation with Hemophan in comparison to low-ﬂux polysulfone [99]. Another, albeit less reliable, parameter for the assessment of thrombogenicity is the residual blood volume (RBV) in the dialyzer postdialysis. This subjective parameter has a number of pitfalls that mostly concern the reproducibility of the method. Although TAT measurements are preferable, RBV values have been reported in the literature. Under conditions of low heparin dosage, residual blood volume was found to be higher with Hemophan than with CA, PMMA and low with Fresenius Polysulfone [100]. In contrast to coagulation factors, platelets adhere more to cationic charges and hydrophobic materials [95]. Therefore, it is not surprising that a signiﬁcant decrease

1.6 Dialysis Membranes and Biocompatibility

in the platelet count after 15 min of dialysis of about 9 % could be found with hydrophobic cellulose acetate dialyzers (Cordis Dow), whereas only insigniﬁcant decreases were observed with cuprammonium rayon (AM50-Bio, Asahi), DEAEcellulose (Hemophan, GFS, Gambro), polyacrylonitrile (AN69 Hospal), and high-ﬂux polysulfone dialyzers (Fresenius Polysulfone F60, FMC) [101]. In another study, a similar ranking in platelet drop was observed with the same polymers but different manufacturers: cellulose diacetate (CA 150 Baxter), Hemophan (Bio-Allegro, Cobe), cellulose diacetate (Acepal 1500, Hospal), and low-ﬂux polysulfone dialyzers (F6, FMC) [38]. One reliable parameter for platelet activation is the expression of the glycoprotein GMP-140 at the platelet membrane surface. The highest GMP-140 expression was found for Cuprophan dialyzers, less expression was observed with cellulose acetate and PMMA, and the lowest expression with polysulfone and polyacrylonitrile dialyzers [102]. This ranking for platelet activation was also found for another parameter, the expression of P-selectin on the platelet surface: Expression was highest with Cuprophan, followed by cellulose diactetate and cellulose triacetate, and was lowest with Hemophan and Fresenius Polysulfone [103]. Although platelets are obviously activated by regenerated cellulose, this activation does not result in or involve activation of the blood-clotting cascade, as is obvious from previously mentioned investigations into the formation of the TAT complex. A number of publications have dealt with theoretical considerations about artiﬁcial surfaces and their thrombogenic potential, but biological processes are too complex to easily predict the thrombogenic behavior from charges, functional groups, or other characteristics of current dialysis membranes. Differences between membranes, and even between dialyzers of different manufacturers containing the same membrane exist. This has been shown in controlled ex vivo or in vivo studies using the same blood for all membranes compared. However, the numbers of reports about membranes correlate with their frequency of use in the market and not necessarily with their thrombogenic behavior. Therefore, although no published data are available for some membranes this should not be interpreted as absence of thrombogenic potential per se.

1.6.3 Complement Activation by Different Dialyzers and Filters

It is accepted today that membranes made of regenerated, unmodiﬁed cellulose are the strongest complement activators among all dialysis membranes. This can be partly explained by the binding of complement factor C3b to the hydroxyl groups on the membrane surface. In fact, either partial substitution of these hydroxyl groups with acetyl groups (cellulose diacetate or cellulose triacetate), DEAE groups (Hemophan), and benzyl groups (SMC), or coating with polyethylene glycol (PEG-grafted regenerated cellulose) resulted in considerable reduction of complement activation compared to unmodiﬁed regenerated cellulose. Surprisingly, the

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degree of substitution does not necessarily correlate with the degree of complement activation, demonstrating involvement of other factors. For example, in Hemophan, less than 1 % of all hydroxyl groups are modiﬁed, but complement activation is at least equal to or even better than that of cellulose acetate membranes where about 60 (cellulose diacetate) to 90–100 % (cellulose triacetate) of all hydroxyl groups are substituted [23]. Binding of regulatory proteins that downregulate the alternative pathway such as ‘‘factor H’’ by Hemophan [104] or ‘‘factor D,’’ a rate-limiting enzyme of the alternative pathway by AN69 [105] and PMMA [106], also plays a role in the extent of complement generation. Factors H and B bind onto cellulose acetate membranes resulting in an accelerated degradation of surface-bound C3b; neither of these factors bind to regenerated cellulose [107]. However, all measures taken to prevent the interaction of C3 molecules with nucleophilic groups, for example, hydroxyl or amino groups, on a polymeric surface reduce activation of the alternative pathway. This can be achieved either by avoidance of such groups as in synthetic membranes or by masking of such groups with acetyl, DEAE, or benzyl groups (as in cellulose acetates, Hemophan, or SMC) or by coating the whole surface with, for example, PEG groups. In order to avoid activation of the classical pathway, none of the IgG, IgM, C1, C2, or C4 complement components should be adsorbed to the polymeric surface [108]. Parameters mostly used for the assessment of complement activation by hemodialysis membranes are the complement factors C3a and C5a. The amount of detectable C3a or C5a in the efﬂuent of a dialyzer depends on the capability of the membrane to adsorb or to ﬁltrate these factors besides its grade of generation: The polyacrylonitrile membrane AN69 generates more C3a than regenerated cellulose, but adsorbs C3a and C5a almost completely [109]. Furthermore, these anaphylatoxins are eliminated during dialysis by high-ﬂux membranes due to their middle molecular weight (MW C3a: 8900; C5a: 11 000) [110,111]. This is also the fact for the regulatory factor D (MW 23 000). This upregulator of the alternative complement pathway is eliminated by glomerular ﬁltration in the healthy individuals and its concentration is, therefore, markedly elevated in ESRD patients. Factor D removal is negligible in low-ﬂux dialysis but can be signiﬁcant in hemoﬁltration [112]. Hence, detectable complement generation may differ between the low-ﬂux and high-ﬂux versions of a particular membrane [113]. Furthermore, binding to receptors of circulating blood cells may lower the concentrations of C3a and C5a in blood. What are the long-term consequences of complement activation in the long-term perspective? Because complement generation peaks 15–30 min after the start of treatment and returns to nearly normal levels until the end of treatment, it is of interest whether this effect is important or detectable in the long-term perspective. Unfortunately, little data exist on this aspect of complement activation. Predialysis C3a levels were reported to increase slowly over a time period of 1 year with the low complement activating membranes, Hemophan and polyamide [114]. In long-term HD patients (around 8 years on dialysis), the C3 activity is altered in such a way that both the alternate and the classical pathways are suppressed after stimulation. This effect lasts more than 4 h after the end of a dialysis session and is more pronounced

1.6 Dialysis Membranes and Biocompatibility

in sera of patients treated with regenerated cellulose membranes (AM-SD18M, Asahi Medical, Japan), than in the sera of those treated with polyacrylonitrile membranes (PAN17DX, Asahi Medical). Therefore, it appears that due to the chronic high stimulation three times a week, membranes made of regenerated cellulose may induce a suppression of complement activation in the long term. The mechanism has not been fully elucidated, but a contribution of increased levels of the regulatory complement proteins, factor H and SP-40,40, has been discussed [115].

1.6.4 Cell Activation by Different Types of Dialyzers and Hemoﬁlters

Essential functions of polymorphonuclear leukocytes are disturbed in ESRD patients and are additionally inﬂuenced by the dialysis procedure, for example, phagocytosis, oxygen species production, upregulation of speciﬁc cell surface receptor proteins, and apoptosis. With the exception of polymorphonuclear leukocyte degranulation, complement and consequently, complement generating dialysis membranes have the greatest impact on functional alterations of these cells. Leukopenia, the most widespread used parameter to assess leukocyte activation, is the assessment of the disappearance of leukocytes from the blood 15–30 min after the start of hemodialysis. This dialysis-induced leukopenia is mainly induced by the overexpression of receptors (CD11b/CD18, CD15s), leading to an increased adhesiveness and aggregation with subsequent sequestration in the lung. Because these activation processes are mostly complement mediated, complement activation and leukopenia directly correlate with each other [116]. Therefore, regenerated cellulose membranes cause the strongest leukopenia, modiﬁed cellulosic membranes cause an intermediate-to-low drop in leukocyte numbers, depending on the type of modiﬁcation, and synthetic membranes also induce a moderate-to-very low drop depending on the polymer. Leukocyte drop is found to be similar to that of polyamide (Polyﬂux 130, Gambro) and polysulfone (Fresenius Polysulfone F60, FMC) after 15 min at approximately 12 %. Increasing the permeability of regenerated cellulose should, theoretically, result in an increased removal rate of complement factors and, consequently, an improved biocompatibility with respect to measurable complement generation and leukocyte drop. Surprisingly, this was not the case with RC-HP 400 – a high performance version of Cuprophan – which induces the same leukocyte drop as low-ﬂux Cuprophan [117].

1.6.4.1 Apoptosis Dialysis membranes promote neutrophil apoptosis both directly and through interaction with monocytes. Reactive oxygen species seem to be strong mediators of this process [118]. In in vitro experiments, low-ﬂux cellulose diacetate induced signiﬁcantly greater apoptosis than low-ﬂux polysulfone [118]. An effect of HD membranes

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Fig. 1.16 The effect of different dialysis membranes on apoptosis of mononuclear cells. During the in vivo study period, bacterial and endotoxin contamination was below detection limit. Cell apoptosis was measured in circulating MNCs isolated from heparinized blood. Source: adapted from Ref. [119].

on apoptosis was also found in vivo. The percentage of mononuclear cell (MNC) apoptosis was high with Hemophan and Cuprophan dialyzers, was moderate with cellulose diacetate, and was relatively low with both high-ﬂux polyacrylonitrile and polysulfone membranes. When seven of these patients were switched from Hemophan to Fresenius Polysulfone, apoptosis decreased markedly only after 8 weeks of treatment [119]. This in vivo study was repeated in vitro with mini dialyzers of the membranes used and with blood from healthy donors in order to eliminate the inﬂuence of uremia. Interestingly, the results were quite similar implicating a role of the type of dialysis membrane in inducing apoptosis beside uremia [119] (Figure 1.16). 1.6.5 Oxygen Species Production – Induction of Oxidative Stress

Dialysis treatment enhances oxidative stress in chronic renal patients who additionally suffer from a chronic deﬁciency in the major antioxidant systems [120,121]. Detectable manifestations of this physiological condition are the increased plasma levels of protein oxidation products (oxidation of plasma protein-associated thiol groups) in HD patients [122]. Oxidative stress may contribute to atherosclerosis, cardiovascular disease, dialysis-related amyloidosis, and anemia [123,124]. Four main factors are held responsible: uremia and the comorbid status of the ESRD patient, antioxidant and trace element losses during treatment, the bioincompatibility of the system, especially of the dialysis membrane, and contamination of blood with trace amounts of endotoxins introduced by microbial contaminated dialysis ﬂuid [125,126] (Figure 1.17).

1.6 Dialysis Membranes and Biocompatibility

Fig. 1.17 Relative amount of IL-1b gene expression in peripheral blood mononuclear cells from healthy volunteers, nondialyzed uremics and chronic hemodialysis patients treated with different dialyzers. Maximal accumulation of IL-1b mRNA was observed in

blood samples taken from the venous line after 5 min of dialysis. IL-1b gene expression could be detected in cells from all HD-patients, but not in healthy volunteers and nondialyzed uremics. Source: adapted from Ref. [144].

The highest increase in reactive oxygen species production during dialysis was reported in complement activating membranes [126,127]: Cuprophan induced a signiﬁcant production of reactive oxygen species 15 and 30 min after the start of dialysis. A lower production was observed with Hemophan, cellulose diacetate and low-ﬂux Fresenius Polysulfone, whereas increases with SMC were not even statistically signiﬁcant [127]. Another clinical study reported low ROS production with polyacrylonitrile (AN69) and EVAL [128]. However, ROS production in response to stimulation by endotoxins is suppressed in patients treated with Cuprophan membranes and normal with low-ﬂux polysulfone or PMMA membranes [126,129]. Apart from measures regarding the improvement of treatment such as the use of more biocompatible membranes and pyrogen-free dialysis ﬂuid, another approach to reducing the possible deleterious consequences of the prooxidant status has been followed: supplementation with antioxidant vitamins such as vitamin E [130]. In addition to oral prescriptions, some clinical experience has been made in the vitamin E-coated membrane Excebrane: a 10 % predialysis increase of plasma vitamin E concentrations (exclusively in HDLs not in LDLs) was observed after a 3-month treatment period with this membrane, whereas plasma levels did not change when conventional biocompatible membranes, AN69, Fresenius Polysulfone, or PMMA membranes were used. This effect is believed to be due to the vitamin E sparing effect by the Excebrane membrane [130]. The b-carotene content of plasma and the LDL and HDL levels were about 26 % higher than when other biocompatible membranes were used, a ﬁnding which can be explained by the secondary protective effect of vitamin E toward b-carotene [131]. However, no improvement in the blood oxidative stress status could be found in this study, which used determinations of thiobarbituric acid-reactive substances and

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antioxidant defenses, like erythrocyte-Cu, Zn-superoxide dismutase and plasma, and erythrocyte glutathione peroxidase [130]. In contrast, another investigation detected an improvement in oxidative stress in the form of a threefold increase in plasma glutathione concentration and a slight increase in erythrocyte numbers after three months of treatment [132]. In this study, the improved antioxidant status (increased vitamin E and glutathione peroxidase concentration in plasma) resulted in a 10-fold increase in plasma levels of arachidonic acid, one of the most abundant polyunsaturated lipids and a useful marker for lipoperoxidation. Leukocyte function (responsiveness to chemical stimuli) was signiﬁcantly improved after one month of treatment with Excebrane, and the number of apoptotic cells was reduced in comparison to dialysis with regenerated cellulose [132]. An increase in oxidative markers, like malondialdehyde, advanced glycation end products, and 8-hydroxydeoxyguanosine was prevented with Excebrane but not during dialysis with Terumo polysulfone [133]. Another study used plasma vitamin C levels and constant ascorbyl free radical (AFR)/vitamin C ratio postdialysis as an index of oxidative stress. After AN69-dialysis, basal vitamin C levels were decreased and the AVR/vitamin C ratio was increased compared to predialysis levels and to dialysis with Excebrane. Both of these oxidative stress parameters remained nearly unchanged compared to predialysis values when the Excebrane membrane was used [134]. Reviewing the available literature, the antioxidant characteristics of Excebrane compared to other membranes from regenerated cellulose or polysulfone are controversial. It seems an interesting new approach to offer a speciﬁc and timely protection against oxygen free radicals at their site of generation. However, the efﬁciency of membrane-associated effects must be compared with other treatment forms, such as oral prescription of vitamin E or other antioxidant therapy, for example, vitamin C supplementation. A comparison of vitamin E-coated membranes with regenerated cellulose plus additional vitamin C infusion revealed that both approaches had equal positive effects on thiobarbituric acid reacting substances [135]. It is further argued that vitamin E deﬁciency is rare in uremic patients because the vitamin is not removed during dialysis. Lower vitamin E levels may be induced via vitamin C loss, due to its low molecular weight (MW 176 Da). Vitamin C loss may be considerably high during dialysis and result in less regeneration of vitamin E after dialysis [135]. In 2005, Asahi Medical further introduced a vitamin E-bonded polysulfone membrane into the dialysis market, and thus proved that it is possible to bind biologically active compounds also to synthetic dialysis membranes and sterilize them without loss of performance. 1.6.5.1 Degranulation of Neutrophils During hemodialysis neutrophil degranulation occurs, which is independent of complement activation but inﬂuenced by intracellular calcium and two neutrophil degranulation-inhibiting proteins: angiogenin (MW 14 000) [136] and complement factor D (MW 23 000) [137]. Both inhibitors are found to be up to 15-fold elevated in end-stage renal disease patients protecting against degranulation and consequently

1.6 Dialysis Membranes and Biocompatibility

degranulation products like lactoferrin [137]. Both molecules are removed only in small amounts by highly convective treatments, but some protein adsorbing membranes like PMMA and AN69 adsorb angiogenin [116] and factor D onto their surfaces [105,106]. AN69 reduced plasma angiogenin levels by 66 % and factor D level by 37 % during HD treatment, whereas Fresenius Polysulfone reduced plasma angiogenin concentration by 36 % and had no inﬂuence on factor D [116]. The removal of inhibitors during AN69 dialysis resulted in a more neutrophil degranulation and lactoferrin release than did treatment with polysulfone [116]. PMMA and regenerated cellulose induced more neutrophil degranulation than Hemophan, high-ﬂux Fresenius Polysulfone, and polyamide [113].

1.6.6 Stimulation of Cytokine Generation by Different Types of Dialyzers and Hemoﬁlters

Predialysis intracellular interleukin-1 (IL-1) and TNFa levels are higher in HD patients than in healthy individuals. Furthermore, both cytokines may be transiently generated during dialysis treatment with nonultrapure dialysis ﬂuid [138–140]. The same has been described for the soluble receptors, IL-1RA and TNFsRp55 [138]. Predialysis IL-1b and TNFa in zymosan-stimulated PBMCs were found to be higher in patients treated with regenerated cellulose than in patients treated with PMMA and in healthy volunteers [141]. As complement is involved in the activation process, cytokine mRNA production and complement activation normally correlate [141,142]. However, a second stimulus is necessary for cytokine protein production and release. In hemodialysis, this is mainly provided by microbial contamination products [143] and, if used, by acetate from the dialysis ﬂuid. Therefore, neither higher IL-1 nor TNFa plasma levels could be observed with membranes from regenerated cellulose [144] when pure, acetate-free dialysis ﬂuid was used. In studies using pure dialysis ﬂuid, increased plasma levels of IL-1 were observed to differ only insigniﬁcantly between normal subjects, nondialyzed chronic renal failure patients, and hemodialyzed patients [143]. Therefore, comparison of results from different clinical studies is only meaningful if the purity of the dialysis ﬂuid is similar. In cases of contaminated dialysis ﬂuid, the permeability of a membrane as well as its ability to adsorb microbial products at its outer surface is key factor for cytokine release [145]. The level of circulating cytokines that can be determined with modern methods may also be affected by patient-speciﬁc factors such as residual renal function, different renal and comorbid diseases, drugs, and different cellular productions. Dialyzer-related inﬂuencing factors are permeability/clearance, ultraﬁltration rate, adsorption of cytokines, and stimulation of generation by different materials [140,146]. A comparison of plasma cytokine levels of circulating IL-1b, IL-6, and IL-10 measured using hemodialysis over a period of 4 months with high-ﬂux polyamide and low-ﬂux Hemophan revealed no differences between the membranes [147].

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In contrast, in vitro experiments with isolated monocytes from these patients reveal a higher state of preactivation of these cells with Hemophan. Preactivated monocytes secrete high amounts of proinﬂammatory cytokines when exposed to a second stimulus such as endotoxin [147]. A second study compared gene expression of IL-1b with different membranes using pure dialysis ﬂuid: No differences in plasma levels among dialysis patients using different membranes, healthy subjects, and nondialyzed ESRD patients were found, but gene expression of IL-1b was threefold higher with regenerated cellulose in comparison to PMMA and polysulfone [144]. These data clearly demonstrate that gene expression of cytokines, but not necessarily their release from cells, is triggered with regenerated cellulose membranes. Therefore, plasma levels of circulating cytokines may remain stable during treatment. 1.6.6.1 The Impact of Membrane Types on LPS-Stimulated IL-1b Secretion Intracellular cytokine levels, such as IL-1b, are increased in hemodialysis patients, especially during dialysis with regenerated cellulose [139]. This chronic stimulation obviously leads to a suppression of the response to biological stimulants such as bacterial products. Lipopolysaccaraides (LPS)-induced IL-1b secretion is reduced in patients treated with membranes made of regenerated cellulose compared to that of healthy controls, nondialyzed chronic renal failure patients, and patients dialyzed with polyacrylonitrile membranes [148]. This functional change in PBMC response is speciﬁc for IL-1b, and it does not happen with TNF-a, is hemodialysis membrane is dependent and reversible. As possible mechanism may refer to increased prostaglandine E2 (PGE2) levels that suppress IL-1b secretion due to the following observed effects: when patients dialyzed with regenerated cellulose were switched to PAN membranes, total cell-associated and secreted IL-1b concentrations remained nearly constant, but the secreted amount after LPS stimulation increased. In parallel, PGE2 synthesis decreased with PAN compared to regenerated cellulose. Evidence for the role of PGE2 is further provided by the fact that with the addition of a PGE2 inhibitor, IL-1b secretion after LPS stimulation improved in PBMCs from HD-patients treated with regenerative cellulose.

1.6.7 The Impact of Large-Pore Dialysis Membranes on the Inﬂammatory Response in HD Patients by Cytokine Elimination

Due to the presence of large pores and their ability to adsorb proteins, high-ﬂux dialyzers or hemoﬁlters may eliminate some cytokines. Here some limitations of extracorporeal removal have to be kept in mind: cytokines are molecules of molecular size ranging from 15 000–30 000, which have a short half-life (in the range of minutes) and may be bound to carrier proteins such as a2-macroglobulin [149]. Therefore, a detectable amount may only be efﬁciently removed when convective treatments like hemoﬁltration with high ultraﬁltration volumes and highly permeable membranes are performed. The adsorption capacity of the membrane for

1.6 Dialysis Membranes and Biocompatibility

cytokines is probably of more importance. In contrast to membranes from regenerated cellulose, membranes made from PAN (AN69) are able to adsorb considerable amounts of IL-1 in vitro [146].

1.6.8 The Effect of Different Dialyzers on the Acute Phase Reaction

IL-6 is the major regulator of the hepatic acute phase response in inﬂammation: it stimulates hepatic synthesis of C-reactive protein (CRP) and serum amyloid A (SAA) up to several 100-fold [150]. Cultured PBMCs from patients dialyzed with Cuprophan spontaneously released more IL-6 than PBMCs from healthy individuals or from patients dialyzed with PMMA membranes [151], SMC, or cellulose diacetate [152]. The same effect was observed with the soluble IL-6 receptor, which probably reﬂects more the biological activity of IL-6 [153]. IL-6 release correlates positively with levels of circulating CRP [150], and hemodialysis patients exhibit elevated levels of IL-6, C reactive protein and serum amyloid A [150]. The reason for this is yet not clear, but the role of dialysis membranes and contaminated dialysis ﬂuid is implicated: IL-6 levels were found to be elevated after the third hour of treatment with cuprammonium rayon membranes (Asahi AM-UP-75) and correlated with the release of acute phase proteins, albeit with different time schedules [150]. Increased concentrations of CRP and secretory phospholipase A2 (sPLA2) were found 24 hafter the start of hemodialysis with this membrane. In contrast, dialysis with polysulfone (F60S, FMC) showed no marked variation at these time points neither for IL-6, CRP nor sPLA2 levels [150].

1.6.9 Activation of the Kinin System by Different Types of Dialyzers and Hemoﬁlters

The main factor determining activation of the kinin system with subsequent generation of bradykinin is the electronegativity of the dialysis membrane [154]. Table 1.4 provides an overview of the zeta potentials –as parameter for membrane electronegativity – of seven frequently used membranes. As is obvious from these in vitro experiments, AN69 and PAN DX membranes are the most electronegative membranes and generate the highest amounts of bradykinin, a mediator for anaphylactoid reactions [154]. When the electronegativity of the polyacrylonitrile polymer is reduced to near zero, as found for the PAN modiﬁcation AN 69ST, in vitro bradykinin generation is reduced 200-fold. This was achieved by coating of the membrane with the polycationic polymer, polyethyleneimine [154]. Bradykinin generation with PAN membranes could be detected in vivo, especially in patients under ACE inhibitor therapy. Bradykinin is normally rapidly degraded in the body by the serine protease kininase II and therefore most patients are symptomfree under AN69 and PAN DX dialyzes. However, if this enzyme is blocked by ACE,

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1 Membranes in Hemodialysis Zeta potential (surface electronegativity) plasma kallikrein activity and bradykinin generation with different membranes in vitro.

Tab. 1.4

Membrane Polyacrylonitrile AN69 Polyacrylonitrile PAN DX Polymethylmethacrylate Cellulose triacetate Cuprophan Fresenius Polysulfone Polyacrylonitrile AN69ST

Zeta potential (mV)

Plasma kallikrein (U/mL)

Bradykinin generation (fmol/mL)

70 5

60 15

32 100 (26 500–41 200)

60 4

80 20

28 983 (22 600–36 150)

25 2 20 2 10 1 5 1 3 1

10 5 <5 <5 <5 <5

130 65 78 62 150

(50–130) (25–100) (25–150) (25–120) (30–450)

Source: adapted from Ref. [154]

inhibitors plasma levels may increase by more than 100-fold and patients will suffer from severe anaphylactoid reactions [50–55,155]. The highest bradykinin generation reported was measured in vivo in sheep under ACE inhibitor therapy and treated with poylacrylonitrile membranes (AN69). Lower

Fig. 1.18 What is the possible future strategy for the treatment of kidney patients by means of extracorporeal blood circuits? Recent investigations have shown that a variety of uremic toxins is found in the larger molecular weight range. Conventional dialysis membranes such as the F60S PSu-membrane are determined by their molecular weight cutoff, that is, the sieving coefﬁcient distribution.

Based on the experience with liver failure therapy, recent observations have shown that adsorber cartridges placed in a secondary circuit may support the removal of largemolecular-weight toxins, given that a ﬁlter with a higher cut-off (e.g., an Albuﬂow ﬁlter with a highly permeable polysulfone membrane) allows for the separation of plasma in the primary circuit.

References

generation was observed when a PAN membrane with lower surface electronegativity (SPAN-Membrana, not available any more on the market) was used [156]. In another study, even symptom-free patients without ACE medication exhibited an approximately 10-fold higher bradykinin generation with PAN (AN69) plate dialyzers than with Cuprophan, Hemophan, and high-ﬂux Fresenius Polysulfone [157]. A very low bradykinin generation by high-ﬂux Fresenius Polysulfone and low-ﬂux Hemophan membranes was conﬁrmed also for patients under ACE inhibitor therapy [158].

1.7 Conclusion

During the last 20 years of hemodialysis therapy, an intense research on membranerelated factors and properties has been performed. Understanding of membrane properties and performance in the clinical setting has considerably improved such that adverse events related to the membrane material have become rare. A general trend can be seen in that cellulosic membranes have lost their dominating role and synthetic membranes have taken the lead. Furthermore, membranes with high-ﬂux properties gain increasing importance due to new concepts of therapy, such as online hemodiaﬁltration. Dialysis membranes of today have become bulk products and are available without any restriction. Future therapies will possibly involve highly permeable membranes combined in series with adsorber cartridges in order to cope with newly identiﬁed uremic toxins of large molecular weights and under hemodiaﬁltration treatments (Figure 1.18).

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of rinsing technique, dialyser storage time, and potting compound. Artiﬁcial Organs, 11, 118–122. Vienken, J. and Ronco, C. (2001) New developments in hemodialysers. Contributions to Nephrology, 133, 105–118. Ronco, C., Brendolan, A., Crepaldi, C., Rodighiero, M., Scabardi, M. (2002) Blood and dialysate ﬂow distributions in hollow-ﬁber hemodialyzers analyzed by computerized helical scanning technique. Journal of the American Society of Nephrology, 13 (Suppl. 1), S53–S61. Ronco, C., Brendolan, A., Crepaldi, C., Rodighiero, M., Everard, P., Ballestri, M., Capelli, G., Spittle, M., La Greca, G. (2000) Dialysate ﬂow distribution in hollow ﬁber hemodialysers with different dialysate pathway conﬁgurations. International Journal of Artiﬁcial Organs, 23, 601– 609. Ronco, C., Scabardi, M., Goldoni, M., Brendolan, A., Crepaldi, C., La Greca, G. (1997) Impact of spacing ﬁlaments external to hollow ﬁbers on dialysate ﬂow distribution and dialyser performance. International Journal of Artiﬁcial Organs, 20, 261– 266. ¨ ber Diffusion. Annals Fick, A. (1855) U of Physics and Chemistry, 94, 59–86. Ward, R. and Schmidt, B. (1989) The impact of erythropoietin on hemodialyzer design and performance. Artiﬁcial Organs, 13, 35–42. Basmadijan, D., Sefton, M., Baldwin, S. (1997) Coagulation on biomaterials in ﬂowing blood: some theoretical considerations. Biomaterials, 18, 1511–1522. Storck, M., Hartl, W. H., Zimmerer, E., Inthorn, D. (1991) Comparison of pump-driven and spontaneous continuous hemoﬁltration in postoperative acute renal failure. The Lancet, 337, 452–455. Delanna, F., Wu¨pper, A., Baldamus, C. A. (1996) Internal ﬁltration – advantage in hemodialysis?Nephrology

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Dialysis Transplantation, 11 (Suppl. 2), 83–86. Gurland, H., Davison, A., Bonomini, V., Falkenhagen, D., Hansen, S., Kishimoto, T., Lysaght, M., Moran, J., Valek, A. (1994) Deﬁnitions and terminology in biocompatibility. Nephrology Dialysis Transplantation, 9 (Suppl. 2), 4–10. Salzmann, E. (1971) Thrombosis in artiﬁcial organs. Transplantation Proceedings, 3, 1491–1496. Dolovich, J. and Bell, B. (1978) Allergy to a product of ethylene oxide gas: demonstration of IgE and IgG antibodies. Journal of Allergy and Clinical Immunology, 62, 30–32. Craddock, P., Fehr, J., Brigham, K., Kronenberg, R., Jacob, H. (1977) Complement and leukocyte-mediated pulmonary dysfunction in hemodialysis. New England Journal of Medicine, 296, 769–774. Craddock, P., Fehr, J., Dalmasso, A., Brigham, K., Jacob, H. (1977) Hemodialysis leukopenia: pulmonary vascular leukostasis resulting from complement activation by dialyser cellophane membranes. Journal of Clinical Investigation, 59, 879–888. Henderson, L., Koch, K., Dinarello, C., Shaldon, S. (1983) Hemodialysis hypotension: the interleukin hypothesis. Blood Puriﬁcation, 1, 3–8. Tielemanns, C., Madhoun, P., Lenaers, M., Schandene, L., Goldman, M., Vanherweghem, J. (1990) Anaphylactoid reactions during hemodialysis with AN69 membranes in patients receiving ACE-inhibitors. Kidney International, 38, 982–984. Tsunoda, N., Kobuko, K., Sakai, K., Fukuda, M., Miyazaki, M., Hiroshi, T. (1999) Surface roughness of cellulose hollow ﬁber dialysis membranes and platelet adhesion. ASAIO Journal, 45, 418–423. Matata, B., Courtney, J., Sundaram, S., Wark, S., Bowry, S., Vienken, J., Lowe, G. (1996) Determination of contact phase activation by the measurement of the activity of

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(1993) Prevention of blood loss in dialysers with DEAE cellulose membranes does not require increased doses of heparin. Nephrology Dialysis Transplantation, 8, 1140–1145. Debrand-Passard, A., Lajous-Petter, A., Schmidt, R., Herbst, R., vonBaeyer, H., Krause, A., Schifﬂ, H. (1989) Thrombogenicity of dialyser membranes as assessed by residual blood volume and surface morphology at different heparin dosages. Contributions to Nephrology, 74, 2–9. Verbeelen, D., Jochmanns, K., Herman, A., Vander Niepen, P., Sennesael, J., DeWaele, M. (1991) Evaluation of platelets and hemostasis during hemodialysis with six different membranes. Nephron, 59, 567–572. Cases, A., Reverter, J., Escolar, G., Sanz, C., Lopez-Pedret, J., Revert, L., Ordinas, A. (1993) Platelet activation on hemodialysis: inﬂuence of dialysis membranes. Kidney International, 43 (Suppl. 41), S217–S220. Cases, A., Reverter, J., Escolar, G., Sanz, C., Sorribes, J., Ordinas, A. (1997) In vivo evaluation of platelet activation by different cellulosic membranes. Artiﬁcial Organs, 21, 330–334. Rauterberg, E., Ritz, E., Schulze, H., Rother, K. (1987) Limited derivatisation of Cuprophan increases factor H binding and diminishes complement activation (Abstract). Kidney International, 31, 243. Pascual, M. and Schifferli, J. (1993) Adsorption of complement factor D by polyacrylonitrile membranes. Kidney International, 43, 903–911. Pascual, M., Schifferli, J., Pannatier, J., Wauters, J. (1993) Removal of complement factor D by adsorption of polymethacrylate dialysis membranes. Nephrology Dialysis Transplantation, 8, 1305–1306. Cheung, A., Parker, C., Wilcox, L., Janatova, J. (1989) Activation of the

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alternative pathway of complement by cellulosic hemodialysis membranes. Kidney International, 36, 257–265. Deppisch, R., Goehl, H., Smeby, L. (1998) Microdomain structure of polymeric surfaces – potential for improving blood treatment procedures. Nephrology Dialysis Transplantation, 13, 1354–1359. Cheung, A., Parker, C., Wilcox, L., Janatova, J. (1990) Activation of complement by hemodialysis membranes: polyacrylonitrile binds more C3a than Cuprophan. Kidney International, 37, 1055–1059. Jorstadt, S., Smeby, L., Balstad, T., Wideroe, T. (1988) Generation and removal of anaphylatoxins during hemoﬁltration with ﬁve different membranes. Blood Puriﬁcation, 6, 325–335. Falkenhagen, D., Brown, G., Boetcher, M., Falkenhagen, U., Schmidt, B., Gurland, H., Klinkmann, H. (1987) Permeation of complement factors through highﬂux dialysers and plasma separation membranes. In Bambauer, R. Malchesky, P. Falkenhagen, D. (eds) Therapeutic plasma exchange and selective plasma separation, Schattauer Verlag, Stuttgart, 215–222. Kaiser, J., Oppermann, M., Gotze, O., Deppisch, R., Goehl, H., Asmus, G., Rohricht, B., vonHerrath, D., Schaefer, K. (1995) Signiﬁcant reduction of factor D and immunosuppressive complement fragment Ba by haemoﬁltration. Blood Puriﬁcation, 13, 314–321. Deppisch, R., Ritz, E., Ha¨nsch, G., Scho¨ls, M., Rauterberg, E. (1994) Biocompatibility perspectives. Kidney International, 45 (Suppl. 44), S77– S84. Meier, P., von Fliedner, V., Markert, M., VanMelle, G., Deppisch, R., Wauters, J. (2000) One year immunological evaluation of chronic hemodialysis in end stage renal

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disease patients. Blood Puriﬁcation, 18, 128–137. Ohi, H., Tamano, M., Sudo, S. (2001) Cellulose membranes suppress complement activation in patients after hemodialysis. American Journal of Kidney Diseases, 38, 384–389. Hoerl, W. (1998) Hemodialysis membranes: interleukins, biocompatibility and middle molecules. Journal of the American Society of Nephrology, 22, 585–590. Hoenich, N. and Stamp, S. (2000) Clinical investigation of the role of membrane structure on blood contact and solute transport characteristics of a cellulose membrane. Biomaterials, 21, 317–324. Nahar, N., Shah, H., Sui, J., Colvin, R., Baskaran, M., Ranjan, R., Wagner, J., Singhal, P. (2001) Dialysis membrane induced neutrophil apoptosis is mediated through free radicals. Clinical Nephrology, 56, 52–59. Martı´n-Malo, A., Carracedo, J., Ramirez, R., Rodriguez-Benot, A., Soriano, S., Rodriguez, M., Aljama, P. (2000) Effect of uremia and dialysis modality on mononuclear cell apoptosis. Journal of the American Society of Nephrology, 11, 936–942. Nguyen-Khoa, T., Massy, Z., DeBandt, J., Kebede, M., Salama, L., Lambrey, G., Witko-sarsat, V., Drueke, T., Lacour, B., The´venin, M. (2001) Oxidative stress and hemodialysis: role of inﬂammation and duration of dialysis treatment. Nephrology Dialysis Transplantation, 16, 335–340. Weinstein, T., Chagnac, A., Korzets, A., Boaz, M., Ori, Y., Herman, M., Malachi, T., Gafter, U. (2000) Hemolysis in hemodialysis patients: evidence for impaired defense mechanisms against oxidative stress. Nephrology Dialysis Transplantation, 15, 883–887. Himmelfarb, J. and McMonagle, E. (2001) Manifestations of oxidative stress in uremia. Blood Puriﬁcation, 19, 200–205.

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dialysis on oxidative stress in uremia. Redox Report, 4, 17–22. Morena, M., Cristol, J., Canaud, B. (2000) Why hemodialysis patients are in a pro-oxidant state? What could be done to correct the pro/antioxidant imbalance?. Blood Puriﬁcation, 18, 191–199. Canaud, B., Cristol, J., Morena, M., Leray-Moragues, H., Bosc, J., Vaussenat, F. (1999) Imbalance of oxidants and antioxidants in hemodialysis patients. Blood Puriﬁcation, 17, 99–106. Himmelfarb, J., Ault, K., Holbrook, D., Leeber, D., Hakim, R. (1993) Intradialytic granulocyte reactive oxygen species production: a prospective cross-over trial. Journal of the American Society of Nephrology, 4, 178–186. Bonomini, M., Sirolli, V., Settefrati, N., Stuard, S., Tropea, F., DiLiberato, L., Tetta, C., Albertazzi, A. (1999) Surface antigen expression and platelet neutrophil interactions in hemodialysis. Blood Puriﬁcation, 17, 107–117. Sirolli, V., Ballone, E., Amoroso, L., DiLiberato, L., Di Mascio, R., Capelli, P., Albertazzi, A., Bonomini, M. (1999) Leukocyte adhesion molecules and leukocyteplatelet interactions during hemodialysis: effects of different synthetic membranes. International Journal of Artiﬁcial Organs, 22, 536–542. Vanholder, R., Ringoir, S., Dhondt, A., Hakim, R., Waterloos, M., VanLantschoot, N., Gung, A. (1991) Phagocytosis in uremic and hemodialysis patients: a prospective and cross sectional study. Kidney International, 39, 320–327. Galli, F., Canestrari, F., Buoncristiani, U. (1999) Biological effects of oxidant stress in hemodialysis: the possible role of vitamin E. Blood Puriﬁcation, 17, 79–94. Bonnefont-Rousselot, D., Lehmann, E., Jaudon, M., Delattre, J., Perrone,

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B., Reschke, J. (2000) Blood oxidative stress and lipoprotein oxidizability in hemodialysis patients: effects of the use of a vitamin E coated dialysis membrane. Nephrology Dialysis Transplantation, 15, 2020–2028. Galli, F., Rovidati, S., Chiarantini, L., Campus, G., Canestrari, F., Buoncristiani, U. (1998) Bioreactivity and biocompatibility of a vitamin Emodiﬁed multilayer hemodialysis ﬁlter. Kidney International, 54, 580–589. Satoh, M., Yamasaki, Y., Nagake, Y., Kasahara, J., Hashimoto, M., Nakanishi, N., Makino, H. (2001) Oxidative stress is reduced by the long-term use of vitamin E-coated dialysis ﬁlters. Kidney International, 59, 1943–1950. Clemont, G., Lecour, S., Cabanne, J., Motte, G., Guilland, J., Hevet, D., Rochette, L. (2001) Vitamin E-coated dialyser reduces oxidative stress in hemodialysis patients. Free Radical Biology and Medicine, 31, 233–241. Eiselt, J., Racek, J., Treﬁl, L., Opatrny, K. (2001) Effects of vitamin Emodiﬁed dialysis membrane and vitamin C infusion on oxidative stress in hemodialysis patients. Artiﬁcial Organs, 26, 430–436. Fett, J., Strydom, D., Lobb, R., Alderman, E., Bethune, J., Riordan, J., Vallee, B. (1985) Isolation and characterisation of angiogenin, an angiogenin protein from human carcinoma cells. Biochemistry, 24, 5480–5486. Schmaldienst, S. and Hoerl, W. (2000) Degranulation of polymorphonuclear leukocytes by dialysis membranes. Nephrology Dialysis Transplantation, 15, 1909– 1910. Pereira, B., Shapiro, L., King, A., Falagas, M., Strom, J., Dinarello, C. (1994) Plasma levels of Il-1b, TNF and their speciﬁc inhibitors in undialysed chronic renal failure. CAPD and hemodialysis patients. Kidney International, 45, 890–896.

139 Haeffner-Cavaillon, N., Cavaillon, J.,

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Ciancioni, C., Bacle, F., Delons, S., Kazatchkine, M. (1989) In vivo induction of interleukin-1 during hemodialysis. Kidney International, 35, 1212–1218. Bingel, M., Lonnemann, G., Koch, K., Dinarello, C., Shaldon, S. (1988) Plasma interleukin-1 activity during hemodialysis: the inﬂuence of hemodialysis membranes. Nephron, 50, 273–276. Schindler, R., Gelfand, J., Dinarello, C. (1990) Recombinant C5a stimulates transcription rather translation of Il-1 and TNF: translational signal provided by LPS or IL-1 itself. Blood, 76, 1631–1638. Varela, M., Kimmel, P., Philips, T., Mishkin, G., Lew, S., Bosch, J. (2001) Biocompatibility of hemodialysis membranes: interrelations between plasma complement and cytokine levels. Blood Puriﬁcation, 19, 370–379. Okusawa, S., Dinarello, C., Yancey, K., Endres, S., Lawley, T., Frank, M., Burke, J., Gelfand, J. (1987) C5a induction of human interleukin 1: synergistic effect with endotoxin or interferon. Journal of Immunology, 139, 2635–2640. Qian, J., Yu, Z., Dai, H., Zhang, Q., Chen, S. (1995) Inﬂuence of hemodialysis membranes on gene expression and plasma levels of interleukin-1. Artiﬁcial Organs, 19, 842–846. Lonnemann, G., Behme, T., Lenzner, B., Floege, J., Schulze, M., Colton, C., Koch, K., Shaldon, S. (1992) Permeability of dialyser membranes to TNF-inducing substances derived from bacteria. Kidney International, 42, 61–68. Lonnemann, G., Koch, K., Shaldon, S., Dinarello, C. (1988) Studies on the abilities of dialysis membranes to induce, bind, and clear human interleukin-1. Journal of Laboratory and Clinical Medicine, 112, 76–86. Girndt, M., Heisel, O., Ko¨hler, H. (1999) Inﬂuence of dialysis with polyamide vs.haemophan dialysers on monkines and complement

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activation during a 4-month longterm study. Nephrology Dialysis Transplantation, 14, 676–682. Lonnemann, G., Barndt, I., Kaever, V., Haubitz, M., Schindler, R., Shaldon, S., Koch, K. (1995) Impaired endotoxin-induced interleukin-1 secretion, not total production of mononuclear cells from ESRD patients. Kidney International, 47, 1158–1167. Schindler, R., Senf, R., Frei, U. (2002) Inﬂuencing the inﬂammatory response of hemodialysis patients by cytokine elimination using large-pore membranes. Nephrology Dialysis Transplantation, 17, 17–19. Schouten, W., Grootemann, M., VanHoute, A., Schoorl, M., VanLimbeck, J., Nube, M. (2000) Effects of dialyser and dialysate on the acute phase reaction in clinical bicarbonate dialysis. Nephrology Dialysis Transplantation, 15, 379–384. Memoli, B., Libetta, C., Rampino, T., DalCanton, A., Conte, G., Scala, O., Ruocco, M., Andreucci, M. (1992) Hemodialysis related induction of interleukin-6 production by peripheral blood mononuclear cells. Kidney International, 42, 320–326. Memoli, B., Minutulo, R., Bisesti, V., Postiglione, L., Conti, A., Marzano, L., Capuano, A., Andreucci, M., Balletta, M., Guida, B., Tetta, C. (2002) Changes of serum albumin and C-reactive protein are related to changes of interleukin-6 release by peripheral blood mononuclear cells in hemodialysis patients treated with different membranes. American Journal of Kidney Diseases, 39, 266–273.

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Cianciaruzo, B., Bisesti, V., Cimmaruta, C., Marzano, L., Minutulo, R., Cuomo, V., Guida, B., Andreucci, M., Rossi, G. (2000) Role of different dialysis membranes in the release of interleukin-6 soluble receptor in uremic patients. Kidney International, 58, 417–424. Renaux, J., Thomas, M., Crost, T., Loughraieb, N., Vantard, G. (1999) Activation of the kallikrein-kinin system in hemodialysis: role of membrane electronegativity, blood dilution and pH. Kidney International, 55, 1097–1103. Schaefer, R., Schaefer, L., Hoerl, W. (1994) Anaphylactoid reactions during hemodialysis. Clinical Nephrology, 42 (Suppl. 1), S44–S47. Krieter, D., Grude, M., Lemke, H., Fink, E., Bo¨nner, G., Scho¨lkens, B., Schulz, E., Mu¨ller, G. (1998) Anaphylactoid reactions during hemodialysis in sheep are ACEinhibitor dose-dependent and mediated by bradykinin. Kidney International, 53, 1026–1035. Verresen, L., Fink, E., Lemke, H., Vanrenterghem, Y. (1994) Bradykinin is a mediator of anaphylactoid reactions during hemodialysis with AN69 membranes. Kidney International, 45, 1497–1503. Mannstadt, M., Touam, M., Fink, E., Urena, P., Hruby, M., Zingraff, J., Uhlenbusch-Ko¨rwer, I., Grassmann, A., Lemke, H., Drueke, T. (1995) No generation of bradykinin with a new polyacrylonitrile membrane [SPAN] in hemodialysis patients treated with ACE-inhibitors. Nephrology Dialysis Transplantation, 10, 1696–1700.

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2 Membranes for Artiﬁcial Lungs Frank Wiese 2.1 Introduction

Membrane oxygenators are used inside heart–lung machines during open-heart surgery or for lung support systems. In the natural lung, oxygen and carbon dioxide are exchanged between the blood and the body environment. The oxygen is needed in the body tissues to sustain life. The carbon dioxide produced during the metabolism has to be removed. Worldwide, oxygenators are used in surgical interventions during which the heart– lung function is interrupted, for example, for cardiovascular bypass, heart valve replacement, or when repairing congestive heart failure [30–32]. Furthermore, extracorporeal membrane oxygenation (ECMO) is used in several conﬁgurations for long-term lung support, for example, in cases of multiorgan failure (MOF) and in different pulmonary diseases [7,33–35]. Implantable membrane devices for longterm lung support are under development [5]. Special newly developed membranes were required for these long-term applications. In general, at the very core of these devices are membranes that separate the patient’s blood from gases and that allow oxygen to enter and carbon dioxide to leave the blood. In 2005, around 1.2 million open-heart operations were carried out, corresponding to a consumption of 3 million km of oxygenation membranes. Membranes for gas exchange are also an important component in membrane bioreactors for biotechnology applications and in hybrid bioartiﬁcial organs (see Chapters 5 and 8).

2.2 History of Blood Oxygenation

Some milestones in the history of the development of blood membrane oxygenators [6,30,37] are listed below: 1812: Le Gallois’ suggestion for the ﬁrst extracorporeal circulation

Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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1876: an isolated kidney was perfused with oxygenated blood for the ﬁrst time 1918: discovery of heparin 1944: Kolff had the idea of constructing an oxygenator (beginning of membrane oxygenation) 1955: development of the ﬁrst membrane oxygenator by Kolff and Balzer 1956: ﬁrst clinical application 1959: ﬁrst heart surgery using a membrane oxygenator 1969: membrane oxygenators became commercially available.

The start of the historical development of extracorporeal circulation is strongly related to Ce´sar Julian Jean Le Gallois (1770–1814). He suggested that a part of the body might be preserved by a mechanical heart replacement and some kind of external perfusion devices. The important monograph ‘‘Experiences sur le principe de la vie’’ was published in 1812 [1] and 1 year later in the United States [2] (see Figure 2.1). Alexander Schmidt for the ﬁrst time perfused an isolated dog kidney with oxygenated blood in 1876 [3]. Many different types of disks, ﬁlms, and bubble oxygenators were developed for extracorporeal circulation experiments. All these early oxygenators exposed blood directly to gas mixtures or oxygen to provide oxygenation and remove CO2. The developments, started in the 1930s, resulted in disposable single-use bubble oxygenators and opened the area of cardiac surgery in the early 1950s. John Gibbon was a pioneer in this development [4,8,30]. In 1955, the DeWall–Lillehei bubble oxygenator was a major breakthrough [36,23]. The simplicity of the basic concept was linked to simple construction and disposability. The principle is: Venous blood is pumped into an oxygenation chamber where oxygen bubbles are dispersed. The gas exchange occurs on the bubble surface. The mixture of blood and gas that emerges from the bubble chamber has to be defoamed in a defoaming compartment. The main disadvantage of bubble oxygenators is that they are basically not biocompatible. The blood–gas interface represents a nonphysiological state resulting, for example, in hemolysis, coagulopathy, and so on. Nevertheless, bubble oxygenators were used until 10 years ago in around 2–3 % of open heart surgery worldwide. 2.2.1 Membrane Oxygenators

In 1944, Kolff and Berk found that venous blood was oxygenated while ﬂowing through a cellophane dialyzer and being in contact with oxygen containing dialyzate. Oxygenation was carried out across the artiﬁcial membrane [24]. This discovery stimulated the development of the use of gas permeable membranes in order to separate the blood phase from the gas phase in an artiﬁcial lung. The ﬁrst membrane oxygenator was reported by Kolff in 1955 [45,46] (Figure 2.2). At this time the available membrane materials were relatively impermeable to the respiratory gases.

2.2 History of Blood Oxygenation

Fig. 2.1 Front page of paper of Le Gallois, describing the basic idea of extracorporeal circulation and oxygenation of blood for the ﬁrst time (1812).

The ﬁrst membrane oxygenator built and used clinically was reported by Clowes and Hopkins in 1956 [8]. It was constructed of ethylcellulose multilayer ﬂatsheet membranes and had a membrane area of 25 m2. These factors limited the clinical applications until the 1970s and early 1980s, although the very ﬁrst thin sheet membranes of polyethylene and polytetraﬂourethylene were available. The development of polydimethylsiloxane with high permeability for oxygen and carbon

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Fig. 2.2 Professor Kolf in 2004 – ‘‘The father of artiﬁcial organs,’’ also invented artiﬁcial coil kidney and the coil oxygenator 50 years ago [47].

dioxide brought a major advance in establishing the technical feasibility of membrane oxygenators in the 1960s and 1970s. The required membrane area could be optimized at less than 6 m2. The real breakthrough came with the development of hydrophobic microporous ﬂat sheet membranes with pore sizes <0.1 mm in the 1980s [26]. Later in the 1980s microporous hollow ﬁber membranes [27] became commercially available. With the introduction of extraluminal ﬂow oxygenators [28,29] membranes, areas of less than 2 m2 were required to treat an adult patient.

2.3 Principle of Gas Transfer

2.3 Principle of Gas Transfer

The transport of gas into the blood is a process of convection and diffusion via the membrane and the blood combined with O2 and CO2 reaction with the erythrocyte hemoglobin. The hemoglobin in the erythrocytes has a speciﬁc chemical gas-binding capacity for O2 and CO2 and assures the adequate transport of both these gases inside the body. The driving force for the gas exchange is the partial pressure difference of these gases between the blood side and the gas side (O2) of the membrane. According to the pressure differences, gas diffuses through the membrane wall from the side of high partial pressure to the side of low partial pressure (Figure 2.3). The maximum concentration of chemically bound O2, called the oxygen capacity, is calculated from the hemoglobin content of the blood and the binding capacity of hemoglobin for oxygen. For a mean hemoglobin content of Hb ¼ 0.12 kg/L and a binding capacity of Ha ¼ 1.34 L/kg, the oxygen capacity will be 0.1608 L O2/L blood. This value will be reached at an O2 partial pressure of approximately 20 kPa ¼ 150 mmHg, which is the ambient partial pressure of oxygen. An additional increase in the O2 content is possible but nonphysiological. The oxyhemoglobin dissociation curve is usually calculated according to the mathematical model of Hill [9] (Figure 2.4). Nowadays the active area of membrane oxygenators is between 0.5 and 2.5 m2. This is around 10 % of that of the natural lung of young children or adults. In order to reach a sufﬁcient gas exchange, the contact time has to be increased and/or the partial pressure differences have to be increased or other suitable conditions have to be adjusted. According to Fick’s law (1) the gas exchange efﬁciency can be increased [30,37]. VO 2 ¼

P1 P2 KF: L

ð1Þ

In Equation (1), VO2 is the amount of oxygen exchanged in a certain time, P1 P2 is the partial pressure difference, K is the diffusion factor (adsorption coefﬁcient, turbulence), F is the surface, and L is the diffusion layer thickness.

Fig. 2.3 Partial pressure difference of oxygen and carbon dioxide across an oxygenation membrane.

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Fig. 2.4 Blood saturation according to Hill.

The transfer of oxygen into the blood is hindered by three main resistances: Rtotal ¼ RB þ RM þ RG :

ð2Þ

In Equation (2), the total resistance to gas transport Rtotal is the sum of the resistance on the blood side RB, the resistance across the membrane RM, and the resistance on the gas side RG (Figure 2.5). For a usual oxygenator type, containing microporous capillary membranes, as described in Section 2.4, the blood boundary layer lies outside the capillaries. The diffusion resistances that opposes the gas exchange consists of the resistance of the gas boundary layer, the membrane, and the blood boundary layer. Here the resistance of the gas boundary layer is inﬁnitely small and approaches zero, and accordingly, can be neglected. The resistance of the membrane is about 1.25 104 cm2 s cmHg/cm3. The resistance in the blood boundary layer is approximately 100 times larger and should be optimized in such a way that this layer becomes very small (see Section 2.6). The mass transfer in relation to the geometry ﬂow pattern and the hydrodynamics of an artiﬁcial lung has been investigated in hundreds of papers [6,10–13].

Fig. 2.5 Mass transfer model for an extraluminal ﬂow oxygenator.

2.4 Membranes and Membrane Properties

2.4 Membranes and Membrane Properties

Nowadays membranes for oxygenators used in open-heart surgery are mostly microporous hydrophobic capillary membranes with pore sizes <0.1 mm and outer diameters between 300 and 500 mm. For special applications, for example, in lung support systems, dense skinned membranes are used because of better long-term blood plasma resistance. The main membrane polymers used are polyoleﬁns such as polypropylene (PP), polyethylene (PE), and poly-4-methylpentene (PMP). Further, the polymers used on a small scale are polyvinylidenﬂuoride (PVDF) and silicon rubber. In ﬂat sheet membrane oxygenators, the blood is separated from the gas by spiral wound membranes or stapled membranes. Bafﬂes placed in the area between the membranes, in order to reduce the boundary layer and increase efﬁciency, require a lot of effort. Therefore, only one type of microporous ﬂat sheet membrane oxygenator is available but without any market relevance. Around 80 % of oxygenation membranes used worldwide are produced by Membrana GmbH in Wuppertal. Other producers of oxygenation membranes are Dainippon Inc. and Terumo. 2.4.1 Microporous Membranes

Gas exchange between the blood side and the gas side in microporous membranes is performed by diffusion and convection of the respiratory gases via the open pores. Thus, the resistance of the membrane is very low for these gas exchanges, as explained in Section 2.3. Most microporous oxygenator membranes are made from polypropylene by temperature-induced phase separation (TIPS) process (see Section 2.5). The sponge-like structure of this membrane is homogeneous and isotropic with open surfaces on both wall sides (Figure 2.6). These membranes achieve reliable, adequate performance (Figure 2.10), and the mechanical stability in terms of tensile strength and elongation at break is combined with high reliability and handling safety during the manufacturing process (see Section 2.6). Today, it is the most widely used oxygenation membrane. The second type of microporous membranes is produced by the ‘‘melt spin stretch process’’ (see Section 2.5), resulting in ﬁbrillar polymer structures (see Figure 2.7). Polyoleﬁns are again suitable for these processes because only polymers with a certain crystallinity are suitable for creating such membrane structures. Performance wise the membranes can reach a comparable level to that of TIPS membranes. Long-term plasma breakthrough characteristics are a little more critical because of the structure. Due to the hydrophobicity of the membrane and the pore size (<0.1 mm), gas molecules can pass through the membrane wall, whereas liquids and corpuscular components are retained. The polar part of the surface energy [15] has to be very small so that no wetting of the pores with aqueous solutions, for example, blood, is possible,

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Fig. 2.6 REM photograph of a microporous PP membrane (OXYPHAN, Membrana, Wuppertal).

and a water droplet on this surface has to create a high-contact angle 908 (see Figure 2.8). For the duration of an open-heart operation and for several days, these open membranes work very reliably and are safe. However, during the treatment, plasma proteins adsorb ﬁrst at the membrane surface and then this protein layer also grows

Fig. 2.7 REM photograph of a microporous PP membrane produced by a melt spin stretch process.

2.4 Membranes and Membrane Properties

Fig. 2.8 Contact angle of hydrophilic and hydrophobic surfaces.

through the membrane pores. This results in an increase in the surface energy of the pore ‘‘skin.’’ Moreover, due to this surfactant impact the pores become wet and plasma breakthrough from the blood side to the gas side occurs [41,42]. 2.4.2 Dense Membranes/‘‘Diffusion Membranes’’

Dense membranes, also called ‘‘diffusion’’ membranes, have no open pores in the membrane wall or the outer skin of the wall [38,44,53]. In long-term ‘‘extracorporeal membrane oxygenation’’ applications such membranes are used to prevent plasma breakthrough. The respiratory gases have to have a good solubility in the membrane polymer. Driven by the concentration difference of the gases on both sides of the membrane, the gases are transported by diffusion through the polymer matrix to the lower concentration. In the past, silicon rubber or silicon-skinned microporous membranes were used because of the high speciﬁc permeability of this polymer. The disadvantage of these membranes was their relatively high wall thicknesses due to mechanical requirements. According to Fick’s law (Section 2.3), the diffusive gas transport is reciprocally proportional to the thickness of the polymer layer. Thus, high membrane areas were required to achieve sufﬁcient gas exchange. To overcome this problem, membranes made of poly-4-methylpentene were developed. The membrane with the best performance is produced by a thermally induced phase separation process (ACCUREL) (Figure 2.9). The membrane is integrally asymmetric, as it has a sponge-like microporous wall with a thin dense outer skin (Figure 2.9). Due to the high speciﬁc permeabilities of PMP to oxygen and carbon dioxide (Table 2.1) and the dense outer skin, which is kept very thin (0.1 mm), OXYPLUS has gas transfer capabilities comparable to common microporous membranes. Figure 2.10 shows a comparison of the gas transfer of the same oxygenator types equipped with either PMP or PP membranes. The plasma breakthrough times for OXYPLUS are very much longer than those of common microporous membranes: in the clinical experience of different working groups the membrane has worked properly in ECMO for up to 42 days without exchanging the device [18,19]. Additionally, the dense outer skin largely prevents entry of gas into the blood and avoids direct blood–gas contact. As a trade off, the transfer of volatile anesthetics is impaired [20] and the anesthetic protocol may have to be amended. The permeability of PMP is higher for oxygen than for nitrogen. This membrane is thus also suitable for gas separation or enrichment. Table 2.2 shows the ﬂuxes of

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Fig. 2.9 Gas transfer rates of two commercially available oxygenators, equipped either with a microporous PP membrane or with a PMP membrane with dense outer skin [16,17].

Gas permeability coefficients of polypropylene and poly-4-methylpentene for oxygen and carbon dioxide [14].

Tab. 2.1

Gas

PP

PMP

Oxygen Carbon dioxide

2.2 9.2

32.3 92.6

Permeability (1 Barrer ¼ 1010 cm3 cm/(s cm2 cmHg).

2.5 Membrane Production

Fig. 2.10 REM photograph of a microporous PMP membrane with dense outer skin (OXYPLUS, Membrana, Wuppertal).

Gas permeabilities of OXYPLUS for different gases and the resulting separation coefficients.

Tab. 2.2

Gas Nitrogen Oxygen Carbon dioxide

Flux (L/m2 min bar) 1.68 5.90 16.90

Separation coefﬁcient related to nitrogen –— 3.53 10.13

different gases through OXYPLUS and the resulting gas separation coefﬁcients. With this membrane, the enrichment of oxygen in air to a concentration of, for example, 25 %, is possible and effective in a single-step process.

2.5 Membrane Production

Oxygenation membranes are mainly produced by two different processes: Temperature-induced phase separation process Melt spin stretch process. Seventy percent of oxygenation membranes worldwide are produced by a special TIPS process, the so-called ‘‘ACCUREL’’ process [48–50] (Figure 2.11). A suitable polymer and a solvent or solvent mixture such as natural seed oils (e.g., soybean and castor) are heated together in an extruder to produce a homogeneously mixed polymer solution. This solution is pumped through a spinneret together with a gas as bore medium to create the right geometry of the hollow ﬁber. The preformed hollow ﬁber geometry is then cooled in an air gap and in a cold spinning oil, initiating phase separation of the solvent and the porous polymer matrix with pores still ﬁlled

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Fig. 2.11 Scheme of the ACCUREL process for manufacturing of membranes by thermally induced phase separation.

with oil. After a washing step, for example, with hot alcohol, to remove the solvent, the membrane is wound onto bobbins. The following makeup steps will be explained in Section 2.6. The membrane structure is adjustable by varying the composition of the polymer solution, the cooling speed, and the process temperatures during spinning and cooling. A homogenous sponge-like structure is formed, as shown schematically in Figure 2.12. In the melt spin stretch process, the membrane pores are produced in a completely different way. First a suitable crystalline or semicrystalline polymer is molten in an

Fig. 2.12 Schematic picture of ACCUREL membrane wall.

2.6 Operational Modes and Membrane Makeup in Oxygenators

Fig. 2.13 Scheme of the stretching process to produce microporous membranes.

Fig. 2.14 Scheme of microporous membrane produced by melt spin stretch process.

extruder. The molten polymer is pressed through a spinneret and on cooling a dense compact capillary a ﬂat sheet is created. By annealing, stretching, and cooling steps with a special temperature and tension proﬁle in several repeated steps (see Figure 2.13), microporous membranes with completely different structures are created. This process has advantages in that no solvent and extraction are required but also disadvantages due to the lower porosity, the limited range of pore size, and the elongated pores (lower selectivity) of the product. The schematic structure of the membrane wall is shown in Figure 2.14.

2.6 Operational Modes and Membrane Makeup in Oxygenators The oxygenator construction and the arrangement of the membranes inside the oxygenator are as important as the performance of the oxygenation membrane itself.

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2.6.1 Microporous Capillary Membranes, Blood Inside

In the 1980s, the blood was pumped inside the capillary membrane (luminal ﬂow) whereas the membrane was ﬂushed with gas from the outside. The disadvantages of these oxygenators, which are no longer in use today, are a large membrane quantity is required, as the gas exchange surface is the smaller inner capillary surface poor biocompatibility large priming volume high production costs high pressure drop. 2.6.2 Microporous Capillary Membranes, Blood Outside

Because of the disadvantages of intraluminal blood ﬂow described above, oxygenators with blood ﬂowing outside the capillaries (extraluminal ﬂow) were developed in the mid-1980s [28,29,37,43]. The structure of these oxygenators with wound capillary membranes or with knitted mats provides good mixing of the blood and thus an improvement in the oxygen transfer rates, especially in the blood boundary layer (see Section 2.3). The membrane surface used can be minimized. More than 95 % of the membrane oxygenators used in hospitals today are oxygenators with blood ﬂow outside the capillaries. The most effective construction uses cross-laid knitted mats (Figure 2.15). With this capillary arrangement, the following features are achieved: ‘‘enhanced mass transfer efﬁciency’’ with ‘‘local laminar mixing’’ (static mixer effect) repeated renewal of the concentration in the boundary layers; therefore, highest efﬁciency (see Section 2.3) uniform distribution of capillaries

Fig. 2.15 Cross-laid knitted membrane mat.

2.6 Operational Modes and Membrane Makeup in Oxygenators

capillaries ﬁxed by warp threads (no movement) suitable for various oxygenator designs (axial, tangential, crossﬂow, etc.) blood ﬁlm thickness (pressure drop) can be adjusted easily easy handling during manufacture high production yield closed capillary ends for safe potting.

The process steps for the production of cross-laid knitted mats are shown in Figure 2.16. Several capillaries are continuously laid down in parallel to a knitting machine and connected with a warp thread at deﬁned distances. During this step multiple mats are cut, the membrane ends are closed, and the mat spools are wound up in parallel. In the next step, two mats are pulled at each deﬁned angle and wound up together.

Fig. 2.16 Scheme of production of cross-laid knitted hollow ﬁber mats.

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Fig. 2.17 Mass transfer coefﬁcient versus membrane angle in a cross-wound mat device.

For each oxygenator setup, a special distance between the membranes, and a special warp thread distance and angle between the cross wound mats are possible and can be optimized (Figure 2.18). In this way one can always create an optimal blood stream. In the case of membrane contactors in technical applications, such constructions are also an advantage. The effects of design and operating variables on module performance have been investigated with respect to oxygen transfer into water. It was found that the membrane angle with respect to the main direction of the liquid ﬂow had a signiﬁcant inﬂuence on oxygenator performance [37]. The mass transfer coefﬁcient increased proportionally to the membrane angle up to about 108 and started leveling off at about 20–258 (Figure 2.17).

Fig. 2.18 Scheme of cross-laid hollow ﬁbers.

2.7 Extracorporeal Circulation

2.7 Extracorporeal Circulation 2.7.1 Cardiodiapulmonary Bypass (CPB)

In Figure 2.19, an extracorporeal circuit called a cardiopulmonary bypass, as used during open-heart surgery, is schematically demonstrated. Venous blood is sucked out by several cannulas from the patient. The blood is pumped, directly or via a venous reservoir, through the oxygenator. If applied, it is mixed with a part of sucked cardiotomic wound blood. The blood ﬂow through the oxygenator is between 3 and 6 L/min. Here the CO2 is removed from the blood and the oxygen is transported into the blood. All the blood from the wound is also brought back to the circuit (Figure 2.19). 2.7.2 Lung Support Systems

If the native lung is diseased, for example, by asthma, emphysema, or acute respiratory distress syndrome (ARDS), or is injured and thus no longer able to supply the body with sufﬁcient amounts of oxygen, two treatment options exist, depending on the severity and comorbidities: lung support with oxygen-enriched air or blood oxygenation (ECMO). For the ﬁrst treatment option, there is a need for a mobile oxygen-enriching device, and for the second option, there is a need for an oxygenation membrane suitable for long-term application (several days to weeks). Some membranes described in Section 2.4.2 are suitable for both tasks (Figure 2.21). In intensive care, artiﬁcial lungs with long-term oxygenation membranes, as described in Section 2.4.2 become a treatment option for acute lung failure (e.g.,

Fig. 2.19 Scheme of an oxygenator in a heart lung machine setup.

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Fig. 2.20 Example of an oxygenator setup with heat exchanger, venous reservoir, and oxygenator and photo of the oxygenator with cross-laid knitted mats inside.

Fig. 2.21 Nova lung oxygenator for an ECMO application.

from virus pneumonia, after blunt trauma [21], following pancreatitis, or in conjunction with sepsis or multiorgan failure) or for lung failure, for example, posttransplant, postoperative, or pediatric cases. An excellent example for a pumpless ECMO device is the Novalung ‘‘membrane ventilator.’’ Two cannulas are connected to large vessels of the patient and by the force of the patients’ heart the blood is

Fig. 2.22 Implantable lung support setup.

References

pumped through the oxygenator to remove carbon dioxide and/or provide oxygen to the body [39] (Figure 2.21). The membrane can also be suitable for use in long-term paracorporeal or implantable artiﬁcial lungs, for example, inside the vena cava [22,25]. In Figure 2.22, an intravascular membrane oxygenator (IVOX) shows the ﬁrst vena caval device with membranes [5,40]. Comparable devices and oxygenation membrane setups, as described in Sections 2.6 and 2.7, are also used in biotechnology (see Chapter 5) and bioartiﬁcial organs (see Chapter 8) [51,52].

References 1 Le Gallois, J.J.C. (1812) Expe´riences sur

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le principe de la vie, nottament sur celui des mouvements du Coeur, et sur le sie´ge de ce principe; Suievies du Rapport fait a` la premie`re classe d l’I´nstitut sur celles relatives aux mouvemens du coeur, D’Hautel, Paris, 134–135. Le Gallois, J.J.C. (1813) Experiments on the Principle of Life, and Particularly on the Principle on the Motions of the Heart and on the Seat of this Principle: Including the Report Made the First Class of the Institute, upon the Experiments Relative to the Motion of the Heart, M. Thomas, Philadelphia, PA, 130–131. Schmidt, A. (1867) Die Atmung innerhalb des Blutes. Zweite Abhandlung. Aus dem physiologischen Institut zu Leipzig. Berichte u¨ber die Verhandlungen der ko¨niglich sa¨chsischen Gesellschaft der Wissenschaften zu Leipzig. Mathematisch–physische Classe, 19, 99–130. Gibbon, J. H. (1937) Artiﬁcial maintenance of circulation during experimental occlusion of pulmonary artery. Archives of Surgery, 34, 1105– 1131. Mallabiabarrena Ormeaetxea, I. (2003) Experimental set up and numerical simulations of intravenous gas transfer devices, Doctoral thesis no. 2873. E´cole Polyteechnique Fe´de´rale de Lausanne, Switzerland. Dierickx, P. (2001) Blood ﬂow and gas transport in artiﬁcial lungs: in numero

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and in vitro analysis, Doctoral thesis. University of Gent. Zwischenberger, J. B. and Alpard, S. K. (2002) Perfusion, 17, 253–268. Clowes, J. J. and Hopkins, A. L. (1956) An artiﬁcial lung depended upon diffusion of oxygen and carbon dioxide through plastic membranes. Journal of Thoracic and Cardiovascular Surgery, 32, 630–637. Sewringhaus, J. W. (1979) Journal of Applied Physiology, 46, 599–602. Diericks, P. W., De Somer, F., De Wachter, D. S., Van Nooten, G., Verdonck, P. R. (2000) ASAIO Journal, 46, 532–535. Mockros, L. F. and Leonard, R. (1985) ASAIO Transactions, 31, 628–633. Vaslef, S. N. and Mockros, L. F. (1994) ASAIO Transactions, 40, 990–996. Wicksramasinghe, S. R., Semmens, M. J., Cussler, E. L. (1992) Journal of Membrane Science, 69, 235–250. Allen, S. M., Fujii, M., Stannett, V., Hopfenberg, H. B., Williams, J. L. (1977) Journal of Membrane Science, 2, 153–163. Wiese, F., Paul, D., Possart, W., Malsch, G., Bossin, E. (1990) Acta Polymerica, 41, 95–98. Prospectus ‘‘Jostra QuadroxD’’, Jostra AG, Hechingen, Germany, 2003. Prospectus ‘‘Medos Hilite7000 LT’’, Medos AG, Stolberg, Germany, 2002. Philipp, A. (2003) Kardiotechnik, 1, 7–13. Horton, S., Thuys, C., Bennet, M. (2004) Perfusion, 19, 17–23.

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(2002) Perfusion, 17, 175–178. Schmidt, F. X., Philip, A., Link, J. (2002) Annals of Thoracic Surgery, 73, 1618–1620. Zwischenberger, J. P. (2001) ASAIO Journal, 47, 316–320. DeWall, R. A. and Lillehei, C. W. (1957) Surgery of Gynecology and Obstetrics, 104, 699. Kolff, W. J. and Berk, H. T. (1944) Acta Medica Scandinavica, 117, 121–134. Zwischenberger, J. P. and Alpard, S. K. (2002) Perfusion, 17, 253–268. Product information (1982) The Journal of Extracorporeal Technology, 14, 18. Sienkewich, M. and Maserko, J. J. (1982) Proceedings of the American Academy on Clinical Perfusion, 3, 13–16. Gassmann, C. J., Balbraith, G. D., Smith, R. G. (1987) The Journal of Extracorporeal Technology, 19, 297–304. Alpha, D., King, E., Bicknell, D. A. (1986) Proceedings of the American Academy on Clinical Perfusion, 7, 32–34. Lauterbach, G. (2002) Handbuch der Kardiotechnik, Urban & Fischer, Mu¨nchen. Tschaud, R. J. (1999) Extrakorporale Zirkulation in der Theorie und Praxis, Papst Science Publischers, Lengerich. Taylor, K. M. (1986) Cardiopulmonary Bypass, Lippincott Williams & Wilkins, Philadelphia, PA. Philipp, A., Foltan, M., Gietl, M., Reng, M., Liebold, A., Kobuch, R., Keyl, C., Bein, T., Mu¨ller, T., Schmid, F.-X., Birnbaum, D. E. (2003) Kardiotechnik, 1, 7–13. Schmid, F.-X., Philipp, A., Link, J., Zimmermann, M., Birnbaum, D. E. (2002) Annals of Thoracic Surgery, 73, 1618–1620. Zwischenberger, J. B., Anderson, C. M., Cook, K. E., Lick, S. D., Mockros, L. F., Bartlett, R. H. (2001) ASAIO Journal, 47, 316–320. Lillehei, C. W. (1955) Postgraduate Medicine, 17, 388–396. Wodetzki, A., Breiter, S., Scheuren, J., Wiese, F. (2000) Maku (Membrane),

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Japanese Membrane Journal, 25, 102– 106. Breiter, S. M., Wiese, F., Wodetzki, A., Schuster, O. (2004) ASAIO Journal, 50, 153. Jegger, D., Revelly, J.-P., Horisberger, J., Boone, Y., Seigneul, I., Jachertz, M., von Segesser, L. K. (2004) ESAO, Aachen. Ex vivo evaluation of a new extracorporal lung assist device. Novalung, ESAO, Wausau. Mortensen, J. D. (1992) Artiﬁcial Organs, 16, 75–82. Steffen, S., Oedekoven, S. B., Henseler, A., Mottaghy, K. (1997) Kardiotechnik, 3, 68–71. Montoya, J. P., Shanely, C. J., Scott, I. M., Bartlett, R. H. (1992) ASAIO Journal, 38, 399–405. Catapano, G., Hornscheidt, R., Wodetzki, A., Baurmeister, U. (2004) Journal of Membrane Science, 230, 131–139. Baker, R. W. (2000) Membrane Technology and Applications, McGrawHill, New York. Kolff, W. J. and Balzer, B. R. (1955) ASAIO Transactions, 1, 39–42. Kolff, W. J. and Efﬂer, D. B. (1956) Cleveland Clinic Quarterly, 23, 69–97. Prospectus Museum ICMT (2006) International Center for Medical Technologies. Castro, A. J. (1981) US Patent 4 247 498. Henne, W., Pelger, M., Gerlach, K., Tretzel, J. (1983) Lysagth M. J. and Gurland H. J. Plasmaseparation and Plasmafractionation, Karger AG, Basel. Hiatt, W., Vitzthum, G. H., Wagner, K. B., Gerlach, K. (1985) Materials Science of Synthetic Membranes, American Chemical Society, Washington, DC. Gerlach, J., Kloppel, K., Stoll, P., Vienken, J., Mulle, C. (1990) Artiﬁcial Organs, 14, 328–333. Gerlach, J. C. (1996) International Journal of Artiﬁcial Organs, 19, 645–654. Krause, B., Go¨hl, H., Wiese, F. (2006) Ohlrogge K. and Ebert K. Membranen, Wiley-VCH, Weinheim.

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3 Membranes for Blood Fractionation/Apheresis Frank Wiese 3.1 Introduction

The use of membranes in extracorporeal systems connected to the human body started in the 1940s with Kolff’s ﬁrst use of an artiﬁcial kidney. Since that time, the application of membranes has been established in different areas (Figure 3.1). Nowadays, the expression, apheresis or plasmapheresis, is used for the removal of blood plasma from cells as well as for the removal of protein components or fractions

Fig. 3.1 Membranes in extracorporeal systems.

Fig. 3.2 Therapeutic hemapheresis classiﬁed according to technologies. Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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from this plasma. The origin is the Greek word aphairesis – ‘‘to separate, to remove, and to take away.’’ In comparison to 120 million artiﬁcial kidney treatments and 1.2 million open-heart operations using artiﬁcial lungs, there are only around 600 000 membrane apheresis treatments carried out per year. Nevertheless, these treatments are an important therapeutic tool for special medical indications for autoimmune diseases or diseases related to blood rheology. Membrane procedures have signiﬁcance for plasma separation and plasma fractionation processes – also called cascade ﬁltration or double ﬁltration – as well as for speciﬁc adsorption processes because of the requirement for cell-free plasma for these procedures. For the functionality and the life-time of these modern adsorption processes, cell-free plasma is a strict requirement (Figure 3.2). Blood plasma donation for medical and pharmaceutical applications is dominated by centrifugal procedures. Procedures such as cytapheresis or cell separation are not discussed in this chapter because of the negligible use of membranes in these procedures.

3.2 History of Plasmapheresis

In the ancient world, the role of blood in health and illness was already known. The name of Hippocrates (460–377 B.C.), the ‘‘Father of Medicine,’’ is linked to the

Fig. 3.3 Hippocrates Hiraclidae (460–377 B.C.) – The ‘‘Father of Medicine.’’

Fig. 3.4 Bloodletting depicted by Peytel.

Fig. 3.5 ‘‘Plasmapheresis’’ described by Abel, Rowntree, and Turner [3].

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Fig. 3.6 John Jacob Abel 1857–1938.

Fig. 3.7 Monsanto hollow ﬁber device as experimental plasma separator 1972 [4].

3.3 Principles of Plasmapheresis

Fig. 3.8 Principle of double ﬁltration ﬁrst described by Agishi [5].

introduction of bloodletting as a preventative method as well as a remedy [1,2]. Although bloodletting should not be confused with plasmapheresis, it is its therapeutic precursor (Figures 3.3 and 3.4). The term plasmapheresis was ﬁrst used by Abel, Rowntree, and Turner in 1914. In their paper ‘‘Plasma Removal with Return of Corpuscles (Plasmapheresis)’’ [3], they investigated the effect of plasma removal from dog blood and reinjected the cells together with replacement solution (Figures 3.5 and 3.6). This was the early beginning of therapeutic plasmapheresis. The ﬁrst hollow ﬁber membrane device for plasma separation was introduced by Monsanto in 1972 [4]. This was the start of therapeutic plasma separation by membranes (Figure 3.7). The ﬁrst double ﬁltration in place of two membrane ﬁlters with different pore sizes was described by Agishi [5]. High molecular weight proteins such as low-density lipoproteins (LDL) and triglycerides were removed and lower molecular weight proteins such as albumin and high-density lipoprotein (HDL) were returned to the body (Figure 3.8).

3.3 Principles of Plasmapheresis

The schematic setup of a plasma separation treatment is shown in Figure 3.9. The blood is removed from the body and ﬁltered in a plasma ﬁlter device. Then the ﬁltered plasma is removed and cells are returned to the body as concentrate, together with a substitution ﬂuid such as an isotonic electrolyte solution, an albumin-isotonic electrolyte solution, or donor plasma [6–9]. Blood is a special streaming medium, and this special characteristic has to be considered by selecting the right process conditions. In capillary modules, the transmembrane pressure (TMP) is usually not higher than 150 mmHg, and the adjusted ﬁltrate ﬂow is between 20 and 35% of the blood ﬂow rate of 50–250 mL/min. In most classical plasma exchange procedures (TPE – therapeutic plasma exchange), plasma is removed and replaced by donor plasma (FFP – fresh frozen plasma). It is used not only for the treatment of several immunological diseases but

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Fig. 3.9 Membrane plasma separation set up.

also for the removal of metabolism products and exogenic and endogenic toxins [6,8]. In several countries, this unspeciﬁc exchange is carried out even now (mostly by the use of centrifuges), but the frequency is decreasing because of the availability of more speciﬁc adsorption methods and the many clinical disadvantages of this method. To save mostly patients’ own speciﬁc proteins such as albumin and IgG and to remove only the poisoning components, for example, LDL from, for example, hypercholesteremic patients, membrane plasma separation is used in combination with secondary membrane ﬁltration called ‘‘plasma fractionation,’’ ‘‘membrane double ﬁltration,’’ ‘‘rheopheresis,’’ or ‘‘cascade ﬁltration’’ (Figure 3.10) [29]. The blood is divided into a cell-concentrated part and a cell-free plasma part in the ﬁrst ﬁltration step. In the second step, the blood plasma is divided into high and low molecular

Fig. 3.10 Membrane double ﬁltration set up (cascade ﬁltration, plasma fractionation).

3.3 Principles of Plasmapheresis Tab. 3.1

Blood as a streaming medium – specific features during application.

Primary ﬁltration

Secondary ﬁltration

Blood viscosity depends on Type of cells Concentration of cells (hematocrit) ‘‘Cross-ﬂow operation’’: plasma-rich layer Hemolysis

Plasma viscosity depends on Type of macromolecules/proteins Concentration of macromolecules/proteins ‘‘Dead-end operation’’: secondary membrane layer Blocking

weight fractions of plasma. The low molecular weight fraction and nearly the complete plasma ﬂuid volume are then recombined with the cell concentrate and returned to the patient. The sieving characteristics of these membranes are described in Section 3.4. The primary ﬁltration step (plasma separation) is done as a ‘‘cross-ﬂow operation,’’ and the secondary ﬁltration as a ‘‘dead-end operation.’’ This has to be considered by selecting the right process conditions to prevent blood damage and to keep the process running. The right shearing proﬁle in the primary ﬁlter prevents hemolysis, and the right pressure in the secondary ﬁler prevents blocking (see Table 3.1 and Section 3.6). The setup of plasma fractionation in combination with adsorbers is shown in Figure 3.11. Again the plasma is separated with a plasma ﬁlter. In the second step, the separated plasma is pumped through protein-speciﬁc adsorbers to remove special proteins such as immunoglobulins or toxins according to the therapeutic objectives. Finally, the treated plasma together with the blood cells is returned to the patient. Different kinds of single-use adsorbers or regenerable double adsorbers for different medical indications are on the market. In Figure 3.11, an example of a double-column setup is shown. While one column is recirculated with plasma and loaded with

Fig. 3.11 Membrane plasma separation combined with plasma fractionation by adsorbers.

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proteins to be removed, the second column is regenerated. The columns are switched from time to time and are applicable for multiple patient treatments.

3.4 Membranes and Membrane Properties

The general proﬁle of the requirements of plasma treatment membranes is summarized as follows [6,7,10,28]: permeable for all plasma proteins/special sieving characteristics high surface porosity and volume porosity to reach high ﬂuxes hydrophilic, spontaneous, and wettable low fouling behavior low protein adsorption smooth blood contact surfaces low hemolytic potential consistency in sieving behavior and ﬁltration performance during treatment good blood compatibility [17–19,22,24] suitable mechanical stability stability against chemicals and all common sterilization procedures (ETO, g-radiation, steam). The plasma fractionation membranes have to have special sieving characteristics for the removal of different protein fractions, depending on the medical indications. Different producers of ﬁlters use different polymers such as cellulose acetate (CA), polypropylene (PP), polyethylene (PE), polysulfone (PSU), polyethersulfone (PES), and polymethylmethacrylate (PMMA) to fulﬁll these requirements as far as possible. Hydrophobic polymers such as PP are hydrophilized during module manufacture or PE is coated with polyethylvinylalcohol (EVAL), and PES is blended with polyvinylpyrolidone (PVP) [31–34]. 3.4.1 Plasma Separation Membranes

Plasma separation capillary membrane ﬁlters from different suppliers are listed in Table 3.2 [14]. All the used membranes have nominal pore sizes around 0.2–0.3 mm (microﬁltration membranes). The module suppliers often do not explain whether the pore size in their data sheets is theoretically calculated from the bubble point or measured by a sieving curve, thus creating some confusion. The size exclusion pore size of 0.2 mm guarantees that all plasma proteins can pass through the membrane whereas all cells including the small thrombocytes with 2 mm diameter are rejected.

3.4 Membranes and Membrane Properties Capillary membrane plasma separation filters from different producers.

Tab. 3.2

Supplier

Filter

Asahi

Plasmaﬂo OP-02

B. Braun Bellco

Gambro Fresenius Kaneka

Kuraray Toray Edwards Dideco

Plasmaﬂo OP-05 Plasmaﬂo OP-08 Hemoselect MPS 02 MPS 05 MPS 07 PF 1000 PF 2000 Plasmaﬂux P1S Plasmaﬂux P2S Sulﬂux FS-03 Sulﬂux FS-05 Sulﬂux FS-07 Plasmacure Plasmax PS-02 Plasmax PS-05 Microplas MPS 05 Hemaplex BT 900

Membrane material

Membrane area (m2)

ID/wall (mm)

Sterilization mode

Polyethylene/ EVAL coating

0.20

330/50

ETO

330/150 300/100

Steam ETO

330/150

ETO

330/150

Steam

340/50

g-radiation

320/65 330/90

g-radiation g-radiation

300/100 330/150

ETO, g-radiation ETO

Polypropylene Polyethersulfone

Polypropylene Polypropylene Polysulfone

Polysulfone PMMA Polyethersulfone Polypropylene

0.50 0.80 0.20 0.28 0.45 0.68 0.16 0.35 0.25 0.50 0.30 0.50 0.70 0.60 0.15 0.50 0.45 0.50

The sieving coefﬁcient (SC) for a special protein in a ﬁltration experiment is calculated according to Equation (1): CF is the concentration of the protein in the ﬁltrate, CBi the ﬁlter input concentration, and CBo the ﬁlter output concentration of the protein in the blood or blood plasma. SC ¼

2CF : CBi þ CBo

ð1Þ

The sieving characteristics of a typical plasma separation membrane are shown in Figure 3.12(a). Nearly 100 % of all plasma proteins pass through the membrane including the lipoproteins with molecular weights >2 000 000 Da. Fibrinogen is slightly reduced. This can be related to the special shape (see Figure 3.13) [11] of this molecule and adsorption effects. The same membrane was also investigated with latex beads of deﬁned different sizes. The rampant sieving curve shows that it has a nominal pore size of 0.2 mm and a narrow pore size distribution (Figure 3.12(b)). These capillary membranes also meet the requirements of the HIMA test, which means that the membrane can also be used for sterile ﬁltration, for example, intravenous ﬁltration, because it is fully bacteria retentive against pseudomonas diminuta. Depending on the membrane polymer used, the composition of the spinning solution and the production process, and the structure and wall thickness of the

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Fig. 3.12 Sieving coefﬁcients of plasma separation membrane tested with blood and latex beads (MicroPES TF10, Membrana GmbH).

Fig. 3.13 Size comparison of noncellular blood components [11].

3.4 Membranes and Membrane Properties

Fig. 3.14 SEM picture of cross sections (a, b) and blood contacting inner surfaces (c, d) of two different plasma separation membranes [25] (a, c) Plasmaphan (Membrana GmbH) and (b, d) Plasmaﬂo (Asahi)).

membranes are different. The SEM pictures in Figure 3.14 illustrate this with examples of two different plasma separation membranes produced from polyoleﬁns PP and PE. The PP membrane has a signiﬁcantly higher wall thickness than the PE membrane and different pore structure of the inside skin, whereas in vivo performance is comparable. 3.4.2 Plasma Fractionation Membranes

Plasma fractionation capillary membrane ﬁlters from different suppliers are listed in Table 3.3. The membranes used in the ﬁlters have different sieving characteristics, depending on the medical indications. For example, the different Evaﬂux membrane ﬁlter types are used for the removal of different protein fractions. Evaﬂux 2A is used for the removal of immunoglobulins IgG and IgA, Evaﬂux 3A for the removal of cytokines, Evaﬂux 4A for the removal of IgM and LDL, and Evaﬂux 5A for the removal of LDL-C (cholesterin) [26]. The sieving coefﬁcient for a special protein in a ﬁltration experiment is calculated according to Equation (1). In Figure 3.15, the sieving characteristics of a plasma fractionation membrane are shown to exemplify the behavior of such a membrane.

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3 Membranes for Blood Fractionation/Apheresis Tab. 3.3

Capillary membrane plasma fractionation filters from different producers.

Supplier

Filter

Membrane material

Membrane area (m2)

ID/wall (mm)

Sterilization mode

Asahi

Rheoﬁlter Cascadeﬂo EC-20 Cascadeﬂo EC-30 Cascadeﬂo EC-40 Cascadeﬂo EC-50 Evaﬂux 2A Evaﬂux 3A Evaﬂux 4A Evaﬂux 5A Albuafe FP 2

Cellulose diacetate EVAL EVAL EVAL EVAL EVAL EVAL EVAL EVAL Polyethersulfone

2.0 2.0 2.0 2.0 2.0 2.0 2.0 2.0 2.0 1.7

220/80 175/40 175/40 175/40 175/40 175/40 175/40 175/40 175/40 200/35

g-radiation

Kuraray

Dideco

g-radiation

g-radiation, steam

The sieving coefﬁcient for albumin is around 0.88. This is important because the albumin has to pass through the membrane so that as much as possible of this protein is given back to the patient. Most molecules with molecular weights up to around 400 kDa, for instance HDL, pass through the membrane with the exception of ﬁbrinogen. Fibrinogen is a stretched long molecule and for this geometric reason is partly rejected (see Figure 3.13). Larger molecules such as alpha-2-macroglobulin or IGM are mostly rejected – the sieving coefﬁcient of IGM is 0.2. The main goal of the treatment is to remove these lipoprotein molecules. In conclusion, this membrane can be used for hypercholesteremic as well as for rheological indications in the clinical applications (see Section 3.7). The SEM picture of the cross section of this membrane (see Figure 3.16) shows an isotropic and homogeneous pore structure and a smooth inner surface. The inner diameter is 200 mm, and the wall thickness is 35 mm.

Fig. 3.15 Sieving coefﬁcients of a plasma fractionation membrane for different blood plasma proteins – FractioPES (Membrana GmbH).

3.5 Membrane Production

Fig. 3.16 SEM picture of the membrane structure of the plasma fractionation membrane FractioPES (Membrana GmbH).

3.5 Membrane Production

Plasma separation membranes are mainly produced by three different processes (see also Section 2.5) (Figure 3.17–3.19): Solvent-induced phase separation (SIPS) process Temperature-induced phase separation (TIPS) process Melt spin stretch (MESS) process. The schematic course of an SIPS process, also called the wet spinning process, is shown in Figure 3.17. From a suitable polymer and a suitable solvent or solvent mixture, a solution is prepared. Here the ‘‘genome’’ of the membrane is created

Fig. 3.17 Schematic depiction of the solvent/nonsolventinduced phase separation process for manufacturing of microporous membranes.

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because the membrane structure is predetermined (preﬁxed) in the solution. The solution preparation, including proper ﬁltration, is a complex and important step. After this step, the solution is pumped into a spinneret for capillary membranes or into a casting box in the case of ﬂat membrane production. To give the capillary membranes the correct geometry and to steer the inner skin pore structure, the bore liquid can be an inert or a composition of solvent and nonsolvent. The next important step is the precipitation, the ‘‘birth’’ of the membrane. The solvent is replaced by the nonsolvent of the coagulation bath and creates the porous polymer matrix. This diffusion process is inﬂuenced by the polymer solution composition as well as by process conditions such as the temperature of the precipitation bath, the spinning speed, the climatic conditions between the spinneret and the precipitation bath, and so on. The coagulation is followed by the washing, drying, and bundle makeup steps until the bundle is ready for insertion into a module housing. In special cases, a postﬁnal extraction is carried out. Most membranes worldwide are produced by wet spinning processes. Plasmapheresis membranes from polyoleﬁns (PP, PE) are mainly produced by two different processes: TIPS MESS. A special TIPS process, the so-called ‘‘ACCUREL’’ process, is described in Figure 3.18 [15,16,38,44]. A suitable polymer and a solvent or solvent mixture such as natural seed oils (e.g., soybean and castor) are heated together in an extruder to produce a homogeneously mixed polymer solution. This solution is pumped through a spinneret

Fig. 3.18 Schematic depiction of the ACCUREL process for the manufacturing of membranes by thermally induced phase separation.

3.6 Operational Modes in Plasmapheresis Procedures

Fig. 3.19 Schematic depiction of the pore building part in a melt spin stretch process for producing of microporous membranes.

together with a gas as bore medium to create the right geometry of the hollow ﬁber. The preformed, hollow ﬁber geometry is then cooled in an air gap and in cold spinning oil, initiating a phase separation of the solvent and the porous polymer matrix, the pores of which are still ﬁlled with oil. After a washing step, for example, with hot alcohol, to remove the solvent, the membrane is wound onto bobbins or into bundles. The membrane structure can be adjusted by varying the composition of the polymer solution, the cooling speed, and the different process temperatures during spinning and cooling. In the melt spin stretch process, the membrane pores are produced in a completely different way. First, a suitable crystalline or semicrystalline polymer is melted in an extruder. The molten polymer is pressed through a spinneret, and on cooling a dense compact capillary or ﬂat sheet is created. By annealing, stretching, and cooling steps with a special temperature and tension proﬁle in several repeated steps (see Figure 3.19), microporous membranes with completely different structures are created (see Figure 3.14(d) and Section 2.5). This process has advantages because no solvent and extraction is required but also disadvantages because of the lower porosity, limited range of pore size, and elongated pores (lower selectivity).

3.6 Operational Modes in Plasmapheresis Procedures

As mentioned in Section 3.3, suitable process conditions have to be adjusted in the clinical application of plasma separation to reach a successful treatment without negative side effects. The main design criteria are the wall shearing rate, the created pressure drop along the ﬁlter housing, and the resulting plasma ﬁltration rate. The wall shearing rate is calculated according to Equation (2): wall shearing rate : gw ¼

4QB ; Npr 3

ð2Þ

where N is the number of capillary membranes with inner radius r and QB is the blood ﬂow distributed to the different capillary membranes. The blood ﬂow changes by reduction of the plasma content over the length of the membrane or the ﬁlter,

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respectively. This has to be considered when calculating the wall shearing rate. Furthermore, the transmembrane pressure is further an important parameter for steering the process in the right direction. The transmembrane pressure is calculated according to Equation (3): TMP ¼

PBi þ PBo PF : 2

ð3Þ

The TMP is calculated by blood input pressure (PBi), blood output pressure (PBo), and ﬁltrate pressure (PF) in the plasma compartment. In a laminar stream, the velocity proﬁle will be distributed as shown in Figure 3.20. The erythrocytes will be forced to go into the central stream. As explained in Section 3.3, plasma separation is carried out in cross-ﬂow ﬁltration mode. In the low-pressure range, the ﬁltration performance increases linearly with increasing transmembrane pressure. Later, a plateau is reached according to the wellknown concentration polarization phenomenon [7,30,36,37] in cross-ﬂow ﬁltration. Blood is not a Newtonian ﬂuid, but nevertheless it is well known that, according to Hagen–Poiseuille’s law, the streaming volume is very sensitive to tube/capillary diameter. Consequently, the blood ﬂow rate, the plasma ﬂux, and the corresponding transmembrane pressure have to be in the right relation to each other [14,31]. As long as the streaming velocity is high enough, the cells will be forced to move into the center of the stream (see Figure 3.20). If the wall shearing rate becomes too low and the membrane pressure is too high, the risk of hemolysis increases. This means that the red blood cells (erythrocytes) are destroyed and hemoglobin is released in the body circuit. The blood plasma becomes

Fig. 3.20 Velocity proﬁle in a capillary membrane [13].

3.6 Operational Modes in Plasmapheresis Procedures

Fig. 3.21 Schematic explanation of hemolytic behavior of different membranes in blood contact depending on transmembrane pressure and wall sheer rate.

red in color resulting in toxic behavior. This condition has to be prevented. The starting phase of the treatment is particularly critical. For this reason, some ﬁlters have to be processed at transmembrane pressures less than 50 mmHg. The user has always to follow the care of cartridge producers’ manual. In Figure 3.21, two membranes A and B are compared with respect to their hemolytic dispositions. In general, the hemolytic tendency decreases with increasing wall shear rate and decreasing pressure. Membrane B shows hemolysis related to lower transmembrane pressures than membrane A. Above the dashed line is the area where hemolysis takes place. The schematic assembling of a plasma separation capillary membrane ﬁlter is shown in Figure 3.22. The requirements according to module design for capillary membrane plasma ﬁlters are not so high as for, for example, dialyzers [7,9,11,35]. Only a convective ﬂow via the membrane wall takes place, and because the plasma is only drained in the ﬁltrate (plasma) compartment of the module, no deﬁned streaming conditions have to be created inside the bundle. Therefore, no ﬁber makeup such as undulation is required. According to the above-explained criteria, the ﬁlter area and the length have to be adjusted to the membrane. According to Table 3.2, ﬁlter areas between 0.15 and 0.8 m2 are usual, depending on the application. A special monitor steers the ﬂows and pressures comparable to a dialysis monitor. Plasma volumes between 2 and 4 L are ﬁltered in one treatment session. Flat membranes are also used in special ﬁltration devices. Regarding the membrane area they are very effective, but need special monitors (Figure 3.23). In the Autopheresis C system produced by Baxter/Fenval, IL, USA, a polyamide (PA) membrane with an area of 70 cm2 is applied to a rotating cylinder. The blood enters the top of the ﬁlter assembly and travels down the deﬁned space between the

85

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3 Membranes for Blood Fractionation/Apheresis

Fig. 3.22 Schematic drawing of the module construction of a capillary membrane plasma separation ﬁlter.

rotating cylindrical membrane and the outer cylindrical housing (stationary cylinder). The plasma ﬁltered through the membrane ﬂows out from the ﬁlters’ central tube (axis). By rotation of the membrane, Taylor vortices are created, spiraling downward. The rotation-induced Taylor vortex enhances the plasma ﬂux [27]. In the device of Milttenyi Biotec GmbH, as described in Figure 3.24, less than 250 cm2 membrane area is required to ﬁlter several liters of blood. A polyethersulfone

Fig. 3.23 Flat membrane ﬁlter used in a rotating plasma separation device. The inner rotating cylinder is upholstered with a microﬁltration membrane of polyamide (Baxter/Fenwall, USA).

3.6 Operational Modes in Plasmapheresis Procedures

Fig. 3.24 Rotating disc plasma separator. Microﬁltration ﬂat membranes of MicroPES (Membrana GmbH) with rotating disk in between.

ﬂat membrane MicroPES is ﬁxed on both sides of the housing where the ﬁltered plasma is drained. A rotating disk in between and at a deﬁned small distance from the membrane ensures the right streaming force. This device is used to produce cell-free plasma, that is, treated in regenerable speciﬁc adsorber columns.

Fig. 3.25 From sieving . . . to adsorption – from diffusion . . . to convection membrane systems in future with combined sieving and selective adsorption feature inside the pores by grafted tentacles inside the pores.

87

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3 Membranes for Blood Fractionation/Apheresis

View to the future: Classical adsorber technology uses beads packed in a column. The beads carry the ligand and are exposed to the plasma or the blood stream. Membrane technology is an alternative way to build adsorption devices. The inherent advantage of membranes lies in the better perfusion of the plasma or whole blood. This is seen in the convective transport of the target molecules to the ligand sites without diffusive limitations. This is in contrast to the bead technology where the adsorption kinetics is limited by the long diffusion pass (Figure 3.25).

3.7 Medical Indications for Blood Plasma Treatment

In Table 3.4, medical indications for plasma fractionation treatments are listed. Nowadays, especially applications where the blood viscosity has to be reduced to improve the microcirculation of blood have been successful in an increasing number of cases [21,22,23,25]. For example, the membrane described in Figure 3.15 is well adapted to this type of application [20]. The high-molecular-weight lipoproteins and triglycerides are rejected completely, and the slight reduction in ﬁbrinogen also improves the viscosity behavior. Other illnesses require immunomodulation of the blood by fractionation of immunoglobulins [8,11,32,38,39]. Special kinds of such treatments are the so-called cryoﬁltrations where plasma proteins are precipitated by cooling the plasma to 4 8C. The precipitated aggregates are then ﬁltered by

Tab. 3.4

Medical indications for plasma fractionation by double filtration.

Indications belonging to top 10 of registry: Hypercholesterolemia Myasthenia gravis (MG) Syst. lupus erythematosus (SLE) Guillian–Barre´ syndrome (GBS) Thromb. thrombocyt. purpua (TTP) Cryoglobulinemia New indications: Age-related macular degeneration (AMD) Sudden hearing loss Diseases with rare incidence: Burger’s arthritis Cyclosporemia Diabetic retinopathy Gromerulonephritis Polymyositis/dermatomyositis Polyneuropathy Small vessel disease Vasculitis Waldenstrom’s macroglobulinemia W. granulomatosus

References

membranes. In the so-called thermoﬁltration special proteins are precipitated by heating the plasma to 42 8C and are then removed by ﬁltration [11,12]. The relatively ‘‘old’’ technology of membrane plasma separation is now undergoing a renaissance because of new indications and the requirement for highly sophisticated speciﬁc adsorber technologies for cell-free plasma [40–43].

References 1 Sawada, K. (1990) Nydegger U. E.

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Therapeutic Hemapheresis in the 1990s, Karger AG, Basel. Kambic, H., and Nose´, Y. (1997) Therapeutic Apheresis, 1, 83–108. Abel, J. J., Rowntree, L. G., Turner, B. B. (1914) Journal of Pharmacology and Experimental Therapeutics, 5, 625–641. Nose´, Y. and Malchesky, P. (1981) Therapeutic Plasmapheresis, 1, 3–14. Agishi, T., Kaneko, I., Hasuo, Y., et al. (1980) Transactions – American Society for Artiﬁcial Organs, 26, 406–411. Bambauer, R. (1997) Therapeutischer Plasmaaustausch und verwandte Plasmaseparationsverfahren, Pabst, Lengerich. Zeman, L. J. and Zydney, A. L. (1996) Microﬁltration and Ultraﬁltration: Principles and ApplicationsMarcel Dekker, New York. Ho¨rl, W. H. and Wanner, C. (2004) Dialyseverfahren in Klinik und Praxis, 6th edn , Georg Thieme Verlag, Stuttgart. Malchesky, P. S. (2001) Therapeutic Apheresis, 5, 270–282. Bo¨hler, J., Donauer, K., Ko¨ster, W., Schollmeyer, P. J., Wieland, H., Ho¨rl, W. H. (1991) American Journal of Nephrology, 11, 479–485. Sueoka, A. (1997) Therapeutic Apheresis, 2, 135–146. Siami, F. and Siami, G. A. (1997) Therapeutic Apheresis, 1, 58–62. Gerthsen, C. and Kneser, H. O. (1971) Physik, 11. Auﬂage, Springer, Berlin. Krause, B., Go¨hl, H., Wiese, F. (2006) Ohlrogge K. and Ebert K. Membranen, Wiley-VCH, Weinheim. Castro, A. J. (1981) US Patent 4,247,498.

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Tretzel, J. (1983) Lysagth M. J. and Gurland H. J. Plasma Separation and Plasma Fractionation, Karger AG, Basel. Nose´, Y., Jamaji, K., Sueoka, A., Yamane, s. (1997) Therapeutic Apheresis, 1, 5–12. Nose´, Y. (1988) Artiﬁcial Organs, 12, 377–378. Malchesky, P. S. (1992) Journal of Clinical Apheresis, 7, 145–146. Valbonesi, M., Carlier, P., Borberg, H. (2004) Journal of Artiﬁcial Organs, 27, 513–515. Borberg, H., Brunner, R., Widder, R. A. (1987) Japanese Journal of Apheresis, 16, 288–292. Beykirch, H., Voigt, H., Wilke, B. (1998) Nieren und Hochdruckkrankheiten, 27, 137–144. Henderson, L. W., Koch, K. M., Dinarello, C. A., Shaldon, S. (1983) Blood Puriﬁcation, 1, 3–8. Rock, W. (1990) Apheresis, Wiley-Liss, New York. Brunner, R. and Borberg, H. (1995) Journal of Artiﬁcial Organs, 794–798. Siami, G. A. and Siami, F. (2002) Therapeutic Apheresis, 5, 315–320. Kaplan, A. A. and Halley, S. E. (1990) Kidney International, 38, 160–166. Andrade, J.D. (1985) Surface and Interfacial Aspects of Boimedical Polymers, Vol. 2, Plenum Press, New York. Baeyer, H., Kochinke, F., Schwertfeger, R. (1985) Journal of Membrane Science, 22, 297–315. Jaffrin, M.J., Gupta, B.B., Ding, L.H., Garreau, M. (1984) TransactionsAmerican Society for Artiﬁcial. Organs, 30, 401–405.

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1, 42–48. Nakaji, S. and Yamamoto, T. (2002) Therapeutic Apheresis, 6, 267–270. Siami, G. A. and Siami, F. S. (2001) Therapeutic Apheresis, 5, 315–320. Gupta, B. B., Ding, L. H., Jaffrin, M. Y., Baurmeister, U. (1991) International Journal of Artiﬁcial Organs, 14, 56–60. SMOLIK, G. (1989) Ph.D. thesis. Technische Universita¨t Mu¨nchen (Mu¨nchen). Mulder, M. (1996) Basic Principles of Membrane Technology, Kluwer Academic, Dordrecht. Strathmann, H. (1979) Trennung von molekularen Mischungen mit Hilfe synthetischer Membranen, Steinkopf, Darmstadt.

38 Klingel, R., Fassbender, C., Fischer, I.,

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Hattenbach, L., Gu¨mbel, H., Pulido, J., Koch, F. (2002) Therapeutic Apheresis, 6, 271–281. Geiss, H. C., Parkhofer, K. G., Donner, M. G., Schwandt, P. (1999) Therapeutic Apheresis, 3, 199–202. Stegmayr, B. G. (2000) Blood Puriﬁcation, 18, 149–155. Stoffel, W., Borberg, H., Greve, V. (1981) The Lancet, 2, 1005–1007. Schneider, K. M. (1998) Kidney International, 53 (Suppl. 64), 61–65. Braun, N. and Bosch, T. (2000) Expert Opinion on Investigational Drugs, 9, 2017–2038. Hiatt, W., Vitzthum, G. H., Wagner, K. B., Gerlach, K. (1985) Materials Science of Synthetic Membranes, 9 American Chemical Society.

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4 Membranes in the Biopharmaceutical Industry Anthony Allegrezza, Todd Ireland, Willem Kools, Michael Phillips, Bala Raghunath, Randy Wilkins, Alex Xenopoulos 4.1 Introduction

Membranes for the biopharmaceutical industry operate in streams containing relatively low concentrations of exceedingly expensive products in a highly regulated industry. These considerations have determined the progress of membrane development for this industry and the manner of choosing and optimizing membrane usage. In this chapter we will describe the present state of microporous, ultraﬁltration (UF) and virus removal membranes, and the developments that led to them. Membrane use will be described from the point of view of a process engineer who has to specify and optimize the various membrane separation and puriﬁcation steps in a biopharmaceutical manufacturing process. From its inception, the biopharmaceutical industry has needed to concentrate and purify the protein drugs it produced. While the ﬁrst membranes used were those available at the time, manufacturers have become more sophisticated and demanding in their requirements for membranes as the industry has grown and matured. In addition, competition among membrane companies has increased. These trends have led to new membrane products speciﬁcally designed for the needs of the biopharmaceutical industry. The response of membrane manufacturers to industry needs for microporous, ultraﬁltration, and membranes for virus removal has gone beyond providing economical ﬂux rate and satisfactory retention properties and has focused on other needs. These include lowered extractables from the membrane device to the permeating stream and the capability to withstand alkalies. More recently, the need for increased capacity has resulted in newer asymmetric and composite membranes and multilayered combinations of membranes and preﬁlters in ﬁltration devices. All membrane ﬁltration processes are designed to remove some species and allow the passage of solvent and other, usually smaller, species. The membrane can remove large impurities and pass the product, or retain the product and pass solvent and salts. Every manufacturing process is different, but some common themes prevail. The

Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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4 Membranes in the Biopharmaceutical Industry

Fig. 4.1 Basic elements of a recombinant therapeutic manufacturing line.

process engineer seeking to develop and eventually optimize a process step must identify the proper membrane through a combination of experience (e.g., manufacturer and reference literature) and laboratory experimentation, and then use scaled-down testing to determine how the membrane will perform. Intermediate sized testing is then used to determine operating parameters, and cleaning and sanitizing regimes. Microﬁltration, ultraﬁltration and virus removal process development all follow these basics steps, with additional speciﬁc requirements for each. Figure 4.1 shows a generic biopharmaceutical manufacturing line and will be referred to throughout this chapter. The shaded boxes denote UF processes and the shaded circle, the virus removal ﬁltration step. All other ﬁltrations are done using microporous membranes. It is evident that microporous steps are numerous in the overall process. With the exception of crossﬂow microﬁltration sometimes used for clariﬁcation, all microporous use is with dead-end disposable cartridges or capsules. Microporous and UF membranes can be deﬁned by what they do, and how they do it. Microporous membranes are capable of removing particles larger than 0.1 mm and being used primarily in dead-end ﬁltration applications with disposable cartridges or capsules. Ultraﬁltration membranes are used for concentrating or diaﬁltering soluble macromolecules that have a size in solution of less than about 0.1 mm and operating continuously in a tangential ﬂow mode for extended periods of time, usually more than 4 hours and for up to 24 hours. We will see, however, that in the virus removal application, specialized UF membranes are used in a dead-end mode.

4.2 Microfiltration Membranes Used in the Biotech Industry

Microporous membranes clarify the streams from the fermenter and the centrifuge output. They are used in the process train for aseptic ﬁltration of media and other streams entering the fermenter and for buffers used in chromatography steps. Microporous membranes aseptically ﬁlter the ﬁnal ﬁlling of dosage containers, perhaps the most critical ﬁltration step in the entire process train. Ultraﬁltration membranes concentrate and diaﬁlter protein solutions. Diaﬁltration (DF) is used to change buffer solutions for subsequent process steps. Flat sheet membranes dominate this area, although hollow ﬁbers are used for some applications. Polyethersulfone (PES) and regenerated cellulose are the two major types of ﬂat sheet membranes. Hollow ﬁbers are primarily PES. Regenerated cellulose ﬁnds particular use where low protein binding of the expensive protein therapeutic drug product to the membrane is desired. For uses where harsh cleaning conditions or where protein binding is less important, PES membranes can be used. Virus removal in the production process for biopharmaceutical drugs require several separate robust steps with additive capabilities. Membrane processes have been proven to be an important part in virus removal and are now found in virtually all biopharmaceutical manufacturing processes. These membranes operate under perhaps the most severe requirements in membrane technology. Not only must they be capable of removing ‘‘3 logs’’ (99.9 %) of virus, but also they must pass almost the entire protein product in the feed stream. As the desire to remove smaller parvoviruses has developed, the difference in size between the protein and the virus has diminished, making meeting these almost contradictory requirements more difﬁcult. In addition, virus removal membranes must be validatable. The most prevalent virus removal membranes presently used are regenerated cellulose hollow ﬁbers, and ﬂat sheet composite polyvinylidene ﬂuoride (PVDF), PES asymmetric, or PVDF symmetric membranes.

4.2 Microﬁltration Membranes Used in the Biotech Industry 4.2.1 Introduction

In the early 1900s, Zigmondy [1] developed a process of making membranes that could be used to retain bacteria. A century later, similar processes are still being used to make the microﬁltration membranes that are used throughout a typical biopharmaceutical process. The more efﬁcient pleated cartridge devices introduced in the 1970s accelerated membrane development as stronger membranes were needed to withstand the pleating process. Pleated cartridges made the process and sterile microﬁltration of liquids more economical. When the ﬁrst biopharmaceutical products started moving from the lab to commercialization in the early 1980s, these membrane devices were available and eased process development. For decades they have served biopharmaceutical manufacturers with little change. A combination of increasing customer

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requirements has led to new microporous membranes and membrane structures. In this section, we will discuss how present-day microporous membranes evolved. This overview will attempt to rationalize why certain materials and speciﬁc morphologies are chosen for certain applications and how one might choose membranes for various applications. Although the original customer emphasis was on the performance of the membranes (ﬂux and rejection), these attributes are now expected. Current emphasis is focused on the robustness of the membrane device, and increased capacity. Robustness refers to the control of variability of a ﬁlter device so that the device operates successfully within the window of process variability. Capacity is the ability to run longer at a reasonable rate and not become plugged. Increasing capacity has been an empirical endeavor at best. Now, newer mathematical models more accurately describe the plugging behavior of microﬁltration membranes. Final ﬁltration applications require a high retention of Mycoplasma or the removal of bacteria from process liquids. Although the guidance of Food and Drug Administration (FDA) is clear, the interpretation of retention assurance is less clear. Some early and more state-of-the-art interpretations of retention assurance will be presented. 4.2.2 Microﬁltration Membranes: Development of Industrial Membranes

Ever since Zigmondy’s invention in 1914, membranes have been made by phase inversion processes. In the Zigmondy process, a thin layer of cellulose nitrate solution is evaporated and the resulting porous ﬁlm used to ﬁlter bacteria while passing other constituents of the feed. Zeman and Zydney [37] give an excellent description of processes currently used to manufacture microporous membranes. In most applications, microﬁltration membranes are deﬁned as those porous membranes that have the ability to retain bacteria but allow proteins to pass through. It was only in the 1960s that through the availability of scanning electron microscopy, it became clear that these ﬁlms consisted of highly interconnected pores uniformly distributed within the ﬁlm. In the 1970s, the need for new membrane materials for pleated cartridges led to polyamide membranes (nylon) from Pall Corporation and AMF (now Cuno/3M), and PVDF membranes from Millipore Corporation. These materials signiﬁcantly improved the mechanical properties of the membranes in part due to slightly lower porosities. However, their strong mechanical properties allowed the fabrication of robust pleated devices which more than made up for any loss in permeability. Introduced as the biotech industry started, these devices are still being used in the majority of the microﬁltration applications. Early microporous structures were web supported with woven or nonwoven fabrics to make up for the inherent weakness of the membrane. Web-supported membranes have the disadvantage that the web material is not compatible with certain posttreatments. For example, polypropylene (PP) is not resistant to gamma irradiation but has a high melting point, whereas polyethylene (PE) is not resistant to high steaming

4.2 Microfiltration Membranes Used in the Biotech Industry

temperatures but is resistant to gamma. PP is therefore often the material of choice for steamable cartridges while polyethylene can be used in gamma-treated cartridges. With the development of more robust membranes in the 1970s for pleated cartridges, a transition occurred from web-supported membranes to unsupported membranes, which are presently used in the majority of the microﬁltration applications. The ﬁrst microporous membranes were symmetric in structure, meaning that they had a uniform pore structure through their depth or thickness dimension. In the mid-1980s, the introduction of a new class of engineering plastics, ﬁrst polysulfone and then the more stable polyethersulfone resulted in asymmetric microporous membrane structures developed at Brunswick (eventually US Filter, now part of Pall Corporation) by Wrasidlo [2]. Further developments using polysulfone polymers came from Kraus [3] (assigned to Gelman, now Pall Corporation and commercialized as Supor1) and Roesink et al. [4] (assigned to X-Flow/Norit) who independently invented the use of hydrophobic– hydrophilic polymer blends to create membranes that were intrinsically hydrophilic. The Kraus patent, based on polyethersulfone, led to structures that are symmetric. AKZO AG [5] (currently Membrana) and Fuji [6] ﬁled patents on preparing asymmetric membranes based on a single-polymer solution that had a retentive portion of the membrane located internally in the membrane, sometimes called an hourglass symmetry from the proﬁle of pore sizes through the membrane cross-section. Despite being available for several years, asymmetric membranes were not successfully introduced into the biotech ﬁltration marketplace until the introduction in 1997 of the Sartopore 2 from Sartorius, a polyethersulfone-based membrane with a relatively low, hourglass-shaped asymmetry. In 2003, Millipore followed with highly asymmetric composite membranes having a superior permeability. In 2005, Pall Corporation introduced its Mach V technology. These membranes resemble in microstructure the original Supor technology as described by Kraus but show an asymmetric structure. Depending on the product, low or medium asymmetries are available. Besides Millipore, Pall, and Sartorius, other companies introduced products for this industry as well, that is, Membrana (DuraPES) and Cuno/3M (BioAssure). Although most recent products are based on polyethersulfone, the major market share in 2006 is still held by symmetric membranes: Durapore (polyvinylidene ﬂuoride) and Ultipor (Nylon 66). 4.2.3 Effect of Membrane Structure on Properties

Figure 4.2 below illustrates the different structures. These are scanning electron photomicrografts of the cross-sections of the three types of membranes. The cartoon below each is used, in some of the following ﬁgures, to help explain membrane properties. Asymmetric membranes have higher permeabilities for a given bubble point compared to symmetric membranes. Symmetric or isotropic membranes can be regarded as having a pore size that remains relatively constant throughout the thickness of the membrane. Thus, if the membrane thickness were to be divided into

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Fig. 4.2 Types of microporous membrane structures.

equal fractions, each would contribute equally to ﬂow resistance. For asymmetric membranes, each equally thick fraction has a different pore size and different ﬂow resistance. For asymmetric membranes, because the retention fraction is only a small part of the total thickness, higher permeabilities result. A new class of membranes, composite asymmetric membranes having two zones within the membrane with different pore size, show an increased ﬂow performance compared to singlepore-size symmetric membranes and to its asymmetric variants. The data shown in Figure 4.3 below illustrates how membrane structure affects properties. The data are typical of commercial membranes used in the biopharmaceutical industry. For each type, as the water bubble point (BP) increases, indicating smaller average pore size, permeability decreases. Bubble point is described in Section 3.1.2.2. Asymmetric membranes with what is considered ‘‘low’’ asymmetry,

Fig. 4.3 Water permeability relationships for three types of microporous membrane structures.

4.2 Microfiltration Membranes Used in the Biotech Industry Tab. 4.1

Commercial microporous membranes by structure and polymer.

Brand name

Vendor

Membrane material

Membrane structure

Ultipor Sartobran Durapore Supor Fluorodyne IIP Millipore Express Sartopore 2 Supor Mach V BioAssure

Pall Corp. Sartorius Corp. Millipore Corp. Pall Corp. Pall Corp. Millipore Corp. Sartorius Corp. Pall Corp. Cuno/3M

Nylon Cellulose acetate PVDF PES PVDF PES PES PES PES

Symmetric, web supported Symmetric, web supported Symmetric Symmetric Symmetric Composite asymmetric Low asymmetry, hour glass Low asymmetry Hour glass

where the ratio of pore size from one surface to the other is approximately 10, have higher permeability than symmetric membranes for equivalent BP. The asymmetric composites have the highest permeabilities. Higher permeabilities are desirable for reducing the footprint of installed membrane devices (less membrane area needed ¼ less capital cost) for applications where membrane pores are not plugged during ﬁltration. For applications where the feed stream plugs the membrane, capacity, that is, amount of liquid that can be ﬁltered, is more important, and permeability is less critical. Asymmetric membranes have signiﬁcantly higher capacity in plugging streams when used with the more open pores at the feed side. This is more pronounced for streams that do not have any preﬁltration. Given this advantage, it is therefore also not surprising that all major vendors of microﬁltration normal ﬂow ﬁltration (NFF) devices are currently introducing devices that have asymmetric membranes. Membranes from suppliers to the biopharmaceutical industry classiﬁed by membrane material and structure are given in the Table 4.1. 4.2.4 Aspects of Cartridge Design

For some steps in the manufacturing train, many commercial devices have a combination of membranes layered in the cartridge, that is, a microporous guarding layer and a sterilizing grade microporous layer (see Table 4.2). Especially in cases where the loading on the guarding layer is not extremely high, using a guarding layer instead of a separate preﬁltration step is highly desirable due to process compression and its associated reduction in footprint and hardware requirements. Pleated cartridges have one or more layers of nonwoven support layers in addition to the membrane. In pleated cartridges, cartridge output is the product of membrane ﬂux and membrane area. Membrane suppliers can make up for lower inherent membrane ﬂux by increasing membrane area in a cartridge, sometimes with very imaginative pleat designs. See, for example, US Patents 5,543,047 and 6,315,130. Pleated cartridge area (nominal 10 in. long by 3 in. diameter) ranges from 1 to 3 m2. How effectively the membrane is used in such designs depends on the packing

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Guarded microporous devices.

Brand name

Vendor

Pore size rating combination

Sartobran Multimedia Durapore Multilayer Durapore Supor EKV Fluorodyne EX Millipore Express SHC Sartopore 2

Sartorius Corp. Millipore Corp. Millipore Corp. Pall Corp. Pall Corp. Millipore Corp. Sartorius Corp.

0.45/0.22 0.2/0.22 0.45/0.22 0.65/0.22 0.2/0.22 0.5/0.22 0.45/0.22

density and the type of support materials used, and the nature of the feed being ﬁltered. Scaling studies in ﬁltration process development for a particular stream will quantitate device efﬁciency. The engineering scaling process (described later in Section 3) starts with membrane disks, moving to small area devices and ﬁnally large area manufacturing scale devices. The purpose is to determine inherent membrane properties at small scale and then to determine any loss in effectiveness for the conditions seen in the process being studied, due to cartridge packing or design. 4.2.5 Membrane Surface Modiﬁcation

In addition to the basic material and structure of the membrane, surface modiﬁcation approaches and chemistries differentiate membranes. The engineering plastics used to make membranes are hydrophobic and exhibit high protein binding (adsorption). Membrane manufacturers have devised many surface modiﬁcation chemistries to produce hydrophilic, low binding membranes. The patent literature contains a wide variety of chemistries that go beyond the production of hydrophilic surfaces to improved caustic stability, low binding capabilities for speciﬁc components and increased temperature resistance. A review [7] by Ulbricht covers much recent research. Caustic stability refers to the desire of biotech producers to clean their system with up to 1N alkali and not have ﬁlter properties changed by such cleaning. In ﬁnal ﬁll operations, low binding is required for the product being ﬁltered. For protein ﬁltration, an ultra low binding surface modiﬁcation is needed. For other active pharmaceutical ingredients (APIs), a low binding characteristic for those components is similarly useful. The surface-modiﬁed membrane must be stable toward different sterilization methods. Typical requirements for steam sterilization require temperatures signiﬁcantly higher than 121 8C. Many sterile ﬁlters have product claims specifying steaming temperatures of up to 136 8C. Stability toward steaming can be judged

4.2 Microfiltration Membranes Used in the Biotech Industry

by an integrity test (see Section 3.1.2.2) of these membrane devices. Such tests give biotech manufacturers proof that the sterility assurance is maintained after steaming and ﬁltration. Typically, this integrity test is performed through either diffusion testing or bubble point testing. Another method that is sometimes used is to steam sterilize the membrane devices, and then determine the ability to obtain a sterile ﬁltrate when challenged with a high bioburden load. Furthermore, steaming may change the surface chemistry resulting in a change in the measured bubble point. Bubble point is directly proportional to the surface energy of the test liquid, inversely proportional to the largest pore size in the most retentive zone in the membrane and proportional to the cosine of the contact angle of the liquid with the membrane surface. If steaming changes membrane contact angle sufﬁciently, interpretation of calculated pore size will be in error. During membrane development, the likelihood of steaming having a deleterious effect on wetting can be probed by measuring the critical surface energy of the membrane surface before and after one or multiple steaming cycles. The critical surface energy can be characterized by identifying the highest surface tension liquid that instantly wets the membrane. Having a membrane with a critical surface energy that is signiﬁcantly higher than 72 dyn/cm is desirable to guarantee that proper wetting occurs. Ideally, the critical surface energy is not changed by steaming or autoclaving. 4.2.6 Sterilizing Filters

Sterilization by use of 0.2 rated ﬁlters in pharmaceutical and biotech processes is more than just using a 0.2 rated membrane. A ﬁlter is qualiﬁed as sterilizing grade based on its ability to completely retain microorganisms under the speciﬁc conditions and with a particular process feed stream composition. Test methods are described in ASTM F-838-05 and HIMA tests. The FDA provides guidance in the Guideline on sterile drug products produced by aseptic processing [8]. Although the overall test conditions are clear, differences in some details still vary between different manufacturers. 4.2.6.1 Retention In the past, retention of sterilizing grade ﬁlters was expressed as the ability to retain more than 107 microorganisms of a certain size. Since retention occurs through size exclusion, good correlations were found between bubble point, a measure of the largest pore size of the membranes, and retention. Results were expressed in log reduction value (LRV) units. LRV ¼ log10

colony forming units found in the filtrate : colony forming units challenged to the filter

Although the LRV gives a good indication what constitutes retention, it fails to describe of what sterility means.

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The authors expect that with the implementation of the FDA’s new initiative [9]: ‘‘Pharmaceutical cGMP’s for the twenty-ﬁrst century: a risk based approach,’’ membrane science in the near future will require a high level of understanding of where certain risks originate. For sterilizing grade membranes, manufacturers will need to better deﬁne how sterility assurance can be deﬁned in a scientiﬁcally sound way. Recently, a paper [10] was presented by researchers at Millipore Corporation in which a statistical-based model was described by which sterility assurance can be calculated. The researchers based the model using fundamental principles and assumptions based on experimental studies of the retention of bacteria with microporous membranes. Rather than focusing on achieving a particular log reduction value, this approach relied on calculating the assurance to obtain complete retention downstream of a membrane ﬁlter as a function of bubble point. The method chosen as most appropriate for the data was regression of logarithm of ﬁltrate colony forming units (CFU) versus bubble point, using retentive outcomes as censored (i.e., partial) information. A linear model of isopropyl alcohol (IPA) bubble point versus retention was ﬁtted. From it, the probability of sterile ﬁltrate was obtained by solving Probability (retention) ¼ exp (10(LC)), where LC is the expected log count at a given bubble point. This assumes that the probability of zero counts (i.e., retention) at a ﬁxed bubble point follows a Poisson distribution with parameter 10^(LC) when testing is done under uniform conditions. Retention conﬁdence was obtained from the regression relationship between bubble point and LRV (Figure 4.4). From this relationship, the minimum bubble

Fig. 4.4 Logarithm of colony forming units found on the downstream side of the membrane as a function of alcohol bubble point for a symmetric and asymmetric membrane. The lowest bubble points with complete retention are indicated with an upward arrow. The calculated minimum speciﬁcations for bubble point with a 99.9 % assurance level that sterile efﬂuent is obtained is indicated with a downward arrow.

4.2 Microfiltration Membranes Used in the Biotech Industry

point of the membrane device is calculated for a conﬁdence level of greater than 99.9 % that a sterile efﬂuent is obtained when challenged with B. revundimonas diminuta at >107 organisms per cm2.

4.2.6.2 Permeability Membrane manufacturers are constantly trying to attain higher ﬂux and capacity while maintaining required sterility assurance. One illustration of how membrane structure affects the balance between properties is shown in Figure 4.5. In this ﬁgure, the bubble point corresponding to a sterile ﬁltrate in B. diminuta testing is given for the three types of membranes. To get the same retention assurance for an asymmetric membrane, a minimum bubble point is deﬁned, which is signiﬁcantly higher than that for the symmetric membrane such as are those currently being used commercially. Despite the higher bubble point, which generally reﬂects a smaller pore size, a signiﬁcantly higher ﬂux is achieved for the asymmetric membranes as shown in Figure 4.5. Besides relying on the minimum bubble point for the sterilizing grade membranes established by the membrane manufacturer, users of these membranes who sterile ﬁlter liquids have to demonstrate that acceptable sterility is obtained in their process ﬂuid. This is described in detail in Section 3. It is known that certain organisms could potentially penetrate through 0.2 sterilizing grade membranes, that is, Mycoplasma. For those cases, a smaller pore size membrane ﬁltration is recommended. Due to the smaller pore size, lower permeabilities and capacities are typically obtained and a larger installed footprint is needed to ﬁlter the same volume of process liquid.

Fig. 4.5 Bubble point corresponding to obtaining sterile ﬁltrate with B. diminuta for different membrane structures.

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4.2.6.3 Capacity In many steps throughout the process, the membrane retains particulate matter and other species. How these particles are retained whether on or within the membrane structure will deﬁne the ‘‘dirt holding capacity.’’ Common practice is to use a standard pore-blockage model or a caking model to predict, based on smallscale testing, what the total capacity is for candidate membranes. This will be described in Section Membrane System Sizing – Capacity. Currently, academic and industrial researchers are studying capacity with more sophisticated models. Ho and Zydney [11] published a paper describing a transition between a standard pore-blockage model and a caking regime. Bolton et al. [12] combined a cake model with a complete pore-blockage model in a single equation to capture this transition. Knowledge of what fraction of the volume ﬁltered is governed by complete blockage versus caking can be used to optimize membranes for its capacity. Further research is underway in an attempt to describe capacity as a function of asymmetry and as a function of preﬁlter layering.

4.3 Practical Membrane Considerations for Sterile Filtration by Microporous Membranes

Biopharmaceutical manufacturing is a membrane intensive industry and in particular is a heavy user of microporous membrane devices. As illustrated in Figure 4.1 and as discussed in this section, microporous devices are primarily used for sterile ﬁltration for input streams and for aseptic vial or other container ﬁlling at the end of the manufacturing process. In the following brief tour, we will start at the fermenter and show where microporous membrane devices are placed in the various manufacturing steps. The fermenter requires multiple ﬁltration devices. Media and associated buffers, which comprise the nutrients for the cells being grown, are sterile ﬁltered and fed into a bioreactor. Hydrophobic sterilization membranes ﬁlter any gas streams entering the fermenter, and are on any vents during processing. Since they will not wet from liquid splashing during the process, there is no opportunity for bacteria to grow through. Except for microﬁltration used for crossﬂow or tangential ﬂow ﬁltration (TFF) clariﬁcation of the cell culture, all microﬁltration unit operations focus on minimizing the risk for bacterial or Mycoplasma contamination of the product. Biopharmaceutical manufacturers select media to maximize the expression levels in the cells. In recent years, there has been a shift from animal derived, serum-based media to plant derived media (i.e., soy digests) due to regulatory concerns with Kreuzfeld–Jacob syndrome and other potential contaminants. A shift to chemically deﬁned media would be even more desirable but expression levels using these media have not been optimized. As this trend continues, membrane ﬁltration of media will become easier and more economical. To maintain sterility in the bioreactor, a pharmaceutical grade 0.2 or 0.1mm rated membrane is used. For media where a high reduction in Mycoplasma is desirable, different 0.1 mm rated devices are recommended. Many media, especially serumbased media, have a plugging character, and in order to obtain process economies, microporous preﬁlter devices are commonly used prior to sterile ﬁlter devices.

4.3 Practical Membrane Considerations for Sterile Filtration by Microporous Membranes

Depending on the choice of expression system, proteins are expressed internally or externally to the cell. For internally expressed proteins, the cell culture is harvested when a predeﬁned yield is achieved. The cells are usually concentrated by a centrifugation step, although crossﬂow microporous membrane devices are being used successfully as well. Once the cells are concentrated, the cells are lysed, most typically in a high-pressure homogenizer. The cell lysate containing the product of interest is then clariﬁed to remove cell debris using depth ﬁlters and ﬁnally sterile ﬁltered into a holding tank. Cell cultures which express protein internally, such as E. coli, are optimized to yield high titer levels of expressed protein, which is typically associated with a low viability of cells. Such cells are more easily ruptured during harvest. As a consequence, there is a high loading of cell debris that needs to be clariﬁed. A second clariﬁcation after the centrifugation or microporous TFF step is performed through depth ﬁltration. These are usually lenticular pad ﬁlters of bound cellulose ﬁbers and ﬁlter aids, such as diatomaceous earth. The number and size of colloids and other particles exiting the depth ﬁlter depends on the attributes of the particular depth ﬁlter being used and the feed stream. If this permeate stream is not clean enough, the downstream sterile ﬁlter will become quickly plugged. An intermediate ﬁlter may be needed. A series of chromatography columns, that is, ion exchange, protein A, or other afﬁnity media, puriﬁes the clariﬁed product. Buffers used in chromatography are usually sterile ﬁltered to remove bioburden and other contaminants in order to increase column life and economies. For protein that is expressed external to the cells (i.e., Chinese Hamster Ovary cell lines, NS0 cell lines), the cells are removed by centrifugation, TFF or depth ﬁltration. The separated product is then processed as above. Figure 4.1 shows that large volumes of buffers are used throughout the process, constituting a considerable usage of membrane ﬁlters. Since these buffers are clean, membrane devices used to prolong column life are chosen primarily for high permeability, to reduce the number of cartridges needed. It is expected that any commercial sterilizing (0.2mm rated) membrane will have the retention needed. Filtration of protein pools with high levels of protein can lead to undesirable protein aggregation. In such cases, microﬁltration membranes are used to remove aggregates before the expensive virus removal ﬁltration. This is described in more detail in Section 6. Once the proteins are concentrated to their ﬁnal level, a ﬁnal sterile ﬁltration step can be used to sterilize the liquids for vial or container ﬁlling. In order not to affect the dosage level, a low protein binding characteristic of the membrane is required. In most ﬁltrations at this stage, permeability and capacity are less important than the conﬁdence that the membrane device can be validated. 4.3.1 Sterile Filtration Process Considerations

Properly done sterile ﬁltration removes all living organisms from a ﬂuid. It is a critical application of membranes to the biopharmaceutical manufacturing process. To

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implement a sterile ﬁltration process, the manufacturer must take great care in material selection, system design, and scale-up to assure that the ﬁltered product is suitable for use. In this section, we will present a standardized methodology to assist in this process. We will cover membrane ﬁlter selection criteria including materials, retention requirements and testing, ﬂow and capacity performance testing, installation, sterilization, integrity testing, and process validation. 4.3.1.1 Filter Selection Generally, ﬁlter selection has four main considerations: compatibility; retention;

and two considerations which determine system sizing ﬂow rate; capacity. A logical approach to ﬁlter selection will consider these four criteria in the order listed. For instance, there is no point in determining if a given ﬁlter has sufﬁcient ﬂow for a given application if the ﬁlter materials are not compatible with the feed stream. Compatibility Compatibility covers how membranes and membrane devices withstand the conditions under which they are operated. Chemical compatibility refers to solvation, swelling, or weakening of the membrane due to contact with the process ﬂuids. Typical modern membranes used for sterilizing ﬁltration include PVDF, PES, polytetraﬂuoroethylene (PTFE used for venting), and cellulosics. PTFE has the broadest chemical compatibility and cellulose acetate is the most limited. Whatever the material, chemical compatibility guides can be used to make an initial assessment of suitability. Guides can be found in manufacturers’ literature or chemical engineering manuals. Chemical compatibility can also be determined experimentally. Using membrane coupons, key membrane properties such as water ﬂux, weight, thickness, and integrity can be measured before and after exposure to the product stream or individual components of the feedstream. Changes in any of these key properties indicate chemical incompatibility that could impact membrane performance. Chemical compatibility also includes other feed-stream membrane interactions such as adsorption of critical feed-stream components. Examples are proteins, preservatives, and APIs. Here again, initial determination can be done by soak tests with membrane disks and quantization done by spectroscopic or other analytical methods. Pitt [13] provides some guidance and test methodologies for protein adsorption. Hem [14] provides some guidance for preservative adsorption. In most cases, however, the only true measure is with the actual formulation processed with actual process parameters. One important reason for this is to be able to understand the effect of nonmembrane components of the ﬁltration device on adsorption and ﬁltration performance.

4.3 Practical Membrane Considerations for Sterile Filtration by Microporous Membranes

Heat resistance of sterilizing grade membranes is another key compatibility question. Live steam is the most common method for sterilizing membranes prior to sterile ﬁltration. Most typical sterilizing membranes have very good thermal compatibility. Typically, the limiting factor is the support materials used in cartridge and capsule devices. Manufacturers’ recommendations should be followed. Gamma irradiation is also typical for ﬁlters that are supplied sterile. PVDF and PES are compatible with the levels of gamma irradiation required for sterilization, typically 25 kGray. PTFE and cellulosics are not gamma compatible. AAMI provides guidelines for qualifying and requalifying gamma sterilization processes [15]. A ﬁnal consideration for materials compatibility is whether the membrane is hydrophobic or hydrophilic. Hydrophilic membranes wet spontaneously with water, and hydrophobic membranes repel water. Generally, hydrophilic membranes are used for aqueous streams, hydrophobic membranes for vent applications or gas ﬁltration. Hydrophobic membranes can also be used to ﬁlter low surface tension ﬂuids such as alcohols, but hydrophilic membranes are suitable and often used to minimize the number of different ﬁlters that must be stocked by an end user. Retention Retention requirements for sterilizing membranes are quite simple – complete retention of microorganisms in the feed stream and complete passage of the product components. Because this separation is primarily based on size, and there is a generally a large difference in size between microorganisms and most proteins or other APIs, membranes with pore size ratings of 0.2 mm or 0.1 mm will provide the necessary bacterial retention and product/passage requirements. When selecting a 0.2 or 0.1mm membrane, it is important to assure there is a certiﬁed sterilizing grade claim. The general deﬁnition of a sterilizing grade ﬁlter is a ﬁlter that when challenged with at least 107 cm2B. diminuta (ATCC19146) will produce a sterile efﬂuent. Standard methods are available (ASTM F-838-05) that detail methods for B. diminuta retention testing. The ASTM method or close derivatives are used by ﬁlter manufacturers for product development and quality control tests. Some manufacturers use the test for membranes and some for membranes and devices. The ASTM test is designed to challenge each pore of the membrane with a single bacterial particle. B. diminuta has long been accepted as a suitable ‘‘worst case’’ organism for this test [8,16]. It is suitable due to the small size, ability to grow monodispersed and at high concentration. It is also desirable for its relative safety. Although the ASTM test is considered a suitable QC release test for membrane manufacturers, it is not sufﬁcient for demonstrating the ﬁlter will provide suitable retention in a speciﬁc drug manufacturing process. To qualify a ﬁlter’s retention in actual production, a modiﬁcation of the ASTM test must be used. The modiﬁed test must take into account process operating parameters, which could impact bacterial retention such as process time, differential pressure, and volume/area. In addition, the modiﬁed test must take into account actual process bioburden and any impact the process ﬂuid may have on organism size. The result is a ‘‘custom’’ retention test performed as a component of the validation process.

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The components of the custom retention test are as follows: Identify native bioburden and compare it to B. diminuta to determine worst case challenge organism. (Over 99 % of existing custom retention validations have been performed with B. diminuta.) Determine if actual drug solution will inhibit microbial growth. This would be the case with antibiotics or preservative containing solutions. A noninhibitory modiﬁed product must then be used. Spiked product or noninhibitory modiﬁed product with test organism. Challenge membrane under process scale down conditions, keeping volume/area, time, and pressure consistent with fullscale process. Acceptance criteria for the test are a challenge level of at least 107 organisms per cm2 membrane area and 0 cfu downstream. The results of this test are included with the process validation documentation submitted to regulatory agencies. Results from the custom retention test, along with documentation that membrane manufacturing processes are validated, and integrity test processes are correlated to retention, are used to demonstrate that a speciﬁc ﬁltration process is sterilizing grade. Membrane System Sizing – Flow Rate Requirements Compatibility and retention deﬁne the type of ﬁlter to be used for a given application. The next step is determining the necessary amount of ﬁlter area. Filter area requirements are driven by two process parameters – ﬂow rate and volume capacity. Examples of process constraints that deﬁne ﬂow rate requirements are as follows: The need to ﬁlter a speciﬁc volume in a speciﬁc period of time (e.g., one shift) or other processing time constraints. Tankto-tank transfers, such as buffer ﬁltration processes, are typically time based. Minimum ﬂow constraints, where the ﬂow through the ﬁlter must be maintained at a certain level in order to maintain adequate process ﬂuid to the next step. Filling operations are an example where low ﬂow means low ﬁll volume and likely process shutdown.

The driving force for ﬂow is a critical component of ﬂow rate determination. Driving pressure can come from either positive pressure applied with gas pressure or a pump. Or driving pressure can come from vacuum as is common in tank vent applications. When the ﬂow requirements and driving force are deﬁned, the ﬁrst estimate of ﬁlter size can be determined using the ﬂow/DP (DP is the pressure drop across the ﬁlter system) information provided by membrane manufacturers for the various

4.3 Practical Membrane Considerations for Sterile Filtration by Microporous Membranes

devices. These curves are given for standard ﬂuids, usually water for hydrophilic ﬁlters and air for hydrophobic. Information for water ﬂow is usually given for 25 8C and 1 cps. Correction tables are available for water at temperature other than 25 8C. A ﬂow/DP curve, adjusted for feed stream viscosity, will provide a reasonable estimate of membrane behavior at the start of the ﬁltration. However, as the membrane plugs or fouls, the ﬂow/DP properties will change. The amount of change varies with feedstream. In some cases, such as gases or puriﬁed water, the ﬂow/DP properties change very little and therefore the manufacturers’ ﬂow/DP curve will not only give an estimate of how the membrane performs at the start of the ﬁltration but also at the end. Sizing ﬁlters for these applications is simply a matter of understanding the ﬂow/DP curve. Membrane System Sizing – Capacity However, there are many feedstreams that have enough retained contaminants to signiﬁcantly change the ﬂow/DP properties of the membrane. Examples include proteins, cell culture media, cell harvest, and so on. In these cases, trials are generally required to determine how the ﬂow/DP conditions will change through the course of the ﬁltration. Capacity trials can be operated in one of two modes. Abbreviated trials run at constant pressure, Vmax trials, can be used when gradual pore plugging is the operative mode. A major advantage of Vmax trials is the ability to screen multiple membrane samples in a short time with minimum volume of feedstock. Constant ﬂow trials can be used when gradual pore plugging is not operative or for simulating processes run at constant ﬂow.

Vmax is a method for predicting the throughput of ﬁlters based on a gradual pore plugging model [17]. Gradual pore plugging occurs when colloids or suspended matter collect on the sides of ﬁlter pores to gradually block them off, until a state of total occlusion is eventually reached. This gradual blocking of the pores results in a distinct ﬁltration pattern. In a Vmax test, the time and volume collected up to that time are recorded at regular intervals. Test apparatus is shown in Figure 4.6. A plot of time/volume versus time is made on linear/linear axes. If the results plot as a straight line, the ﬁlter is considered to plug by the gradual pore plugging model and the formulas of Vmax can be applied to predict ﬁlter life. If the results do not plot as a straight line, it indicates the ﬁlter is plugging by some other model, such as cake formation. In these cases, Vmax should not be used. A traditional ﬂow decay method, in which the ﬁlter is actually run with the feed stream or close surrogate until completely plugged, should be used. In the gradual pore plugging model, ﬂow decay follows the equation Q ¼ Qi ð1 kV=2Þ2 ;

ð1Þ

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Fig. 4.6 Vmax test apparatus.

where Q is the ﬂow rate (L/min), Qi is the initial ﬂow rate (L/min), k is the constant, and V is the volume ﬁltered (L). At inﬁnite time, Q ¼ 0, and Equation (2) results Maximum volume filtered ¼ Vmax ¼ ðk=2Þ 1:

ð2Þ

When time/volume versus time is plotted and a straight line results, it follows the equation t=v ¼ At þ B;

where t is the time, v is the volume ﬁltered up to that time, A is the slope, and B is the y-intercept. Solving for V gives the following equation: V ¼ t=At þ B ¼ 1=ðA þ ðB=tÞÞ:

Letting time go to inﬁnity, B/t becomes zero and V ¼ Vmax ¼ 1=A:

Therefore, the inverse of the slope of the plot of t/v versus t equals the maximum volume that can be ﬁltered by the test ﬁlter. B is the y-intercept, or the value of t/v when time is zero. 1/B has the dimensions volume per time, a ﬂow rate. Therefore, the inverse of the y-intercept is the ﬂow rate through the test ﬁlter at zero time or the initial ﬂow rate. Rearranging Equation (1) with Vmax = 2/k gives Q=Q i ¼ ð1 V=Vmax Þ2 :

This equation can be used to determine the percentage of Vmax that has been used when the ﬂow rate has dropped to a percentage of the initial ﬂow rate. The process development engineer will use Vmax to test and compare multiple membrane options in a short time with small volumes of feed solution. Vmax results

4.3 Practical Membrane Considerations for Sterile Filtration by Microporous Membranes

can be used to make initial choices with a small investment in time and product solution. Since the trial does not simulate all the process parameters, such as process ﬂux, time, differential pressure, and volume/area, it is limited as to scaling accuracy. For scaling purposes, it is highly recommended that process simulation trials be run before ﬁnalizing the system size. The trial should simulate as closely as possible the actual process mode (constant pressure or constant ﬂow), actual process time, volume per area, and membrane ﬂux. Process simulation trials can be performed with membrane coupons or small-scale capsules when sufﬁcient quantities of feed stock are available. An alternative to Vmax testing is the constant ﬂow Pmax test. Pmax is conducted with a positive displacement pump feeding the test ﬁlter at constant ﬂow. The differential pressure is recorded and trended as a function of the volume ﬁltered. The advantage of Pmax tests is that they often more closely simulate the actual process. This is especially true when process ﬂux and maximum process operating pressure is adhered to. In addition, unlike Vmax, the test is run until the ﬁlter is plugged. The plugging volume is not predicted by extrapolation. Another advantage of Pmax test is it is not dependent on any speciﬁc plugging mechanism. The disadvantage of Pmax tests is it could potentially take a long time and large volume to run, as in cases where ﬁlter fouling is very low. Pmax test apparatus is shown in Figure 4.7.

Fig. 4.7 Pmax test apparatus.

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Fig. 4.8 A typical pleated cartridge ﬁlter showing the pleat components. Cartridges are inserted into permanently installed housings.

4.3.1.2 Device Selection There are two major categories of devices used in NFF – cartridges (Fig. 4.8) and capsules (Fig. 4.9). Cartridges supplied by membrane manufacturers are inserted into housings with an o-ring seal at the process site. Capsule ﬁlters include a plastic housing with the membrane permanently bonded inside. Cartridges generally will have lower cost per membrane area. Capsules can provide lower operating costs by minimizing labor requirements. Sterilization Considerations When deciding between cartridges or capsules, the ﬁrst key consideration is sterilization method. If live steam is to be used, a cartridge in stainless steel housing will provide the highest degree of safety. Steam in place (SIP) for capsules is not recommended for the vast majority of typical capsules. Capsules with SIP claims require extensive safety precautions and must be operated with great care. On the other hand, autoclave sterilization is very common and suitable for both cartridge and capsule type devices. Typically temperatures from 121 to 126 8C are used. Gamma irradiation is the sterilization method of choice for ‘‘pre-sterilized’’ devices because it is highly lethal to all organisms. There is also the advantage of being able to sterilize large numbers of devices in their ﬁnal packaging, eliminating issues of package exposure to steam or moisture.

4.3 Practical Membrane Considerations for Sterile Filtration by Microporous Membranes

Fig. 4.9 An all in one capsule ﬁlter. The pleated membrane/ support is sealed into the capsule and the entire device is disposed of after use.

All sterilization methods must be qualiﬁed for effectiveness and membrane compatibility. AAMI [11] provides guidelines for gamma irradiation dose qualiﬁcation and requaliﬁcation intervals. Implementation Successful implementation of the chosen membrane, device and sterilization method includes installation, sterilization, integrity testing, and running the actual ﬁltration. An initial consideration for installation is to assure that there is sufﬁcient space above and around the ﬁlter installation to raise the housing dome, attach and detach tubing and piping to inlet and outlet. Before ﬁnalizing system installation, the process development engineer must assure that the process scale system will perform as predicted from small-scale testing. Typical mistakes include sizing based on membrane pressure drop, then encountering ﬂow limiting piping restrictions at full scale. For example, a common mistake is to size a process based on pressure at source, with the membrane device

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being operated at a signiﬁcant distance or head difference from the source, resulting in less driving force than design. The choice of sterilization method will have a signiﬁcant impact on implementation. Easiest to use are disposable ﬁlter assemblies supplied pre-sterilized. These can be obtained complete with whatever tubing, valves, containers, and so on, are required for the application. All that is required is attachment to other processing equipment. Next easiest to implement are autoclaved assemblies of the membrane device and associated tubing, and so on. These are assembled by the user on-site, wrapped in protective material, and autoclave sterilized. Ideally, the assembly will be as complete as possible prior to autoclaving in order minimize the number of aseptic connections required post-sterilization. Most difﬁcult to operate are steam-in-place operations. First, there are safety considerations as operators must handle live steam lines. Second, the steam inlet pressure and membrane differential pressure must be carefully controlled to avoid damaging the membrane or device. Finally, as discussed in the microporous membrane section, the membrane must be able to withstand steaming without a change of properties. Integrity testing is performed to prove the membrane is free of oversized pores that would compromise retention. Most common are nondestructive tests based on capillary theory. The theory and practice of integrity testing have been covered previously many times [18]. Nondestructive integrity tests are based on capillary forces as deﬁned in the bubble point equation BP ¼ 4 kgcosu=d;

where BP is the bubble point, k is the shape correction factor, g is the surface tension, u is the contact angle, and d is the pore diameter. This shows the relationship between pore size and the pressure required to force liquid out of membrane pores to initiate bulk gas ﬂow. This is the fundamental basis for all nondestructive membrane integrity tests. ‘‘Bubble Point’’ tests measure the pressure required to initiate bulk gas ﬂow. ‘‘Diffusion’’ tests measure the gas ﬂow as a pressure below the expected bubble point pressure. Excessive gas ﬂow at a pressure below the expected bubble point indicates the presence of oversized pores or defects. Successful integrity testing is dependent on completely wetting the membrane pores with test ﬂuid. The most common test ﬂuid for hydrophilic ﬁlters is water. But because water has a relatively high surface tension, it can be difﬁcult to get water to fully penetrate all the pores. False integrity test failures due to poor wetting are common and time consuming. As discussed above, a common problem with hydrophilic modiﬁcations is they can be removed or reversed by process exposure. Live steam will reduce the surface energy of many modiﬁed polymers. In some cases, hydrophilic membranes are rendered hydrophobic by steam sterilization. The result is limited ﬂow or, more commonly, false integrity test failure. Manufacturers’ recommendations for steam sterilization should be followed to minimize this problem. Before ﬁnal membrane is

4.4 Ultrafiltration and Virus Filtration Membranes for Biopharmaceutical Applications

speciﬁed into the process, trials should be run that simulate all process conditions, including integrity testing and sterilization. The information presented in this section is common to all sterile ﬁltration applications in the biotech industry. A process engineer who is designing a speciﬁed one or more sterile ﬁltration steps will ﬁnd that following the considerations given here and cartridge manufacturers’ recommendations will result in a well functioning process.

4.4 Ultraﬁltration and Virus Filtration Membranes for Biopharmaceutical Applications

Filtration membranes are generally classiﬁed based on their size exclusion properties. At the largest pore size, we have microﬁltration membranes (typical pore sizes in the range of 0.05–10mm). Ultraﬁltration and virus ﬁltration membranes, the subject of this section, have pore sizes approximately in the range of 1–100 nm, corresponding to solute molecular weights in the 1000–1 000 000 Da. Although there is overlap in the pore size range of ultraﬁltration and virus ﬁltration membrane groups, their application uses are very different. As a result, different membrane types have emerged to dominate each application. For ultraﬁltration, regenerated cellulose and PES membranes are the two main types. Flat sheet membranes dominate the market, but PES hollow ﬁber devices are also available. For virus ﬁltration, PVDF membranes are used, in addition to PES and regenerated cellulose. Both hollow ﬁber (regenerated cellulose) and ﬂat sheet devices are available. We will cover ultraﬁltration and virus ﬁltration membranes separately. 4.4.1 Ultraﬁltration Membranes

UF is widely employed in biopharmaceutical manufacturing to concentrate and diaﬁlter biological molecules, generally proteins [19]. Biopharmaceutical manufacturers use these process steps to attain a speciﬁed product concentration range and buffer composition that is optimal for subsequent processing such as chromatography or formulation. 4.4.1.1 Membrane Suppliers The principal suppliers of UF membranes for biopharmaceutical applications today are GE Healthcare [20], Millipore [21], Pall [22], and Sartorius [23]. The ﬁrst UF membranes (nonwoven supported regenerated cellulose – YM series) were introduced commercially by Amicon – now part of Millipore Corporation – in the 1970s. Millipore introduced nonwoven based PES membranes soon thereafter. Filtron – now part of Pall Corporation – followed in the 1980s with Omega membranes (also nonwoven based PES). Alpha PES membranes are also available from Pall for applications where antifoam agents are a problem. The 1990s saw the introduction of several new membranes. Millipore started offering cellulose (PL-series) and Biomax (PES) membranes, a response to YM and Omega membranes, respectively.

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4 Membranes in the Biopharmaceutical Industry Manufacturers and brand names of ultrafiltration membranes for biopharmaceutical manufacturing.

Tab. 4.3

Supplier GE Healthcare GE Healthcare Millipore Millipore Pall Pall Sartorius Sartorius

Cellulose

Ultracel PLC (5–1000 kDa) Ultracel PL and YM (1–300 kDa) Regen (10–50 kDa) Hydrosart (10–100 kDa) Cellulose triacetate (5–20 kDa)

PES A/G HF (1–750 kDa) Kvick (5–100 kDa) Biomax (5–1000 kDa) PT (10–300 kDa) Omega (1–1000 kDa) Alpha (10–30 kDa) PESU (5–700 kDa) PESU (5–700 kDa)

The novelty of Biomax membranes was that they were the ﬁrst macrovoid-free membrane structures. Sartorius introduced cellulose Sartocon Hydrosart as a response to YM and PL, with the additional claim of sodium hydroxide cleanability. Sartorius also offers PES and cellulose triacetate membranes. Millipore also introduced cellulose composite membranes, replacing the traditional nonwoven support with a microporous membrane. The ﬁrst hollow ﬁber devices were offered by A/G Technology – now part of GE Healthcare – are made of PES and also claimed a macrovoid-free structure. The most recent offering is Kvick PES ﬂat-sheet membranes by GE Healthcare Biosciences AB. Table 4.3 summarizes the key suppliers and membranes, along with the NMWCO offered. 4.4.1.2 Membrane Selection As is obvious from the above survey, PES and regenerated cellulose are the two dominant types of ultraﬁltration membranes. Regenerated cellulose is often selected as a result of its low protein binding properties. This is especially important for expensive protein therapeutic drugs. Low protein binding reduces fouling and beneﬁts a process with improved consistency, easier cleaning, and improved yield. PES membranes are used in applications where harsher cleaning chemicals are used or where low protein binding is less important. UF membranes are typically reused several times, and for that reason cleanability and solvent compatibility is a key consideration in membrane selection. Separate, but related, considerations apply for membrane storage, as the devices need to be stored in appropriate preservative solutions, in many cases, alkali, between uses.

Membrane selection for a given application is based on the following factors: retention of the product protein; retention consistency; process ﬂux and overall process economics; scalability; mechanical robustness; chemical compatibility for cleaning/storage.

4.4 Ultrafiltration and Virus Filtration Membranes for Biopharmaceutical Applications

Selection of a particular pore size is usually done empirically and should always be conﬁrmed in the user’s particular application. A rule of thumb for selecting membrane nominal molecular weight cut-off is to select one third to one ﬁfth of the product MW. For example, membranes rated at 30–50kDa should be used for retention of a monoclonal antibody that has an MW around 150kDa. Even tighter membranes should be used if anything above 0.01 % passage of the product is unacceptable. 4.4.1.3 Membrane Structures Flat sheet ultraﬁltration membranes are typically used, cleaned, and sanitized, sometimes stored, through many cycles. To obtain the required mechanical stability and robustness for multiple uses, most UF membranes are made with a nonwoven fabric support. Nonwoven polyoloeﬁn supports have been used for a long time, as they offer robustness and simplicity, as well as an open, permeable substructure. Composite membranes (e.g., Millipore Ultracel PLC) are instead cast on microporous polyethylene membrane supports. The advantages are reduced microdefects that could be caused by the rough surface of nonwovens and more uniformity and ﬂatness of the UF layer, resulting in high retention consistency. In addition, there is improved adhesion of the UF layer to the substrate, resulting in a more mechanically robust membrane. The original UF membranes had large macrovoids. More recent membranes claim a macrovoid-free structure. Although macrovoid-containing membranes have been successfully used for many years, macrovoid-free structures provide higher assurance of retention and improved consistency. They have become the norm in most current applications. Figure 4.10 shows examples of typical membrane structures. 4.4.1.4 Characterization Ultraﬁltration membranes are characterized in terms of a nominal molecular weight cut-off rating, NMWCO. The rating corresponds to the retention to a certain level

Fig. 4.10 SEM images of typical ultraﬁltration membranes: macrovoids on nonwoven substrate (left), macrovoid-free on nonwoven substrate (middle), composite, that is, macrovoid-free on microporous substrate (right).

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(typically 90 %) of a marker molecule. Each membrane manufacturer has their own characterization methods, and they should not be directly compared without knowledge of the methodologies used. Originally, manufacturers characterized all UF membranes with a protein retention test. Each membrane was challenged with one or more protein molecules, and speciﬁcations were set at a minimum required retention. Given the variability in purity of commercially available proteins, the lack of relevance to speciﬁc protein formulations used in the industry and the low robustness of a ‘‘minimum speciﬁcation’’ approach, alternative methods were introduced. A mixed dextran solution with molecules covering a broad MW range from 1 to 1 000 000 kDa was proposed and has been adopted by the UF community [24]. Normal Molecular Weight Cutoff (NMWCO) is typically deﬁned as the dextran MW where 90 % of dextran molecules are retained, a value known as R90. The advantage of such tests is that they provide a common platform for testing a whole family of UF membranes, which would otherwise require testing with different proteins. In addition to retention properties, water permeability, back pressure (to test integrity of the membrane–support interface) and air diffusion (to test for the presence of large defects) are also usually measured and reported for UF membranes. 4.4.1.5 Devices Essentially all ultraﬁltration membranes for bioprocessing are used in TFF devices, with the exception of some virus removal devices. In TFF, also referred to as ‘‘crossﬂow’’ ﬁltration, ﬂuid ﬂows across (tangentially) the ﬁlter membrane surface, with a small fraction of the ﬂow permeating the membrane (perpendicular to the surface). The sweeping action of the tangential ﬂow clears the surface of deposits thereby minimizing membrane fouling and maximizing process ﬂux and product yield. The common device (module) formats used for biopharmaceutical applications are ﬂat sheet membranes used in plate and frame-type cassettes or spiral wound module formats, and hollow ﬁber modules (Figure 4.11). Cassettes can have an open channel or a turbulence promoting screen. Screened cassettes have become the dominant module format in the biopharmaceutical

Fig. 4.11 Commercial ultraﬁltration modules: hollow ﬁbers (left), spiral-wound (middle), cassettes (right).

4.4 Ultrafiltration and Virus Filtration Membranes for Biopharmaceutical Applications

industry, as they offer high mass transfer efﬁciency due to their controlled ﬂow channels, resulting in high ﬂuxes at low tangential ﬂows. Open-channel cassettes are used for high solids or high viscosity feeds. A very important attribute of cassettes is their ability to offer linear scale-up for reliability and speed of implementation. Spiral wound modules are still being used in high-volume pharmaceutical or industrial biotech applications. When cost is the main concern, the high membrane packing density of spiral modules is an important beneﬁt. Hollow ﬁber devices are preferred when viscous ﬂuids are being processed and an open-channel device is needed. 4.4.2 Virus Filtration Membranes

Membrane processes have proven to play an important part in virus removal and are now found in virtually all biopharmaceutical manufacturing processes. Virus removal for biopharmaceutical drugs requires several steps with additive capabilities. Virus membranes have severe requirements, as they must remove 99.9 % of virus particles while passing almost the entire protein product in the feed stream [25]. As smaller viruses (e.g., parvoviruses) have become important, the difference in size between the protein and the virus has diminished, making size-based separations even more difﬁcult. The additional requirements (compared to UF) have led membrane development in a slightly different direction. This section will only focus on membranes that remove viruses by size exclusion. Membrane adsorbers also claim some virus removal capabilities through electrostatic interactions, but they will not be covered here. 4.4.2.1 Membrane Suppliers The most prevalent virus removal membranes presently used are regenerated cellulose hollow ﬁbers and ﬂat sheet polyvinylidene ﬂuoride (PVDF) and PES membranes. The principal suppliers of virus membranes today are Asahi [26], Millipore [27], Pall [28], and Sartorius [29]. Filter manufacturers classify virus clearance ﬁlters into two broad categories based on the removal needs of the biotech industry – ﬁlters that are capable of removing viruses 50 nm or larger (retroviruses) and ﬁlters that can remove both small (20nm parvoviruses) and large viruses. Asahi Kasei Corporation (Japan) manufactures Planova ﬁlters, which were among the ﬁrst membranes speciﬁcally designed for virus ﬁltration. They are dead-end hollow ﬁber units with regenerated cellulose membrane made through the cuprammonium process (i.e., different from regenerated cellulose UF membranes). Pall Corporation uses hydrophilic PVDF for its Ultipor DV50 and DV20 membranes. Sartorius produces the Virosart CPV ﬁlter, which features a PES membrane. For TFF removal of small viruses, Millipore offers membranes with hydrophilic PVDF in two ratings, Viresolve 70 and Viresolve 180 membranes, designed to pass proteins with 70 kDa or 180 kDa molecular weight, respectively, while retaining virus particles. For NFF removal of small and large viruses, Millipore offers NFP and NFR

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4 Membranes in the Biopharmaceutical Industry Tab. 4.4

Commercial virus removal membranes and brand names.

Supplier Large viruses only Asahi (TFF/NFF) Millipore (NFF) Pall (NFF) Small viruses Asahi (TFF/NFF) Millipore (TFF) Millipore (NFF) Pall (NFF) Sartorius (NFF)

Cellulose

PES

PVDF

Planova 35N, 75N NFR Ultipor VF DV50 Planova 15N, 20N V70, V180 NFP Ultipor VF DV20 Virosart CPV

ﬁlters, respectively. NFP uses membrane similar to V180, while NFR features a hydrophilic PES membrane. Table 4.4 above summarizes available virus membranes, separating them according to the polymer used, the viruses removed and the ﬁltration mode they are used in. 4.4.2.2 Membrane Structures There are two broad families of virus membrane structures, skinned asymmetric membranes (Millipore) and symmetric membranes (Asahi, Pall, Sartorius). Early UF membranes of appropriate pore size were examined for virus removal, but it was found that the pore size distribution was too broad to provide the necessary removal [30]. Development of virus membranes, such as Viresolve V180, then focused on eliminating macrovoids in the UF layer. The pore size in the ultraﬁltration layer increases gradually from the skin layer progressing towards the large-pore microﬁltration layer. The UF skin provides the selectivity needed to exclude viruses, while the thicker microporous support layer provides mechanical support for the membrane. The advantage of a thin retentive membrane is signiﬁcantly higher permeability. Alternative approaches used a much thicker, dense retentive layer that traps virus particles in its multilayer, interconnected pore structure [31]. Asahi Planova, for example, has a 35 mm-thick ﬁltration layer to assure high viral retention. Pall’s Ultipor DV membranes are symmetric without a distinct skin layer. Commercial membranes have average pore sizes 10 nm and larger in size, so that viruses, in some cases, as small as 20 nm may be sterically excluded from the pores while still permitting transmission of the desired protein product. 4.4.2.3 Devices Both tangential ﬂow (TFF) and normal ﬂow (NFF) ﬁltration devices are used for viral clearance. In both cases membrane units are typically disposable and designed for a single use, in order to avoid the strict validation requirements that would be linked to cleaning a potentially infectious particle. Historically, TFF systems were more

4.4 Ultrafiltration and Virus Filtration Membranes for Biopharmaceutical Applications

Fig. 4.12 Commercial virus membrane modules. (Courtesy Pall Corporation, Millipore Corporation, Asahi Kasei Medical, and Sartorius AG)

common, following UF experience. Recent improvements in NFF ﬁlters along with their reduced complexity of operation and easier validation have led to a predominance of NFF ﬁlters. All NFF ﬂat-sheet ﬁlters contain two to three layers of membrane in order to increase virus removal assurance. TFF has been covered in the UF section. In NFF, also referred to as ‘‘dead-end’’ ﬁltration, ﬂuid ﬂows perpendicular to the ﬁlter membrane surface, with the ﬂow driven either with a pump (constant ﬂow) or by pressure (constant pressure). NFF virus ﬁlters are available in a variety of formats and ﬁlter areas to satisfy strict requirements for scalability. In virus ﬁltration, scalability affects not just process economics but also validation of virus spiking studies. Small-area ﬁlters are available for process development and virus validation studies, while for pilot and production scale operations, ﬁlters with membrane area up to 6 m2 are available. Additionally, multiple ﬁlter cartridges or module assemblies can be conﬁgured to achieve even larger ﬁltration areas. Figure 4.12 shows various commercial devices in use. 4.4.2.4 Membrane Selection Criteria for selecting a virus ﬁlter tend to be product- and process-speciﬁc [32]. A typical biotech process might contain more than one virus ﬁltration steps. To effectively evaluate the impact of these variables, in-house testing is typically performed. Performance related criteria for selecting a virus ﬁlter include the following: virus retention capabilities; protein product transmission;

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product throughput requirements; overall process economics; process compatibility; thermal and hydraulic stress resistances; extractables.

Compatibility with cleaning solutions is not an issue for virus ﬁlters, as they are typically disposable. Detailed validation packages, including spiking studies, are typically required by users of these ﬁlters. 4.4.2.5 Characterization The key parameter needed to characterize a virus ﬁlter is its ability to retain viruses of a certain size. This parameter is normally expressed in terms of a log reduction value, LRV, deﬁned as follows: LRV ¼ logðCfeed =Cpermeate Þ;

where Cfeed and Cpermeate are the virus concentrations in the feed and permeate streams, respectively. For smaller viruses (such as parvovirus) membrane ﬁlters provide LRVs above 3–4. LRVs above 6 are normally seen for larger viruses (such as retrovirus). These parameters are typically determined during viral spiking studies, where appropriate virus preps are used to challenge the membranes. Membrane manufacturers have speciﬁc virus removal claims on their ﬁlters, but users generally perform a separate spiking study, using their own product formulations. Regulatory agencies currently require validation of virus removal steps using actual viruses. For assurance of ﬁlter integrity, membrane manufacturers practice particle challenge tests, using bacteriophage [33] or gold nanoparticles [34]. Compared to viruses, the advantages of bacteriophages are increased prep purity, faster more sensitive assays and safer operation. Gold particles offer much better control of size and uniformity, and simpler detection. Particle challenge tests can be performed in conditions that simulate more or less actual process conditions. The presence of fouling streams, for example, might affect retention. Proper selection of the stream is important to generate meaningful data. To completely validate virus removal within a speciﬁc process run, novel liquid porosimetric integrity tests have been developed that rapidly and nondestructively predict the intrinsic viral retention capabilities of virus-retentive membranes. The CorrTest method [35] is essentially a ratio of two membrane permeabilities measured at preselected operating conditions using a pair of mutually immiscible ﬂuids, one of which is employed as a membrane wetting agent and the other used as an intrusion ﬂuid. The CorrTest method can be used to verify pre- and post-use membrane integrity and has been shown to correlate well with bacteriophage fX-174 retention. Standard air diffusion tests are also performed on devices to assure against the presence of gross leaks. In contrast to particle tests – that give information on the pore

4.5 Applications of Ultrafiltration Membranes in Biopharmaceutical Manufacturing

size distribution of the membrane – air diffusion tests identify large defects on the membrane surface. In addition to integrity, adequate passage of the product of interest – or an appropriately sized marker molecule – can be veriﬁed in a separate test. The optimization of virus retention and product passage properties can yield a membrane that is best suited for this application.

4.5 Applications of Ultraﬁltration Membranes in Biopharmaceutical Manufacturing

Biopharmaceutical manufacturers use ultraﬁltration to concentrate and condition biological molecules, generally proteins [35–39]. Ultraﬁltration membrane processes are used to provide optimum concentration ranges and/or buffer compositions for subsequent processing steps. An ultraﬁltration process operates in a tangential ﬂow (also known as crossﬂow) mode to retain a biological molecule while letting small molecular components such as buffer and water to freely pass through the membrane. In biopharmaceutical processing, ultraﬁltration is primarily operated in a batch or a fed-batch mode. Process operation times are typically between 3 and 8 hours. Manufacturing scale system sizes range from being as low as 5 m2 of membrane ﬁlter area for specialty therapeutics to as high as 300 m2 for high-dose monoclonal antibody applications. Figure 4.1 (see page 92) shows the typical location of ultraﬁltration steps in the downstream puriﬁcation of a monoclonal antibody. There are broadly three different areas where manufacturers use ultraﬁltration in downstream biopharmaceutical applications: 1. Post-harvest concentration: The ultraﬁltration step is used to reduce volume and adjust the buffer conditions prior to loading of the product on the ﬁrst chromatography step. Low feed concentrations (0.2–2 g/L), high process volumes (1000– 20 000 L), and higher level of impurities (DNA, lipids, host cell protein, etc.) characterize this area. Bioburden and fouling play an important role in process design in this process step. A typical process will concentrate the feed stream 5–10-fold with two- to threefold diaﬁltration. 2. Intermediate buffer exchange: This step is very similar to step 1 above except that the feed is more concentrated (5–30 g/L) and process volumes (200–1000 L) and impurity levels are lower. Typically, the feed is concentrated between 5 and 10 fold with three- to ﬁvefold diaﬁltration. 3. Formulation: The emphasis on the formulation UF step is to carefully achieve a predetermined product concentration and formulation buffer conditions. The product is very pure at this step and may be concentrated up to 180–200 g/L. Diaﬁltration

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volumes are also relatively high, ranging between 8 and 12 depending on the extent of buffer exchange required. The engineer faced with developing and operating an ultraﬁltration process operation will want to optimize the key elements of process consistency and robustness, the ability of the process to operate under normal operating condition variations, product yield, product purity, ease of use, scalability, validation, and process economics. The manufacturing process may impose constraints in the form of ﬁtting a process to an existing equipment, limitations on buffer or cleaning chemical usage, meeting ﬂoor space requirements and product quality considerations at operating conditions (pumping, process duration, etc.). 4.5.1 Ultraﬁltration Theory

There are several excellent treatments of ultraﬁltration theory in the literature [36,37,41–43]. In this section, we examine the relationships among critical parameters that derive from the commonly employed theories to explain ultraﬁltration behavior. In the polarization theory, the product concentrations may be related to the permeate ﬂux and tangential ﬂow parameters as follows: Cw Cp J ¼ exp ; Cb Cp k

where Cw, Cb and Cp are the concentrations of the product at the membrane surface, bulk feed, and permeate, respectively, J is the ﬂux or permeate ﬂow per unit area, and k is the mass transfer coefﬁcient, which is affected by tangential ﬂow rate. An empirical gel model is obtained from Equation (1) by taking Cp ¼ 0 and setting Cg (or gel concentration) ¼ Cw to get J ¼ k lnðCg =Cb Þ:

Equation (1) indicates that the ultraﬁltration performance parameters, k and Cg, may be obtained from a plot of J versus ln Cb, k is obtained from the slope, and Cg from the x-intercept. Finally, there is also an osmotic pressure model that relates process ﬂux to transmembrane pressure and osmotic pressure, P, resulting from concentration of protein at the membrane wall. The osmotic model is expressed as J ¼ L pðTMP DPÞ:

The method for determining the ultraﬁltration performance parameters, k and Cw, are detailed in [19].

4.5 Applications of Ultrafiltration Membranes in Biopharmaceutical Manufacturing

Fig. 4.13 Characteristic ﬂux versus TMP curve in an ultraﬁltration process.

Figure 4.13 shows a typical ﬂux versus transmembrane pressure (TMP) curve. The region at low ﬂux/low TMP is called the linear region, where the ﬂux is linearly dependent on TMP. The region at high ﬂux/high TMP is called the polarized region and in this section, the ﬂux is independent of TMP. In between these two regions lies what is termed the ‘‘knee’’ of the ﬂux curve. Increasing pressure beyond the knee gives diminishing return in improving ﬂux. The ﬂux versus TMP curve is also dependent on crossﬂow and protein concentration – a higher ﬂux is generally realized at higher crossﬂow rates or lower protein concentrations for a given TMP. The intrinsic membrane sieving, also called passage, is deﬁned as si ¼ (Cperm/ Cwall), while intrinsic membrane retention or rejection is deﬁned as ri ¼ 1 si. The intrinsic sieving is inherent to the membrane and solute, while an observed sieving as so ¼ (Cperm/Cbulk) varies with polarization. 4.5.2 Typical Ultraﬁltration Process

A typical schematic of an ultraﬁltration process is shown in Figure 4.14. The process starts with ﬁlling the protein product solution in the feed tank, recirculating this feed using the feed pump, and removing virtually product-free permeate as a waste stream. As the process proceeds, the tank volume drops and the concentration of retained product in the tank increases. This concentration step proceeds until the product concentration meets a certain target concentration. The buffer condition is then modiﬁed using what is called the diaﬁltration step where a different buffer, or diaﬁltrate, is added at the same rate as permeate is withdrawn (also known as constant volume diaﬁltration). During diaﬁltration, the tank volume and retained protein

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Fig. 4.14 Ultraﬁltration process schematic.

concentration remains constant (for r ¼ 1). Diaﬁltration is complete when the buffer is changed to the desired composition. The solution then proceeds to the next step. The relationship between product concentration and volumetric reduction factor X ¼ (initial volume)/(ﬁnal volume), the number of diaﬁltration volumes, N ¼ buffer volume added)/(ﬁxed retentate volume) and membrane rejection, r is CRN ¼ C0 e½½ð1rÞNþrlnX :

Note that for freely passing solutes (r ¼ 0), >4 diavolumes are needed to achieve a speciﬁcation of <1% of the original buffer components. The corresponding equation for retentate product yield may be written as YieldRetentate ¼ eð1rÞðNþlnXÞ :

Permeate loss is expressed as 1 YieldRetentate. Note that a membrane with 1 % sieving (99 % retention) can have process yield losses much higher than 1 % because the protein is repeatedly cycled past the membrane during the entire process with losses at every pass, that is, the higher the N, the more is the yield loss. A high yielding process (<1 % total product loss) typically requires sieving of <0.1 % (retention of >99.9 %). 4.5.2.1 Process Development and Optimization [38] The process parameters that govern successful development, design and operation of an ultraﬁltration process are as follows: 1. membrane type and module characteristics; 2. operating parameter optimization; 3. processing plan; 4. system design implementation. Membrane Type and Module (Device) Format Membrane selection for a given application is based on factors such as chemical compatibility, mechanical robustness, and retention consistency for the product protein, as well as process ﬂux and

4.5 Applications of Ultrafiltration Membranes in Biopharmaceutical Manufacturing

overall economics (cost). Experience with vendors as well as their record of quality manufacturing practices (ISO, cGMP) is also a consideration that inﬂuences membrane selection. Operating Parameter Selection and Optimization A systematic process development methodology is required to achieve a robust, consistent and optimized TFF step. Candidate membranes are often screened using low membrane area (50–100 cm2) linearly scaleable cassette devices – volume requirements for these devices are between 50 and 500 cc. For more detailed optimization investigation, a 0.1 m2 device is often used with process volumes ranging from 1 to 15 L. Pilot-scale processing, which is often carried out to generate material for product stability and characterization studies, is used to conﬁrm operation scalability and robustness. Volumes at pilot-scale range from 20 to 2000 L depending on the application. A typical process development sequence is described below: 1. Module installation and conditioning: Flush and measure normalized water permeability (NWP). Verify integrity using air diffusion. Precondition membrane devices by ﬂushing with buffer.

2. Initial fouling characterization: Add product solution and measure ﬂux and passage versus time in permeate recycle mode (permeate and retentate recycled to feed tank). This is done to evaluate membrane fouling behavior and allow for membrane conditioning in product solution. (Note: The initial conditioning veriﬁes that the module is integral and also establishes as baseline permeability for the device. The fouling characterization study should show asymptotic approach to steady-state ﬂux and retention. Continued decline in performance indicates a membrane compatibility issue.) 3. Flux and crossﬂow excursion in initial feed: In a total recycle mode (permeate and retentate returned to feed tank), measure ﬂux, protein passage, and turbidity versus TMP (10–60 psi) and crossﬂow (or DP of 10–30 psi). (Note: Vary the TMP (10–60psi) at constant crossﬂow (say P ¼ 10 psi); repeat the procedure at different crossﬂows (or P ¼ 20 psi, 30 psi, etc.).) 4. Concentration mode: Measure ﬂux, passage, turbidity, retentate concentration, temperature, and permeate volume versus time while removing permeate to achieve concentration. 5. Flux and crossﬂow excursion in ﬁnal retentate: Measure ﬂux, passage, and turbidity versus TMP (10–50 psi) and crossﬂow (P of 10–30psi) in total recycle mode.

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6. Diaﬁltration mode: Repeat step 5 while withdrawing permeate and adding diaﬁltrate at the same rate to the feed tank to achieve target impurity concentration – this mode of diaﬁltration is referred to as constant volume diaﬁltration. 7. Product recovery: Recover retentate product by depolarizing and using a plug-ﬂow ﬂush. 8. Measure yield. 9. Membrane CIP: Flush with buffer, then WFI, add 0.1N NaOH solution, recirculate for 45 min at 20 8C in permeate recycle mode, and measure NWP. NWP should be the same as the initial value. If not, recleaning is required. Additional development will include repeatability evaluation, cleaning optimization, and further exploration of particular conditions. Selection of Optimum TMP and Crossﬂow A typical ﬂux excursion behavior is shown in Figure 4.15. The optimum TMP is located at the ‘‘knee’’ of the ﬂux curve to obtain reasonable ﬂux and avoid formation of aggregates. If there is a difference in the ﬂux curves between the starting solution and the diaﬁltered solution, a conservative approach for polarization would be to select the lower of the two TMP values.

Fig. 4.15 General ﬂux trends in ultraﬁltration.

4.5 Applications of Ultrafiltration Membranes in Biopharmaceutical Manufacturing

Fig. 4.16 Process data for example scale-up calculations.

Choice of crossﬂow rate needs further consideration. Higher crossﬂow rates tend to result in higher ﬂux and reduced polarization, but this also leads to larger pump size with a concomitant increase in holdup volume and pump passes (potential for product degradation). For linearly scalable modules, the following approach may be used for guidance in crossﬂow selection. Figure 4.16 provides data for an example calculation. In this example, a device with 0.1 m2 membrane area is operated at two crossﬂow rates and the permeate ﬂux plotted versus TMP. The optimum point for each rate is chosen, and the permeate ﬂux recorded. For this example, the crossﬂow feed rates are Q1 ¼ 0.65 L/min Q2 ¼ 0.39 L/min Crossﬂow feed rates normalized for membrane area are termed feed ﬂux JFQ1 ¼ 6.5 L/min/m2 JFQ2 ¼ 3.9 L/min/m2 The permeate ﬂux at the optimum point for each case is JfQ1 ¼ 150 L/m2/h JfQ2 ¼ 86 L/m2/h We assume V/T, permeate volume/process time is equal. Then, since A ¼ Membrane area ¼ V/T permeate ﬂux, A1/A2 ¼ (1/150)/(1/86) ¼ 0.57 The area at the higher crossﬂow rate is 57 % of the lower crossﬂow rate. Now we calculate the ratio of feed rates required at each feed crossﬂow. Pump rate ¼ feed flow flux area ¼ JFQ A; and the ratio is 6:5=3:9 0:57 ¼ 0:95:

In this example, the difference between pumping rates is not signiﬁcant, so one would chose the higher crossﬂow rate, because less membrane area would result in

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lower capital costs and smaller system size. In other cases, the shapes of the permeate ﬂux versus TMP curves may dictate otherwise. It is also important to note that this method applies to systems that scale linearly (i.e., crossﬂow per unit area is the same for all system sizes). For nonlinear system scaling, for example, some hollow ﬁber or spiral wound systems, the procedure for selecting crossﬂow rates is based more on experience than any analytical method. Flux typically drops as protein concentration increases, so choose an average ﬂux for the above calculation. For a robust scale-up, always incorporate a safety margin into the membrane area requirements to account for lot-to-lot variability effects on membrane permeability, feedstock fouling characteristics, and batch volumes. Typically, a safety margin of at least 20–50 % extra membrane area is used, but this could increase or decrease depending on the expected variability in the process. 4.5.3 Processing Plan Optimization

In the development of a processing plan, the engineer has to consider choice of the mode of operation, diaﬁltration strategy, process time, temperature, and an assessment of process robustness vis-a`-vis lot-to-lot variability associated with feed and membrane. 4.5.3.1 Mode of Operation Most UF processes in the biopharmaceutical industry are operated in a batch mode (as opposed to continuous mode). In a batch mode, the entire product feed is recirculated within the system (no new feed is added to the system); permeate is continuously removed from the system during this time resulting in an increase in product concentration. The retentate, at its desired concentration, is available at the end of the processing time. When large concentration factors (>15–20) need to be achieved, a variation of batch process called fed batch is often used (Figure 4.17). The conﬁguration provides some ﬂexibility in processing a variety of batch volumes in a single skid. For a

Fig. 4.17 Fed-batch conﬁguration.

4.5 Applications of Ultrafiltration Membranes in Biopharmaceutical Manufacturing

Fig. 4.18 Fed-batch process concentrations.

fed-batch operation, the retentate is returned to a smaller tank, not the large feed tank. Feed is added to the small retentate tank as permeate is withdrawn. The smaller retentate tank can allow a smaller working volume without foaming. Figure 4.18 shows the retentate concentration over the course of the process as a function of the tank ratio ¼ VFeed tank/VRetentate tank. Notice that the fedbatch process operates at a higher overall concentration compared to the batch process and consequently at a lower average ﬂux. Typical fed-batch tank ratios are chosen to be <4–5 to allow for process ﬂexibility, while minimizing loss of operating ﬂux. 4.5.3.2 Diaﬁltration Mode/Strategy Constant-volume diaﬁltration is commonly used control mode for adding diaﬁltration buffer. To perform a constant-volume DF, buffer is added to the recycle tank at the same rate that ﬁltrate is removed. The total volume of retentate remains constant throughout the process. Another important aspect of the diaﬁltration process is its placement (strategy) in the overall process. The optimum point for placement of diaﬁltration may be derived from the gel polarization model [44] as (Figure 4.19). Copt ¼ Cg =e;

where Cg is the concentration at which the ﬂux drops to zero and e ¼ 2.718, which is the base of the Naperian logarithm. For monoclonal antibody processing, Cg values of 200–300 g/L have been reported, yielding an optimum diaﬁltration concentration of 70–110 g/L. Until recently, ﬁnal antibody concentrations have generally been in the 5–20 g/L range,

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Fig. 4.19 Optimum diaﬁltration point – polarization model.

where it is convenient to diaﬁlter to the ﬁnal formulation concentration. That is, it is simpler to concentrate to 20 g/L and diaﬁlter at this concentration (i.e., you would not take the concentrate to the optimum range of 70–110 g/L in order to diaﬁlter followed by dilution to come back to 20 g/L). However, with a recent push towards high formulation concentrations, in the range of 150–200 g/L, diaﬁltration at the optimum diaﬁltration concentration in process applications is increasingly becoming a reality. Since, for example, if the ﬁnal desired concentration is 150 g/L, one can pause during the concentration step at 70–110 g/L (which is the optimum diaﬁltration concentration) in order to carry out the diaﬁltration, and once diaﬁltration is complete, further concentrate to the desired ﬁnal concentration of 150 g/L. A more general approach for selecting the optimum diaﬁltration point is to use the DF optimization parameter approach. In this approach, one calculates the diaﬁltration optimization parameter, DF ¼ C J, at each protein concentration (from the plot of the ﬂux vs. protein concentration curve). A plot of the DF optimization parameter versus protein concentration is constructed to ﬁnd the optimum protein concentration where the value of the DF optimization parameter is maximized (Figure 4.20). It is important to plot these data with the protein in both the initial and ﬁnal buffers, since ﬂux can often change signiﬁcantly with different buffers. If the optimum is very different for the two buffers, it is most conservative to choose the optimum based on the buffer curve which results in the lower value for the product concentration. Finally, the goal of a diaﬁltration step is to reduce buffer or contaminant species from a product in the retentate. Since increasing the number of diavolumes that are performed adversely impacts yield, it is best to minimize diavolumes consistent with obtaining correct ﬁnal composition. Figure 4.21 illustrates the relationship between contaminant removal and number of diavolumes. The maximum effective diavolume is about 14 before the limits of the system or the chemistry are reached (system deadlegs, mixing effects, etc.).

4.5 Applications of Ultrafiltration Membranes in Biopharmaceutical Manufacturing

Fig. 4.20 Optimum diaﬁltration point – diaﬁltration optimization parameter approach.

4.5.4 Scale-up Considerations

Once a process has been developed at lab bench scale, it must be translated to industrial scale and validated, and this can present unanticipated surprises and challenges. Often, there is little opportunity to perform intermediate-scale runs due to time and material constraints. In addition, material from the initial industrial-scale runs is usually required for time-sensitive clinical or market supply. Therefore, accurate and dependable scale-up is critical for the success of a process. The simplest way to ensure accurate and predictable translation of product yield and purity from bench to industrial scale is to use linear-scale techniques.

Fig. 4.21 An example of contaminant removal by diaﬁltration.

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To linearly scale a TFF process [45], all ﬂuid dynamic and membrane module parameters must be kept constant between scales of operation. Fluid dynamic parameters, which are set by the user, include the ratio of feed volume to membrane area, feed rate per membrane area, ﬁltrate ﬂux, and retentate and ﬁltrate pressures. Membrane module parameters are inherent in the speciﬁc ﬁlters that are chosen, and include membrane material and pore size, turbulence promoter, channel height, channel length, and feed and ﬁltrate ﬂow geometries. Therefore, true linear scaling can only be done with modules from the same supplier. 4.5.4.1 Process Implementation Considerations Once a protein processing procedure has been developed, it must be integrated into a complete process. The typical sequence of steps in an ultraﬁltration/diaﬁltration process are set-up and pre-use cleaning, integrity and permeability testing, pre-use equilibration, protein processing, product recovery, post-use cleaning and testing and storage (Figure 4.22). The integrity test helps to ensure that the installed membranes have not sustained any damage during storage and handling. An integrity testing is normally carried out prior to startup and after each post-use cleaning. The typical integrity test for an ultraﬁltration system involves measuring air diffusion across the wetted membranes at a speciﬁed pressure. The diffusive ﬂow is compared against the speciﬁcation provided by the membrane manufacturer. The air ﬂow integrity test allows for identiﬁcation of problems such as macroscopic holes in the membrane, cracks in the seals, or improperly seated modules.

Fig. 4.22 Typical operational sequence for an ultraﬁltration process.

4.5 Applications of Ultrafiltration Membranes in Biopharmaceutical Manufacturing

It is also common practice to measure and compare the permeability (also referred to as normal water permeability or NWP) of the membrane system to assess its cleanliness from run-to-run. The normal water permeability is calculated as NWP ¼ ðflux temperature correction factorÞ=TMP:

The units of NWP are [L/m2/h psi/g]. Product recovery strategy is a key component of the process operating sequence. Product recovery is the process of removing the product from the TFF system into a vessel appropriate for storage or further processing. The bulk of the product, which is typically in the recycle tank, is pumped out using the feed pump. However, some liquid remains held up in the piping and the modules. A welldesigned system has minimal deadlegs in the piping and is sloped to a collection port at the lowest point in the piping to improve draining. Beyond simply draining the system, however, one of the following methods can be used to increase product recovery: low-pressure air blowdown; buffer displacement; buffer ﬂush; buffer recirculation. Overall product recovery in the ﬁnal retentate should be >95 % and yields of >99 % are common. After each membrane use and product recovery, the system assembly is cleaned using standard cleaning protocols [46,47]. Cleaning/sanitizing chemicals that are commonly used in ultraﬁltration processes include 0.1–0.5 N NaOH, 200–400 ppm sodium hypochlorite, 0.1–0.3 N phosphoric acid, peroxiacetic acid, and so on. If another protein processing run does not immediately follow the cleaning, recirculate an appropriate storage solution through the TFF assembly to prevent organism growth. The membrane modules must remain ﬁlled with storage solution until the next run to prevent drying of the membranes. 4.5.4.2 Process Robustness A process implemented at large scale must be robust if it is to be successful. A robust TFF process will perform well within the lot-to-lot feedstock and membrane variations that it encounters. While developing a process, it can be very useful to test performance at the extremes of these variations, as far as possible. These performance parameters characterize a well-designed, robust ultraﬁltration process:

Yield (overall) Process ﬂux consistency Product retention (membrane) NWP recovery (run-to-run)

95–98 % 10 % run-to-run 99.9 % 20 % cellulose membranes 20–35 % for PES membranes

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Typical ﬂux

Typical sizing

Product and process speciﬁc; for example, 30–120 LMH for 30 kDa membrane with MAb Product and process speciﬁc; for example, 5–10 m2/KL for MAb

4.5.5 System Considerations To implement a complete TFF process, the piping and equipment associated with the membrane modules must be selected and a method for controlling the process parameters at their set-points must be chosen.

4.5.5.1 Equipment Options In addition to the membrane modules and holder, at minimum a working TFF operation requires a recycle vessel, a feed pump, a retentate control valve, and pressure sensors for the feed and retentate lines. Many systems also include feed and ﬁltrate ﬂow meters, a ﬁltrate pressure sensor, and sensors for temperature, pH, conductivity, or UV absorbance. Most TFF systems used for protein processing are operated in a sanitary manner, requiring sanitary ﬁtting.

4.5.5.2 Process Control Options Throughout a TFF process, as protein is concentrated or exchanged into different buffers, the process parameters need to be adjusted so that they remain at their set points. Several methods of process control are used to accomplish this. The tangential ﬂow can be controlled to maintain either constant pressure drop; constant crossﬂow rate.

Controlling to constant pressure drop is a simple and easy option in the sense that it requires the use of only feed and retentate pressure gauges. The exact feed ﬂow is not known. Since pressure drop is uniquely related to crossﬂow (P ¼ AQn), a constant pressure drop is assumed to connote a constant ﬂow rate. However, in many cases a constant pressure drop may not always ensure that the ﬂow rate is maintained constant. Differences in the internal membrane module resistance, pump slip, ﬂuid viscosity changes, and so on. during a process may contribute to changes in the ﬂuid crossﬂow without impacting pressure drop. Controlling to constant crossﬂow rate is a more robust way of controlling the process. A ﬂow meter is required in the feed or retentate ﬂow line for this control option. The applied pressure can be controlled (by modulating the retentate valve) to maintain a constant Retentate pressure; TMP;

4.6 Practical Aspects of Virus Filtration Process Design and Implementation

ﬂux; Cwall (protein concentration at the membrane surface); mixed mode control.

Our process development engineer, having reached this point, will ﬁnd that by following the systematic methodology presented here will have an excellent understanding of the biopharmaceutical manufacturing process that results. Our engineer will be able to produce the necessary SOP’s, validation documents and operating instructions from detailed experience. In operation, the variance in day-to-day running can be expected to be controlled based on the knowledge gained during scale-up and implementation.

4.6 Practical Aspects of Virus Filtration Process Design and Implementation

Recombinant products such as monoclonal antibodies are expressed by mammalian, bacteria or yeast cells in fermenters, and within ascites ﬂuid, or ﬂuids from transgenic mammals. Mammalian cell lines may contain endogenous viruses that are generated in the bioreactor. Endogenous retroviruses are expressed because the retroviral genome is integrated into the cell line and cannot be screened out during the creation of the Master Cell Bank. This causes retrovirus like particles, or RVLPs, to be produced within the bioreactor. Products can also become contaminated by adventitious viruses, which enter the process streams. Although regulatory agencies do not mandate the inclusion of speciﬁc viral clearance technologies, they do mandate the safety standard for ﬁnal doses, and require full validation of viral clearance steps by the manufacturer. Regulations and good manufacturing practice (GMP) require two orthogonal and robust steps for endogenous viral clearance [48]. Orthogonal steps rely upon different mechanisms to achieve virus clearance. For adventitious viruses, these requirements typically result in the use of multiple viral clearance steps [49]. Filtration removes viruses based on size exclusion, since the pores of the ﬁlter are smaller than the virus. Virus ﬁltration is considered a robust operation because the removal efﬁciency is insensitive to normal variations in process conditions. Consequently, most well-designed downstream processes include a virus ﬁltration step. 4.6.1 Membrane Selection

There are many commercially available virus ﬁlters, each with their own strengths and weaknesses. Filter selection should be based upon the nature of the application (product and process) and the performance of the ﬁlter as demonstrated in both qualiﬁcation and validation studies. A systematic methodology for developing a robust virus removal step includes the following:

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virus ﬁlter selection; process design and optimization; process sizing and simulation; virus validation studies; manufacturing implementation.

Virus clearance ﬁlters are classiﬁed into two broad categories to meet the removal needs of the biotech industry – ﬁlters that are capable of removing viruses 50nm or larger (retroviruses) and ﬁlters that can remove both small (20nm Parvoviruses) and large viruses. NFF virus ﬁlters are available in a variety of formats and ﬁlter areas. Small-area ﬁlters are generally available for process development and optimization studies as well as virus validation studies. For pilot and production scale operations, commercial vendors offer individual virus ﬁlters ranging in size from 0.07 to 6m2. Additionally, multiple ﬁlter cartridges or module assemblies can be conﬁgured to achieve even larger ﬁltration areas. Commercial ﬁlters that are currently available for virus removal applications in the normal ﬂow ﬁltration mode are summarized in Table 4.5.

Tab. 4.5

Vendors and brand names of commercially available virus filtration products.

Company type

Product name

Virus

Sizes

Millipore FF

NFP

>4 log X-174 bacteriophage

Scale-down 3.5 cm2; process modules 0.08–1.5 m2

Millipore TFF

NFR Viresolve 70

>6 log retrovirus >4 log polio; >7 log retrovirus

Sartorius NFF

Viresolve 180 Virosart CPV

>3 log polio; >6 log retrovirus >4 log PP7 bacteriophage; >6 log retrovirus

Pall NFF

DV20

>3 log PP7 bacteriophage; >6 log PR772 bacteriophage

Asahi TFF/NFF

DV50 Planova 15N

>6 log PR772 bacteriophage >6.2 log parvovirus; >6.7 log poliovirus

Planova 20N

>4.3 log parvovirus; >5.4 log Encephalomyocarditis >5.9 log Bovine viral diarrhea virus; >7.3 HIV

Planova 35N

Scale-down 150 and 1000 cm2;process modules 0.75–1.4 m2 Scale-down module 5 and 20 cm2; process module through 0.7–2.1 m2 Scale-down 14 and 140 cm2; process modules 0.07–6 m2 Scale-down modules 10 and 100 cm2; process modules 0.12–4.0 m2

4.6 Practical Aspects of Virus Filtration Process Design and Implementation

Performance-related criteria for selecting a virus ﬁlter include virus retention capabilities, protein product transmission/product recovery from the ﬁltration step, product throughput (rate) requirements, and overall process economics. These criteria tend to be product speciﬁc, and to effectively evaluate the impact of these variables, in-house testing is typically performed. Less obvious considerations for selecting a virus ﬁlter revolve around process compatibility and system integration issues. All materials of construction must be chemically compatible both with the protein product as well as all relevant processing conditions. Additionally, thermal and hydraulic stress resistances, extractibles, and the effects of cleaning/sterilization/sanitization on the membrane devices should be evaluated to determine if they are consistent with the proposed implementation scheme. Additional information on these topics can be found in the PDA Technical Report No 41 on Virus ﬁltration [49] or from the various vendors listed in Table 4.4. To properly optimize a virus ﬁltration process and establish process robustness, one must consider all processing variables that impact virus retention (LRV), product recovery, and product throughput. Additionally, from an economic point of view, process optimization is extremely important. For large volume processes, such as monoclonal antibodies, virus ﬁltration can be one of the most expensive unit operations. Virus ﬁltration is more expensive than sterile ﬁltration due to both higher ﬁlter costs and lower product throughputs. Table 4.6 provides a typical range for cost and performance parameters for virus and sterile ﬁlters. Virus ﬁltration can either be run in a NFF mode or a TFF mode [50]. Historically, TFF systems were more common, but recent improvements in NFF ﬁlters have lead to their predominance. Because of this predominance of NFF approaches, TFF systems will not be discussed further here. In NFF, also referred to a ‘‘dead-end’’ ﬁltration, ﬂuid ﬂows perpendicular to the ﬁlter membrane surface. NFF processes can be run either under constant ﬂow operation or constant pressure operation. Constant ﬂow is more common in manufacturing settings as most people would use pumps to drive the ﬁltration and it is simpler to run at a constant pump setting.

Tab. 4.6 Comparison of virus filter costs, process fluxes, and capacities compared to sterilizing grade filters.

Filter type Virus Parameter

Sterile

NFF 20 nm virus ﬁlter

NFF 50 nm virus ﬁlter

TFF virus ﬁlter (>20 nm)

Unit ﬁlter cost ($/m2) Typical process ﬂux (L/m2/h/psi) Typical design capacity (L/m2)

200–300 200–500 >2000–4000

>2000–4000 0.5–6 60–500

>1000 20–30 800–1500

>2000–4000 5–15 250–500

NFF, normal ﬂow ﬁltration and TFF, tangential ﬂow ﬁltration.

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During operation, protein products or other components can accumulate at the membrane surface or adsorb to internal surfaces. These two fouling mechanisms will reduce the hydraulic permeability of the membrane and may impact virus retention. Fouling results in a decrease in ﬁltrate ﬂow rate with time for constant pressure operations or an increase in upstream pressure with time for constant ﬂow operations. 4.6.2 Process Development and Optimization

A schematic representation of the experimental set-up to conduct normal ﬂow ﬁltration experiments at constant pressure is shown in Figure 4.23. The scale-down test may be carried out in either a constant pressure or a constant ﬂow mode. The constant pressure set-up is often simpler, and the testing is easier to execute in that the set-up does not require a pump to drive the ﬁltration process. The typical steps employed in a scale-down NFFprocess evaluation are described in the following general test protocol. System set-up. Water ﬂush. Installation check: Flush the ﬁlter with water and carry out a pressure hold test to conﬁrm installation integrity of the devices. (This step is optional for process development and is typically carried out in virus validation studies.)

Fig. 4.23 Experimental set-up for virus removal testing.

4.6 Practical Aspects of Virus Filtration Process Design and Implementation

Buffer conditioning. Product ﬁltration: During ﬁltration, ﬁltrate volume (V) collected is measured and recorded at various ﬁltration times (t). Filtration time may vary between 45 and 120 minutes. Assay ﬁltrate sample for product concentration. Recovery: Assay buffer ﬂush sample for product concentration and calculate product recovery. Installation check. Calculate ﬁlter capacity and initial ﬂux: The ﬁlter capacity and the initial ﬂux are generally obtained by ﬁtting the experimental data (V vs. t) to the gradual pore plugging model. (Vmax – see Section 3.1.1.4.) Calculate minimum required ﬁlter area: Once the ﬁlter capacity and initial ﬂux values are calculated, the minimum required ﬁlter area can be calculated from the Vmax model. This value should only be used as a comparative tool for different devices during process optimization studies. The ﬁnal design ﬁlter area will be determined during process simulation and virus validation studies.

Optimization of a virus ﬁltration process involves evaluating the effect of a variety of process parameters to arrive at optimum conditions that would ensure robust, consistent, and scalable operation. Figure 4.24 represents a generic approach to optimization schematically. Some of the key process development parameters that impact process performance are described in more detail below. There are typically three locations within the downstream process train where a normal ﬂow virus ﬁltration step is implemented within a given downstream process. These locations are as follows: following the low-pH inactivation step; following the intermediate chromatographic operation; or after the ﬁnal chromatography step. Since protein concentration, impurity concentration, and process volumes vary dramatically throughout the downstream process train, it should come as no surprise that the actual ﬁltration requirements are highly dependent upon where in the process the virus ﬁltration step is located. The beneﬁt of increasing ﬁlter capacity and ﬂow at lower product concentrations is offset by an increase in process volume. The interplay of these two competing effects can often result in a feed concentration that minimizes required ﬁltration area [51]. An optimum feed concentration may exist that maximizes ﬁltration performance (minimizes ﬁltration area (m2) and maximizes productivity (g/m2/h)). For high concentrations (>10–15 g/L), it may be advantageous to dilute the product to improve ﬁlterability. The effect of ﬁltration pressure is often best determined by conducting an excursion study to evaluate ﬁlter capacity and ﬂow as a function of pressure. It is customary to evaluate pressure effects in the 10–50 psi range. In general, higher

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Fig. 4.24 General approach to optimization of virus ﬁltration.

operating pressures increase the average process ﬂux and decrease the required ﬁlter area [40]. The magnitude of this impact is dependent upon several factors, including feed solution condition, feed concentration, impurity proﬁle, and virus ﬁlter. 4.6.3 Capacity

Preﬁltration of the feed solution can have a dramatic impact on ﬁlter performance. Preﬁltration removes various impurities or contaminants such as protein aggregates, DNA and other trace materials. While larger size impurities can be removed by preﬁltering with a 0.2mm or 0.1mm rated microporous membrane, smaller

4.6 Practical Aspects of Virus Filtration Process Design and Implementation

impurities such as protein aggregates that may only be marginally larger in size compared to the protein product, are not easily amenable to size-based removal methods. Preﬁltration through adsorptive depth ﬁltration has been observed to provide signiﬁcant protection for certain virus removal ﬁlters [52]. The impact of preﬁltration can be quite dramatic; with up to 10-fold reductions in required ﬁlter area sometimes achievable. As these ﬁlters work by nonspeciﬁc multimode adsorption, product recovery should be conﬁrmed to ensure good yield. For some protein solutions, the freeze–thaw of a material can have a signiﬁcant impact on ﬁltration performance. In fact, in some instances, it has been observed that the required ﬁlter area is ﬁve- to sixfold higher when measured using material that has been previous frozen compared to fresh feed [51]. While the actual puriﬁcation process may not have a freeze–thaw step, feed samples required for virus validation testing are often conveniently submitted in a frozen form due to material stability/ availability considerations. In such situations, if freeze–thaw is observed to produce an adverse impact of ﬁltration performance, a preﬁlter is often used to restore performance similar to the unfrozen material. As discussed in Section 3.1.1.4, a gradual pore-plugging model is described by the Vmax model. Filter sizing is impacted by the ﬁlter capacity, Vmax, the initial ﬂow rate, Qi, and the batch time, tb. For typical processing times less than 4 hours, a higher ﬂux membrane with a corresponding lower capacity often results in lower ﬁltration area compared to a high-capacity/low-ﬂux ﬁlter. For processing times greater than 18 hours, the high-capacity/low-ﬂux ﬁlter would result in a process with lower ﬁlter area. It should be noted, however, that shorter processing times have the added beneﬁt of allowing for the possibility of in-line processing with other puriﬁcation steps as well as mitigating potential product stability issues. 4.6.4 Small-Scale Simulation

Once the optimum ﬁltration conditions have been determined, it is recommended that a simulation study be performed. This would initially be performed at the small scale (3.5–14 cm2), then repeated at a larger scale as the process is scaled-up. This would involve running the ﬁltration to the desired end point, which may be a speciﬁc ﬁltration time, volume/area ratio, or percent ﬂux decline. One of the outcomes of successful process development is a process that is robust and easy to implement when scaled-up to manufacturing. In order to demonstrate scalability of the process, it is recommended that pilot-scale studies be conducted using devices containing 100–1000 cm2 of ﬁlter area. This scale represents a 10–300fold scale-up from the initial simulation studies. 4.6.5 Pilot-Scale Studies

The objectives of the pilot-scale studies are twofold. A ﬁrst objective is to obtain conﬁrmation that the process parameters (process loading, time, ﬂux or pressure,

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and yield) are with predicted ranges and estimated bounds. Second, the pilot-scale studies are used to obtain information on the entire operation (installation, ﬂushing, sterilization/sanitization, integrity testing, process, product recovery, etc.) so as to enable drafting of SOPs and batch records for cGMP manufacturing. Information obtained during the pilot-scale studies can also be used to establish appropriate performance limits for water permeability (NWP), integrity testing, and other secondary operations related to the virus ﬁltration. 4.6.6 Virus Validation Studies

The purposes of the virus validation studies are to conﬁrm the LRV claims for the ﬁltration step and to verify the ﬁlter sizing established in the scale-up phase. These tests are run at a small scale, maintaining critical parameters such as pressure, ﬂux, and loading capacity at their commercial operation values while mimicking other operating procedures such as pre-processing WFI ﬂush-outs and buffer equilibration. The tests are typically run concurrent with the manufacturing-scale consistency/validation batches. Due to the handling and assay requirements for virus studies, the tests are typically conducted at specialized labs. The ﬁlter is challenged with representative feedstock containing a virus spike. The concentration of the virus in the feed and the pooled permeate is measured to calculate the LRV. Parallel control assays are run to correct for virus losses due to artifacts such as dilution, concentration, ﬁltration, and storage of samples before titration. Typically, manufacturers will place a lower limit of LRV 3 on the (log) reduction factors that will be combined to yield the overall reduction factor for the manufacturing process. To accurately represent manufacturing settings, the test feedstock must be identical to the commercial-scale feedstock. Shipping or storage constraints may require freezing feedstock, which can result in protein aggregates. Aggregates can cause premature ﬁlter plugging that may alter scaling parameters such as loading capacity. The problem can be obviated by either removing aggregates with microﬁltration or generating fresh feedstock at the site of the virus spiking study. The viruses used in the spiking studies depend on the speciﬁcs of the process and virus contaminants. The regulations recognize ‘‘relevant’’ model viruses that represent endogenous viruses, and nonspeciﬁc model viruses to validate ‘‘general viral clearance’’ for adventitious contamination. Murine Leukemia Virus (MuLV) is the generally accepted RVLP model for endogenous virus tests. If use of a relevant virus is not possible, the manufacturer chooses the best speciﬁc model virus to serve as a model for the relevant virus. To satisfy the ‘‘general viral clearance’’ objective, the study sponsor will generally evaluate two or three additional viruses. The nonenveloped parvoviruses (20 nm) are often accepted as a worst case for ﬁltration. The test thus comprises a four- or ﬁve-virus panel that represents viruses of different genomes (DNA and RNA), sizes and surface properties (enveloped and nonenveloped).

4.6 Practical Aspects of Virus Filtration Process Design and Implementation

Regulatory guidance states that ‘‘the amount of virus added to the starting material for the production step that is to be studied should be as high as possible’’ [52]. However, so as not to unacceptably alter the product composition, the volume of spike should be kept below 10 % and typically below 5 %. The guidance also voices concerns over virus aggregation that could be induced by deliberately concentrating the virus. The use of aggregated virus could lead to underestimation of inactivation effectiveness and overestimation of size-exclusion effectiveness [54]. The common practice is to use size-based preﬁltration, such as a 0.22 or 0.45 microporous membrane to remove virus aggregates from a spiked feed stream prior to performing the clearance study. Asahi uses a Planova 35 or 70 as pre-ﬁlter. Impurities contained within the virus spike may also foul the membrane, preventing the tests from reaching important scaling parameters such as loading capacity. Methods of generating highly pure virus preparations are increasingly being used to prevent fouling due to spike impurities [55,56]. Minimizing the amount of virus spike used while still being able to reach the desired LRV target is another approach. Virus retention has been observed to decline with fouling for a variety of ﬁlters [25,51,55,57,58]. If the virus spike required to achieve the target LRV causes excessive fouling, alternative validation methods may be used to determine LRV at higher throughput values [55]. These alternative methods may be used for validation after consulting with the appropriate regulatory agencies. 4.6.7 Implementation

Once the virus clearance step has been optimized and virus validation studies completed, an implementation strategy is required for robust process operation. After determining the ﬁlter capacity (L/m2) required for a process during process simulation/scale-up and virus validation studies, the ﬁlter area required for processing a given batch volume can be calculated. Various ﬁlter conﬁgurations are made available by manufacturers to facilitate large-scale implementation. Normal ﬂow virus ﬁlters are operated either in constant pressure mode or in constant ﬂow mode. Typical factors to be considered during large-scale virus ﬁltration system design include the following: Minimum and maximum batch volume. Minimum and maximum ﬂow rate; it is important to consider ﬂow rates during pre- and post-use water ﬂush and for post-use system cleaning/sanitization. Maximum operating pressure and differential pressures across the preﬁlter and the virus ﬁlter. If in-line dilution is needed to maintain constant feed concentration, appropriate dilution and mixing hardware. Minimum and maximum concentration and appropriate instrumentation to span the range.

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Appropriate hardware and connections to enable the ﬁlters and the system to be steamed, autoclaved or chemically sanitized. Filter housing conﬁguration – individual ﬁlters in parallel or a multiﬁlter housing. System hold-up volume versus validated post-process buffer rinse. Pre- and post-use integrity tests.

In the case of normal ﬂow virus clearance ﬁlters, pressure vessels are typically employed for constant pressure operation. However, when very large process volumes are involved, it may be easier to use a pump with a pressure feedback loop to carry out constant pressure ﬁltration. A typical sequence of operations in a virus ﬁltration process includes the following steps, very similar to the protocol used for ultraﬁltration process scale-down testing. Filter installation and ﬂushing – typically used to reduce extractables. Follow manufacturer’s directions and validation package. Measurement of NWP – Conﬁrm ﬁlter is within established ranges. Sterilization/sanitization – This step must integrate to the downstream processing philosophy of the user. Some virus ﬁlters are available pre-sterilized. Consider the Manufacturer’s guidelines for autoclaving/SIP treatment of the virus ﬁlter. Pre-use integrity testing. Buffer pre-conditioning. Processing and product recovery. Post-production integrity testing. To ensure that virus clearance is consistent with manufacturer’s claims and results obtained during virus validation studies, ﬁlter integrity should be checked both preand post-use. To facilitate this, ﬁlter manufacturers have developed a variety of destructive and nondestructive physical integrity tests that are related to virus retention, which were discussed in Section 4.2.5. Ultimately, the objectives of properly designed physical integrity testing are threefold: to conﬁrm that the virus removal ﬁlter is properly installed; assurance that the ﬁlter is free from gross defects and damage; conﬁrmation that the ﬁlter removes viruses consistent with both manufacturers’ speciﬁcations and end-user virus validation studies. Filter manufacturers should be able to provide evidence that integrity test methods and acceptance criteria correlate to retention of viruses in the targeted size range under standard conditions. The complexity of integrity testing virus ﬁlters should not be overlooked when selecting the virus removal ﬁlter for manufacturing. Key integrity test considerations include performance, safety, logistics, validation, and regulatory support [32].

4.7 Membrane Adsorbers

Currently available integrity tests for virus removal ﬁlters can generally be classiﬁed into three categories [50,59]. The ﬁrst category is a particle challenge test, the second type is a gas–liquid porosimetry test and the third type is a liquid–liquid porosimetry test. A more detailed summary of the various tests along with troubleshooting techniques can be found in the PDA TR41 [32]. While only nondestructive tests can be used pre-use, either type of test can be used for post-use testing. Nondestructive tests are either gas–liquid or liquid–liquid porosimetry tests. In general, a gas–liquid porosimetry test such as diffusion test or pressure hold test is recommended to complement liquid–liquid porosimetry test or particle challenge test to check for gross defects in the system. Pre-use integrity testing can be performed either before or after sterilization/ sanitization. Post-sterilization integrity tests are particularly useful since they ensure that the ﬁlters are not damaged during the sterilization process. However, in an aseptic process, one must maintain system sterility during ﬁlter wetting and integrity testing steps. Prior to protein processing, a buffer ﬂush is generally recommended in order to displace WFI with the appropriate buffer. The buffer ﬂush can be carried out using the conditions that are employed during protein ﬁltration (same DP, TMP or ﬁltrate ﬂux). About 10 L of buffer per m2 of ﬁlter area is a reasonable volume of buffer. After the buffer ﬂush, the system is ready for protein processing. The protein product should be processed using the process conditions and operating window established during the scale-down optimization studies and virus validation studies. In the case of normal ﬂow ﬁlters, protein recovery may be enhanced with a buffer rinse. The buffer rinse can be carried out using the conditions that are employed during protein ﬁltration (same DP or TMP or ﬁltrate ﬂux). Flush volume depends on the upstream volume of the system and desired protein yield. About 10 L/m2 is a reasonable ﬂush volume for a well-engineered system.

4.7 Membrane Adsorbers

Historically, membrane devices have puriﬁed products in ﬂuid streams by size-based ﬁltration. More recently, the use of membranes functionalized with speciﬁc ligands has started to gain widespread use for the adsorptive puriﬁcation of biotherapeutics. While bead-based chromatography is widely employed and effective, membrane chromatography has been heralded as a technology potentially suited for large-scale applications due to its ability to integrate capture and puriﬁcation steps for processing large amounts of product in relatively short times [60]. In bead-based chromatography, most of the available surface area for adsorption is internal to the bead. Consequently, the separation process is inherently slow since the rate of mass transport is controlled by pore diffusion. To minimize this diffusional resistance and concomitantly maximize dynamic binding capacity, small diameter beads can be employed. However, the use of small diameter beads comes at the price of increased column pressure drop. Consequently, the optimization of preparative

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chromatographic separations often involves a compromise between efﬁciency/ dynamic capacity (small beads favored) and column pressure drop (large beads favored). In contrast, membrane-based chromatographic systems (also called membrane adsorbers), have the ligands attached directly to the convective membrane pores, thereby eliminating the effects of internal pore diffusion on mass transport. Additionally, the use of microporous membrane substrates with a tight membrane pore size distribution coupled with effective ﬂow distributors can minimize axial dispersion and provide uniform utilization of all active sites [61,62]. Consequently, mass transfer rates of membrane adsorber media may be an order of magnitude greater than that of standard bead-based chromatography media [62], allowing for both high efﬁciency and high-ﬂux separations. Since single or even-stacked membranes are very thin compared to columns packed with bead-based media, reduced pressure drops are found along the chromatographic bed, thus allowing increased ﬂow rates and productivities. The necessary binding capacity is reached by using membranes of sufﬁcient internal surface area, yielding device conﬁgurations of very large diameter to height ratios (d/h) [60]. Properly designed membrane adsorbers have chromatographic efﬁciencies that are 10–100 better than standard preparative bead-based resins. Consequently, to achieve the same level of separation on a membrane adsorber, a bed height 10-fold less can be utilized. Bed lengths of 1–5 mm are standard for membrane adsorbers, compared to bed heights of 10–30 cm for bead-based systems. Due to the extreme column aspect ratios required for large-volume membrane adsorbers, device design is critical. To maintain the inherent performance advantages associated with membrane adsorbers, proper inlet and outlet distributors are required to efﬁciently and effectively utilize the available membrane volume. 4.7.1 Membrane Chemistries

Membrane adsorbers are commercially available in a variety of chemistries ranging from standard ion-exchange chemistries (strong and weak anion and cation exchangers) to hydrophobic interaction chemistries, reversed phase chemistries, and afﬁnity chemistries [61]. Membrane adsorber devices are traditionally available as single sheets, stacked disks, radial ﬂow systems, pleated devices, or hollow ﬁbers [62]. Device sizes range from small-volume devices containing <1 mL for scouting experiments and method development to devices containing several liters of media for large-scale preparative separations. Table 4.7 summarizes the currently commercially available membrane adsorber products. Sartorius’s Sartobind family of single-use disposable membrane adsorbers comprises 15 layers of a nominal 3mm cellulosic membrane [23]. Small-scale devices designed for method development employ a syringe ﬁlter format with a bed volume of approximately 2.1 mL. Larger manufacturing scale devices employ a radial ﬂow cartridge with membrane volumes ranging from 7 mL up to 540 mL. Sartorius also has a family of multiple-use membrane chromatography radial ﬂow

4.7 Membrane Adsorbers Tab. 4.7

Commercially available membrane adsorbers.

Company

Membrane

Available chemistries

Small-scale device

Large-scale device

Sartorius AG Nominal 3 mm (www.sartorius.com) cellulosic Pall Corporation (www.pall.com)

Ion-exchange 15-layer stack 15-layer afﬁnity activated (2.1 mL) radial ﬂow specialized (7–2100 mL) Nominal 0.8 nm Ion-exchange 16-layer stack 16-layer pleated polyethersulfone (0.35 mL) capsule (5–5000 mL)

units with device volumes ranging from several milliliter to 2100 mL. These devices can also be run in series or parallel to achieve the necessary performance. Pall’s Mustang family of disposable membrane adsorbers comprises 16 layers of a nominal 0.8mm polyethersulfone membrane [22]. For method development, small 16-layer coins (0.35 mL bed volume) contained in a stainless-steel housing are available. Manufacturing scale devices comprise a 16-layer pleated capsule with membrane volumes ranging from 5 mL to 5000 mL. To process even larger volumes of ﬂuid, these devices can also be conﬁgured in series. Although Sartorius AG and Pall Corporation are currently the only two companies with commercially available membrane adsorbers, due to the relative infancy of membrane adsorber technology, we can expect new competitors and new technologies to continuously emerge. 4.7.2 Current Applications

Early work in the area of membrane chromatography was primarily focused on the bind–elute puriﬁcation of various proteins [62–66]. However, recent advances in the theoretical and experimental understanding of the performance of membrane chromatography has helped to focus the current applications of membrane adsorbers to primarily two areas – ﬂow-through applications aimed at removing trace impurities and bind–elute puriﬁcation of extremely large molecules essentially excluded from the internal pore structure of commercially available beadbased resins. Examples of ﬂow-through applications include polishing applications aimed at removing low levels of nucleic acid, host cell protein, endotoxin, virus, and product aggregates and pre-capture applications aimed at removing impurities that either foul the Protein A column or create product stability issues upon product elution. Examples of large-molecule bind–elute puriﬁcation include gene therapy vectors such as viral vectors and plasmids, as well as extremely large protein molecules. 4.7.2.1 Flow-Through Polishing Due to the lower binding capacity of membrane adsorbers compared to bead-based resins, universal adoption of membrane adsorber technology has been slow, even

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though the high ﬂux advantages provided by membrane adsorbers would lead to higher productivity [67]. For trace-impurity removal applications, adsorptive capacity is not a signiﬁcant limitation. Rather, the adsorptive binding capacity of bead-based columns used in this application are typically three to four orders of magnitude larger than required since the columns are normally sized to achieve a desired ﬂow rate (i.e., the process is throughput limited). Since membrane-based systems have a distinct ﬂow rate advantage and sufﬁcient capacity for binding trace levels of impurities and contaminants, membrane adsorbers are ideally suited for this application. Flow-through anion-exchange membrane adsorbers have proven to be an extremely powerful polishing step designed to remove trace impurities from monoclonal antibodies solutions. Under neutral pH conditions, many viruses, nucleic acids, endotoxins, and host cell proteins are negatively charged and will bind strongly to an anion-exchange membrane. In contrast, many monoclonal antibodies are typically basic in nature (pI value >7), and, as such, will not interact with an anion-exchange membrane adsorber. This difference in charge has been successfully exploited to effect the separation of various trace impurities from monoclonal antibodies [68–70]. Reported LRVs for endotoxin (4 LRV), nucleic acids (6 LRV), and several mammalian viruses (2–6 LRV) are consistent with those obtained with standard bead-based resins. To highlight the performance advantages of membrane adsorbers, Phillips et al. has shown that for efﬁciently designed membrane adsorbers, 6 LRV clearance of virus can be achieved at residence times less than 0.4 seconds [69]. Residence times on the order of several minutes are typically required to achieve this degree of separation with bead-based chromatographic systems. Zhou et al. [70] have conducted an economic analysis and concluded that singleuse disposable membrane adsorbers can be an economically viable alternative to standard bead-based separations. Although the cost of single-use membrane adsorber media is typically higher than bead-based media that is reused for hundreds of cycles, the use of membrane adsorbers often result in lower hardware costs, signiﬁcantly lower buffer usage, and savings in validation studies (column packing, column reuse). Primarily due to savings in buffer, Zhou et al. conclude that membrane adsorber technology is economically preferred. 4.7.2.2 Flow-Through Precapture Protein A column chromatography is routinely used for the capture and puriﬁcation of monoclonal antibodies from cell culture harvest streams. Product elution from Protein A columns is generally done at low-pH conditions where the possibility of coeluted impurities precipitating is real. The consequences of this precipitation may include clogging of downstream sterilizing-grade ﬁlters and fouling of the Protein A column, both of which are detrimental to the robustness and process economics of the downstream puriﬁcation. Adsorptive-based removal of these impurities prior to Protein A column chromatography has been shown to minimize the possibility of precipitation during product elution. Shukla et al. [71] have exploited the adsorptive properties of depth ﬁlters and

4.7 Membrane Adsorbers

anion-exchange resins prior to Protein A capture chromatography to minimize the extent of precipitation. Lepore et al. [72] have shown that the use of anion-exchange membrane adsorbers also work well for this application, with signiﬁcantly higher loadings compared to standard bead-based resins. Lepore et al. also conducted an economic analysis indicating that the use of membrane adsorbers is economically advantageous. 4.7.2.3 Large Molecule Bind–Elute Puriﬁcation For very large molecules, bead-based chromatography resins have been shown to display very low dynamic capacities that decrease signiﬁcantly with increasing linear velocity [73]. The low dynamic binding capacities are mostly attributable to the inaccessibility of the resin pores to the larger molecules. The dynamic capacity for membrane adsorbers has been shown to be both signiﬁcantly larger than highly porous bead-based resins and essentially independent of ﬂow velocity. These qualities make membrane adsorbers ideally suited for the puriﬁcation of very large molecules. The puriﬁcation of plasmids and viral vectors for gene therapy applications are uniquely suited to take advantage of the properties of membrane adsorbers. Pora [74] describes the use of a Q membrane adsorber for the capture and puriﬁcation of plasmids. Using a 260 mL membrane adsorber, approximately 71 L of clariﬁed lysate containing 1.5 g pDNA was processed. The step yield was approximately 95 % and the total cycle time was 24 minutes. The use of bead-based resins most likely would have required signiﬁcantly higher bed volumes and much higher cycle times to effect this separation. Han et al. have successfully used various ion-exchange membranes for the adsorption of Aedes aegypti densonucleosis virus (a mosquito speciﬁc parvovirus with pI around 5.6) [75]). They conclude that membrane adsorbers may be ideally suited for virus capture since nearly all of the ligands are available for interaction. 4.7.3 Future Trends

Membrane adsorbers will continue to ﬁnd niche markets that are capable of exploiting the inherent advantages of membrane chromatography compared to bead chromatography – namely high efﬁciency due to minimal mass transfer effects (fast separations), large external surface areas (high-binding capacities for very large molecules currently excluded from the porous structure of bead-based resins), and pre-packed disposable devices (ease-of-use considerations). Anion-exchange membrane adsorbers have been successfully employed for the efﬁcient removal of several impurities from monoclonal antibody streams. Other applications that require the removal of trace level of impurities from a product stream can obviously exploit the advantages of membrane chromatography. To effect these separations, however, may require the development of novel membrane chemistries capable of achieving the desired selectivity between the impurity and product molecule. It is quite likely that new membrane adsorber chemistries will be developed as these new polishing applications are identiﬁed.

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Currently, one of the most visible trends in biotechnology manufacturing is disposability. As entire downstream puriﬁcation trains begin to migrate toward complete disposability, membrane chromatography begins to play a more important role. Although complete disposability may never be a reality on the manufacturing scale, the ﬂexibility and rapid change-out capabilities associated with disposable manufacturing may be ideally suited for early phase manufacturing of clinical product and contract manufacturing organizations. Additionally, disposable manufacturing may be a requirement in the area of personalized medicines where minuscule batches are manufactured for the treatment of a single patient. In these instances, membrane chromatography would play a critical role in the puriﬁcation of these products. Finally, a focus on membrane adsorber device conﬁguration could have an important effect on the future of membrane chromatography. Historically, membrane adsorbers were incorporated into traditional ﬁlter conﬁgurations. Although this was ﬁne for polishing applications, the ﬂow distribution and hold-up volumes were not properly designed to make membrane adsorbers competitive in the bind– elute puriﬁcation markets. Improvements in these areas could signiﬁcantly increase the potential markets available for membrane chromatography.

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Filter and method of producing same. US Patent 1,421,341. Wrasidlo, W. J. Asymmetric membranes. US Patent 4,629,563. Kraus, M. et al. Filtration membranes and method of making same. US Patent 4,900,449. Roesink, H. D. W. et al. Process for the preparation of hydrophilic membranes and such membranes. US Patent 4,798,847. Walch, A. et al. Microporous asymmetrical hydrophilic membrane made of a synthetic polymer. US Patent 4,720,343. Sasaki, J. et al. Asymmetric microporous membrane containing a layer of minimum pores below the surface thereof. US Patent 4,933,081. Ulbricht, M. (2006) Advanced functional polymer membranes. Polymer, 47, 2217–2262. Guidance for Industry Sterile Drug Products Produced by Aseptic Processing – Current Good Manufacturing Practice, U.S.

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Department of Health and Human Services, Food and Drug Administration September 2004. FDA report. Pharmaceutical cGMPs for the 21st Century – a risk based approach, ﬁnal report, September 2004 (www.fda.goc/cder/gmp/ gmp2004/GMP_ﬁnal report2004.htm). Blanchard, M. (2007) Quantifying sterile membrane retention performance. BioProcess International 5 (5) 44–51. Ho, C. and Zydney, A. L. (2000) A combined pore blockage and cake ﬁltration model for protein fouling during microﬁltration. Journal of Colloid and Interface Science, 232 (2), 389–399. Bolton, G.,La Casse, D., Kuriyel, R. (2006) Combined models of membrane fouling: development and application to microporous and ultraﬁltration membranes. Journal of Membrane Science, 277 (1-2), 75–84. Pitt, A. M. (March 1987) The Nonspeciﬁc Protein binding of

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Polymeric Microporous Membranes. Journal of Parenteral Science and Technology. Bin, T., McCrosky, L., Kulshreshtha, A. K., Hem, S. L. (2000) Adsorption of esters of p-hydroxybenzoic acid by ﬁlter membranes: mechanism and effect of formulation and processing parameters. Pharmaceutical Development and Technology, 5 (1), 95–104. ANSI/AAMI/ISO 11137–1994/5, Sterilization of health care products – Requirements for validation and routine control – Radiation sterilization. PDA Technical Report No. 26 – Sterilizing Filtration of Liquids, 1998. Scott, F. E. (1989) Principles of integrity testing hydrophilic microporous membrane ﬁlters, parts 1 and 2. Pharm. Tech. Badmington, F., Wilkins, R., Payne, M., Honig, E. S. (1995) Vmax testing for practical microﬁltration train scale-up in biopharmaceutical processing. BioPharm, 8 (7). Ghosh, R. (2003) Protein Bioseparation Using Ultraﬁltration: Theory, Applications and New Developments, Imperial College Press, London. http://www1.amershambiosciences. com/APTRIX/upp00919.nsf/content/ ProteinConc_HomePage. http://www.millipore.com/ eHandbook.nsf/docs/ufhome. http://www.pall.com/ datasheet_biopharm_ 33901.asp. http://www.sartorius.com/ index.php?id=845. Tkacik, G. and Michaels, S. (1991) A rejection proﬁle test for ultraﬁltration membranes and devices. Bio/ Technology, 9, 941. Carter, J. and Lutz, H. (2002) An overview of viral ﬁltration in biopharmaceutical manufacturing. European Journal of Parenteral Science, 7 (3), 72–78.

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5 Membrane Applications in Red and White Biotechnology Stephan Lu¨tz, Nagaraj Rao 5.1 Introduction

In living systems, membranes and membrane processes play a crucial role in the survival, growth, metabolism, defense, and reproduction processes. As our understanding of the role played by membranes gradually increases, the development of new application areas becomes a logical consequence. Several of these newer areas have been covered in other chapters of this book. In this chapter, we will focus on the application of membranes in red and white biotechnology. ‘‘Red’’ biotechnology refers to medical applications of biotechnology, starting from diagnostics and ending with therapy. It also covers the biotechnological production of pharmaceuticals and diagnostics. ‘‘White’’ biotechnology, by deﬁnition, is the application of nature’s toolset to industrial production. Membranes are made up of natural materials (such as tissues) or synthetic materials (such as certain polymers). They are permeable to certain substrates in solution. The movement of molecules across a membrane is regulated in both directions, giving the membrane its unique properties and wide applicability. In Figure 5.1, commonly used ﬁltration processes in biotechnology and the retention properties of membranes are shown. In reverse osmosis, particles, macromolecules, and low molecular mass compounds such as salts and sugars are separated from a solvent, usually water. The feed solution often has high osmotic pressure, and this must be overcome by the hydrostatic pressure applied as the driving force. Thus, microﬁltration, ultraﬁltration, nanoﬁltration, and reverse osmosis differ from each other in the size of the particles being separated. Interestingly, membrane processes are ﬁnding application in biotechnology at practically every stage of production [1–3]. This includes sterile dosage of substrates into the bioreactor, bubble-free aeration, and retention of biocatalysts using a variety of techniques, as well as product concentration and recovery based on different unit operations. These membrane processes are shown schematically in Figure 5.2.

Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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Fig. 5.1 Types of ﬁltration and retention properties of membranes.

5.2 Types of Membrane Processes in Red and White Biotechnology 5.2.1 Bubble-Free Aeration

In the case of biotransformations where aeration is required for the supply of oxygen, gas sparging is the preferred method because of the large mass-transfer rates and operational simplicity. However, gas sparging can damage cultured animal cells, since they are more sensitive to shear stress caused by vigorous mixing and gas sparging. In such cases, bubble-free aeration is achieved by the use of silicon membrane tubing, which allows the gas to diffuse into the medium without the

Fig. 5.2 Schematic representation of membrane processes used in biotechnology.

5.2 Types of Membrane Processes in Red and White Biotechnology

formation of bubbles. Another advantage of such tubing is the fact that its permeability for carbon dioxide is greater than that for oxygen, because of which there is no accumulation of carbon dioxide in the medium. 5.2.2 Filtration Processes

In the case of membrane processes using ﬁltration techniques, the pore size of the membranes plays a deciding factor in contributing to the efﬁciency of the process. With the advent of membranes with fairly well-deﬁned pore sizes, it is now possible to carry out separation processes for the retention of biocatalysts, cofactors, salts, and solvents. The importance of biocatalyst retention on the economics of a biotransformation can be judged from Figure 5.3. When the retention factor R is only 95 %, there is a rapid decrease in the relative concentration of the biocatalyst over a certain number of residence times. When the retention factor is greater than 99.99 %, there is no signiﬁcant loss in the biocatalyst concentration in the reactor, and a very large number of residence times can be achieved in the continuous biotransformation process. 5.2.3 Dialysis and Electrodialysis

In the case of dialysis, one or more solutes are transferred from one solution, called the ‘‘feed,’’ to another solution, called the ‘‘dialysate,’’ through a membrane down their concentration gradient. When pressure is employed besides the concentration gradient for separation, the process is called pervaporation.

Fig. 5.3 Importance of high retention of biocatalyst on membranes.

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In electrodialysis, the separation of components of an ionic solution occurs in a cell consisting of a series of anion- and cation-exchange membranes. These are arranged in an alternate manner between an anode and a cathode to form individual electrodialysis cells. During the process of electrodialysis, there is an increase in the ion concentration of one type in one type of compartment and is accompanied by a simultaneous decrease in the concentration in the other type of compartment. 5.2.4 Adsorption of Microorganisms

In adsorption, material accumulates on the surface of a solid adsorbent having a multiplicity of pores of different sizes. Adsorption is a common operation employed for the puriﬁcation of biological products. Various physical, chemical, or physicochemical forces may be involved in the adsorption phenomenon. Activated carbon and ion-exchange resins may be used, depending on the nature of the material to be adsorbed. In ion-exchange adsorption, the adsorbents have ionic groups with easily dissociable counter-ions. In afﬁnity adsorption, proteins may be absorbed biospeciﬁcally on the basis of their interaction with a complimentary tertiary structure. Microorganisms and enzymes are also adsorbed onto inert surfaces in order to carry out biotransformations.

5.3 Examples of Membrane Processes in Biotechnology 5.3.1 Bubble-Free Gassing 5.3.1.1 Hydrogen Several naturally occurring enzymes use hydrogen as a substrate. For example, the hydrogenase I obtained from the hyperthermophilic archaeon Pyrococcus furiosus (PfH2ase), a microorganism found in the Volcano regions of Italy, can split hydrogen heterolytically and catalyze regio- and enantioselective hydrogenation. The limiting factors for this process are the solubility of hydrogen and the hydrogen transfer rate. In nature, the solubility of a gas is increased by an enormous increase in the membrane surface, as exempliﬁed by the bronchi, bronchioles, and alveoli of the lungs. Based on these principles, polymeric membranes are employed to introduce a gas into a bioreactor. The polymeric membrane functions as a gas distributor and is able to control the gas–liquid interface area and the mass-transfer coefﬁcient independently of each other [4]. The pressure can also be divided into two partial pressures, one for the gas phase and one for the liquid phase. For the regeneration of cofactors, sources of hydrogen that have been studied in biotransformations include alcohols such as isopropanol, sugars such as glucose6-sulfate, formate, and molecular hydrogen. Continuous cofactor reduction using biotransformations with molecular hydrogen as the source of the reducing agent is

5.3 Examples of Membrane Processes in Biotechnology

Fig. 5.4 Bubble-free H2-aeration in an enzyme membrane reactor for continuous reduction of cofactor NADPþ.

possible when the biocatalyst accepts hydrogen as a substrate. The enzyme PfH2ase has been used in batch and continuous experiments in order to reduce the cofactor NADP+ to NADPH (Figure 5.4) [5,6]. Mertens et al. have described the potential applications of PfH2ase and similar hydrogen-activating hydrogenases for the production of NADPH, biosensors, and biofuel cells [7]. 5.3.1.2 Oxygen The bubble-free introduction of oxygen into a bioreactor is usually done with the help of silicone tubings that pass through the bioreactor. Mammalian cells have been immobilized on macroporous carriers made from small glass beads of up to 0.7 mm diameter. These are then suspended and cultivated in a continuously operated ﬂuidized bed reactor. The introduction of oxygen occurs via a silicone tube that passes through the reactor (Figure 5.5). As a result, the entire length of the ﬂuidized bed reactor is provided with oxygen, irrespective of the axial position of the reactor. Depending on the oxygen transfer rate and the oxygen consumption rate, the length of the tube is calculated. Efﬁcient mixing of the gas in the liquid helps in reducing the radial oxygen gradient. Maximum cell densities of up to 3.3–50 107 cells/mL have been achieved using this principle. Pharmaceutically important proteins, including antibodies and immunoglobulins, can be manufactured using this method [8,9]. The cultivation of CHO-2DS cell lines in a protein-free medium with bubble-free aeration, accompanied by improved cell growth and better production of human prothrombin, has been reported. [10]. Bubble-free aeration was found to be advantageous. BHK-21 cells have been fermented on a technical scale in protein-free culture media in a membrane-airlift bioreactor [11]. The production of recombinant human interleukin 2 was studied both with and without bubble-free aeration. By modifying the SM1F7 culture medium by addition of the surfactant Pluronic F-68 (BASF), the tolerance of the cells toward hydrodynamic stress was increased. The defoamer did

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Fig. 5.5 Bubble-free introduction of oxygen through a silicone tube for cell culture.

not adversely affect the cultivation. The cultivation of the BHK-21 cells could be carried out on a 1000-L scale without bubble-free aeration. This implies that a decision has to be taken on a case-to-case basis if bubble-free aeration is a must or not. The human tumor antigen and type I transmembrane protein Mucin MUC1 plays an important role in the diagnosis of cancer cells. A continuous perfusion culture process by the fermentation of CHO-K1 cells, using bubble-free oxygenation, has been developed [12] . The optimum pO2 value was found to be at 40 % air saturation. The cell activity could be maintained at a level above 85 %. A space–time yield (STY) of 100 mg/L/day could be achieved for MUC1-IgG2a. The CHO-K1 cells could be transferred into the serumfree suspension culture with the help of the ProCHO4-CDM medium. Studies on the control of bubble size, which has a direct bearing on the performance of mammalian cells, have been carried out by Nehring et al. [13]. Keeping the bubble size small releases lesser energy when the bubbles are burst. Microbubbles with 100–500 mm diameter can be generated in an aqueous media by using specially designed hydrophilic materials such as porous ceramics. The ceramic gas bubbling system consisted of a porous ceramic pipe ﬁlled with titanium dioxide and ﬁxed on to a steel structure. An early bubble-free reactor for the cultivation of mammalian cells has been described by Schulze and Stahl [14]. This has been modiﬁed for use in the cultivation of the phototropic microalgae Scenedesmus communis by Gorenﬂo et al. Light energy was provided to the microorganisms by ﬁtting the bioreactor with a new system of irradiation of light. A cylinder made of stainless steel mesh was designed. The pore size was twice the size of the cells (10 mm in diameter). As a result, the cells were retained outside the cylinder and internal aeration was found to be optimal [15]. Bubble-free aeration often helps in stabilizing the biocatalyst. The extracellular laccase of the white-rot fungus Pyconporus cinnabarinus (DSM 15225) was immobilized on DEAE-Sephadex after cultivation. The bubbling of oxygen through a silicone

5.3 Examples of Membrane Processes in Biotechnology

tube in the enzyme membrane reactor used to carry out the biotransformation stabilized the enzyme in continuously run experiments [16]. 5.3.2 Membranes for Cell Retention 5.3.2.1 Higher Cells/Red Biotechnology The fermentation of higher cells, such as hematopoietic stem cells, poses more challenges than the cultivation of fungi and bacteria. This is because of the unique requirements of the higher cells, including sensitivity to shear stress and the nature of cell division. In an interesting study exploiting the physical characteristics of cells and imitating the naturally occurring hematopoietic microenvironment, Meissner et al. have shown that it is possible to replicate bone marrow stromal cells by a three-dimensional cocultivation in porous microcarriers in a ﬁxed-bed reactor [17]. A membraneseparated cocultivation method for the separation of hematopoietic and stromal cells was employed, instead of a conventional coculture in microcarriers. During the cocultivation, the two types of cells were separated with the help of a porous membrane. The membrane pore size allowed the diffusion of the secreting stromal growth factors as well as an intercellular contact through the membrane (Figure 5.6). As a result, various degrees of expansion were observed for different colonyforming and burst-forming cells. For very early progenitor cells (CFU-GEMM) and later progenitor cells (CFU-GM and BFU-E), expansion degrees of 4.2-fold, 7-fold, and 1.8-fold were observed. Using similar principles, cultivation of progenitor cells (colony-forming cells, CFC, and cobblestone area forming cells, CAFC) was found to be more efﬁcient. The population of the CAFCs, for example, could be increased 39 times in the membrane method described above, while only a sevenfold increase was observed in the suspension cultures [18]. The production of IgG2a monoclonal antibodies by hybridoma cells has been carried out using a perfusion culture system along with a ceramic membrane module. In this so-called stirred ceramic membrane reactor system, consisting of 10 vertical, cylindrical tubes with an active surface area of 400 cm2, the volumetric productivity increased sevenfold [19].

Fig. 5.6 Membrane-separated cocultivation of hematopoietic and stromal cells.

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5.3.2.2 Whole-Cell Biotransformation With the advent of simpler genetic engineering techniques, ‘‘tailor-made’’ microorganisms (so-called ‘‘designer bugs’’) are now being increasingly developed for speciﬁc biotransformations. Ernst et al. have expressed a regio- and enantioselective (R)-alcohol dehydrogenase from Lactobacillus brevis and a formate dehydrogenase from Mycobacterium vaccae N10 into Escherichia coli. This functional overexpression enabled the reduction of keto compounds into hydroxy compounds, for example, the reduction of methyl acetoacetate to (R)-3-hydroxybutanoate [20]. The speciﬁc cell productivity could be signiﬁcantly improved by the introduction of a recombinant cofactor regeneration system in the whole-cell biocatalyst. No byproduct formation was observed, in addition. Current trends indicate that such ‘‘designer bugs’’ will be increasingly used in biotransformations. The retention of microorganisms on membranes during biotransformation has been utilized to manufacture chiral diol, (2R,5R)-hexanediol. In this process, resting cells of Lactobacillus keﬁri DSM 20587 were retained over an ultraﬁltration membrane with a molecular weight cut-off (MWCO) of 400 kDa. The reduction of the diol was carried out at 30 8C under anaerobic conditions (Figure 5.7). The selectivity for the diol was 78 % and the enantio- and diastereoselectivity were found to be 99 %. A space–time yield of 64 g/L/day was achieved. Due to the high selectivity of the process, downstream processing and product puriﬁcation became much simpler [21]. By using genetically modiﬁed E. coli coexpressing genes of LbDH and FDH from M. vaccae, (R)-methyl-3-hydroxybutanoate with an enantiomeric excess of >99 % was obtained, starting from methyl acetoacetate in this whole-cell biotransformation. Reaction engineering enabled the lowest biocatalyst consumption in a continuous reactor (0.9 gWCW/g). In comparative studies with enzyme-coupled systems, it was concluded that reaction engineering techniques allow lower biocatalyst consumption in whole-cell bioreductions. Whole-cell immobilization further reduced the biocatalyst consumption [22]. In an unusual application, whole cells of Pseudomonas putida type A1 were retained in a hollow-ﬁber membrane reactor in order to degrade gas-phase toluene biocatalytically with a high oxygen availability. A porous, hydrophobic polyethylene membrane functioned as a carrier for the active bioﬁlm. Toluene and other components diffused through the lumen of the membrane into the aqueous phase and were biocatalytically oxidized in this phase containing nutrient and microorganism. Incoming toluene concentration between 86 and 97 % could be degraded when the incoming air had a toluene concentration between 0.85 and 4.3 kg/m3/day [23]. 5.3.3 Membranes for Enzyme Retention

The use of membranes for retention of enzymes in bioreactors forms a signiﬁcant part of red and white biotechnology. Since enzymes usually are large molecular

5.3 Examples of Membrane Processes in Biotechnology

Fig. 5.7 Manufacture of (2R,5R)-hexanediol with cell retention.

weight compounds, with molecular weights ranging from 10 to 150 kDa or more, membranes with deﬁned pore sizes (or MWCO) ﬁnd a wide range of applications for enzyme retention during the biotechnological production of ﬁne chemicals. A typical example is the industrial production of amino acids by the acylase process using enzyme membrane reactors by the German ﬁrm of Degussa. This process is in use for over two decades now. In addition to proteinogenic amino acids such as alanine, methionine (Figure 5.8), valine, and tryptophan, nonproteinogenic amino acids such as O-benzylserine, norleucine, and nor-valine are manufactured on an industrial scale. Starting from the racemic acetate mixture, only one enantiomer is hydrolyzed by the catalyst [24,25]. By establishing a direct phase contact in the pores of a hydrophobic, microporous membrane, a hydrophobic product can be extracted from the stream coming out of a bioreactor. Thus, an extraction step can be directly coupled to a biotransformation step, thereby enabling the removal of poorly soluble products and unreacted educts (Figure 5.9). Kruse et al. have used this method to couple an extraction step to an enzyme membrane reactor with continuous cofactor regeneration [26]. A commercially available hydrophobic membrane (Celgard membrane in a LiquiCel lab module) functions as the extractor and extracts the hydrophobic product into

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Fig. 5.8 Racemic resolution of N-acetyl-methionine using acylase.

the organic phase. The regeneration of the hydrophilic cofactor takes place in the aqueous phase unhindered. The residence time of the cofactor is decoupled from that of the substrate. Several ketones have been reduced enzymatically, based on these principles, to the corresponding, hydrophobic chiral alcohols. An NADþ-alcohol dehydrogenase from Rhodococcus erythropolis was employed. There was an increase in the total turnover number by a factor of 20–25. In the above example, an extraction step has been used after a biotransformation step to remove a hydrophobic product from the loop. In the case of substrates that are poorly soluble in aqueous media, attempts have been made to introduce the so-called

Fig. 5.9 Coupling of a membrane-based extraction step to a biotransformation process.

5.3 Examples of Membrane Processes in Biotechnology

emulsion membrane reactors prior to the biotransformation step. By using a hydrophilic ultraﬁltration membrane, it has been shown that an emulsion of 2-octanone can be separated into its phases and the aqueous phase is saturated with the organic substrate. This aqueous phase is fed into the enzyme membrane reactor where the actual biotransformation takes place. The enantioselective reduction of 2-octanone to (S)-2-octanol is shown in Figure 5.10. The biotransformation is catalyzed by a carbonyl reductase obtained from Candida parapsilosis (CPCR) and is accompanied by continuous reduction of the cofactor NADþ. By bringing the product stream in contact with the emulsion reactor, the product is extracted. Production of (S)-octanol to the extent of 21.2 g/L/day and with an enantiomeric excess of >99.5 % could be achieved over a period of four months in a continuous experiment [27]. A variety of methods are now available to modify the carrier surfaces for immobilization of biocatalysts. Similarly, the surface of membranes can be chemically modiﬁed in order to change its properties. Plasma-induced graft polymerization of poly(a-allyl glycoside) has been used by Deng et al. to increase the biocompatibility and hydrophilicity of a polypropylene hollow-ﬁber microﬁltration membrane. A two-phase membrane reactor using the lipase from Candida rugosa as the biocatalyst, immobilized on the membrane, was used for the hydrolysis of olive oil. Under optimized reaction conditions, a volumetric reaction rate of 0.074 mm/L/h could be attained [28]. The importance of chiral amines as the building blocks to manufacture pharmaceuticals and agrochemicals is growing. Transaminases are being considered as

Fig. 5.10 Coupling of an emulsion membrane reactor with an enzyme membrane reactor for biotransformation of poorly soluble substrates.

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potential biocatalysts for their synthesis. In one type of biotransformation, however, the reaction is accompanied by the oxidation of one of the enantiomers to the corresponding ketone, which causes strong product inhibition. To overcome this problem, the enzyme membrane reactor has been coupled with a hollow-ﬁber membrane contactor. The racemic substrates a-methyl benzylamine and 1-amino tetraline yielded the products with enantiomeric purities of 98 and 95.5 %, respectively, when the (S)-speciﬁc v-transaminase from Vibria ﬂusialis JS17 or Bacillus thuringensis JS64 was used [29]. Similarly, hollow-ﬁber contactors have been used for the production of enantiomerically pure monoesters from meso diesters. A pig liver esterase was used, and the polysulfone ultraﬁltration hollow-ﬁber module had a MWCO value of 30 kDa. The asymmetrical pores of the ﬁbers served to immobilize the enzyme. The reaction and the separation steps occur simultaneously in this module. The loss of enzyme activity was found to be low [30]. Baccatin III is a precursor in the synthesis of the antitumor compound paclitaxel. It has been continuously produced in an enzyme membrane reactor, starting from 10-deacetyl baccatin III by using an acyl transferase of plant origin, overexpressed in E. coli. The substrate is obtained by methanolic extraction of the European yew needles or cuts. The product was selectively separated from the biocatalyst in the bioreactor and extracted in the solvent phase, made up of diisopropyl ether present in an integrated module. The equilibrium of the reaction is shifted favorably by using this method. A polysulfone membrane with a MWCO of 10 kDa was used. It was observed that a reduction in the biotransformation temperature from 15 to 10 8C improved the biocatalytic stability [31]. This example demonstrates once again the importance of membrane processes in red biotechnology applications. Miniaturization of the enzyme membrane reactor has led to the development of micro enzyme membrane reactors with volumes of <200 mL. Such reactors function as practical tools for education, research, and product development. The quantities of biocatalyst and chemicals required for the experiments are signiﬁcantly reduced [32]. 5.3.4 Membranes for Cofactor Retention

In many cases, the catalytic activity of an enzyme comes into force only in the presence of a cofactor, often called a coenzyme. While metallic cofactors are cheap to obtain, cofactors such as NADþ and NADPþ are extremely expensive. For this reason, cofactor regeneration plays a crucial role in the economics of a biotransformation. Wichmann and Vasic-Racki have reviewed the advantages and disadvantages of the different strategies of cofactor regeneration [33]. Closely related to cofactor regeneration is the retention of the cofactor in the biocatalytic reactor, so that it is not subjected to the unit operations that the product stream has to undergo subsequently. Cofactor retention by the use of membranes has been studied. Thus, for example, in the continuous reduction of trimethyl pyruvate to L-tert-leucine, catalyzed by the enzyme leucine dehydrogenase, the total turnover number (ttn) of the cofactor NADH is dependent on the type of membrane employed.

5.3 Examples of Membrane Processes in Biotechnology

Fig. 5.11 Dependence of total turnover number of cofactor NADH on type of retention membrane.

When an ultraﬁltration membrane was used in an enzyme membrane reactor, the ttn was found to be 2325 (Figure 5.11). However, when a nanoﬁltration membrane was used, the ttn increased to 7920. Rosmarinicacid,anesterbetweencaffeicacidandachiraldihydroxyphenyllactic acid (‘‘hydroxy-DOPA’’), has been investigated intensively for its anti-inﬂammatory properties. Rosmarinicacid has been producedby planttissue culture in signiﬁcant quantities. Biotransformation reactions also allow the synthesis of the naturally occurring product. Thus, 3,4-dihydroxyphenyl pyruvic acid has been reduced to (S)-3,4-(dihydroxyphenyl) lactic acid using a lactate dehydrogenase and a formate dehydrogenase, the latter for continuous cofactor regeneration in a manner similar to the one described above. The (S)-3,4-(dihydroxyphenyl) lactic acid has been converted to rosmarinic acid in a second biotransformation using caffeoyl coenzyme A transferase [34,35]. The continuous regeneration of the cofactor NADPH during the Baeyer–Villiger oxidation of 4-methyl cyclohexanone to 5-methyloxepan-2-one has been described by Rissom et al. [36,37]. The product could also be synthesized in a repetitive batch mode (Figure 5.12). The chiral lactone had an enantiomeric purity of >99 % [38]. 5.3.5 Application of Dialysis and Electrodialysis in Biotransformations

White biotechnology offers an interesting alternative to the synthesis of pyruvates. Pyruvates are chemically synthesized at high reaction temperatures by oxidation reactions using heavy metal catalysts, starting from tartaric acid, propylene glycol, or lactic acid. In the case of biotransformations, a high extracellular pyruvate concentration greater than 500 mmol/L causes an inhibition of the microbial pyruvate synthesis. Therefore, the product has to be separated, as soon as it is formed, in an integrated process. This has been achieved by the integration of an electrodialysis step by Zelic et al. [39]. The fermentation using the designer bug E. coli YYC2O2

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5 Membrane Applications in Red and White Biotechnology

Fig. 5.12 Monooxygenase catalyzed Baeyer–Villiger oxidation.

Idh2A:Kan yields pyruvate in a concentration of more than 900 mmol/L under optimized conditions. Amino acids and peptides are interesting product groups for investigations using electrodialysis, because of their charged nature at different pH values except the isoelectric point. By combining electrodialysis with ultraﬁltration, acid and base bioactive peptides have been separated simultaneously by Poulin et al. [40]. Peptides from a b-lactoglobulin (b-lg) hydrolysate obtained from whey were separated by using an ultraﬁltration membrane stacked in an electrodialysis cell. The process, called electrodialysis with ultraﬁltration (EDUF), showed a high selectivity. Starting from 40 peptides present in the raw hydrolysate, only 13 were recovered in the separated adjacent solutions. In addition, among these 13 migrating peptides, three acidic– anionic peptides migrated only into one compartment, while three basic–cationic peptides migrated only into the other compartment, independent of the pH of the hydrolysate. The ACE-inhibitory peptide b-Ig 142–148 showed the highest migration value of 10.75 %. In addition to cationic and anionic membranes, a cellulose ester UF membrane with a MWCO of 20 kDa was used in the above experiments. 5.3.6 Application of Pervaporation and Stripping in Biotransformations

Substrate-coupled cofactor regeneration (e.g., the oxidation of 2-propanol to acetone) is interesting for NADPH-dependent enzymes as no efﬁcient enzyme-coupled

5.3 Examples of Membrane Processes in Biotechnology

system is yet available. Although higher conversions can be achieved by increasing the 2-propanol concentration, this also leads to the formation of more coproduct acetone. In such situations, it is essential to remove the coproduct as much as possible, preferably in situ, in order to achieve higher conversions. During the bioconversion of 5-oxo-hexanoic acid ethyl ester to the (S)-hydroxy ester, 2-propanol was reduced to acetone and was removed by employing stripping or pervaporation. The biotransformation was catalyzed by the NADPH-dependent enzyme carbonyl reductase from C. parapsilosis (Figure 5.13). Although both methods gave similar results with regard to product formation, pervaporation offers several advantages over stripping on a technical scale with regard to foaming and emission values [41]. The alcohol dehydrogenase from L. brevis (LbADH) is dependent on the cofactor NADPH for its catalytic activity. A two-phase bioreactor has been used to carry out the reduction of tert-butyl-6-chloro-3,5-dioxo-hexanoate (CDHE) to tert-butyl-(S)-4chloro-5-hydroxy-3-oxo-hexanoate (CHOE). During the biotransformation, a byproduct, tert-butyl-(4-oxo-4,5-dihydrofuran-2-yl)-acetate is formed due to the elimination of HCl in aqueous media. In order to avoid this, the CDHE is dissolved in methyl-tertbutyl ether (MTBE) and the enzyme and 2-propanol are dissolved in the buffer. The biphase emulsion reactor used gave the best results for the LbADH catalyzed reduction of CDHE with respect to selectivity, substrate concentration, ttn, and enzyme consumption, as well as downstream processing. Efﬁcient separation of acetone also improved the space–time yield [42].

Fig. 5.13 Pervaporation for acetone removal in a biotransformation process.

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5.3.7 Nanoﬁltration and Ultraﬁltration in Biotechnology

Several of the examples described in the previous sections employ either nanoﬁltration or ultraﬁltration either during the biotransformation step or for subsequent downstream processing. The protective action of neuraminic acid-containing glycoconjugates in human milk has been well documented. This has led to the synthesis of various molecules based on neuraminic acid for possible therapeutic applications. The performance of a diaﬁltration membrane process for desalting of the permeate stream during the enzymatic synthesis of CMP-Neu5Ac and GDP-Man has been studied. The nucleotide sugars could be retained by a nanoﬁltration membrane in a cross-ﬂow module. Only salts were allowed to pass through. The puriﬁcation steps yielded CMP-Neu5Ac in >95 % purity and >90 % yield. GDP-Man gave values of 95 and 88 %, respectively [43,44]. In another well-established application of red biotechnology, the industrial production of recombinant tissue-type plasminogen activator (rt-PA) and other proteins, starting from mammalian cells, employs tangential ﬂow ﬁltration using UF membranes for product isolation. Up to 5000 L/h of the medium can be processed with protein yields generally >99 %. Cell vitality and cell density were maintained during the process. Since the UF membrane had a pore size of only 0.2 mm, the ﬁltrate could be collected directly into containers after passing through a sterile ﬁlter. The latter could also be steamed in place. Scaling-up of the process followed a linear pattern [45]. Recombinant HIV-1 transcription–transactivator protein (Tat-protein) from a bacterial lysate could be separated easily by using a nylon membrane, modiﬁed chemically with an acyl anhydride and subsequent covalent coupling of avidin and biotin. In this afﬁnity microﬁltration technique, the product mTat primarily contained monomeric forms of the oligopeptidase sequence. While the membrane process led to a fourfold improvement in protein recovery, nonspeciﬁc protein adsorption was also observed [46]. 5.3.8 Bioelectrochemical Applications

Electrochemical oxidation, ultraﬁltration, extraction, and distillation have been combined in a process to carry out the oxidative separation of racemic 1-phenyl1,2-ethane diol. The cofactor NAD+ was regenerated by anodic oxidation of NADH using the mediator 2,20 -azinobis(3-ethyl-benzothiazolin-6-sulfonate (ABTS). An electrochemical, extractive enzyme membrane reactor was developed for this purpose [47]. In further studies, the cofactors NADþ and NADPþ have been electrochemically reduced by an indirect method, using (pentamethylcyclopentadienyl-2,20 -bipyriddine aqua)-rhodium(III) as the mediator. It was observed that for the reduction reaction, adsorption of the mediator on the ﬁxed-bed graphite cathode was essential. Stronger adsorption led to lower mediator consumption. Using a coated titanium mesh as the anode, values of ttn of up to 400 were achieved for the mediator. In order to hinder the reoxidation of the product, a cation exchange membrane was placed between the electrodes (Figure 5.14).

5.3 Examples of Membrane Processes in Biotechnology

Fig. 5.14 Indirect electrochemical cofactor reduction.

Selectivity and conversion values achieved were practically quantitative. The space–time yield was between 500 and 1000 g/L/day [48]. The ﬁrst asymmetric electroenzymatic oxidation of thioanisole to (R)-methylphenyl sulfoxide with an enantiomeric purity greater than 98.5 % has been carried out by Luetz et al. with in situ electrochemical generation of hydrogen peroxide [49]. The hydrogen peroxide functioned as the oxidizing agent for the enzyme chloroperoxidase from Caldariomyces fumago. A dialysis tube membrane separates the counter electrode (a platinum wire) from the working electrode (a carbon felt). The divided cell (Figure 5.15) gave productivity values as high as 30 g/L/day when the experiments were carried out on a 300 mL scale.

Fig. 5.15 Divided cell for electrochemical hydrogen peroxide generation and biotransformation.

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In situ regeneration of cofactors such as NADH using electroenzymatic pathways offer attractive, low-priced alternatives to the regeneration methods described in earlier sections of this chapter. Considerable efforts are under way to achieve this. The in situ electroenzymatic regeneration of NADH, during the formation of lactate from pyruvate, has been studied in ﬁxed-bed membrane reactors. The lipoamide dehydrogenase enzyme and the electron mediator methyl viologen were used. Porous graphite electrodes were coated with a cation exchange membrane. Substrate, product, enzyme lactate dehydrogenase, and coenzyme were present in the same solution. Yield of lactate improved considerably as a result of the electroenzymatic regeneration of NADH [50].

5.4 Summary From the wide variety of applications and principles described above, it is evident that membranes have a major role to play in industrial red and white biotechnology. This is further fuelled by the fact that properties of membranes can be modiﬁed in a better fashion now. In addition, a variety of polymeric materials are available for these applications. Mechanical and chemical stability can be improved to suit individual process requirements. The range of membrane-based applications in red and white biotechnology will continue to show an upward trend.

5.5 Acknowledgment

The authors are grateful to Prof. Dr C. Wandrey for ongoing support and discussions. References 1 Luetz, S., Rao, N., Wandrey, C. (2005)

2

3

4

5

Membranen in der Biotechnologie, Chemie Ingenieur Technik, 77, 1669. Liese, A., Seelbach, K., Wandrey, C. (2006) Industrial Biotransformations, 2nd edn , Wiley-VCH, Weinheim. Luetz, S., Rao, N., Wandrey, C. (2006) Membranes in Biotechnology, Chemical Engineering & Technology, 29, 1. Bommarius, A., Krimmer, H.-P. Reichert, D. et al. (2003) Volume gassing, German Patent DE 101 63 168 of 03.07.2003. Mertens, R., Greiner, L., van den Ban, E. C. D. et al. (2003) Practical application of hydrogenase I from Pyrococcus furiosus for NADPH generation and regeneration, Journal of Molecular Catalysis B: Enzymatic, 24/25, 39.

6 Mertens, R., Wandrey, C., Liese, A.

(2005) Reaktionstechnische Aspekte der enzymatischen Herstellung von NADPH im Enzym-Membran-Reaktor, Chemie Ingenieur Technik, 77, 609. 7 Mertens, R. and Liese, A. (2004) Biotechnological applications of hydrogenases, Current Opinion in Biotechnology, 15, 343. 8 Luellau, E., Dreisbach, E., Grogg, A. et al. (1992) Immobilization of animal cells on chemical modiﬁed siran carrier in Animal Cell Technology: Developments, Processes and Products, (eds. Spier, R. E.Grifﬁths, J. B. Macdonald, C.), Butterworth-Heinemann, London, 469. 9 Noll, T., Biselli, M., Wandrey, C. (1996) Wirbelschichtreaktor und BiomasseMonitor – ein leistungsfa¨higes

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6 Membranes in Controlled Release Nicholas A. Peppas, Kristy M. Wood, J. Brock Thomas 6.1 Introduction

Membranes have become increasingly important for use in separation processes including microﬁltration, ultraﬁltration, gas permeation, pervaporation, dialysis, and reverse osmosis. In the biomedical industry, synthetic polymer membranes are used in a variety of applications including hemodialysis, blood oxygenators, and controlled release of therapeutics [1,2]. Rose and Nelson [3] were responsible for the ﬁrst introduction of membranes for controlling drug release using osmotic pumps. Folkman and Long [4] developed silicon rubber controlled release delivery systems for anesthetics and cardiovascular treatments. With the creation of ALZA, membrane-based controlled release drug delivery systems (CR-DDS) became a commercial realization, creating both academic and industrial interest that energized the ﬁeld. The OROS oral drug delivery technology was originally introduced to the market by ALZA in 1973 as a gastrointestinal transport system. The OROS systems utilize osmotic pressure [5] in combination with membranes that control water permeation, drug release, or pushing force. OCUSERT was approved a year later for the control of elevated intraocular pressure. This CR-DDS utilizes a thin multilayer membrane device to control the release of the therapeutic pilocarpine at constant rate over a 7-day period. ALZA further reﬁned their use of membranes by developing transdermal drug delivery system of which the ﬁrst was approved by the FDA in 1981 for the prevention of nausea and vomiting associated with motion sickness. These drug delivery systems (DDS) employ a backing layer, a drug containing reservoir, a ratecontrolling membrane, and an adhesive layer that is capable of controlling rate of release for periods of 1 day up to 1 week. Membranes are utilized in CR-DDS to control the rate of permeation of a therapeutic effectively allowing the rate of release of the drug to be ﬁnely tuned according to the parameters associated with the polymer membrane [6]. The permeation of the drug species through the membrane has been described by two theories: solution-diffusion model and pore-ﬂow model. The solution-diffusion

Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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model is the most commonly accepted theory to describe the phenomena of CR-DDS employing polymer membranes. This model describes the drug molecules dissolution in the membrane and diffusion through the membrane due to a concentration gradient. The polymer parameters responsible for controlling the drug diffusion coefﬁcient are degree of crystallinity and size of crystallites, degree of crosslinking, degree of swelling, and molecular weight of the polymer. By tailoring the properties of the membrane at the molecular lever, one can effectively design materials that possess speciﬁc release kinetics [6,7]. The kinetics can be designed so as to provide a sustained release of drug delivery in the desired application of the membrane. This review evaluates the release kinetics of the drug species through membranes and describes how the various parameters associated with the polymer membrane affect this release. The traditional types of membrane-controlled drug delivery systems – diffusion-controlled, osmotically controlled, swelling-controlled, and chemically controlled [8] – and current applications of membranes in CR-DDS will also be discussed.

6.2 Controlled Release Kinetics

Dissolving or dispersing an active pharmaceutical ingredient (API), whether it is a small molecule, peptide, protein, or other bioactive agent, in a membrane is a method in which to control the release. Membranes can also act as rate-controlling layers covering reservoirs of therapeutic agents. During the development of these CR-DDS, it is necessary to develop and use mathematical models to describe the release kinetics. Although the models are based upon diffusion equations, they are often referred to kinetic models or expressions due to their time-dependent behavior of drug release. CR-DDS have been classiﬁed according to the physical mechanism that acts to control the release of the therapeutic. Based on the mechanism of transport, membrane-based DDS can be classiﬁed as diffusion-controlled, swelling-controlled, osmotically controlled, and chemically controlled systems. The mathematical models developed to describe the release kinetics [9] are used to (a) predict the rate of drug release from and the transport of drug through the membrane and (b) describe the mechanisms by which drug is transported. Mechanisms of the diffusion phenomena in these CR-DDS are accurately elucidated with expressions that relate a mathematical model and the physical parameters that describe the membrane being used for controlled release. 6.2.1 Diffusion in Membrane-Controlled Release

Two forms of Fick’s law of diffusion are often employed to describe drug release. The Equations (1) and (2) are for one-dimensional diffusion where ci and ji are the

6.2 Controlled Release Kinetics

concentration and mass ﬂux of drug i, respectively; x and t are position and time of release; and Dip is the drug diffusion coefﬁcient through the membrane: ji ¼ Dip

dci ; dx

ð1Þ

@ci @ 2 ci ¼ Dip : @t @x 2

ð2Þ

Certain assumptions are made to describe the release kinetics using Equations (1) and (2), and these include (a) that the geometry of interest is in the form of a thin, planar system, (b) that the diffusion coefﬁcient is independent of drug concentration, and (c) that ji is the drug ﬂux with respect to the mass average velocity v. Solutions to the Fickian diffusion equations are obtained when sufﬁcient knowledge of the initial and boundary conditions pertaining to the controlled release environment are provided [9]. These solutions allow for the determination of concentration proﬁles from the normalized drug concentration, (c/co), versus dimensionless position, (x/d), as a function of dimensionless Fourier time, (Dipt/d2), where co and d are the initial concentration and slab thickness, respectively. One can also evaluate the drug release rate, (dMt/Adt), by differentiating the above equations with respect to position and then evaluating the derivative at the membrane–dissolution medium interface (Equation (3). dMt @ci : ¼ Dip Adt @x x¼interface

ð3Þ

Finally, the total amount of drug release per cross-sectional area, (Mt/A), is given by integrating the above expression over the experimental release time. Mt ¼ A

ðt

dMt dt ¼ 0 Adt

ðt D 0

@ci @x

:

ð4Þ

x¼interface

Solute transport through membranes has been extensively studied in recent years. Much of the work, which characterized the polymer structure and the diffusing solutes, has led to theories used to describe solute transport. In the area of diffusion of drugs and proteins, am Ende et al. [10] studied the solute transport through ionic hydrogels as a function of mesh size and environmental conditions (i.e., pH and ionic strength) and determined that each factor plays a very important role in solute transport. They also concluded that hydrogels may be tailor-made for a release of a speciﬁc drug, protein, or peptide. Other studies showing the effect of pH on drug transport from ionized hydrogels were done by Brannon-Peppas and Peppas [11]. They found that pH-dependent hydrogels could be prepared to exhibit zero-order or near zero-order release; this behavior could be attributed to the effect of the pH on the relaxation, swelling, and release mechanism of the hydrogel.

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In biological systems, another important factor is polymer–solute interactions when both the polymer and solute are ionized. Collins and Ramirez [12] studied these interactions and determined that polymer–solute interactions decrease the transport of the solute. Gudeman and Peppas [13] also studied these interactions using wellcharacterized interpenetrating networks of PVA and PAA by varying the content of PAA (the ionic component) in the membrane. Some of the early work in solute transport through hydrogels was done by Renkin [14] who studied solute diffusion through porous cellulose membranes based on Fick’s law: dN dc ¼ DA : dt dx

ð5Þ

Here dN/dt is the solute diffusion rate, D is the solute diffusion coefﬁcient in the solvent, A is the apparent diffusion area, and dc/dx is the concentration gradient across the membrane. Renkin [14] performed diffusion experiments and measured the rate of diffusion for a variety of solutes through inert, porous membranes. He compared his experimental results to predictions based on the theory proposed by Pappenheimer and collaborators [15] and found that they were in close agreement. He concluded that solute diffusion was a function of pore and solute size. Yasuda et al. [16,17] studied the relationship between salt rejection and water ﬂux of nonionic [16] and ionic [17] membranes using reverse osmosis. In the nonionic study [16], they prepared a variety of membranes and monitored the water ﬂux and salt rejection until equilibrium was reached. They found that the size and solubility of ions in the membrane were responsible for transport depletion, related water permeation, v, and salt rejection, Rs, in the following manner: Rs ¼

P2 RT vþ ðD p Dsˇ Þ P1 y 1

1 :

ð6Þ

Here, P1 is water permeability, P2 is the salt permeability, y1 is the molar volume of water, and ðD p Dsˇ Þ is the effective pressure. The difference in this study and the ionic study is that ionic polymers were used in the study and that there is a different relationship between Rs and K1. Equation (7) gives the relationship for ionic membranes. K1 ¼ A expfBRS g

ð7Þ

Here, A and B are constants, and the equation is independent of ionic charge and nature or morphology of the membrane. Yasuda et al. [16] found that the principle behind the salt rejection of ionic membranes is the difference in the transport volumes for mobile co-ions and water. The repulsive forces between a ﬁxed ion and a mobile co-ion decreased the transport volume, creating a transport depletion of salt ﬂux relative to water transport.

6.2 Controlled Release Kinetics

Quinn and Anderson [18] studied hydrodynamic equations governing transport in microporous systems (r 1 mm) accounting for Brownian motion and steric restrictions. They showed that a one-dimension diffusion–convection analysis could be used for such systems and developed a series of equations to account for the effect of the pore wall on the solute–solvent drag. Peppas and Reinhart [19] developed a theory based upon the free volume theory for a three-component system (water, solute, and polymer). It predicted the dependence of the solute diffusion coefﬁcient on solute size, mesh size, degree of swelling, and other structural characteristics of the hydrogels. Equation (8) is the Peppas–Reinhart equation: DSM MC MC k2 rs2 ¼ k1 exp : DSW Qm1 Mn MC

ð8Þ

Here, DSM and DSW are the diffusion coefﬁcients of the solutes in the membrane and water, respectively. The ratio of the two is known as the normalized diffusion coefﬁcient. Also MC is the molecular weight between crosslinks, k1 and k2 are structural parameters of the polymer–water system, MC is the critical molecular weight between crosslinks at which diffusion could not occur, Mn is the molecular weight of the polymer before crosslinking, rs is the Stokes hydrodynamic radius of the solute, and Qm is the degree of swelling of the membrane. A very important criterion for this theory is that the diffusion occurs through highly swollen membranes. Using well-characterized, amorphous PVA membranes [20], they were able to validate their theory using experimental data. Prausnitz and collaborators [21] used Monte Carlo simulations to develop a modiﬁed size exclusion theory based on statistical distribution of chains in the network. However, the theory does not consider ionic interactions or the effects of side groups on the structure, but focuses on chains in a large region of space. The intention of their theory is to provide a general understanding of partially ionized polyelectrolytes where other theories may be built upon it. The ﬁelds of membrane-based controlled release, which are three-dimensional crosslinked polymer networks capable of imbibing signiﬁcant amounts of water, have been widely used in such applications because of their biocompatibility with the human body and because many hydrogels exhibit characteristics similar to natural tissue. Selection of membranes used in such processes depends on the characteristics of the polymer and the solute. Membranes have several important characteristics that play an important role in solute diffusion. These include ionization of the membrane, degree of swelling, and speciﬁc mesh or pore size. Membranes have functional groups along the polymer chain that react to the external environment (e.g., temperature, ionic strength, and pH of the swelling agent). The response of the hydrogel may be a reaction or increase in the mesh size of the hydrogel. For this work, the increase in mesh size is studied as a function of pH. The degree of swelling is also a very important parameter because it describes the amount of water that is contained within the hydrogel at equilibrium and is a function of the

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network structure, crosslinking ratio, hydrophilicity, and ionization of the functional groups. It is calculated from swelling studies and can be used to determine the molecular weight between crosslinks and the mesh size of the hydrogels. The mesh or pore size is the space available for transport. Generally, the mesh size or pores can lead to classiﬁcation of hydrogels as nonporous, microporous, or macroporous. Nonporous hydrogels have a mesh size between 20 and 100 A in diameter. Microporous hydrogels have a pore size between 0.01 and 0.1 mm, where transport occurs by a combination of diffusion and convection. Macroporous hydrogels have pore size greater than 0.1 mm; transport may occur by convection. These pore sizes determine the size of the solute permitted to diffuse through the membrane. The characteristics of the solute are as important as those of the membrane. The size, shape, and ionization of the solute affect its diffusion through the membrane. In the case of ionization, if the membrane and the solute are ionized, interactions may occur that may hinder or assist in the diffusional process, depending on the charges on the membrane and solute. If the charges are the same, the membrane repels the charges of the solute and does not hinder, and in some case assist, transport. If the charges are opposite each other, interactions between the membrane and the solute may take place, hindering transport. 6.2.2 Physical Parameters of Controlling Release

Membranes are prepared from uncrosslinked or crosslinked, hydrophobic or hydrophilic polymers that can be moderately or completely swollen in water or in buffered or physiological solution. These networks may be produced by a chemical reaction of monomers or by physical entanglements of chains. They may contain functional groups such as amino, hydroxyl, or carboxyl groups along the polymer chain. Incorporation of such functional groups increases the hydrophilicity of the membrane leading to a hydrogel. Drug release from the polymer membranes can be controlled and tailored by adjusting several different physical parameters of the polymer including, but not limited to, molecular weight, crystallinity, crosslinking, porosity, swelling, and branching.

6.3 Membranes and Solute Transport 6.3.1 Characterization of Membranes

In the presence of water, the functional groups become ionized and cause the hydrogel to swell due to charge repulsion. This swelling process may be described in terms of the equilibrium swelling ratio, Q, which increases as the swelling of the network increases.

6.3 Membranes and Solute Transport

Fig. 6.1 Equilibrium degree of swelling of ionic hydrogels as a function of solvent pH.

The swelling behavior of the hydrogel is affected by the pH of the swelling solution as depicted in Figure 6.1. Upon ionization, the charges repel each other causing the network to swell. In ionic hydrogels, swelling occurs when there is an increase in pH from a value below to a value above the pKa of the network. For cationic hydrogels, swelling occurs when there is a decrease in pH from a value above the pKa to a value below the pKa of the network (Figure 6.2). In addition to pH, the concentration of the functional groups on the network and the ionic strength of the swelling agent affect the swelling of hydrogels. By increasing the concentration of the functional groups, the swelling of the network increases. Increasing the ionic strength of the swelling agent decreases swelling in the network. As the ionic strength increases, the ions of the swelling agent counterbalance the mutual repulsion of the functional groups in the membranes. This behavior was observed by Khare and Peppas [22] in studying poly(2-hydroxyethyl methacrylate-comethacrylic acid) and poly(2-hydroxyethyl methacrylate-co-acrylic acid) using dynamic and equilibrium swelling studies. Chou and associates [23] and Kuo and associates [24] also observed this behavior in poly(2-hydroxyethyl methacrylate-comethacrylic acid), N,N-dimethyl aminoethyl methacrylate, and poly(hydroxyl ethyl methacrylate). Through swelling studies similar to the ones used by Khare and Peppas [22] and Chou et al. [23], Hariharan and Peppas [25] were able to develop a

Fig. 6.2 The effect of changing the pH of the swelling agent on ionic hydrogel membranes.

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6 Membranes in Controlled Release

model showing the effects of pH, ionic strength of the solution, and concentration of ionizable groups on polymer. The ionic nature of hydrophilic networks makes them ideal for uses in separation and controlled release applications. In separation processes, the external environment of the hydrogel can be adjusted to produce the desired solute diffusion. In controlled release systems, drugs can be delivered to a speciﬁc site in the body where the pH is at a certain value or range of values (e.g., in the gastrointestinal tract). The hydrogel would act as a carrier and release the drug at a desired rate once it has reach the site of release. The ionic nature, as well as other structural parameters, of the hydrogel has been studied to prepare hydrogels that can be tailor-made for a speciﬁc application. 6.3.2 Solute Transport in Network Membranes 6.3.2.1 Structural Parameters of Membranes Swelling studies can be used to determine the structure of membranes for controlled release. The membranes are prepared and their polymer volume fraction in the relaxed state is calculated using Equation (9). After each membrane has swollen to equilibrium, the polymer volume fraction of the swollen polymer is calculated using Equation (10). y2;r ¼

Vd ; Vr

ð9Þ

y2;s ¼

Vd : Vs

ð10Þ

Here, Vd, Vr, and Vs are the volumes of the polymer sample in the dry, relaxed, and swollen states, respectively, and y2,r and y2,s are the polymer volume fractions of the relaxed and swollen polymer gel, respectively. The volumes are calculated using Equations (11) through (12), which utilize the weights of the dry polymer, Wd, the relaxed polymer, Wr, and the swollen polymer, Ws, in air and heptane. Vd ¼

Wa;d Wh;d ; rh

ð11Þ

Vr ¼

Wa;r Wh;r ; rh

ð12Þ

Vs ¼

Wa;s Wh;s : rh

ð13Þ

6.3 Membranes and Solute Transport

Here rh is the density of heptane. The densities of the swollen and dry polymer can also be calculated using Equations (14) and (15). rswollen ¼

rdry ¼

Wa;s ; Vs

Wa;d : Vd

ð14Þ

ð15Þ

The equilibrium swelling ratio of a membrane can be affected by the ionic strength, temperature, and pH of the swelling agent. Data taken from swelling studies can be used to calculate the equilibrium swelling ratio: Q¼

1 : y2;s

ð16Þ

6.3.2.2 Determination of Molecular Pore Sizes There are several methods to accurately determine the pore size of membranes. While mercury porosimetry and BETmethods continue to be the two most important methods to analyze macroporous membranes, molecular analysis can be used for nonporous membranes. For example, in such membranes one can ﬁrst determine the average molecular weights between crosslinks, MC , in the case of crosslinked membranes or the average molecular weights between entanglements in uncrosslinked membranes. The molecular weights between crosslinks, MC , are calculated from the swelling data using Equation (17) as discussed by Peppas and Merrill [26]. ðy=v1 Þ½lnð1 y2;s Þ þ y2;s þ xy22;s 1 2 : ¼ Mc;e Mn y2;r ½ðy2;s =y2;r Þ1=3 ð1=2Þðy2;s =y2;r Þ

ð17Þ

Here, Mn is the number-average molecular weight of the polymer before crosslinking, y is the speciﬁc volume of the polymer, V1 is the molar volume of the solvent, y2,r is the volume fraction of the polymer in the relaxed state, y2,s is the volume fraction of polymer in the swollen state, and x is the interaction parameter of the polymer–solvent system in water. A theoretical MC value can be calculated from knowledge of the nominal crosslinking ratio X as follows: MC;t ¼

Mr : 2X

Here, Mr is the molecular weight of the repeating unit of the polymer.

ð18Þ

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Brannon-Peppas [27] developed equations to calculateMC for cationic and anionic membranes. Starting from the Peppas–Merrill equation, she accounted for ionization of the polymers. 2 h i h i V1 Ka y2;s 2 ¼ lnð1 y2;s Þ þ y2;s þ x1 y22;s pH 4I 10 þ Ka y " # V1 2MC y2;s 1=3 1 y2;s þ 1 y2;r : yMC Mn y2;r 2 y2;r

ð19Þ

Here, I is the ionic strength, pH is the pH of the buffer solution in which the membrane was swollen, and Ka is the dissociation constant. The membrane mesh size, j, deﬁnes the linear distance between consecutive crosslinks. Indirectly, it indicates the diffusional space available for solute transport and can be calculated using Equation (20). 1=3

j ¼ y2;s

1=2 2MC Cn ð Þ 1: Mr

ð20Þ

Here, Cn is the Flory characteristic ratio, and l is the carbon–carbon bond length. The crosslinking density of the membranes is calculated using Equation (21). rx ¼

1 yMC

ð21Þ

6.4 Applications in Drug Delivery

Hydrophilic membranes are predominantly hydrogel membranes that swell in water or biological ﬂuids without dissolving. The swelling characteristics are the result of crosslinks (tie-points or junctions), permanent entanglements, ionic interactions, or microcrystalline regions incorporating various chains. In the last twenty-ﬁve years, hydrogel membranes have been researched as prime materials for pharmaceutical applications, predominantly as carriers for delivery of drugs, peptides, or proteins. They have been used to regulate drug release in reservoir-type, controlled release systems or swellable systems [28,29]. Hydrogel membranes are characterized as neutral, anionic, or cationic networks. Their swelling behavior is governed by a delicate balance between the thermodynamic polymer–water Gibbs free energy of mixing and the Gibbs free energy associated with the elastic nature of the polymer network. In ionic hydrogels, the swelling is governed not only by the thermodynamic mixing and elastic–retractive forces but also by the ionic interactions between charged polymer chains and free ions. The over swelling behavior and the associated drug release kinetics are affected by the osmotic force that develops as the charged groups on the polymer chains are neutralized by mobile counterions. Electrostatic repulsion is also produced between

6.4 Applications in Drug Delivery

ﬁxed charges and mobile ions inside the gel, affecting the over swelling of the ionic gel. The equilibrium swelling ratios of ionic hydrogels are often an order of magnitude higher than those of neutral gels because of the presence of intermolecular interactions including coulombic, hydrogen-bonding, and polar forces. This means that drug or protein transport in ionic gels may be signiﬁcantly faster than in neutral gels [30]. Drug, peptide, or protein release through equilibrium-swollen membranes has been analyzed using classical Fickian diffusion theories. Drug diffusion is expressed in terms of hydrodynamic theories that consider the frictional characteristics of spherical solute drugs as they diffuse through the mesh of a gel. Derivation of drug diffusion equations through the mesh is based on the Fickian equation of drug ﬂux with additional terms for convection, wall partitioning, and restrictions due to the tortuosity of the diffusional drug path. Various theories for drug transport in gels have been based on the Eyring theory of rate processes; they utilize a free volume approach to describe the probability that a drug molecule will pass across the available mesh area. A successful model to describe drug transport in neutral hydrogels was derived by Peppas and Reinhart [19]. This model relates the normalized diffusion coefﬁcient to the drug size, the equilibrium degree of swelling, and the gel mesh size. Swollen membranes include a wide range of new materials that have been considered for such applications. For example, chitosan is an extremely promising carrier for drug delivery and has been studied either in the pure form or as a copolymer with other important polymers, for example, poly(ethylene glycol) (PEG). For example, Saito et al. [31] studied graft copolymers of the above for drug delivery applications. PEG continues to be an important carrier for drug delivery mostly in the ˜ a´n-Lo´pez and Bodmeier form of homopolymeric or copolymeric hydrogels. Remun [32] studied the process conditions for the formation of chitosan-based carriers and examined various applications of such systems either in colonic delivery or as mucoadhesives. Again, blends or copolymers with PEG or larger molecular weight poly(ethylene oxide) (PEO) were found to exhibit improved release properties. Recently there has been increased research in the preparation and characterization of membranes responsive to changing environmental conditions. Some of the environmental conditions that can affect hydrogel swelling include pH, ionic strength, temperature, and drug concentration. Many properties contribute to the swelling of ionic hydrogel membranes. An increase in the ionic content of the gel increases the hydrophilicity leading to faster swelling and a higher equilibrium degree of swelling. Gutowska et al. [33] prepared and tested a series of extremely versatile pH-sensitive gels that can be used for a wide range of pharmaceutical applications. Such hydrogels were shown to be the prime candidates for release of proteins, among other applications, including insulin and calcitonin. For example, environmentally sensitive hydrogels have been studied as possible controlled insulin release. Recent advances in the development of ionic membranes have concentrated on several aspects of their synthesis, characterization, and behavior. Major questions that have been addressed in recent work include (i) synthetic methods of preparation of hydrophilic polymers with desirable functional groups, (ii) synthetic methods of

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preparation of multifunctional or multiarm structures including branched or grafted copolymers and star polymers, (iii) understanding of the criticality and the swelling– syneresis behavior of novel anionic or cationic polymers, (iv) synthesis and characterization of biomimetic hydrogels, (v) understanding of the relaxational behavior during dynamic swelling, and (vi) modeling of any associated drug dissolution [7]. Hydrogel membranes may undergo volume-phase transition with a change in the temperature of the environmental conditions. The reversible volume change at the transition depends on the degree of ionization and the components of the polymer chains. This type of behavior is related to polymer phase separation as the temperature is raised to a critical value known as the lower critical miscibility temperature or lower critical solution temperature (LCST). Networks showing lower critical miscibility temperature tend to shrink or collapse as the temperature is increased above the LCST. Recently, in our laboratory, we composed hydrogel membranes of lightly crosslinked N-isopropyl acrylamide (NIPAAm) and methacrylic acid (MAA), which were synthesized and characterized for their sensitivity to external conditions and their ability to control release of two antithrombotic agents, heparin and streptokinase. PNIPAAm is noted for its sharp change in swelling behavior across the lower critical solubility temperatures of the polymer, while PMAA shows pH-sensitive swelling due to ionization of the pendant carboxylic groups in the polymer. Hydrogel copolymers of NIPAAm and MAA with appropriate composition were designed to sense small changes in blood stream pH and temperature to deliver antithrombotic agents, such as streptokinase or heparin, to the site of a blood clot. Experiments were performed to show that hydrogels with certain compositions could show both temperature- and pH-sensitivity. Numerous other temperature-sensitive membranes have been used for drug delivery including responsive, microporous hydroxypropyl cellulose gels modiﬁed so as to exhibit LCST close to room temperature and copolymers based on NIPAAm, butyl methacrylate, and acrylic acid (AA) with distinct LCST points [34,35]. Star polymers have been synthesized and characterized in the past twenty years. Star and dendritic polymers may be used in a variety of pharmaceutical applications. They can act as drug delivery membranes. These star chains could carry enzymes on the short arms, while long arms with a nonreactive functional group could provide protection for these enzymes. An alternative possibility would be to use star polymers with long arms carrying a cell recognition moiety to adhere to a speciﬁc site with shorter arms carrying a cytotoxic compound. In both these examples, the longer and shorter arms carry two different reactable groups to permit coupling of different molecules to the arms [28]. The main components of a star polymer are the core or foundation site from which the diverging branches of the star structure start. One or more branching arms emanate from this core site, each one incorporating a further branch point. Finally, a terminal functionality is observed for each of the branches, usually having a reactivity that allows it to further react in the dendritic structure. Star-shaped polymers of PEO have been prepared by anionic polymerization on a ‘‘core’’ of crosslinked divinyl benzene and are presently studied for drug delivery

6.4 Applications in Drug Delivery

applications, particularly as hydrogels. These polymeric systems are particularly promising because they can serve as micro- or nanoparticulate carriers for drug delivery system development. The use of functionalities at the end of the star arms allows immobilization of a wide range of biologically active agents. Such agents can be antibodies, antigens, and polysaccharides forming interesting conjugates for medical applications. Reaction of PEG at the end of the arms can lead to improved ‘‘stealth’’ particles that can be used in the body avoiding the reticulo-endothelial system. Various enzymes can be immobilized. Thus, heparinase can be immobilized to give nanoparticle star polymers, which can be used for postoperative blood puriﬁcation and removal of heparin. Fibrinoyltic enzymes such as streptokinase and urokinase can be immobilized and used for lysis of thrombi in the blood. Urease can be immobilized to produce microparticles, which will be useful in blood puriﬁcation by hemoperfusion. The advantages of such enzyme-immobilized star polymer particles are enhanced stability, ease in separation and reuse, ability to prepare enzyme free products, retention of a controlled enzyme microenvironment, low or no immunogenic response, and low cost–high purity products. A wide range of hydrophobic membranes has been studied in drug delivery applications. They include the widely used ethylene vinyl acetate copolymer series that has been used in many successful products, such as the early transdermal and ocular therapy systems. Other hydrophobic polymers include polyethylene, polypropylene, various types of hydrophobic cellulose derivatives such as ethyl cellulose, among others. Recent work in the ﬁeld is concerned with the formulation and evaluation of delivery devices based on highly crosslinked acrylates. These materials, prepared in loaded form by an extremely rapid, solvent-free photopolymerization process, are multiacrylate-acrylic acid copolymers having nominal crosslinking ratios ranging from 60 to 100 %. Release rates are therefore substantially lower than those exhibited by typical hydrogel formulations. By manipulation of the structure of the PEGcontaining crosslinking monomer and of the AA feed composition ratio in the comonomer–solute mixture, systems possessing a wide variety of delivery characteristics are attainable. Additionally, the presence of ionogenic AA moieties along the polymer backbone allows for environmental sensitivity and complexation-mediated release under appropriate conditions. Biodegradable membranes, by deﬁnition, change their chemical and potentially physical form upon contact with the biological environment. There are two distinct stages to the biodegradation process, especially in bulk degradation. The ﬁrst stage is restricted to the random cleavage of molecular linkages. The resulting decrease in molecular weight produces some change in mechanical properties and morphology, but no weight loss [36]. The second stage involves a measurable weight loss in addition to chain cleavage. It begins when the molecular weight of the polymer has decreased to the point that chain scission produces oligomers that are small enough to solubilize and diffuse out from the network. These oligomers are necessarily released into the adjacent tissue and therefore should be biocompatible. The predominant means by which polymers degrade is by hydrolytic degradation, enzymatic degradation, or a combination of the two. The degree to which each type of

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degradation contributes to the degradation of a speciﬁc polymer may vary during the degradation process and is dependent upon many factors, especially the environment in which the degradation is taking place. The degradation is on a biological time scale of hours to months, not an environmental time scale of years to decades. The products produced from this degradation will be resorbed by the body or completely eliminated by normal biological pathways. Polymers that degrade by hydrolysis include poly(lactic acid), poly(glycolic acid), poly(caprolactone), polyanhydride, poly(ortho ester), and polycyanoacrylate. Many poly(amino acids) such as poly(L-lysine), poly(L-arginine), poly(L-aspartic acid), poly(L-glutamic acid), and poly[N-(2-hydroxyethyl)-L-glutamine] are enzymatically degradable [37]. After degradation has occurred, it is essential that the products of the degradation be resorbed or eliminated from the body. For formulations that have been administered orally, the elimination of the polymers follow natural processes, but biodegradable polymers used parenterally or in other invasive techniques require more analysis and design to ensure fully biodegradable formulations. In order for this elimination to occur, the degradation byproducts, often small molecular weight polymer chains, must ﬁrst be solubilized before they can be eliminated. For the more water-soluble polymer fragments, this is a fairly straightforward process. Other materials must ﬁrst be ionized before solubilization can occur. There are three main classiﬁcations of degradable linkages that apply to biodegradble polymers as a whole. The category depends upon the location of the biodegradable linkage: (i) the polymer backbone, (ii) at crosslinks, and (iii) at pendant chains. Biodegradable formulations may include one or more of the mechanisms mentioned. Furthermore, the biodegradation means described earlier (i.e., hydrolysis or enzymatic degradation) can be used in any of the types. Degradation of the backbone and/or crosslinking agent usually is designed to lead to solubilization of the material. Release of a pendant group, however, may not lead to solubilization of the polymer matrix as the main, and potentially crosslinked, chain network will still be in place (Mathiowitz et al., 1999). From the previous analysis it is clear that membranes continue to have a prominent position in the ﬁeld of controlled release. References 1 Peppas, N. A. (2000) Intelligent

3 Rose, S. and Nelson, J. F. (1955) A

hydrogels and their biotechnological and separation applications, In Gu¨ven, G. ed. Radiation Synthesis of Intelligent Hydrogels and Membranes for Separation Purposes, IAEA, Vienna, 1–14. 2 Gehrke, S. H. and Lee, P. I. (1990) Hydrogels for drug delivery systems, In Tyle, P. ed. Specialized Drug Delivery Systems, Marcel Dekker, New York, NY, 333–392.

continuous long term injector. Australian Journal of Experimental Biology and Medical Science, 33, 415–419. 4 Folkman, J. and Long, D. M. (1964) The use of silicone rubber as a carrier for controlled therapy. Journal of Surgical Research, 4, 139–142. 5 Langer, R. and Peppas, N. A. (2003) Advances in biomaterials, drug

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(1996) Effect of formulation and process variables on the formation of chitosan-gelatin coacervates. International Journal of Pharmacology, 135, 63–72. Gutowska, A., Bark, J. S., Kwon, I. C., Bae, Y. H., Cha, Y., Kim, S. W. (1997) Squeezing hydrogels for controlled oral drug delivery. Journal of Controlled Release, 48, 141–148. Serres, A., Baudys, M. and Kim, S. W. (1996) Temperature and pH-sensitive polymers for human calcitonin delivery. Pharmaceutical Research, 13, 196–201. Yakushiji, T., Sakai, K., Kikuchi, A. et al. (1998) Graft architectural effects on thermoresponsive wettability changes of poly(N-isopropylacrylamide)modiﬁed surfaces. Langmuir, 14, 4657– 4662. Brannon-Peppas, L. (1997) Polymers in controlled drug delivery. Med. Plast. Biomaterials, 4, 34–44. Brannon-Peppas, L. (1995) Recent advances on the use of biodegradable microparticles and nanoparticles in controlled drug delivery. International Journal of Pharmaceutics, 116, 1–9.

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7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane Dimitrios F. Stamatialis 7.1 Introduction

The prolonged constant drug level in the body is the ultimate goal of every drug administration system. In fact, the constant drug level in the blood and the bypassing of the hepatic ‘‘ﬁrst-pass’’ metabolism are challenges for every delivery system. Figure 7.1 presents the time variation of the drug concentration in the blood. The traditional medical forms (tablets, injection solutions, etc.) provide delivery with peaks, often above the required dose that may even be toxic for the patient. The intravenous infusion can achieve constant drug delivery but requires hospitalization and supervision of the patients. As a result the cost of the treatment is high and the patient compliance is low. Historically, the skin has been viewed as an impermeable barrier that protects the body from external factors. Nevertheless, the introduction of drugs through it via creams and gels has been used for a long time. In the past years, the skin has been considered as a port for topical or continuous systematic administration of drugs. In fact, for drugs with short half-lives, the transdermal drug delivery (TDD) provides a continuous administration, rather similar to that provided by an intravenous infusion. However, unlike the latter, the TDD is noninvasive and no hospitalization is required. The drug level in the blood stays within the required acceptable limits (Figure 7.1). The TDD via a patch (Figure 7.2) can provide continuous drug release through intact skin into the blood stream for a long time and the delivery can be simply stopped by removing the patch. The contact with the skin is usually achieved via an artiﬁcial membrane. As we will see later the membrane can have or not have a regulatory role in the drug delivery. In this chapter, we will ﬁrst give a short description of the skin structure, which is one of the most important biological membranes, and of the fundamentals of skin permeation. Skin is a great protective barrier. However, some innovative methods have been already used to enhance TDD. The most promising ones (and rather close to broad commercialization) will be discussed more in detail.

Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

Fig. 7.1 Variation of the drug concentration in the blood during drug delivery.

7.2 Human Skin – Fundamentals of Skin Permeation 7.2.1 Human Skin Structure

The skin is the largest organ of the human body. It covers approximately 2 m2 of surface area and receives about one third of all the blood circulating through the body. Besides the protection against entry of foreign agents, it also plays a role in the blood regulation and acts as a body thermostat. It is not our intention to present the skin structure and properties in great detail in this chapter. The reader can ﬁnd such information in other excellent dedicated literature [1–4]. We will however focus our attention on those properties important to TDD. The skin is a rather complicated multilayer organ. It can be basically described in terms of two basic tissue layers: the dermis and the epidermis (see Figure 7.3).

Fig. 7.2 Schematic illustration of a transdermal drug delivery patch in contact with the skin.

7.2 Human Skin – Fundamentals of Skin Permeation

Fig. 7.3 A cross section of the human skin. (Source: Reprinted from Ref. [1], with permission from Elsevier, 2001.)

The dermis (thickness of 100–200 mm) forms the bulk of the skin and consists of connective tissue elements. It provides physiological support for the epidermis via blood and lymphatic vessels as well as nerve endings [5]. The epidermis, the top layer of the skin (thickness of 100–110 mm), is composed of epithelial cells held together by highly convoluted interlocking bridges. These bridges are responsible for the skin integrity. The epidermis comprises several physiologically active tissues and a physiologically inactive top layer: the stratum corneum (SC, 10 mm thick) that is exposed to the external environment (see Figure 7.3). The drugs can potentially pass through the skin either via the intact SC and/or via the hair follicles and sweat ducts (see Figure 7.3). In fact, the average human skin contains 40–70 hair follicles and 200– 250 sweat ducts per square centimeters of area. However, both these appendages occupy only 0.1 % of the skin; therefore, the SC is the main barrier to drug transport. 7.2.2 Stratum Corneum – Main Drug Barrier

The drug transport through the skin evolves in a series of steps in sequence. The drug is ﬁrst absorbed in the SC, diffuses through it and through the lower layers of epidermis, and then passes through the dermis and ﬁnally into the blood circulation. The rate of penetration through the SC controls the drug delivery because the drug transport through the deeper layers as well as through the vessel walls happens at a higher rate. The control of delivery by the SC is due to its unique structure. The human SC has about 18–21 cell layers of ﬂattened, mainly dead keratinized, metabolically inactive

193

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7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

cells (corneocytes). The individual corneocytes are 20–40 mm in diameter, in contrast to 6–8 mm for the basal cell and may differ in size or packing depending on the body site or their location within the SC [2]. These cells are formed and continuously replenished by the slow upward migration of cells produced at the lower layers. The SC contains only 20 % wt H2O in contrast to normal 70 % in the lower physiologically active layers. The SC is composed of closely packed cells. The intercellular spaces are narrow and occupied by other substances. In fact, the SC has been described as a wall with bricks and mortar [1,6,7] (Figure 7.4). The corneocytes of hydrated keratin are the ‘‘bricks’’ impended in the ‘‘mortar,’’ consisting of lipid bilayers of ceramides, fatty acids, cholesterol, and cholesterol esters. This arrangement creates a tortuous path that most of the drugs should follow in order to pass through the SC. As we will see later, some of the techniques aim to increase the drug transport through the skin by either increasing the intercellular space or even completely disrupting the structure of the SC. As it was mentioned earlier, the SC is partially hydrated (20 % wt H2O). When it is in contact with water, it absorbs water slowly and when is fully hydrated, it absorbs ﬁve to six times its weight of water. The water diffusion coefﬁcient (Dw) through the SC is comparable or slightly smaller than that of the biological membrane and in the order of 1010 cm2/s. It is important to note that the thickness of the SC and the Dw differs in different body sites (Table 7.1, adapted from [8]). For example, the SC of the palm is much thicker than in other sites but also highly water permeable (see Table 7.1). In any case, the Dw through the SC is three to four

Fig. 7.4 Drug permeation through SC. (Source: Reprinted from Ref. [1], with permission from Elsevier, 2001.)

7.2 Human Skin – Fundamentals of Skin Permeation Water diffusivity through the SC at different sites.

Tab. 7.1

Skin site

Thickness (mm)

Dw (cm2/s, 1010)

15.0 10.5 13.0 49.0 400.0

6.0 3.5 12.9 32.3 535.0

Abdomen Back Forehead Back of hand Palm

orders of magnitude lower than in the much thicker (200 mm) dermis (Dw,dermis ¼ 2.106 cm2/s). 7.2.3 Drug Transport Through the Skin 7.2.3.1 Passive Diffusion General – Transport Mechanism The drug delivery through the skin due to the drug concentration difference between the patch reservoir (CD,res) and the skin (CD,skin) is called drug passive diffusion and can be described by Fick’s law: JDPD ¼

kD DD ðCD;res CD;skin Þ ; ‘

ð1Þ

where JDPD is the steady state drug ﬂux through the skin, kD is the drug partition coefﬁcient in the skin, DD is the diffusion coefﬁcient of the drug through the skin, and ‘ is the skin thickness. When the CD,res >> CD,skin, then Equation 1 can be written as JDPD ¼

kD DD CD;res ¼ KDPD CD;res ; ‘

ð2Þ

where KDPD is the passive drug permeability coefﬁcient through the skin. The drug passive diffusion increases by using the maximum drug amount that can be dissolved in the drug reservoir (the maximum solubility of the drug in the reservoir), CD,S: PD JD;max ¼ KDPD CD;S :

ð3Þ

Drugs with high partition into the skin and high diffusion coefﬁcient are good candidates for a TDD system. In fact, Potts and Guy [9] suggested an empirical relation to predict the drug permeability through the skin: log½KDPD ðcm=hÞ ¼ 2:7 þ 0:71logkD;oct 0:0061ðMWÞD :

ð4Þ

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The kD,oct is the partition coefﬁcient of the drug in octanol/water and can be used as an indication of the drug partition into the skin. Because the drug has to go through lipophilic and hydrophilic layers, it should be soluble satisfactorily in both oil and water. One should, however, note that drugs with high octanol/water partition usually have low solubility. Therefore, for each drug an optimum in partition and solubility should be achieved. In addition, the lower the molecular weight of the drug, (MW)D, the higher is its diffusivity through the skin. Equation 4 is often used to predict skin permeability, and the reader can ﬁnd other proposed models and empirical relations elsewhere [9]. Currently, the transdermal delivery mostly involves potent drugs of low molecular weight (clodinine, estradiol, fentanyl, nicotine, nitroglycerin, testosterone, scopolamine, and others) that are active at low blood concentrations (order of ng/mL or less) [10]. In order to expand to include a broader range of drugs, various methods have been developed and will be described later in the chapter. Measurement of Drug Passive Diffusion Most of the studies for a TDD system start with an in vitro evaluation of the delivery. Skin (full skin or only SC) is isolated from animal and/or human sources and used in laboratory experiments. Figure 7.5 presents a piece of human skin obtained from a patient who had undergone a cosmetic surgery. The isolation and treatment of the skin is an important and labor intensive procedure requiring a series of steps [11]. The skin is cleaned with ethanol and dermatomed from the rest of the tissue (see Figure 7.5). Depending on the planned experiments, the SC can be separated from the epidermis and then tried in N2 atmosphere. When access to normal skin samples is not possible, artiﬁcial skin substitutes often used in burn surgery and the treatment of chronic wounds [12] may be used for in vitro tests.

Fig. 7.5 Photo of human skin sample.

7.2 Human Skin – Fundamentals of Skin Permeation

Fig. 7.6 The Franz diffusion cell.

The in vitro experiments are performed using various types of devices. One of the most common devices is the ‘‘Franz’’ diffusion cell (Figure 7.6). The drug permeates from the donor compartment through the skin and is detected in the receptor compartment. The temperature is kept constant at 37 8C with circulating water. Figure 7.7 presents a typical example of the passive diffusion of timolol (TM) [13] through pig SC. TM is a nonselective beta-adrenergic blocking agent that is used in the management of hypertension, angina pectoris, myocardial infraction,

Fig. 7.7 Passive diffusion proﬁle of TM through pig SC.

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7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane Tab. 7.2

Some examples of passive transdermal drug delivery.

Drug – permeant

Skin type

References

Amino acids – dipeptides Dextran – morphine Dihydroergotamine Thyrotropin-releasing hormone Tetrapeptide Piroxicam Nalbuphine Fentanyl Diclofenac Nicotine Theophylline Naltrexone Timolol

Porcine skin Human skin – in vivo Rabbit skin Mouse skin Human skin Human SC Human SC Rat skin Human skin – in vivo Human skin – in vivo Human skin Porcine skin Human and porcine skin

[15] [16] [17] [18] [19] [20] [21] [22] [23,24] [25,26] [27] [28] [29–32]

and glaucoma. It undergoes extensive ﬁrst-pass hepatic metabolism, and its elimination half-life is 2–2.6 h [14]. Table 7.2 presents a selection of scientiﬁc literature concerning in vitro passive diffusion transdermal delivery [15–32]. Transdermal products and short patent overview will be presented later in the chapter. It is important to note that if the results of the vitro study are promising, then in vivo study follows. For this, often animals (mice, hairless rats, guinea pigs) and human volunteers are used. Detailed discussion about these experiments is beyond the scope of this chapter. 7.2.3.2 Iontophoresis General – Transport Mechanism The concept of the TDD via a patch is widely known and used in daily life. Nevertheless, the skin is a signiﬁcant drug barrier and for most drugs the skin permeability is low. The delivery can be assisted by electrical energy. Iontophoresis implies the use of small amounts of physiologically acceptable electric current to drive charged drug molecules into the body [5]. The iontophoresis device comprises two patches containing the two electrodes – one the anode and the other the cathode and the power supply (Figure 7.8). The drug formulation (drug dissolved in either liquid or gel reservoir) is placed in patch – electrode having the same charge as the drug (in Figure 7.8, at the anode). The other electrode/patch contains only reference electrolyte or gel. The two patches are placed on the skin and connected to the power supply. Figure 7.9 presents an iontophoresis patch in a more detailed manner. The drug is driven into the skin by electrostatic repulsion. In addition, a bulk ﬂuid ﬂow or volume ﬂow occurs in the same direction as the ﬂow of the counterions. This phenomenon, which accompanies iontophoresis, is called electroosmosis.

7.2 Human Skin – Fundamentals of Skin Permeation

Fig. 7.8 Iontophoresis principle.

The steady state ﬂux of a charged drug during iontophoresis comprises three parts: the ﬂux due to passive diffusion (JDPD ), the ﬂux due to electromigration (JDEM ), and the ﬂux due to electroosmosis (JDEO ): JDtotal ¼ JDPD þ JDEM þ JDEO :

ð5Þ

In most of the iontophoretic drug delivery systems, JDPD is very low and often negligible. Nevertheless, when it can be measured, it is used to obtain the passive drug permeability (KDPD , Equation 2). The electromigration can be described by the equation [33]: JDEM ¼

iD ; zD AF

Fig. 7.9 Schematic illustration of an iontophoretic patch.

ð6Þ

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7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

where, iD is the drug ionic current ﬂow, zD is the charge of the drug, A is the surface area, and F is the Faraday constant. The drug current ﬂow is related to the applied current, I, via the equation iD ¼ tD I;

ð7Þ

where tD is the transport number of the drug and represents the fraction of the total current transported by the drug. The tD shows the importance of the presence of competitive ions in the drug for the drug delivery – the higher the tD, the higher the drug delivery efﬁciency. By combination of Equations 6 and 7, we get

JDEM ¼

tD I tD I ¼ : zD AF zD F A

ð8Þ

The ratio, I/A, is the current density. Based on Equation 8, the drug transport should increase proportionally to the applied current density. Later, we will see that this does not always hold. The electroosmotic ﬂux, JDEO , is the bulk drug ﬂow occurring when a voltage difference is applied across the charged skin. Rein [34] has proven 80 years ago the electroosmosis ﬂow in human skin, and recent studies conﬁrmed the phenomenon in skin from various sources [35–38]. The electroosmosis occurs always in the same direction as the ﬂow of the counterion and may assist or hinder the drug transport. The electroosmosis increases in importance as the size of the drug ion increases. For small ions, the drug ﬂux increases mostly because of electromigration. For bigger drugs, such as peptides and proteins, the electroosmosis might be the dominant transport mechanism. An excellent recent review [39] describes some examples using hairless mouse and human skin.

Factors Affecting Iontophoretic Drug Delivery Figure 7.10 shows a continuous ﬂow through transport cell, proposed by Van der Geest et al. [40] to measure the iontophoretic drug transport. The drug is placed in the donor compartment (in Figure 7.10, at the anodal chamber), and the electric current is applied via the two electrodes connected to the power supply. A few minutes (2–3 min) after the current application, the skins polarizes and its electrical resistance drops sharply [31]. The drug permeates through the skin and is collected by the ﬂow through solution simulating the blood. The reference electrode compartment (in Figure 7.10, at the cathodal chamber) contains only reference electrolyte. Figure 7.11 presents a comparison of the TM delivery between passive diffusion and iontophoresis through pig SC. The iontophoresis TM permeability is about four times higher than the passive diffusion [31]. During the ﬁrst few minutes of the current application, the electrical resistance of the SC drops sharply from about 16 kV cm2 to about 3 kV cm2 and stays at the lower levels for several hours of iontophoresis.

7.2 Human Skin – Fundamentals of Skin Permeation

Fig. 7.10 The continuous ﬂow through transport cell. (Source: Adapted from Ref. [13].)

Drug Concentration In passive diffusion, besides the selection of the suitable drug (based on its MW and partition into the skin), it is very important to achieve high drug loading in the patch to maximize the drug delivery. In iontophoresis, the drug concentration still has a great inﬂuence on the delivery. Usually, if the delivery is performed under constant current density, the drug delivery increases as the drug concentration in the patch is high. However, often the relationship is not linear. The drug transport number seems to be reaching a plateau at high concentrations. An

Fig. 7.11 Typical result of the TM transport through pig SC under passive diffusion and iontophoresis. (Source: Adapted from Ref. [31].)

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7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

interesting ﬁnding is that in some cases the delivery is independent of the drug concentration in the patch or even decreases at high concentrations. These discrepancies are often related to the competition of the drug with other species of the background electrolyte for the current or to the accumulation of the drug to the skin (‘‘skin fouling’’). The reader can ﬁnd several examples concerning speciﬁc drugs elsewhere [33]. Electric Current Based on Equation 8, the iontophoretic drug ﬂux is directly proportional to the applied current density. In fact, this has often been found in the literature. In some cases, however, the iontophoretic transport can reach a plateau suggesting again the saturation phenomena. For the same drug, the ﬂux level can be inﬂuenced by the drug formulation due to competition phenomena. High amount of background electrolyte can result in lower drug efﬁciency. Often in the literature, especially during in vitro experiments, high current densities are applied. Nowadays, most scientists agree that 0.5 mA/cm2 is the maximum acceptable current density producing minimal skin irritation and/or damage [13,40,41]. It is always very important to consider this fact when designing a TDD system. High currents can often achieve high drug delivery. However, the level of the current, the time of application, and the potential health risks should always be taken into consideration. It is important to note that recent studies indicated that constant conductance alternating current (CCAC) iontophoresis can enhance drug delivery as much as direct current (DC). In fact, the results of CCAC delivery seem to show less inter- and intrasample variability than that of the DC iontophoresis [42–44]. It seems that the CCAC iontophoresis can provide more controlled TDD. In iontophoresis besides the amount of the applied current, the type of electrodes too has an important role. The conventional electrodes used are classiﬁed as inert (metals such as stainless steel, platinum, carbon, or aluminum) or reversible (Ag/AgCl, see Figure 7.10). The inert electrodes do not take part in the electrochemical reaction, but they cause electrolysis of water leading to pH shifts and consequently to skin irritation and perhaps to variations in drug delivery and stability [5]. The Ag/AgCl electrodes can avoid these problems but participate in the electrochemical reactions (therefore they are called ‘‘active’’ electrodes). The use of Ag anode electrode requires the presence of Cl for the electrochemical reaction:

Anode : AgðsÞ þ ClðaqÞ ! AgClðsÞ þ e ; E ¼ þ0:22V; Cathode : AgClðsÞ þ e ! AgðsÞ þ ClðaqÞ ; E ¼ þ0:22V: In most cases NaCl is used to provide the Cl. The presence of Naþ can have a signiﬁcant impact on the drug delivery because it competes with the positively charged drug for the electrical current. In the AgCl cathode, Cl ions are released. To have electroneutrality, cations, which also compete for the current, should be released from the skin. As a result, the drug delivery efﬁciency can be very low. The Ag/AgCl electrodes can also cause precipitation of peptides and protein drugs. In recent work, Stamatialis et al. [45] found precipitation of salmon

7.2 Human Skin – Fundamentals of Skin Permeation

Fig. 7.12 Loss of sCT due to the electrical current application using pure Ag/AgCl electrodes and Ag/AgCl agarose bridged electrodes. (Source: Adapted from Ref. [45].)

calcitonin (sCT, a drug used in the therapy of hypercalcemia, postmenopausal osteoporosis, and the treatment of Paget’s disease of bone) in contact with the Ag electrode. To avoid the direct contact, the electrodes were separated from the sCT solution by agarose salt bridges. Figure 7.12 shows that the agarose bridges avoid sCT losses. The pH The pH of the drug formulation can have a signiﬁcant effect on the iontophoretic drug delivery. One should select the pH when the drug is highly charged to enhance the electromigration. Often, however, a compromise should be achieved between the pH of the drug and the drug stability and solubility, and skin irritation. For example, recent studies have shown that the degradation of sCT is very low at low pH [45–47]. In fact the lowest sCT degradation is found at pH 3–3.3. At such low pH, however, the risk of skin irritation and/or skin damage is high. Therefore, as a compromise, minimum pH 4 was selected to achieve both low drug degradation and skin irritation [45]. Besides the skin irritation, high amounts of Hþ and OH should be avoided to improve the drug current efﬁciency. Finally, the pH of the formulation can also inﬂuence the electroosmosis by changing the skin charge. All the above-discussed parameters affecting iontophoresis delivery have been investigated extensively in the literature. Table 7.3 presents a selection of scientiﬁc papers concerning delivery through various skin types in vitro and in vivo [11,29,31,32,37,40,43,48–66].

203

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7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane Tab. 7.3

Some examples of iontophoresis transdermal drug delivery.

Drug – permeant

Skin type – condition

References

Timolol maleate Salmon calcitonin Timolol maleate Apomorphine Morphine HCl Various dipeptides Triptorelin Amitriptyline HCl Tetraethylammonium chloride Nicotine Deﬁbrase Tacrine HCl Botox Ketamine Catechins from tea Amino acids Heparin Amikacin Dopamine agonist 5-OH-DPAT Vapreotide acetate Nalbuphine, sebacoyl Antihistamines

Porcine SC – in vitro Rat skin – in vivo Human SC – in vitro Human skin/SC – in vitro Human SC – in vitro Human SC – in vitro Porcine skin Cadaver human skin Human epidermis Human skin Human epidermis Rat skin Human skin – in vivo Human skin – in vivo Porcine skin Porcine skin – in vitro Rat skin Rabbit skin Human SC/skin – in vitro Porcine skin – in vitro Human skin – in vitro Human skin – in vivo

[31,32] [48] [29] [11,40,49,50] [51] [52] [37] [53] [43] [54] [55] [56] [57] [58] [59] [60] [61] [62] [63] [64] [65] [66]

7.2.3.3 Electroporation General – Transport Mechanism Electroporation is the structural perturbation of lipid bilayer membranes caused by the application of high voltage pulses. For TDD, high voltage pulses (100–1000 V) for short time period (100 ms–1 s) are used [67]. In contrast to iontophoresis that acts primarily on the drug (by electromigration and electroosmosis, causing structural changes in the skin polarization, decrease of electrical resistance) as secondary effects [41]), electroporation acts directly on the skin, causing changes in the permeability of the tissue. This involves the creation of ‘‘pores’’ (or aqueous pathways) for the drug transport through the SC that are rather small (<10 nm), cover small surface area [68,69], and are generally short lived (ms to s) [70–72]. The parameters affecting the electroporation TDD are Drug properties: charge/MW/hydrophilicity Drug formulation: pH/competitive ions/viscosity Electrical: Pulse waveform/voltage/number/length.

7.2 Human Skin – Fundamentals of Skin Permeation

The effects of the drug properties and drug formulation on the delivery are similar to those of passive diffusion and iontophoresis (and have been discussed earlier). Here we will focus our attention on the electrical properties. Pulse Speciﬁcations Two types of pulse waveforms are applied: the decaying and square wave pulses. The exponential decaying pulses depend on the skin resistance and the electroporation system (electrode, conducting medium). This can be a drawback for clinical application because the reproducibility of the delivery would not be always possible. The square pulses, however, remain constant during application and therefore, their effect and drug delivery could be controlled better. The drug delivery is generally enhanced when the number or the duration of the pulses increases [73–78]. The enhancement is steeper at low voltages than at high voltages. Electrode Material and Design In most cases inert Pt electrodes are used because the Ag/AgCl cannot be used for the application of high voltages for a long time. The design of the electrode is of great importance, too. The parallel plate electrodes, which are the simplest conﬁgurations to create a uniform ﬁeld [79,80], can cause nerve stimulation and/or skin irritation and burning. The meander electrode design localizes the ﬁeld on the superﬁcial skin layer [80] and seems to be a better option. Safety Issues Jadoul et al. [41] have recently reviewed the effects of electroporation and iontophoresis on the SC. Some changes of the SC, which are partly reversible, occur for both methods: Disorganization of the lipid bilayers of SC Increase of skin hydration Decrease of skin resistance; a larger decrease observed in electroporation.

In contrast to iontophoresis that is generally considered a safe method, the safety issues of electroporation are still under debate. It is rather difﬁcult to deﬁne the conditions that will be acceptable for clinical use. Many parameters such as pulse voltage, pulse length, pulse frequency, duration of treatment, electrode size, electrode design, and skin site are important [67,73]. The sensation and pain thresholds exhibit a rather complex dependence on the current [68]. Tests on human volunteers indicate that both sensation and pain depend on the volunteer and besides physiological, often psychological issues are important. Generally, the levels of sensation increase at high current/charge, pulse rate, and pulse length. The quality of sensation is determined by the frequency, ranging from warmth to tingling to vibration. Quite often muscle stimulation and blocking might occur.

205

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7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

Deﬁnitely, in the development of a drug delivery system, a balance between effective drug delivery and safety should be reached following carefully the safety guidelines and standards of responsible organizations [68]. The in vitro electroporation experiments can be performed using devices similar to those used in iontophoresis (Figure 7.10). In the scientiﬁc literature, several important studies show the signiﬁcant effect of electroporation on the transdermal drug delivery (see Table 7.4 [30,36,78,81–94]). Usually, two different experimental protocols are used: intermittent application of more ‘‘short’’ (<10 ms) high voltage pulses and less number (<15) or ‘‘long’’ (100–500 ms) medium voltage pulses (50–250 V) [73,74]. Figure 7.13 shows a typical example of a TM permeation proﬁle under passive diffusion and electroporation. In this case, passive diffusion was followed for 2 h, then 10 pulses of 400 V during 10 ms were applied, and the diffusion after electroporation was evaluated for 6 h. Owing to pulses, pores are created into the skin that cause signiﬁcant increase in the drug transport. The combination of electroporation with iontophoresis was found to be very effective to improve transdermal delivery of big drug molecules [36,94,96]. In this case, after the pulse application and the creation of the pores on the skin, constant current iontophoresis is applied. Besides, electroporation has been used in combination with ultrasound [97] and chemical enhancers [98–100]. More details about these techniques will be described later in this chapter.

Tab. 7.4

Some examples of electroporation transdermal drug delivery.

Drug/penetrant

Skin type

References

Insulin Salicilic acid Methotrexate Piroxicam Carboxyﬂuorescein Timolol Atenolol Human parathyroid hormone Nalbuphine Buprenorphine Sodium nonivamide acetate Fentanyl Buprenorphine Physostigmine Terazosin HCl FITC-dextran Sulforhodamine Salmon calcitonin

–— Rat skin Pig skin Porcine skin Porcine skin Human SC Human SC Porcine skin Mouse skin Human skin Human skin Human epidermis Human epidermis Human skin Rat skin Rat skin Human skin Human epidermis

[81] [82] [83] [84] [84] [30,36] [36] [85] [78] [86] [87] [88] [89] [90] [91] [92] [93] [94]

7.2 Human Skin – Fundamentals of Skin Permeation

Fig. 7.13 TM ﬂux from the carbopol gel 1 % (w/w) through pig SC under the electroporation protocol (passive diffusion – pulse application – postelectroporation diffusion) [95].

7.2.3.4 Other Methods Skin Penetration Enhancers The TDD can be improved using skin penetration enhancers (otherwise called sorption promoters or accelerants) [101]. These compounds penetrate into the skin, interact with skin components, and reversibly decrease the skin resistance to transport. The penetration enhancers should [102] Be nontoxic, nonirritating, and nonallergenic Have no pharmaceutical activity on the body Act fast in a predictable and reproducible way Not cause irreversible changes in the skin Not cause loss of body components Be compatible with the drug and other components (buffer, etc.) Be cosmetically acceptable.

An effective penetration enhancer of all the above requirements should improve the drug passive skin permeability (KDPD , see Equation 2) that it may increase the DD or improve the partitioning of the drug into the skin (kD). In a recent excellent review, Williams and Barry [101] summarized the potential mechanism of action of the enhancers: Denaturation or modiﬁcation of intercellular keratin of SC Effects on the dermosomes that are responsible for the cohesion between corneocytes

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7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

Modiﬁcation of the intercellular lipid domains and therefore reduction of the resistance of bilayer lipids Alteration of the solvent nature of SC and therefore increase of drug partition and/or drug diffusivity.

In the literature, one can ﬁnd several methods to prepare drug formulations with penetration enhancers. Table 7.5 presents a short overview of the mostly used enhancers. Some details about their action are included, too [103–128].

Ultrasound-Assisted TDD – Sonophoresis Ultrasound is deﬁned as sound having a frequency beyond 18 kHz and has already been used in medicine, for example, in physical therapy, in dentistry or for diagnostic purposes. There are basically three types of conditions [129]: High frequency (3–10 MHz) – for diagnostic purposes Medium frequency (0.7–3.0 MHz) – for therapeutic purposes (physical therapy) Low frequency (18–100 kHz) – for lithotripsy, liposuction, cancer therapy, dentistry, transdermal delivery, and others.

The ultrasound energy penetrates the body tissue and is absorbed by the tissue. The effects on the tissue can be Thermal: Owing to ultrasound the temperature of the tissue increases. The rise varies with the ultrasound intensity and exposure time. Cavitational: Owing to ultrasound-induced pressure, gaseous cavities are formed in the tissue. The cavities can collapse causing signiﬁcant changes in the surrounding tissue. Acoustic streaming: As a result of the sound waves, a onedimensional ﬂow current that affects the surrounding tissue develops. The ultrasound can increase the TDD – a method called sonophoresis – from a small percentage to several orders of magnitude. The mechanism of the improved TDD has been studied extensively in the past 20–25 years. Two excellent recent reviews discuss in detail the developments in this ﬁeld [129,130]. The most dominant mechanism suggests that the ultrasound interacts with the intercellular lipids of the SC that is somewhat similar to the action of the penetration enhancers. Mitragotri et al. [130,131] have evaluated the effects of the various ultrasound-related phenomena to the skin structure and drug delivery. The low-frequency ultrasound seems to be more effective. For example, an ultrasound of 20 kHz of intensity 225 mV/cm2 applied for 100 ms every second can increase the drug delivery up to 1000-fold higher than a therapeutic ultrasound [131]. Mitragotri et al. [130,131] suggested that the cavitation effects cause disorder to the lipids of SC, increasing the water transport through it. The drugs use these aqueous channels to penetrate at high rates. The

Result

Effective at low concentration (0.1–0.5 %), interact with lipid domains of SC Pyrrolidones (N-methylPartition well in the SC and 2-pyrrolidone (NMP) or alter the solvent nature 2-pyrrolidone (2P)) [111,112] of the skin Fatty acids [113–115] Interact and modify the lipid Oleic acid [116,117] domains of SC Alcohols, fatty Increase drug solubility in the patch, alcohols, glycol [118–121] increase drug partition into the skin, solvent drug dragging Anionic, cationic: interact with Surfactants anionic, cationic, intercellular keratin and swell the SC nonionic (sodium lauryl sulfate – Tween 80) [122–125] Nonanionic: only minor interaction Modify solvent nature of SC, Essential oils, terpenes, improving drug partition terpenoids (menthol, Increase drug diffusivity eucalyptus) [126,127] Prepare vehicles to Phospholipids [128] carry drug through the skin Occlude skin surface, increase tissue hydration

Dimethyacetamide (DMAC), dimethyl formamide (DMF) [106–108] Azone [109,110]

Sulfoxides [103–105]

Problems

Erythema and damage of SC at high concentrations. Bad odor in the mouth DMF can cause irreversible damage to the skin

No

Anionic, cationic: irritants potentially damaging to the skin No

No

Promote both hydrophilic and hydrophobic drugs Liposomes fuse with SC lipids and collapse increasing the drug partition to SC

No

No

Promote both hydrophilic and hydrophobic drugs

Promote both hydrophilic and hydrophobic drugs Promote both hydrophilic and hydrophobic drugs

Enhances skin transport of steroids, Not clear antibiotics, and antiviral agents (hydrophilic and hydrophobic) Promote both hydrophilic Cause short-lived erythema and hydrophobic drugs and toxic reactions

Action Increase SC humidity, swelling High DD, KD for hydrophilic and hydrophobic drugs of corneocytes Change intercellular keratin conformation, Promote both hydrophilic distort packing geometry of lipid domains and hydrophobic drugs Change intercellular Promote both hydrophilic keratin conformation and hydrophobic drugs

Water

Skin penetration enhancers for TDD delivery.

Enhancer

Tab. 7.5

210

7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

thermal effects induced by the ultrasound seem to have a relatively lower impact on the drug delivery. Merino et al. [132] showed that about 25 % of the enhancement of the mannitol delivery can be attributed to the increased temperature. Most of the enhancement is due to cavitation effects. In low-frequency sonophoresis, the drug enhancement is determined by four main parameters: frequency, intensity, duty cycle, and application time. The intensity, duty cycle, and time can be combined to a single parameter, the total energy delivered to the skin: Energy ¼ intensity time:

ð9Þ

Generally, at a certain frequency the drug enhancement increases beyond a certain energy dose. At each frequency, there is a threshold intensity below which no delivery enhancement occurs [130,133]. Beyond this threshold, the delivery increases strongly until a second threshold is reached: the decoupling intensity. Beyond the decoupling intensity, no further drug delivery enhancement occurs. The ﬁrst threshold is really low at low ultrasound frequency probably due to the cavitations that are more pronounced at low frequency. For example, for porcine skin, the threshold is 0.11 W/cm2 at 19 kHz and 2 W/cm2 at 93.4 kHz [130]. It is, however, important to note that at very low frequency, the drug enhancement is indeed strong but seems to be localized to certain areas [133]. The effect seems to be more homogenous beyond an optimum frequency of around 60 kHz [130]. Finally, sonophoresis has already been used successfully in combination with skin penetration enhancers to deliver mannitol [134] and iontophoresis to deliver heparin [135]. Table 7.6 presents more examples of sonophoresis drug delivery literature [131,132,136–149]. Microprojection/Microneedle Patch The microprojection or microneedle patch is a recent development for the TDD. It consists of a microneedle array that is applied to

Tab. 7.6

Some examples of sonophoresis transdermal drug

delivery. Delivered drug

References

Insulin Mannitol Glucose Heparin Inulin Morphine Caffeine Lidocaine Ketoprofen Salicylic acid/corticosterone

[136–141] [132,142,143] [132,143] [144] [143] [145] [145,146] [147] [148] [131,149]

7.2 Human Skin – Fundamentals of Skin Permeation

Fig. 7.14 Micromachined hollow microneedles: (a) SEM picture of a 350 mm high microneedle and (b) array of needles with a pitch of 555 mm. (Source: Reprinted from Ref. [154], 2003IEEE.)

the skin. The needles penetrate the skin and create superﬁcial pores through which the drug can be delivered [150]. Such patch has been used in conjunction with passive diffusion and iontophoresis to deliver therapeutical amounts of antisence oligodeoxynucleoside (ODN) to guinea pigs in vivo [150]. Martanto et al. [151] have also delivered insulin successfully to diabetic hairless rats. Alternatively, the needles have been coated with a model protein antigen ovalbumin (OVA). After the needle penetration through the skin, the OVA was delivered to hairless guinea pigs in vivo [152]. Besides, Chabri et al. [153] manufactured silicon-based arrays and delivered genes to human breast skin in vitro. Gardeniers et al. [154] have recently developed a method to fabricate hollow microneedle arrays in silicon (see Figure 7.14). The patch containing these needles has improved the diclofenac delivery by a factor of 750 with respect to passive diffusion. Recently, Park et al. [155] prepared microneedle patches of biocompatible and biodegradable polymers such as polyglycolic acid (PGA) and their copolymers. The ﬁrst results showed that calcein and bovine serum albumin (BSA) can be delivered effectively through human cadaver epidermis [155]. The polymeric needles provide additional safety in comparison to silicon or stainless steel ones. If the needles break into the skin, they will safely degrade in time without risks to the patients.

Magnetophoresis Magnetophoresis is an approach to enhance the delivery of a diamagnetic drug across skin using a magnetic ﬁeld. The concept seems to be attractive, but it has not been exploited yet. The ﬁrst results using benzoic acid showed that the delivery can be enhanced with respect to passive diffusion. The drug ﬂux increases with the applied ﬁeld strength [156].

211

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7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

7.3 Transdermal Drug Delivery System – Structure/Design

In the past two to three decades the TDD systems have become part of our daily life. The ﬁrst ones developed and the most popular ones are still based on passive drug delivery. In the past few years, however, the iontophoresis TDD systems are slowly being introduced. Yet, these systems have not found broad application outside the hospital though the prospects are positive. In this section, we will separate the TDD systems into two main categories: The passive systems: traditional delivery due to drug concentration difference between the patch and the skin The active systems: where an extra driving force such as electric current (iontophoresis/electroporation), ultrasound (sonophoresis), or others is used. The materials and basic technologies for the construction of the patches are the same for both systems. Therefore, it will be described only once for the passive systems. Of course, the active systems require extra components to provide the driving forces (electrical current, ultrasound, and others). These elements will be described separately.

7.3.1 Passive TDD Systems 7.3.1.1 Types The technologies developed to provide controlled passive drug delivery can be classiﬁed into two main categories: 1. Membrane or reservoir systems: In this case, the drug is incorporated into a reservoir (liquid or gel) placed between a drug impermeable layer and a membrane (see Figure 7.2). The device also includes an adhesive layer on the external surface of the membrane to achieve a ﬁrm contact with the skin. The drug release from this system can be controlled by varying the reservoir composition and the drug permeability through the membrane (by tailoring the material, porosity or thickness) and/or through the adhesive. Several successful commercial TDD systems are based on this design (see Table 7.7, adapted from [157,158]). 2. Matrix systems: In this case, the drug is incorporated (dissolved and/or distributed) into a polymer matrix (see Figure 7.15). There is no membrane and the adhesive layer is added when the matrix itself is not adhesive. Table 7.7 presents some examples of systems of this type, too.

7.3 Transdermal Drug Delivery System – Structure/Design Tab. 7.7

Some commercial passive TDD systems.

Trade name

Company

Type

Drug

Action

Nitroderm Nitrodur Deponit Nitrodisc Frandol- Tape

Alza/Ciba Key/Schering Lohman/Schartz G.D. Searle Nitto Electric Ind.

Reservoir Matrix Matrix Matrix Matrix

Antianginal Antianginal Antianginal Antianginal Antianginal

Catapres Duragesic Kimete Patch Transderm – Scop Estraderm Minitran

Alza/Boehringer Ing. Alza/Ivers/Jansen Myun Moon Alza/Ciba Alza/Ciba 3M

Reservoir Reservoir Reservoir Reservoir Reservoir Reservoir

Niconil Nicoderm Nicotrol

Elan Alza Cygnus

–— –— –—

Nitroglycerin Nitroglycerin Nitroglycerin Nitroglycerin Isosorbide dinitrate Clonidine Fentanyl Scopolamine Scoparamine Estradiol Glyceryl trinitrate Nicotine Nicotine Nicotine

Antihypertensive Narcotic analgesic Antimotion sickness Antimotion sickness Hormonal Antianginal Antinicotinic Antinicotinic Antinicotinic

7.3.1.2 Materials The major parts of TDD systems are composed of polymers. The drug impermeable layer, the drug reservoir, the pressure adhesive layer and the artiﬁcial membrane are all prepared from polymers. The range of the polymers used is really broad; natural polymers (gelatin, starch, etc.), semisynthetic (hydroxyl propyl cellulose, nitrocellulose, cellulosic), synthetic (polysiloxane, polybutadiene, polyisoprene, silicone rubber, polyesters, polyurethane, polyethylenevinylacetate polyacrylamide, polyvinylalcohol, polysulfone, polymethylmethacrylate, etc.). The reader can ﬁnd excellent overview of the polymers used elsewhere [157,159,160]. In the membrane (or reservoir) system, the membrane is generally the part in direct contact with the skin and acts as the interface between the drug reservoir and

Fig. 7.15 Matrix TDD system.

213

214

7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

the skin to give optimal control for the transdermal drug delivery. The membrane should have the following requirements: It should be made of biocompatible material to avoid skin irritation. It should control the drug delivery (the permeability of the drug should be lower through the membrane than through the skin). In this case the transdermal bioavailability of the drug becomes independent of any possible intra- and/or interpatient variability in skin permeability. The drug adsorption on it should be low. The issue of the drug-controlled delivery will be addressed in more detail in the following section. 7.3.1.3 Skin or Device-Controlled Delivery The issue of skin or device-controlled delivery has been discussed in the scientiﬁc community for a long time. In the transdermal drug delivery by a patch (see Figure 7.2), the total permeability (KD,total) of the drug through the membrane and the skin is given by 1 1 1 ¼ þ ; KD;total KD;mem KD;skin

ð10Þ

where KD,memb and KD,skin represent the permeability of the drug through the membrane and the skin, respectively. Depending on the ratio of KD,memb and KD,skin, the delivery may be primarily skin-rate controlled or primarily membrane-rate controlled. When the ratio KD,memb/KD,skin is less than 0.2, the delivery is considered to be membrane controlled. When the ratio KD,memb/KD,skin is larger than 5, it is considered to be skin-rate controlled. If the ratio KD,memb/KD,skin is in between 0.2 and 5 the systematic dosage received is controlled by both the skin and the membrane. As we have seen earlier, the passive drug transport can have great intra- and interpatient variability. Moreover, the delivery of some drugs should have a very welldeﬁned therapeutic window (i.e., where a dose above a certain limit is ineffective or even toxic to the patient). To ensure that the drug delivery through the patient skin is always the same, the device should have a major control on the delivery. Based on this fact, a lot of scientists have worked on patches where the drug delivery was designed to be much lower than through the skin. However, for most drugs the safety issues are not that critical. Moreover, the passive TDD is so low that even great variability between patients cannot cause safety problems. In these cases, in order to deliver as much drug as possible, the device should not impose any restriction and/or control on the drug delivery. It should only be used as a storage of the drug. Most of the modern TDD systems have no controlling membranes. Often, devices with quite different release kinetics when in contact with buffer solutions may perform quite similarly when in contact with skin in vitro and in vivo. In these cases, the skin controls the delivery, and possible variability in the actual delivery

7.3 Transdermal Drug Delivery System – Structure/Design

arises solely from variations in patients’ skin permeability. It is important to note that often in the literature for TDD, the term ‘‘rate controlling’’ is unspeciﬁed, if not misleading. A characteristic example has been presented in the work of Hadgraft et al. [161]. The authors studied the delivery of nitroglycerin from four different commercial ‘‘rate controlling’’ TDD systems having different drug loading and surface areas: Transderm – Nitro: drug ¼ 10 mg, area ¼ 20 cm2 Nitrodur II: drug ¼ 80 mg, area ¼ 20 cm2 Deponit: drug ¼ 32 mg, area ¼ 32 cm2 Minitran: drug ¼ 36 mg, area ¼ 13.3 cm2. All these systems were designed and conﬁrmed to deliver [161] in vitro 10 mg in 24 h. However, the differences in their speciﬁcations do not allow the estimation of their control in drug delivery. Guy and Hadgraft revisited this study [162] and calculated the fractional rate control of the device by the equation: Mdevice ðfractional control by deviceÞ ¼

Qtotal ; Qdevice

ð11Þ

where Qtotal and Qdevice represent the amount of the drug transported in a given period of time under steady state conditions through the combination of device and skin and the device alone (in mg/cm2), respectively. The fractional control by the skin is then given by MS ¼ 1 Mdevice. Thus, Mdevice ¼ 1 implies that the delivery is controlled entirely by the device (often the membrane); however, when Mdevice < 1, the skin is contributing to the control process. Guy and Hadgraft found [162] that only Deponit plays a signiﬁcant role in the control of nitroglycerin delivery (Mdevice ¼ 0.87). For the others, the results were mixed: Transderm – Nitro/Mdevice ¼ 0.45; Nitrodur II/Mdevice ¼ 0.13 and Minitran/Mdevice ¼ 0.28. One of the key parameters to be matched in TDD is the amount of drug absorbed by the skin on time. Moreover, for safety reasons, the drug loading should be as close as possible to the amount absorbed by the skin. 7.3.1.4 Commercialization – Patents In the ﬁeld of passive transdermal delivery, several companies have been active in the past 15–20 years. Since the beginning of the 1990s, Alza, Merck, and Ciba/Novartis are the ‘‘big players.’’ In recent years, 3M presents a dynamic proﬁle in the area as well, and it has marketed a series of products used in transdermal patches such as membranes, impermeable backing layers, adhesives, and so on. In the patent literature, one can ﬁnd hundreds of inventions concerning passive TDD. In these patents, the application of a rate controlling dense or microporous membrane for the drug delivery plays a rather important role. The claims mainly address the use of speciﬁc membranes. These membranes either exclusively control the drug delivery or partially control the drug transport together with other components of the invention, such as the adhesive layer. The membranes are either commercial or laboratory-made. They cover a big list, including hydrophobic and

215

216

7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

hydrophilic materials. They may be symmetric or asymmetric, and their pore size may vary from 1 nm to 100 mm but preferably lower than 5–10 nm. Table 7.8 presents a selection of recent patents for passive TDD systems. 7.3.2 Active TDD Systems 7.3.2.1 Types Iontophoresis and electroporation are the most widely used active systems. In the past decades, they have been broadly tested for drug delivery in the clinic. In iontophoresis, a broad range of drugs has been tested in vitro, and in vivo such as steroids, lidocaine, nonsteroidal inﬂammatory drugs (NSAIDs), histamine, antihistamines, antibiotics, and others [5]. In electroporation, mostly the big drug molecules have been delivered, such as heparin, calcitonin, luteinizing hormone–releasing hormone (LHRH), and others [5]. 7.3.2.2 Materials – Devices The basic elements of the active patches are common to those of the passive patches. Polymers are also used for the drug impermeable layers and the skin contacting membrane. Buffer solutions and/or gels are used for the drug formulation, and so on. Of course, in iontophoresis and electroporation, speciﬁc attention is given to the construction of the patches due to the presence of the electrodes for the current application. Often to avoid direct contact with the drug, the electrodes are placed in a separate compartment created by a membrane. The separate compartment is ﬁlled with conductive gel for the current conduction. The basic drug formulation composition varies between manufacturers, but typically is a highly conductive gel, too. In iontophoresis, mostly Ag/AgCl electrodes are used. The electroporation electrodes are mostly manufactured from inert material (Pt) and have a meander structure for better voltage distribution and patient safety. Figure 7.16 presents some examples of commercially available iontophoresis electrodes. In almost all patches – electrodes, an artiﬁcial, nonrate-limiting membrane, is used in contact with the skin. The drug delivery is regulated via the external driving forces, namely the current (iontophoresis), the voltage pulses (electroporation), the ultrasound (sonophoresis) and so on. In these cases, inter- and intrapatient variability in drug – skin permeability is much lower than in the passive systems. In iontophoresis, the electrodes are connected to a relatively small (‘‘walkman’’ or ‘‘discman’’ size) power supply. In the United States, several companies manufacture such devices. Typical examples of such the devices are Dupel (Empi, St Paul, MN), Phoresor II (Iomed, Salt Lake city, UT), Iontophor II (Dynatronics, Salt Lake city, UT), and others [5]. Figure 7.17 presents an example. The electroporation power supplies have been primarily manufactured for genetic manipulation of living cells and have been adapted for TDD. These devices are much bigger in size than the iontophoresis devices because of the need of the capacitor. A partial list of manufacturers in the United States includes Genetronics Inc. (San Diego, CA), CytoPulse Sci. (Columbia, MD), Bio-Rad

7.3 Transdermal Drug Delivery System – Structure/Design Tab. 7.8

Selection of patents on passive TDD.

Patent

Inventor/owner

Short description

DE 19738643/1999 US 4951657/1990

Mueller – Lohman Pﬁster et al.

US 5869089/1999

Risheng Wu

US 5585111/1996

Peterson

US 4913905/1990

Fankhauser – Ciba

US 5904930/1999

Fisher et al.

US 5284660/1994

Lee et al.

US 5462745/1995

Enscore et al.

CZ 287678/2001

Bracher et al.

WO 0130316/2001

Dreyer – 3M

WO 0003698/1999

Unihart Corp

US 5613958/1997

Kochinke et al.

EP 0916339/1989

Pharmacia et al.

WO 0023644/2006 US 078601/2006 US 121102/2006

3M Noven Pharma C.C. Ming

WO 058287/2005

B. Stefan et al.

Regulation of scopolamine delivery by polyethylenevinylacetate (PEVAc) membrane Dense membrane of polydimethyl siloxane (PDMS) and polyurethane (PU) for TDD Nuclear track microporous membrane made from alpha particles or modiﬁed EVAC for TDD Dense membrane (Cotran 9702, PEVAc, etc.) or microporous (Cotran 9710) controls TDD Membrane (dense or microporous) controls TDD. Material could be hydrophobic and hydrophilic, and the membrane could be symmetric or asymmetric Delivery of tamoxifen. Microporous membrane (polypropylene, PEVAc, and silicone) controls the drug delivery to a varying degree. When skin penetration enhancer is applied, the membrane controls TDD Membrane (of hydrophilic or semihydrophilic polymer) impermeable to drug in a dry state and permeable in a hydrated or wet state. The patch, wetted manually by sweat or just prior to use Drug membrane ﬂux no greater than the drug skin ﬂux (dense Cotran 9702, low-density polyethylene membrane). Rate control is achieved by the thickness of both membranes – adhesive Testosterone delivery. The reservoir is separated from the skin by the rate-controlling adhesive layer Drug in adhesive reservoir layer the thickness of which controls the drug delivery Concentration of solubilizers, penetration activators, or providing a membrane control the delivery An adhesive layer containing plasticizer type enhancer (10–40 %). Membrane (optionally used) having pore size between about 0.01 and about 3.0 mm and being nonrate-controlling membrane Adhesive and/or preferably nonporous membrane controls nicotine delivery Reservoir TDD system Transdermal system for delivery of estradiol Transdermal system for delivery of estrogen and/or progestin Transdermal system for delivery of hormones

217

218

7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

Fig. 7.16 Iontophoresis electrodes from Iomed. (Source: Iomed, Salt lake City, UT, USA, 2005, printed with permission.)

(Richmond, CA), and others [5]. Figure 7.18 presents an example of power unit for iontophoresis and electroporation (printed with permission from Moor Instruments Ltd, Devon, UK). 7.3.2.3 Commercialization – Patents Commercialization of iontophoresis/electroporation patches for self-use by the patient is still not achieved. The drug delivery is still rather complex for the patient and the cost of the technology high. Besides, regulatory approval of such systems seems to be more difﬁcult than the passive systems. The manufacturers should present to the regulatory authorities, besides information about the drug, a detailed description of the electronics of the patch and the equipment, as well as of the drug delivery protocols (e.g., current density or voltage, time of application, etc.). In principle, the patches – electrodes can be ﬁlled with the drug formulation of choice and can be broadly used. The drug delivery protocols, however, should be drug speciﬁc. Iontophoresis is already approved in the United States, for the topical delivery of lidocaine and epinephrine for local analgesia (Iomed, USA). Devices available on the market for delivery of local anesthetics and corticosteroids include Phoresor II

Fig. 7.17 Phoresor II auto iontophoresis device from Iomed. (Source: Iomed, Salt lake City, UT, USA, 2005, printed with permission.)

7.3 Transdermal Drug Delivery System – Structure/Design

Fig. 7.18 Iontophoresis and electroporation devices from Moor Instruments Ltd. (Source: Moor Instruments Ltd, Devon, UK, printed with permission.)

(Iomed), Empi Dupel (Empi, USA), Life-Tech Iontophor (Life-Tech, USA), and Henley Intl Dynaphor (Henley Intl, USA). In addition, devices for iontophoresis of pilocaprine are on the market, including the CF Indicator (Scandipharm, USA), among others. Several companies have been trying to commercialize miniature patch systems. A partial list includes the Alza Corp. (USA), Becton Dickinson (USA), Fournier (France), Hisamitsu Pharm. (Japan), and Cygnus (USA). In addition, companies such as Thera tech and Genetronics have developed electrotherapeutic devices, electrodes, and so on applied in transdermal delivery systems with a combination of various driving forces (such as iontophoresis þ electroporation, or iontophoresis þ sonophoresis, or iontophoresis þ magnetophoresis/magnetoporation). Table 7.9 presents a selected list of patents describing active TDD systems. The iontophoresis and electroporation systems seem to dominate the ﬁeld. In most of the patents the control of the drug delivery is left to the applied driving forces. The role of the membrane is rather limited to either a holder or support of the drug reservoir and/or to prevent contact of the drug solution and the electrode. In these cases, its porosity and pore size are just enough to avoid uncontrolled leakage of the drug to the skin surface and skin ﬂuids to pass to the drug reservoir. However, there are inventions that provide rate control by the membrane to avoid: (i) Dependence of the drug transported by iontophoresis through the skin on the skin electrical resistance (therefore the transdermal drug bioavailability becomes dependent on any possible intra- and/or interpatient variability in skin permeability). In patent US 5167616, the iontophoretic ﬂux of metoclopramide coming from the individual’s back through skin could be 50–150 % higher than through other skin sites because of lower electrical resistance of the skin at the back site. (ii) There are safety features to prevent excessive drug electrotransport if the patch is applied on damaged skin or on body surface that has somehow been compromised. (iii) If the permeability of the drug through the membrane material is much higher than the permeability of the drug through skin, the accumulation of a large quantity of the drug on the skin surface could cause skin irritation. In EP 0847775A1/Mori-Hisamitsu Pharm., and US 6167301/Flower et al., the above-mentioned disadvantages have been extensively addressed, and an

219

220

7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

iontophoresis system that measures the resistance of the skin and in response to adjusts the output voltage (drug delivery) is proposed. In several other patents the membrane-controlled drug delivery has also been proposed (see Table 7.9). Finally some patents describe the combination of driving forces, for example, in patent WO 0035533/Hofmann – Genetronics, the combination of iontophoresis and electroporation for the delivery of drugs and genes is described. In US 5947921/Johnson et al., the application of chemical and/or physical enhancers and ultrasound for transdermal drug delivery is proposed, and an increase of permeability of magnetic particles is claimed if a magnetic ﬁeld is applied in combination with ultrasound. Tab. 7.9

Selection of patents on active TDD.

Patent

Inventor/owner

Short description

US 5667487/1997

Henley

US 5362308/1994

Chien et al.

EP 0748636A2/1996

Yanai et al. – TakedaHisamitsu Pharm

WO 0066216/2000

Kontturi et al.

EP 0813887A2/1997

Higo-Hisamitsu Pharm

EP 931 564/1999

Theeuwes et al.

US 5445607/1995

Venkateshwaran et al. – Thera tech

Application of ultrasound + iontophoresis. Drug in a porous material: polyethylene, paper, cotton silicone, PTFE, ceramic A membrane separates two gel layers. The top gel layer contains ion exchange resins and contacts the electrode. The bottom layer contains the drug (proteins and peptides including calcitonin) Iontophoresis delivery of protein and peptides. Low protein adsorptive membrane (hydrophilic or hydrophilized hydrophobic material) acts as a support for the drug (dry or in solution). An ion exchange membrane protects the drug from the electrode. The applied current could be constant or pulsed Iontophoresis delivery. An ion exchange membrane is used to separate the electrode from the drug solution-gel and a microﬁltration membrane is placed in contact with skin Current (constant or pulsed) delivers protein and peptides through animal skin. A patch contains an electrode–conductive gel reservoir and a separate drug loaded membrane No drug ﬂux is obtained when current is off and no linear increase of drug-ﬂux when the current is on Skin doesn’t control the drug delivery due to application of a penetration enhancer. The proposed rate controlling membrane should be electrically sensitive. No drug permeation during passive diffusion but linear increase of drug permeability with the current is obtained

7.4 Conclusions and Outlook 221

7.4 Conclusions and Outlook

The passive TDD technology is well established and has ‘‘won the hearts and minds’’ of the patients. The application and the market of the passive patches are expected to grow further in the next years. New products that are easier to use with better quality materials will be developed. The main expansion of TDD is expected to come from the active systems. Using active systems such as those based on iontophoresis and electroporation, the transdermal delivery of several more drugs with better patient compliance (reduced side effects, etc.) will become possible. Slowly these emerging technologies will overcome the technical difﬁculties and ﬁnd broader applications. One issue that is always in the spotlight is the need for rather bulky, heavy, and complicated power supplies and other components. However, the latest

Fig. 7.19 Iontophoresis patches containing (a) both electrodes on the same patch from Iomed (Source: Iomed, Salt lake City, UT, USA, 2005, printed with permission.) and (b) both electrodes and built-in battery, from Travanti Pharma Inc. (Source: Travanti Pharma Inc., St Paul, MN, USA, 2005, printed with permission.)

222

7 Drug Delivery Through Skin: Overcoming the Ultimate Biological Membrane

developments in miniaturization of systems might give a signiﬁcant help on this issue. The miniaturization of the components and the development of suitable microcomputer for regulation and control of delivery will make these technologies safer and therefore available outside the hospital, too. Figure 7.19 presents some example of patches containing both electrodes and/or the battery built in making their application much easier. Currently, the majority of the active systems are more expensive than conventional drug delivery systems. However, as the technology solves the technical problems, they will become cheaper. Nevertheless, even if the costs are relatively high the patient beneﬁts from these technologies may be able to justify the extra costs.

References 1 Barry, B. W. (2001) European Journal 2 3

4

5

6

7 8 9 10 11

12

13

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8 Application of Membranes in Tissue Engineering and Biohybrid Organ Technology Thomas Groth, Zhen-Mei Liu 8.1 Introduction 8.1.1 Application of Membranes in Blood Detoxiﬁcation

Membranes have found a multitude of applications in biotechnology and medicine where controlled exchange of solutes between different phases or compartments is required [1–4]. Particularly the application of membranes for blood detoxiﬁcation (hemodialysis) has been well established since several decades [1,2]. Nowadays, membranes for hemodialysis represent the largest market segment in the ﬁeld of membrane production [3]. Applications of membranes for blood detoxiﬁcation such as hemodialysis to treat kidney failure are linked to transport of solutes to remove (toxic) metabolites, salts, and water from the organism. On the other hand, the undesired loss of substances, such as proteins, must be prevented [2,4]. The transport process is basically controlled by the molecular cut-off of membranes not only to remove toxins and excess of salts and water but also to avoid any passing of essential components from blood, such as serum albumin and other proteins [4,5]. Elimination of substances from human blood can be also accomplished by speciﬁc or selective adsorption of substances onto the surface of membranes [6]. This has been achieved by the exploitation of the afﬁnity of molecules to membrane bulk materials (e.g., by hydrophobic interaction) [6,7] or to speciﬁc surface-attached functions and ligands [8,9]. Quite novel developments also aim to use imprinting technologies speciﬁcally to adsorb molecules from liquids, which may be interesting for applications in blood detoxiﬁcation as well [3,10,11]. However, it must be emphasized here that in the majority of biotechnological and biomedical applications of membranes, nonspeciﬁc adsorption and adhesion processes have to be minimized [1,4,5]. This is necessary to prevent biofouling, which may cause blocking of membrane pores by adsorbed proteins and attached cells from the surrounding liquids [1]. A further requirement is the prevention of the undesired activation of coagulation and inﬂammation during blood contacting applications, which represents

Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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life-threatening risks for the patients [2,4,5]. The control over adsorption and adhesion processes is basically obtained by using hydrophilic polymers for membrane formation, such as cellulose and its derivatives [12], blending or grafting of hydrophobic polymers with hydrophilic polymers, for example, poly(vinylpyrollidone) [13,14], using copolymers consisting of hydrophobic and hydrophilic co-monomers [15,16], and coating or grafting of membrane surfaces by hydrophilic macromolecules such as poly(ethylene glycol) or phosphatidylcholine polymers [17,18]. Also, the binding of bioactive molecules like heparin has been used to avoid surface-induced adsorption and activation phenomena at least in blood-contacting applications [19,20]. Overall most techniques to prepare membranes for blood detoxiﬁcation and oxygenation aim to reduce efﬁciently the adsorption of proteins and thus adhesion of cells from blood. 8.1.2 Requirements to Support Adhesion and Function of Cells

By contrast, applications in tissue engineering and biohybrid organ technology require that tissue cells adhere on material surfaces [21]. Adhesion of cells is a prerequisite for a multitude of cellular functions such as movement, growth, differentiation, and survival [22]. Cell adhesion is not a simple on–off event. It has been shown that the strength of adhesion controls the degree of cell spreading and hence the shape of cells on a substratum [23]. The shape of cells, however, seems to be a regulator of cellular functions [23,24]. For example, it has been shown that spreading of cells from connective tissue such as ﬁbroblasts promotes their growth and function in terms of extracellular matrix (ECM) synthesis [25,26]. Figure 8.1 shows as an example the effect of surface chemistry and wettability on the shape and functional activity of human dermal ﬁbroblasts. It is visible that a methyl-terminated, nonwettable surface inhibits adhesion and spreading of cells (Figure 8.1a). As a consequence, the synthesis of a ﬁbronectin matrix is greatly reduced (Figure 8.1c). In contrast, an amine-terminated, wettable surface promotes spreading of cells (Figure 8.1b) and ﬁbronectin matrix synthesis (Figure 8.1d). However, epithelial cells such as hepatocytes tend to dedifferentiate if they spread too strongly on a substratum while they maintain their functional activity when attached moderately on a substratum having a round morphology [27,28]. Figures 8.2 and 8.3 show as an example human hepatoblastoma cells – a hepatocyte cell line cultured on a highly adhesive support (Figure 8.2a) in comparison to a less adhesive substratum (Figure 8.2b) on which cell spreading is greatly reduced. The copolymer in Figure 8.2a possesses primary amine functions and is called aminoethylmethacrylate (AEMA), while the copolymer in Figure 8.2b contains N-vinylpyrollidone (NVP) as co-monomer and is named NVP 20 (see also Section 2.2). Please also note that cells make stronger intercellular contacts on the less adhesive membrane, which contains NVP 20. If the functional activity is investigated as shown in Figure 8.3, it is visible that the detoxiﬁcation capacity of cells, which is measured by the conversion of 7-ethoxycoumarin (ECOD), is lower on the highly adhesive substratum (AEMA) in comparison to the less adhesive NVP 20. From the literature it is further known

8.1 Introduction 229

Fig. 8.1 Effect of surface chemistry on shape, adhesion formation and functional activity of dermal ﬁbroblasts. Cells on methylterminated self assembled monolayers (SAM, a and c) have reduced spreading and lack focal adhesions (a), which is corroborated by reduced functional activity of cells indicated by the absence of ﬁbronectin matrix formation (c). Only spots of ﬁbronectin have been deposited on the surface. In contrast, amine-terminated SAM (b and d) promote

adhesion and spreading of cells with formation of well-developed focal adhesions, which are visible as white streaks in the cell periphery (b). This is accompanied by the formation of an abundant extracellular matrix of ﬁbronectin (c). Cells and matrix were stained with ﬂuorescence-labeled antibodies against vinculin (a and c) and ﬁbronectin (b and d) and visualized by ﬂuorescence microscopy. Bars in upper panel have a length of 100 mm and in the lower panel of 50 mm.

that a complete failure of cellular attachment particularly for epithelial cells results in the onset of apoptosis, a mechanism by which cells undergo controlled cell death [29]. It has been shown that cell adhesion is linked to the presence and conformation of speciﬁc attachment proteins on material surfaces [30]. Adhesive proteins such as ﬁbrinogen, ﬁbronectin, vitronectin, and von Willebrand factor are normal components of blood plasma and may adsorb on biomaterial surfaces [31]. The ECM, which surrounds cells in tissues, is composed of structural proteins, like collages, adhesive

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Fig. 8.2 Comparison of C3A hepatoblastoma cell attachment on polymer membranes composed of highly adhesive amine group containing copolymer AEMA (a) and a hydrophilic low adhesive copolymer NVP 20 (b) by staining cells for vinculin. Note the large cell size on AEMA and the reduced spreading on NVP 20. Size of scale bar is 50 mm.

proteins, and glycosaminoglycans [32]. These proteins provide attachment to cellular receptors mainly integrins and deliver signals important for surivival, growth, and differentiation [33,34]. The physicochemical properties of the material surface such as surface energy (wettability) and electrical surface potential (zeta potential), which are

Fig. 8.3 Comparison of metabolic function of hepatoblastoma cells in terms of P450 activity indicated by the ability of cells to convert 7-ethoxycoumarin. It is shown that cells cultured on the less adhesive substratum NVP 20 have a signiﬁcantly higher activity than cells on more adhesive substratum AEMA.

8.2 Application of Membranes in Tissue Engineering

dependent on the chemical composition have an impact on the adsorption of proteins from surrounding liquids, such as blood or tissue ﬂuids [35,36]. Accordingly, cell growth and function were found to be strongly related to the wettability of materials [37–39]. Thepresenceofpolarorchargedfunctionalgroups,suchasaminoorcarboxylic groups, has been identiﬁed as promoting, while apolar groups like methyl may inhibit cellular attachment, growth, and function [40,41]. Hence, the tailored synthesis of copolymers or the modiﬁcation of biomaterial surfaces to obtain speciﬁc quantities of functional groups can be used as a tool to optimize the biocompatibility of materials. It is also known that cell substratum interactions depend on the underlying topography [42]. Even quite small topographical changes have been shown to cause certain cellular responses [43]. For example, structures in the range of 100 nm are sufﬁcient to induce guidance of cells [44,45]. There is also evidence for the inﬂuence of surface porosity on cell surface interactions [46–48]. It was demonstrated that changes in surface topography created by variation of pore size in the range of 0.1–0.8 mm supported corneal epithelia outgrowth, in contrast to the inhibitory effect of pores in the range of 0.9 mm or greater [49]. Earlier studies have also indicated that skin outgrowth can be promoted on membranes with pores in the range of 0.025–1.2 mm, while tissue growth was inhibited on surfaces having pores larger than 3–8 mm [46]. The overall chemistry and topography of material surfaces can be used to control the adhesion and function of cells. This is particularly important when low-cost applications are envisaged, which shall allow a later biomedical application of the biomaterial.

8.2 Application of Membranes in Tissue Engineering 8.2.1 Introduction to Tissue Engineering and Membrane Applications

Tissue engineering has been deﬁned by Langer and Vacanti as a novel discipline aiming at the restoration of tissue structure and function [50]. This shall be achieved by the cooperation of three components. First, some type of scaffold that shall guide tissue regeneration gives mechanical support to cells. The scaffold is merely a degradable material, which shall dissolve in a timescale necessary for cells to replace it by their own extracellular matrix components (e.g., newly formed bone). The scaffold possesses typically certain porosity to allow the colonization with cells, neovascularization, and transport of oxygen, nutrients, and waste products [51]. Second, the scaffold may be loaded or covered with bioactive molecules, which stimulate anchoring, growth, and differentiation of cells [52]. However, this is not an absolute requirement. The material or its surface may have some inherent bioactivity, which stimulates cells (e.g., calcium phosphates for osteoblasts) [53]. Third, the scaffold must be colonized by cells, which can occur in vitro or in vivo (in situ). The cells are preferentially autologous, obtained by biopsies and seeded on the scaffold. Another possibility is that cells colonize the scaffold by migration and ingrowth from surrounding body ﬂuids or tissues [54]. In most tissue engineering applications,

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Fig. 8.4 Scanning electron micrograph showing the establishment of a multilayer of keratinoctyes on a polymer membrane.

degradable scaffold materials are an absolute requirement. However, there are also applications when nondegradable materials are possible. This holds in cases when guidance of tissue regeneration is required as it is known for bone regeneration (see Section 2.3) or speciﬁc arrays of cells, such as for an artiﬁcial retina that must be maintained for a long time in a speciﬁc spatial arrangement [55]. A possible disadvantage of using nondegradable materials is that a second surgical intervention may be needed to remove the material after healing. Membranes can be prepared from a variety of materials including degradable and nondegradable polymers [1–4]. (Flat) membranes are interesting in applications, when two-dimensional tissues shall be replaced. This holds for regeneration of epithelia, which represent single or multilayers of cells covering the outer and inner surface of our body. Figure 8.4 shows as an example of such a structure a multilayer of keratinocytes cultured on a polymer membrane. Membranes may also provide guidance for cell growth in certain applications or can be used in a tubular setting either made of hollow ﬁbers or (rolled) ﬂat membranes to provide guidance for regeneration of bone, nerves, blood vessels, and other tubular tissues. Table 8.1 gives a general overview about applications of membranes in tissue engineering. The following sections will highlight some established or potential applications of membranes in the ﬁeld of tissue engineering. 8.2.2 Tissue Engineering of Skin

Treatment of deep-skin defects such as burns or chronic skin ulcers in elderly or diabetic patients represents a serious problem in medical care. Medical therapies aim to restore the normal skin structure, which consists of an underlying dermal part with ﬁbroblast and other types of cells, and also blood vessels embedded in extracellular matrix material overlaid by the epidermis as a nonvascularized multilayered epithelium [71,72]. This includes the restoration of the dermal part as a prerequisite for proper regeneration of the epidermis [72]. Application of autologous split skin grafts is still the best medical standard but not feasible if the wound size is too large or the constitution of the patient does not allow a further skin injury at the donor site [73,74]. First attempts to repair deep-skin defects by tissue engineering were directed to the regeneration of the epidermis. Rheinwald and Green developed more than 30 years

8.2 Application of Membranes in Tissue Engineering Tab. 8.1

General overview on membrane applications in tissue engineering.

Membrane sources

Applications

Reference

Poly[(ethylalanato)1.4(imidazolyl)0.6 phosphazene] ﬁlm Polylactide ﬁlm Polylactide membrane Polyurethane membrane Polyurethane membrane Polyurethane membrane Polylactides and polyurethanes membrane mixture Expanded polytetraﬂuoroethylene Collagen membrane Collagen membrane Collagen membrane Polyglycolide membrane Chitosan membrane Heparin–chitosan complex membrane Genipin-crosslinked gelatine membrane

Peripheral nerve repair

[56]

Reduction of posterior adhesion Diaphyseal bone regeneration Potential for ventricles Potential for skin repair Potential for pancreas Skin repair

[57] [58] [59] [60] [61] [62]

Periondontal tissue regeneration Skin repair Periodontal tissue regeneration Repair injured urinary tract Bone regeneration Skin repair Skin repair Skin repair

[63] [64] [65] [66] [67] [68] [69] [70]

ago a technique to prepare cultured epidermal autografts (CEA) in vitro, which are derived from skin biopsies of the patient [75]. The cells grow in vitro up to conﬂuence as a multilayered keratinocyte sheet, which has to be separated from the culture ﬂask by enzymatic treatment. However, the CEA are quite fragile because they lack any mechanical support. Procedures for their production are quite labor consuming and require at least three weeks. Unfortunately, the healing of these autografts is often worse than that of split skin grafts. Even months after transplantation, the epidermis can still get lost due to blistering [72,74]. Ideally, wound dressings, which are used for temporary coverage of skin injuries have to adhere to the wound site and must be porous enough to allow oxygenation and removal of wound exudates. Furthermore, they must prevent dehydration and infection of the wound [74,76]. It has been recognized that polymer membranes meet the majority of these requirements. If keratinocytes could be cultured on membranes they would combine the properties of a wound dressing and at the same time being a carrier for cell transplantation. So far, a limited number of approaches have been put forward to apply membranes in tissue engineering of epidermis. One of the examples is the application of a polyurethane membrane (HydroDerm, HD Innovative Technologies, Ltd), which has been developed purposely as a wound dressing [77]. Polyurethanes have been known for their good biocompatibility in blood-contacting applications. Also, their mechanical properties can be varied depending on their chemical composition from stiff to elastic materials [78]. Therefore, they are applicable as wound dressings [77]. In the study described here, keratinocytes were seeded in vitro on the polyurethane membrane. It was observed that cell growth was delayed when compared to standard tissue culture polystyrene. However, under in vitro culture conditions, the formation of a multilayered epidermis was observed, indicating that the

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membrane supports the differentiation of cells [77,79]. If the membrane was placed upside-down at early pre-conﬂuent stages of cell growth on the wound of an animal model, then a partial reformation of an epidermis was obtained. To achieve this, cells must be able to transmigrate from the membrane surface to the wound bed to form a basal layer of keratinocytes [80]. The delayed growth of keratinocytes might be attributed to the chemical composition of the polyurethane membrane, which was not especially adopted to promote keratinocyte attachment and growth. To adjust the chemical composition of membranes to the requirements of speciﬁc cells in different applications, we have developed a number of membrane types, which are based on copolymers of acrylonitrile [81–83]. The co-monomers applied here were acrylonitrile, hydrophilizing, nonionic N-vinylpyrollidone (NVP 5, 20, and 30 mol%), aminoethylmethacrylate or aminopropylmethacrylate (APMA, both about 1 mol%), and sodium methallylsulfonate (NaMAS, about 2 mol%). Copolymers were compared with the poly(acrylonitrile) (PAN) homopolymer. Membranes were prepared by phase inversion and possessed porosities in the ultraﬁltration range. Interestingly, membranes prepared from copolymers AEMA or APMA promoted growth of keratinocytes in vitro [83]. Figure 8.5 shows a comparison of scanning

Fig. 8.5 Comparison of cultures of keratinocytes on PAN (a and c) and APMA (b and d) with scanning electron microscopy after 7 days of culture with low (a and b) and high (c and d) magniﬁcation. Note that the aminefunctionalized copolymer APMA supports rapid monolayer formation (b) with close cell–cell-contacts (d) compared to PAN where no monolayer of cells was established (a and c).

8.2 Application of Membranes in Tissue Engineering

electron micrographs of keratinocytes cultures on PAN (a and c) and APMA (b and d). It is visible that cells on APMA cover the entire membrane surface and have close intercellular contacts. In contrast, keratinocytes cultured on PAN did not reach a conﬂuent monolayer. Also, intercellular contacts between the cells are less well expressed. Moreover, it was possible to show in organotypic coculture models that keratinocytes placed upside-down on a dermal equivalent (composed of a collagen gel with embedded ﬁbroblasts) were able to transmigrate from the membrane into the artiﬁcial wound bed and started to grow and differentiate there, which means to build up a neo-epidermis [82]. Figure 8.6 shows a cross section through membranes with neo-epidermis and dermal equivalent stained with hematoxilin-eosin. It is shown that the APMA membrane (a small and c high magniﬁcation) promotes the formation of a neoepidermis in contrast to a hydrophilic NVP membrane (c small and d high magniﬁcation). To explore further the possibility that polymer membranes with amine groups may support growth and differentiation of keratinocytes, we developed a blend made of poly(ether imide) (PEI) and poly(benzimidazole) and prepared membranes by phase inversion method [83]. PEI has the advantage of good membrane formation properties and high-thermal stability [84], which makes this polymer interesting for biomedical applications. It was found in this study, that the membrane could support the growth of keratinocytes and may be used as a temporary carrier of keratinocytes as well. During our studies we also discovered the possibility

Fig. 8.6 Organotypic culture of keratinocytes immobilized on APMA (a and c) and NVP 5 (c and d) membranes placed upside-down on a collagen gel with embedded ﬁbroblasts for 14 days. Constructs were ﬁxed and cryosections were stained with hematoxylin/

eosin. It is visible that APMA supports the formation of a neoepidermis (a and c) indicated by the thick blue layer beneath the membrane compared to NVP 5 (b and d). (a and b) Magniﬁcation 10, (c and d) magniﬁcation 20.

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to functionalize PEI by a simple wet chemical procedure using different diamines or polyamines [85]. It was observed that covalent binding of poly(ethylene imine) (PETIM) of lower molecular weight may provide advantages to keratinocyte culture, such as rapid establishment of a pre-conﬂuent cell layer on the surface and subsequent release of the cell layer from the membrane support [86]. While the ﬁrst organotypic culture models have shown advantageous effect of the polymer compositions and also ﬁrst animal experiments have been carried out, clinical studies are still necessary. Beside synthetic polymers, natural polymers or combinations with polyesters have also been used to prepare membranes for the transplantation of keratinocytes. The advantage of using natural polymers or polyesters, such as poly(e-caprolactone) (PCL), is that materials are degradable, which avoids the necessity to remove the material at later stages of medical treatment. Hyaluronic acid (HA) is a major component of the extracellular matrix in connective tissues and plays a promoting role in wound healing [87]. However, HA is water soluble and forms hydrogels, which have poor mechanical properties. Therefore, HA has been esteriﬁed with benzyl groups and crosslinked, which decreases its solubility in water and improves its mechanical stability [88]. This material has been used to prepare different kinds of scaffold materials for engineering of skin, cartilage, and other tissues. Thereby, esteriﬁed hyaluronic acid (HYAFF) has been used as a delivery system for keratinocytes. The commercial product is produced by Fidia Advanced Biopolymers and denoted as HYAFF-11 or Laserskin. It is a membrane that additionally contains a regular pattern of large pores generated by a laser. Keratinocytes are seeded on both sides of Laserskin and may migrate through the pores to cover the underlying dermal structures of the skin. Laserskin is applied mainly for the treatment of nonhealing skin ulcers with good success [89]. Also, membranes made of blends from collagen and PCL are used for tissue engineering of epidermis and dermis. One of the examples is a membrane based on a collagen scaffold, which is prepared by a freeze drying process and subsequent impregnation with PCL solution [90]. The resulting scaffold represents a microporous membrane with large pores in the range of about 40 mm. It was shown that both dermal ﬁbroblasts and keratinocytes were able to grow on this scaffold. A speciﬁc coculture model was developed where the membrane separates the ﬁbroblast building the dermal component and keratinocytes making the epidermis [90,91]. Such a construct can be used in the treatment of deep-skin defects when not only epidermis but also the dermal part of the skin must be replaced. Recently, pure PCL membranes have also been suggested as a tool for tissue engineering of skin [92]. These membranes were prepared by solution casting, subsequent biaxial stretching, and laser modiﬁcation to obtain large pores for transmigration of cells. First experimental work with cell cultures could show that these membranes can be applied as culture vehicle for both ﬁbroblasts and keratinocytes [92,93]. 8.2.3 Tissue Engineering of Bone

Bone (also called ‘‘osseous tissue’’) serves multiple functions in the body, including support of body structure, protecting internal organs, acting as major calcium

8.2 Application of Membranes in Tissue Engineering

reservoir and facilitating movement (in conjunction with muscles). One of the most common bone diseases is bone fracture. Autologous bone grafting remains the gold standard for the treatment of bone defects. The main advantages of autologous bone grafting are its viability, immunocompetence, osteogenic activity, and high incorporation rate. However, donor site morbidity, long operative time and bleeding, risk of infection, rapid graft resorption, and limited availability are common restrictions of this procedure [94]. The applications of polymers for bone tissue repair and reinforcement, for regeneration of cartilage associated with bones, for helping in partial replacement of bones by metallic parts, and as carriers of antibiotics to the infected bone tissues have led to the development of bone tissue engineering [95]. By the implantation of dental and orthopedic devices and bone-substitute materials, bone tissue engineering has been practiced since a long time in the repair of bone fractures [96]. In 1956, Bassett et al. ﬁrst used a physical barrier to prevent the migration of unproﬁtable cells to the wounded region and to enable the proliferation of useful cells [97]. Since 1960s, polymeric membranes have been used in experimental animals to facilitate healing of bone defects [98]. In 1982, Nyman et al. attempted the ﬁrst utilization of guided tissue regeneration (GTR) for the periodontal surgery [99]. The basis of GTR technique depends on the use of a mechanical barrier placed between the root surface and the full-thickness periodontal ﬂap. After the placement of a Millipore ﬁlter membrane, a histological analysis demonstrated new cementum formation and the insertion of Sharpey’s ﬁbers on the previously diseased root surface. Dahlin et al. [100] originally applied this technique to bone regeneration in a bone loss area to establish the concept of ‘‘osteopromotion.’’ Based on Dahlin’s technique, Buser et al. [101] proposed the term of ‘‘guided bone regeneration’’ (GBR), which aims to promote bone augmentation by a barrier membrane. GBR membranes also have an important function that encourages bone growth. GBR membranes are occasionally utilized with dental implant [102] or bone-grafting materials [103]; a typical membrane set-up for bone defects is illustrated in Figure 8.7. A number of membrane forms and materials have been applied in animals and humans by GTR or GBR technique to repair bone diseases. The below table is a list of membranes adopted in bone tissue engineering (Table 8.2). As an effective barrier in GTR or GBR, utilized membranes can be divided into two categories: resorbable and nonresorbable. The most popular nonresorbable membrane utilized for GTR or GBR is extended polytetraﬂuoroethylene (ePTFE) [105], which is a biological inert material and can be safely applied clinically. Many studies have demonstrated successful bone regeneration by using ePTFE alone [123,124] or in combination with ﬁlling material for bone augmentation and regeneration [125]. In a bilateral mandibular defect, the use of high-density PTFE membranes to facilitate guided bone regeneration in the rat was examined [126]. On both the medial and lateral aspects of the mandible, the experimental side was covered with PTFE membrane, while the opposite side served as a control. At 10 weeks postoperatively, complete ossiﬁcation was noted on the PTFE treated side, while the control sides displayed little osseous regeneration.

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Fig. 8.7 Schematic presentation of membranes set-ups in a critical-size segmental bone defects in the sheep tibia. A. single external microporous membrane; B. two microporous membranes placed externally and internally in the defect as a ‘‘tube-in-tube’’ construct. The microporous

membrane used here is poly(L/DL-lactide) 80/20 %. M: membrane; OD: osteoperiosteal defect in the diaphysis of tibiae which was then stabilized with a bilateral AO external ﬁxator. Revised from Gugala, Z. and Gogolewski S. (2002) Injury, International Journal of Care Injured, 33, S-B-71–6.

Tab. 8.2

Overview on use of membranes in bone tissue engineering.

Year

Scientists

Membrane

Reference

1960 1961 1982 1984 1987 1987 1987 1988 1992 1992 1993 1995 1995 1999 1999 2000 2004 2005 2005

Lindhorne Goldhaber Nyman et al. Nyman et al. Yaffe et al. Pitaru et al. Blumenthal et al. Magnusson Gogolewski et al. Wanzhang et al. Vuddhakanok et al. Jansen et al. Nishimura et al. Ishikawa et al. Veronese et al. Wang et al. Lee et al. Desai et al. Liao et al.

[98] [104] [99] [105] [106] [107] [108] [109] [110] [111] [112] [113] [114] [115] [116] [117] [118] [119] [120]

2005

Fujihara et al.

2006

Kuo et al.

Polyethylene tube Millipore ﬁlter Millipore ﬁlter Gore-tex membrane (ePTFE) Collagen membrane Collagen membrane Collagen membrane Poly-L-lactic acid Polyurethane Calcium alginate ﬁlm DL-PLGA membrane PLA/hydroxyapatite membrane Cementum-impregnated gelatine membrane Alginate membrane Polyphosphazene membrane Polyesterurethane membrane Polycarbonate membrane Alumina membrane Hydroxyapatite/collagen/PLGA composite membrane Polycarprolactone/CaCO3 composite nanoﬁber Chitosan membrane

[121] [122]

8.2 Application of Membranes in Tissue Engineering

Despite the successful applications of nonresorbable membranes in bone tissue engineering, they must be removed by a second operation. This additional surgical trauma is a negative factor both to the patient and to the newly formed tissue. Hence, resorbable membranes have been developed for bone tissue engineering to avoid a second surgical intervention, and have become widely used in conjunction with GTR and GBR [127,128]. There are two types of resorbable membranes, namely synthetic polymers and natural biomaterials. Collagen membranes have been introduced as resorbable material in 1987 [106]. Since that time numerous applications of collagen membranes in GTR have been developed. For example, Taguchi et al. [127] used a collagen membrane for guided bone regeneration in rat maxillae. Standardized artiﬁcial bone defects (1.0 1.0 2.5 mm3) were covered with a collagen membrane (Bio-Gide, composed of porcine type I and type III collagen ﬁbers). After four weeks it was visible that membrane-associated and cavity-derived bone had completely ﬁlled the defect. The Bio-Gide membrane appeared to have so-called osteoconductivity, resulting in a well-augmented alveolar ridge. While collagen membranes have excellent cell afﬁnity and biocompatibility, they are normally weak in strength and difﬁcult to manipulate. Therefore, polylactides or copolymers of glycolide and lactide been found widespread applications for bone ﬁxation [94,109,112]. However, their acid degradation products may lead to an inﬂammatory response. Moreover, the cell afﬁnity of these polymers is rather low if compared to a collagen membrane. Hence, to meet the demands of resorbable membranes in GTR or GBR both in mechanical property and biocompatibility, composite membranes were developed. For example, a three-layered degradable composite membrane was prepared, which contained one layer of 8 % nano-carbonated hydroxyapatite/collagen/poly (lactic-co-glycolic acid) (nCHAC/PLGA), one layer of pure PLGA membrane, and one layer of 4 % nCHAC/PLGA [121]. Results revealed that the addition of nCHAC increased the biocompatibility and osteoconductivity, while the composite membrane had sufﬁcient mechanical strength. It was concluded that polymer-nanoparticle composite membranes are promising materials for bone tissue engineering application due to their superior mechanical properties, improved durability, and surface bioactivity [129]. For membranes used in bone tissue engineering, the healing efﬁciency depends on the surface chemical and physical properties of membranes. Usually the chemical properties are related to surface chemical groups that control surface free energy, surface charge, and hence the ability to support attachment and function of cells (see Section 2.2). Gogolewski et al. [130] tested the effects of polylactide membranes on the healing process of rabbits’ diaphyseal segmental defects. The polylactide membranes were made either of poly(L/D-lactide) or poly(L/DL-lactide). After one year, they found in most cases that both membranes led to complete bone regeneration in the defects. Obviously, the differences in quantities of stereo isomers of lactides did not have a signiﬁcant effect on osteoblasts. They further investigated the effects of plasma treatment of polylactide membranes on cellular response of osteoblasts. The membranes were subjected to low-temperature oxygen, ammonia, or sulfur dioxidehydrogen plasma treatments. The results revealed that for all treated membranes, the attachment and growth of osteoblasts were greater in comparison to untreated

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polylactide. The treatment with ammonia plasma, which generates different kinds of amine and amide groups on the surface, was found to be most efﬁcacious [131]. On the other hand, surface properties of membranes are also related to the surface morphology such as texture of surface, presence of pores, pore structure, pore size, and pore distributions. The effect of surface topography on osteoblast cultures was tested using different polyesterurethane membranes [117]. The membranes prepared had either a fairly dense and smooth structure (membrane A), or a rough and sunken surface supported by a porous sublayer (membrane B), or a porous surface with particles that were supposed to be generated by the polymer crystallization (membrane C). Rat osteoblasts culture experiments revealed that on membrane C there was a greater number of cells with better spreading and a ﬂattened appearance, which may be due to the presence of particles on the membrane surface. Degradation experiments further revealed that the degradation of membranes was obviously affected by the membrane’s surface morphology; rough surface could decrease the degradation rate, especially for those with particles on the surface. Due to its relatively fast degradation during early stages of healing, membrane A was found to be not suitable for the repair of bone defects [117]. For polycarbonate membranes with different pore sizes (pore diameter from 0.2 to 8.0 mm), on the other hand, the response of MG63 osteoblast-like cells was quite different. For membranes with pore a diameter of 0.2–1.0 mm, cells adhered and spreaded easily. When pore size increased, cells became spherical in shape, which was unfavorable for their functional activity [118]. The overall membrane porosity seems to be also a useful tool to manipulate osteoblast adhesion, proliferation, and function. Until now the effect of membranes in bone regeneration is not completely understood. They may depend on the location of bone defects, their size, the vascularization, and also the type of animal. One model, namely the rabbit radius model, has suggested that membranes have a number of functions namely to concentrate and store the osteogenic components (growth factors) released from the bone ends and bone marrow in the ‘‘medullary cavity’’ formed by the membrane (i), to protect the site against the massive ingrowth of soft ﬁbrous tissues into the defects (ii), and to serve as a guiding scaffold migrating osteogenic cells into the defect (iii) [132,133]. 8.2.4 Further Tissue Engineering Applications of Membranes

Besides the abundant research on membranes for tissue engineering of skin and bone, a broad range of other applications of membranes have also been envisaged. Table 8.3 gives an overview on membrane applications besides skin and bone tissue engineering. For example, membranes were also tested for regeneration of components of the nervous system [134,135]. A typical example is the gelatine composite membrane prepared from tricalcium phosphate and glutaraldehyde-crosslinked gelatine [136]. Several neurotrophic factors such as nerve growth factor (NGF) and brain-derived neurotrophic factor (BDNF) were covalently immobilized using carbodiimide. The effects of these membranes on the regeneration of the sciatic nerve in rats across a

8.2 Application of Membranes in Tissue Engineering Tab. 8.3

Further applications of membranes in tissue engineering.

Scientists

Membranes sources

Application

Chen et al. Qin et al.

Gelatine-tricalcium phosphate Polylactide/polytrimethylene carbonate N,O-Carboxymethyl chitosan Crosslinked hyaluronate/collagen Polyaniline/gelatine

Nerve guidance channel [134] Preventing postoperative adhesion [138]

Zhou et al. Shih et al. Li et al.

Hsiue et al. Gelatine Riboldi et al. Polyesterurethane Vernon et al. Collagen Zhu et al. Collagen-g-poly (DL-lactide-co-glycolide)

Preventing postoperative adhesion Preventing peridural adhesion Potential for cardiac or neuronal tissue constructs Neural retinal transplantation Potential for skeletal muscle tissue engineering Cell alignment or cell perfusion Esophagus tissue engineering

Reference

[142] [143] [147] [148] [149] [150,151] [152,153]

critical size defect of length 10 mm were investigated. The gastrocnemic muscles begin to degenerate when the sciatic nerve is severed. Because they regain their mass proportional to the amount of re-innervation [137], the weight ratio of reinnervated gastrocnemic muscle has been used as a means to evaluate peripheral nerve repair. The results showed that the average size of newly grown nerve axons was the greatest for membranes immobilized with NGF, while the BDNF-immobilized membrane had a much higher gastrocnemic muscle weight ratio than the blank membrane. Although the effects of different growth factors immobilized onto gelatin membranes seemed to be quite complex, which could not be simply illustrated by the sciatic function index, these types of membranes may provide a solution for peripheral nerve repair. Postoperative adhesion formation (e.g., the formation of adhesions between gut and the surrounding tissue) is a serious surgical problem, which may lead to complications for the patient such as infection and chronic pain [138–140]. Polymeric membrane covers maybe a useful method to prevent adhesions after visceral surgery [141–143]. Using a blend membrane prepared from polylactide and poly(trimethylene carbonate) (PLA/PTMC), intestine studies were conducted on the ascending colon in rabbits’ abdomen, which was covered by the blend membrane to investigate their effects on adhesion prevention. Histological observations showed that the PLA/ PTMC blend membrane could act as a physical barrier before their degradation. The distribution of adhesion scores in the intestinal models revealed that PLA/PTMC blend membranes were able to prevent postoperative adhesion formations [138]. A novel research direction is the combination of electrical conductive polymers such as polypyrrole or polyaniline with biodegradable materials [144–147]. Zhang et al. [145] prepared a blend membrane from polypyrrole and polylactide and tested its in vivo tissue response by subcutaneous implantation in rats. Both alkaline phosphatase activity and acid phosphatase activity were used to evaluate the inﬂammatory response to the implanted membranes. The alkaline phosphatase activity was similar to that of the pure PLA membranes, indicating that the addition of polypyrrole

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did not intensify the inﬂammatory response. Acid phosphatase activity for materials containing polypyrrole was similar or even lower than that of the membrane without polypyrrole. The histological ﬁndings were in agreement with the enzymatic results, indicating that the polypyrrole and PLA composite membranes provoked no abnormal tissue response [145]. Cultures of ﬁbroblasts on the composite membranes were conducted based on the hypothesis that the composites can be used to modulate the cellular by direct electrical stimulation [146]. Membranes experienced mediumrange electric current stimulations, which promoted cell growth in comparison to nonstimulated controls. Hence, these composite membranes may be a potential substratum for the arrangement of electrical responsive cells or can be used as a scaffold for electrically stimulated tissue regeneration. Another inherently conductive polymer, polyaniline was blended with gelatine. The blend was used for electrospinning of nanoﬁbers to prepare a microporous membranous scaffold [147]. Cultures of H9c2 rat cardiac myoblast cells showed that the scaffold supported attachment, migration, and proliferation of cells. This implies the possible application of this membrane-like material based on nanoﬁbers for tissue engineering of functional cardiac, cardiovascular, and neuronal tissue constructs. A further example of application of membranes in tissue engineering is the applications of a gelatine sandwich-like membrane as implant in neural retinal transplantation [148]. Using two gelatine membranes sandwiched by a retina cell sheet, the retina-gelatine membrane construct was implanted into rabbits. Histological results revealed that the newly developed implant survived in a proper orientation of the retinal layer, which indicates the potential of this approach for tissue engineering of retina. Membranes were also tested for tissue engineering of skeletal muscle [149]. Collagen membranes with different morphologies were tested for their ability to control muscle cell alignment [150,151]. A poly(L-lactideco-caprolactone) membrane was also used for porcine esophageal cells cultures; the mitochondrial activity assay, and cell morphology indicated that this kind of membrane may be potential scaffold materials for esophagus repair [152,153]. Overall, there is a multitude of established and future applications of membranes in the emerging ﬁeld of tissue engineering.

8.3 Membranes in Biohybrid Organ Technology 8.3.1 Organ Failure and Biohybrid Organ Technology

The term biohybrid (artiﬁcial) organ was ﬁrst introduced by Chick et al. in 1975 to denote a device containing living cells separated from the body by a synthetic material to substitute the function of an organ or speciﬁc tissues [154]. Today both terms bioartiﬁcial organ and biohybrid organ are used. Biohybrid organs can resemble an implantable device for the treatment of chronic diseases, such as secretion of insulin in diabetes [155], human growth factor in dwarﬁsm [156], and Parkinson’s disease

8.3 Membranes in Biohybrid Organ Technology

[157,158]. These applications usually require that cells are located in a compartment separate from the surrounding tissue to maintain the structural integrity of the immobilized cell mass. Moreover, these cells are typically allogenic or xenogenic in origin, which requires their isolation from the immune system of the recipient. Membranes have been identiﬁed as useful tools providing protection from host immune responses but also exchange of substances and oxygenation of immobilized cells [159]. The requirements to replace a speciﬁc tissue or organ cover a large range, which is determined by the speciﬁc activity (e.g., secretion of hormone) per cell and quantity of the agent required by the body [159]. Because of the relatively low quantities of hormones needed by the body, a low amount of cells of about 106–107 is necessary for replacement of most endocrine functions. This corresponds to a cell volume of about 10 mL. In contrast, implantation of Langerhans islets cells to treat diabetes requires already 109 cells, which matches to about 1 ml volume. Full replacement of liver function however needs between 1010 and 1011 cells, which is more than 100 mL cell suspension [159]. It is quite clear that replacement of endocrine functions can be still achieved by an implantable device while liver or kidney replacement therapies require extracorporal machines. Different conﬁgurations of implantable devices have been developed, such as microcapsules, hollow ﬁber, or planar diffusion chambers with ﬂat membranes [159]. A majority of developments aim to encapsulate the cells into hollow spheres, which can be implanted, for example, into the peritoneal cavity. Microcapsules are advantageous regarding the exchange of solutes due to their large surface to volume ratio. However, their application can be problematic if the device must be later retrieved. To overcome this limitation hollow ﬁber membranes have been also used to encapsulate cells with endocrine functions. The hollow ﬁbers were introduced into blood vessels or in the lumbal region of the backbone, from where they can be also removed. While the current review is focused on application of membranes in biohybrid organs for blood detoxiﬁcation, others have reviewed application of cells in implantable devices for the replacement of endocrine functions in more detail [159,160]. Table 8.4 gives an

Survey on implantable biohybrid organs with type of organ, type of membrane and membrane geometry.

Tab. 8.4

Type of organ

Type of membrane

Heart Liver Liver Pancreas Pancreas Pancreas

Polyurethane Alginate/poly-L-lysine/alginate Regenerated cellulose Alginate Alginate Alginate-cellulose sulfate complexed with poly(methylene-co-guanidine) Polyacrylonitrile-sodium methallylsulfonate copolymer Alginate-polylysine

Pancreas Parathyroid

Membrane geometry

Reference

Flat membrane Microcapsules Hollow ﬁber Hollow ﬁber Microcapsules Microcapsules

[59] [161] [162] [163] [164] [165]

Hollow ﬁber

[166]

Microcapsules

[167]

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8 Application of Membranes in Tissue Engineering and Biohybrid Organ Technology

overview on implantable biohybrid organs with type of organ, geometry of the device and type of membrane. The separation of toxins from blood is crucial for the treatment of end-stage renal disease (ESRD), acute renal failure, and also fulminate liver failure. The current majority of extracorporal support and replacement therapies of these organs is based on physical processes that comprise dialysis, ﬁltration, and adsorption [2,4–7]. However, kidney and liver represent organs with multiple functions. Kidney has basically a ﬁltration function, but also a re-adsorptive and endocrine function that cannot be replaced by artiﬁcial organ technology [168]. Likewise, the numerous tasks of the liver, such as conversion of toxins generated by the metabolism, removal of xenobiotics, protein synthesis, and so on, cannot be substituted by hemoﬁltration or adsorption technologies alone [169,170]. Hence, it is not surprising that the conventional organ replacement therapies are associated with an increased morbidity and mortality, which is known from long-term treatment with hemodialysis in ESRD and particularly in acute renal or in fulminant hepatic failure [171–173]. To date, organ transplantation seems to be the most successful therapy for acute and chronic failure of internal organs like heart, kidney, liver, lung, and so on. However, there is an increasing gap between the number of donor organs available and the patients on the waiting list. Because of the shortage in donor organs, the advent of biohybrid organs may be a solution of the described problem because they may support failing organs, which lead to organ regeneration in acute liver and renal failure or may bridge the time to ﬁnd a suitable transplant for the patient. Kidney and liver mainly consist of connective tissue, endothelial, and epithelial cells. Epithelial cells that represent the functional units, such as proximal tubule cells in the kidney or hepatocytes in the liver, have a polar organization where the cytoplasm is separated into an apical and a basolateral region. Main morphological characteristics of epithelial cells are the presence of tight junctions and microvilli on the apical cell membrane, which are both related to transepithelial transport processes [174]. Tight junctions separate the basal and apical region controlling the exchange of solutes along or against concentration gradients. Microvilli on the apical cell surface increase the surface area for the uptake of substances and are an important feature of many epithelia as well. Typical epithelia have a planar structure with an underlying complex of basal membranes composed of different extracellular matrix proteins, such as collagen IV, ﬁbronectin, laminin, and so on [175]. Figure 8.8 shows the typical morphology of kidney epithelial cells with presence of a tight junction between cells and microvilli on the apical region obtained by transmission electron microscopy (TEM). Biohybrid liver and kidney resemble a combination of immobilized epithelial cells on a suitable type of carrier, which is in many cases a polymer membrane in a bioreactor (see, e.g., [176]). Organ cells in biohybrid organs have to make an intimate contact with the surface of the membrane but also to develop close cell–cellconnections, which is a prerequisite for their survival and high functional activity (see also Section 1.2). On the other hand the blood, which has to be detoxiﬁed will contact the other side of the membrane and may not become activated by the synthetic material. Finally as a third important requirement, the transport properties

8.3 Membranes in Biohybrid Organ Technology

Fig. 8.8 Transmission electron micrograph of kidney epithelial cells with typical ultrastructural features such as microvilli on the apical cell surface (arrows) and tight junctions (asterisk) between cells.

of the membrane must be adjusted to allow a sufﬁcient exchange of oxygen, small molecular weight solutes, such as toxins, smaller proteins, but must also protect the immobilized organ cells from the immune system of the recipient [159,177,178]. Figure 8.9 shows a simpliﬁed set-up of a biohybrid organ to emphasize the complex requirements to membranes in this application. 8.3.2 Biohybrid Liver

Acute (or fulminant) hepatic failure after intoxication or viral infections is characterized by defective blood protein and clotting factor synthesis, gluconeogenesis, urogenesis with impairment of plasma detoxiﬁcation, neurological complications (often associated with cerebral edema), and ﬁnally multiorgan failure [179]. Current therapies based on a multitude of measures including hemodialysis or

Fig. 8.9 Schematic set-up of biohybrid organs for blood detoxiﬁcation highlighting the crucial role of membranes as support for epithelial cells, as contacting surface for blood components and barrier controlling the exchange of solutes between the different compartments.

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8 Application of Membranes in Tissue Engineering and Biohybrid Organ Technology

hemoﬁltration have a very poor outcome with survival rates as low as 20 %. It has been stated that orthopic liver transplantation is the best medical therapy with a survival rate higher than 90 % after one year [179]. In light of the fact that access to donor organs is limited for a speciﬁc patient due to not only availability of organs but also immunological mismatch between donor and recipient, development of biohybrid liver support systems bears great hope to bridge the time to ﬁnd a suitable transplant or to support even regeneration of the damaged liver [179]. Biohybrid liver support systems replace the function of the failing liver through the actions of hepatocytes cultured in a bioreactor module. Different concepts of biohybrid liver devices have been realized, which have been reviewed in more detail recently [176,181]. Among them membrane-based systems may provide certain advantages due to their ability to separate donor hepatocytes from the immune system of the recipient (i), to provide a substratum for cell attachment (ii), and the large surface-to-volume ratio allowing a compact bioreactor design (iii). Critical issues, however, are the ability of membranes to provide adequate attachment for hepatocytes and their ability for bidirectional mass transfer. The latter requirement must be fulﬁlled to provide nutrients to sustain cell viability and export of certain cellular products. Moreover, it is known that many toxins in blood are albumin bound and contact hepatocytes directly after passage through the fenestrated endothelium in the liver [176]. At present the majority of membranes used in biohybrid liver support systems have been those developed for other biomedical applications such as hemodialysers or oxygenators [170,182]. These membranes were selected because of their relative good biocompatibility in contact with whole blood or plasma. The molecular cut-off of these membranes is chosen to prevent exposure of cells in the bioreactor to components of the immune system, such as immunoglobulins (about 160 kDa) or factors of the complement system (>200 kDa), but alternatively permeation of serum albumin (about 70 kDa) shall be still possible. Therefore, some groups have selected membranes with a molecular cut-off of around 100 kDa [176]. Conversely, also membranes with a lower cut-off were used to prevent the entrance of cellular proteins from donor cells into the blood stream of the patient, which is important in cases when porcine or cancer hepatocytes are applied in the bioreactor [180,183,184]. Few groups have selected membranes with microﬁltration properties showing that immunoisolation of porcine cells is not an absolute requirement [185]. Table 8.5 gives an overview on biohybrid liver support systems, which underwent clinical testing and application. It must be noted that the majority of membranes used in biohybrid liver systems have not been optimized regarding the promotion of attachment and function of hepatocytes. The majority of current biohybrid liver devices are only used for treatment sessions of 6–8 hours, each patient receiving multiple sessions. The main reason for this limitation is that the primary hepatocytes have unstable function, and lose viability and essential enzyme activities with time in culture [186]. Particularly, quite hydrophilic membranes based on cellulose and its derivatives have shown to be less suitable for primary hepatocytes [182]. Membranes with limited compatibility for hepatocytes can be coated with collagen or ﬁbronectin, which improved the functional activity and survival of hepatocytes [187]. The development of membranes with

8.3 Membranes in Biohybrid Organ Technology Tab. 8.5 Overview on biohybrid liver systems using hollow fiber membranes undergoing clinical

testing. Name of devices/ company

Type of cells

Type of membrane

Cut-off/ pore size

Reference

ELAD Amphioxus Cell Technology HepatAssist Circe Biomedical MELS/Charite´ Humboldt Univ

C3A hepatoblastoima cells Porcine hepatocytes

Cellulose acetate

70 kDa

[192]

Polysulfone

0.2 mm

[193]

Porcine or human hepatocytes

100 kDa oxygenation 80 kDa

[188]

BLSS/Excorp Biomedical Inc.

Porcine hepatocytes

Combination of polyamid, polypropylene, polysulfon Polysulfone

100 kDa

[191]

polymer compositions, which speciﬁcally support the attachment, function, and viability of cultured hepatocytes, would signiﬁcantly improve the length of treatment time possible for patients with acute liver failure. Catapano and co-workers have shown that cell adhesion and urea synthesis increased with increasing wettability and surface roughness of polypropylene membranes [188,189]. The chemical composition of the membranes also markedly inﬂuences the properties of the cultured cells, and in general, adhesion and liver function is better maintained on moderate hydrophilic than on hydrophobic membranes [28,190]. A few groups have also tested other polymer membranes than those developed for hemodialysis or blood oxygenation successfully, such as polyetherketone coated with proteins as a substratum for hepatocytes [191]. We have developed a range of copolymers based on acrylonitrile, which have been already introduced in Section 2.2 and investigated the adhesion, growth, and functional activity of C3A hepatoblastoma cells as well. This cell line has been used also in clinical studies with the extracorporeal liver assist device (ELAD) mentioned above [184]. We found that particularly hydrophilic copolymers, which contained 20 mol% N-vinylpyrollidone (NVP 20), allowed moderate attachment of these cells and supported cell–cell adhesions, which seems to stimulate the functional activity of cells [192]. During these studies we also explored the impact of pore size in the ultraﬁltration range on C3A cell attachment, growth, and function. We used membranes based on NVP 20 and found an inverse relationship between initial cell attachment, which was higher on membranes with larger pores, while cell growth was enhanced on membranes with smaller pores [193]. Figure 8.10 shows examples of this work as a comparison of membrane surface morphology between 6 and 12 nm (a and d) pore size visualized by scanning electron microscopy (SEM). It is obvious that membranes with 12 nm pores possess a higher surface roughness. Studies on cell adhesion with staining of viable cells show that the size of aggregates was bigger on membranes with 12 nm (e) than on 6 nm pores (b). Moreover, it was also observed that intercellular connections visualized by immunostaining of E-Cadherin and confocal laser scanning microscopy were more expressed on membranes with 12 nm (f) than 6 nm (c) pores, which indicates a

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8 Application of Membranes in Tissue Engineering and Biohybrid Organ Technology

Fig. 8.10 Comparison of C3A hepatoblastoma cell growth and morphology on membranes with different pore size made of NVP 20. Upper panel shows membranes with a mean pore diameter of 6 nm (a), vital staining of cells after 4 days of culture (b) and expression of intercellular junctions (E-Cadherin) visualized by immunoﬂuorescence microscopy (c).

Lower panel shows membranes with mean pore diameter of 12 nm (d) and vital staining of cells (e) and expression of intercellular junctions. Note the larger size of aggregates and the more abundant expression of E-Cadherin (arrows) on 12 nm membranes. Scale bars (a and d) 0.2 mm, (b and e) 100 mm, and (c and f) 20 mm.

better function of these cells. However, the functional activity measured by the activity of liver-speciﬁc P450 enzymes was maximal at intermediate pore size [193] as shown in Figure 8.11. The blood compatibility of these copolymers is also of interest for the potential application in biohybrid liver systems. It was found in parallel investigations that introduction of N-vinylpyrollidone improves the hemocompatibility of membranes considerably, when compared to polyacrylonitrile as standard membrane for hemodialysiers [194]. In further studies with primary hepatocytes, we also explored the effect of other co-monomers on survival and functional activity of cells. Here, a polyacrylonitrile membrane containing acrylamido-2-methyl-propansulfonic acid as co-monomer was found to be the most compatible option for maintaining primary hepatocyte morphology and function. It maintained viable, functional cells for at least 16 days in culture [195]. Overall, tailoring membrane properties regarding chemical composition and porosity has a great potential to improve further the functional activity and to prolong the survival of hepatocytes in the biohybrid liver systems.

8.3 Membranes in Biohybrid Organ Technology

Fig. 8.11 Comparison of growth of C3A hepatoblastoma cells in dependence on mean pore diameter of NVP 20 membranes. Please, note that initial number of cells after two days is higher on membranes with larger than on smaller pores. However, cell numbers are higher after 7 days on membranes with smaller pores indicating better growth on smoother surfaces.

8.3.3 Biohybrid Kidney

Despite the progress in medical technology, patients with acute kidney failure still have a poor survival prognosis [196]. Also, patients having end-stage chronic failure treated with hemodialysis suffer from an increased morbidity and have also a limited life expectancy [197,198]. Reasons for the increased morbidity and mortality of these patients are numerous uremic toxins that accumulate in the body [199]. Kidney transplantation is currently the best therapy, but the limited number of donor organs is the cause for long waiting lists of patients. Since the beginning of the twentieth century researchers worked on a device to replace the function of the kidney. In fact, the kidney was the ﬁrst organ whose function was substituted by an artiﬁcial organ – the hemodialysis [200]. However, hemodialysis therapy is not a long-term solution since about 50 % of end-stage renal failure patients develop dialysis-related amyloidosis [201] and other serious health problems. The main reason for this situation is that hemodialysis can only partly replace the ﬁltration function of the kidney glomerulus but does not substitute many other renal functions, such as transport, metabolic, endocrine, and immune activities [201]. A biohybrid kidney must overcome these drawbacks combining the previously established functions of artiﬁcial kidney, such as ﬁltration and adsorption, with the activity of cells derived from the kidney tubule system [202]. In 1987 Aebischer et al. suggested for the ﬁrst time to combine hollow ﬁber membranes with kidney tubular cells to allow for the transport of substances across the cell-attached membranes [203]. This concept was named bioartiﬁcial kidney [204]. The principal set-up of a biohybrid kidney has been realized by Humes et al. in 1999 [205], which is shown in Figure 8.12. It basically consists of a commercial

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Fig. 8.12 Principal set-up of a biohybrid kidney as implemented by Humes and coworkers that combines a synthetic hemoﬁlter, which generates an ultraﬁltrate in series with a tissue-engineered proximal tubule device for metabolic, endocrine and immunological functions. This ﬁgure is a reproduction from [214].

hemoﬁltration device, which generates an ultraﬁltrate like the kidney glomerulus. This primary urine in analogy to the human kidney is transported to a proximal tubule device (RAD), which consists of a hollow ﬁber bioreactor containing kidney epithelial cells. These cells provide the same active transport processes as in the renal tubule along with a multitude of cellular metabolic reactions. They produce some reabsorbate, which ﬂows across the membrane and mixes with the blood of the patient. It has been anticipated by the group of Humes through preclinical and clinical studies that the biohybrid kidney improves survival in acute renal and also multiorgan failure due to sepsis [206,207]. Until now polysulfone (PSU) hollow ﬁber membrane modules have been used as bioreactor for the biohybrid kidney device [208,209] although it has never been explored in detail which type of commercial membrane is the most biocompatible for kidney epithelial cells. Saito et al. compared the biocompatibility of cellulose acetate and polysulfone for a porcine kidney epithelial cell line (LLC-PK) in a recent paper and observed that polysulfone had superior properties [210]. In the same study the

8.3 Membranes in Biohybrid Organ Technology

authors stated that human kidney epithelial cell did not attach and grow that well on polysulfone as they did on a polyimide membrane. Obviously, because of reduced ability of polysulfone to support attachment, growth, and function of kidney epithelial cells, the group of Humes used ProNectin-L to coat the membrane surface. ProNectin-L is a synthetic protein sharing the cellular attachment domain of laminin, which is a component of the kidney tubular basement membrane [201]. Also, other authors have used extracellular matrix proteins to improve the biocompatibility of commercial membrane materials for kidney epithelial cells [210,211]. However, coating with extracellular matrix proteins is an additional step during the preparation of the bioreactor and increases the risk of infections. Therefore, we have looked for the ability of different polymers to support adhesion, growth, and function of kidney epithelial cells [212,213]. Among them we used polyacrylonitrile and its copolymers introduced in Sections 2.2 and 3.2 and compared them with PSU and a commercial Millipore ﬁlter, which represents a blend of cellulose acetate and cellulose nitrate. Investigations with ﬂat membranes were performed with MDCK cells, which represent a canine kidney epithelial cell line. The cell proliferation assays revealed a signiﬁcantly reduced cell attachment and proliferation on AEMA, NaMAS, and PSU in comparison to the other membranes. The morphology of cells on the various ﬂat or hollow ﬁber membranes was studied with SEM, while the ultra structure of cells and their contact with membranes were studied with TEM. It was observed that conﬂuent layers of cells existed on most of the membranes except on PSU and NaMAS where the cells were not uniformly distributed. TEM of MDCK cells cultured for 8 days showed typical morphological features of functional active epithelial cells, such as tight junctions and microvilli on most of the membranes except on PSU and NaMAS, where the cells formed much less microvilli, which indicates a reduced functional activity of these cells. Figure 8.13 shows a comparison

Fig. 8.13 Scanning electron micrographs of kidney epithelial cells grown on the inner surface of hollow ﬁbers made of polysulfone (a) and NVP 5 (b). Note that on polysulfone only single cells are visible, while on NVP 5 after 7 days a conﬂuent layer of cells has been established.

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Fig. 8.14 Transmission electron micrographs of kidney epithelial cells cultured between PSU and NVP 5 membranes. It is visible that cells make very close contact to the NVP 5 membrane and develop numerous microvilli as feature of functional active epithelial cells. For details look at [213].

of cell growth of kidney epithelial cells in contact to PSU and NVP 5 membranes. It is visible that MDCK cells grow better and make up a conﬂuent layer on NVP 5 (b) in contrast to PSU (a), where poor attachment and growth was observed on the inner surface of a hollow ﬁber membrane. On NVP 5 cells also make very close contact with the membrane surface and develop typical features of epithelial cells, such as tight junctions and microvilli as shown in Figure 8.14. The functionality of kidney epithelial cells was estimated by transepithelial resistance (TR) measurements and impedance analysis that are validated methods to estimate the properties of epithelial sheets, such as tight junction formation and active transport. Values for functional epithelia are normally around 2000 V cm2, which is reached on standard Millicells (MC) membranes [212]. Figure 8.15 shows the results of TR measurements. It is visible that not only materials with a lower content in NVP but also AEMA 0.1 mol % provided conditions where the TR values were comparable to that of normal epithelia. In contrast, membranes with acidic sulfonate groups such as AMPS and also PSU to some extent do not support the functional activity of kidney

Fig. 8.15 Transepithelial resistance of MDCK cells cultured 7 days on different supports (MC ¼ Millicell – reference material). It is shown that cells on NVP 5 and NVP 20 have resistance values comparable to the reference material Millipore, which indicates a high functional activity of cells on these membranes.

8.4 Summary and Conclusions 253

epithelial cells. The same investigations were also carried out with LLC-PK cells and yielded comparable results regarding growth, morphology, and functional behavior on the different membrane materials [214]. In summary of the investigations, it can be stated that copolymers with low quantities of basic functionalities such as NVP 5 and AEMA 0.1 were the most suitable supports, while copolymers with acidic groups such as AMPS, polysulfone, and NaMAS did not support kidney epithelial cells. This is in line with the ﬁnding of others that coating with extracellular matrix proteins is necessary to improve the cytocompatibility of polysulfone for epithelial cells [206–211].

8.4 Summary and Conclusions

Membranes have found a multitude of applications in medicine. This holds not only for conventional artiﬁcial organ therapy for blood detoxiﬁcation and oxygenation such as hemodialysis, hemoﬁltration, and cardiopulmonary bypass. Membranes have been also implemented in the ﬁeld tissue engineering basically as guiding structures providing support for cell attachment and also helping to accumulate growth factors and cytokines in the wounded area to promote healing of defects. For helping cell attachment, growth, and functionality, the chemical composition of the membrane surface as well as their surface topography play a crucial role. Membranes have been also used in tissue engineering applications simply as a scaffold to place materials for augmentation of tissues such as bone. While the transport properties of membranes in tissue engineering applications have not been paid speciﬁc attention so far, this is the case in biohybrid organ technology. Here membranes have several functions particularly when in intimate contact with blood for detoxiﬁcation purposes. Hence, membrane materials must be blood compatible on one hand side and cell compatible on the other. This cannot be achieved so easily because blood contact requires a noninteracting surface while cell contact needs a surface that is able to support adsorption of proteins and adhesion of cells. A tailored design of copolymer composition [192,195,213,214], the surface modiﬁcation of one side of the membrane with speciﬁc ligands [194] or the application of speciﬁc membrane formation techniques resulting in bilayer membranes from different polymers [215] can be a solution to the problem. The permeability or pore size of the membrane may also have several effects such as protection of xenogenic or allogenic cells from the immune system of the recipient, control of exchange of solutes important for the metabolic activity of cells, and also creating a speciﬁc topography that has an inﬂuence on the behavior of immobilized organ cells. So far, a large number of attempts have been put forward to produce implantable devices basically for the replacement of endocrine functions. On the other hand, support of blood detoxiﬁcation has been achieved by extracorporal biohybrid organs to replace liver and kidney function particularly in acute failure of these organs. Overall, it may be concluded that membranes represent an essential component in many tissue engineering applications and biohybrid organ technology. However, it may be stated for certain

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8 Application of Membranes in Tissue Engineering and Biohybrid Organ Technology

applications that the full potential of membranes has not yet been exploited due to the restriction to standard polymers and other reasons.

8.5 Acknowledgments

The authors gratefully acknowledges the cooperation with Wolfgang Albrecht, Gregor Boese, Nathalie Faucheux, Gu¨nter Malsch, Dieter Paul, Klaus Richau, Michael Schossig, and Barbara Seifert from GKSS Research Centre, George Altankov and Natalia Krasteva from Bulgarian Academy of Sciences, Ulrich Gross and Frederique Fey-Lamprecht from Free University Berlin, Volkmar Thom from Sartorius AG, Mathias Ulbricht from University Essen-Duisburg, Hans-Ju¨rgen Stark and Norbert Fusenig from German Cancer Research Centre Heidelberg. I am very thankful to Mrs. Ruth Hesse for her excellent technical assistance in a number of experimental studies mentioned in this review. Also, the ﬁnancial support of some of our own studies reviewed herein by BMBF (grant FKZ 01GN0119) and the Commission of European Union (BE 97-4326) is gratefully acknowledged.

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9 Membranes in Bioartiﬁcial Pancreas – An Overview of the Development of a Bioartiﬁcial Pancreas, as a Treatment of Insulin-Dependent Diabetes Mellitus1 Ana Isabel Silva, Anto´nio Norton de Matos, I. Gabrielle M. Brons, Marı´lia Clemente Velez Mateus 9.1 Introduction 9.1.1 Diabetes and Its Treatment

Diabetes mellitus is an autoimmune chronic disease characterized by high blood glucose levels and long-term complications including micro- and macroangiopathic lesions leading to retinopathy, neuropathy, and nephropathy. The insulin secretion is either impaired or entirely destroyed in the pancreas of a diabetic patient, as a result of which the blood glucose concentration climbs to levels far exceeding its normal range. The cells that produce insulin, b cells, are located in structured cell clusters named islets of Langerhans, which make up for 1–2 % of the total mass of the pancreas [1]. There are two types of primary diabetes (type I and type II), differing in etiology, pathophysiology, and treatment. Type I occurs when the pancreas ceases to produce insulin due to a massive immunological attack on the islets by the patient’s own immune system. This causes the destruction of all b cells in the islets in the pancreas; thus, type I diabetes patients depend upon exogenous insulin therapy for life. Type II diabetes is a different, milder type of diabetes, which occurs when the pancreas does not produce enough insulin or when the insulin produced cannot be used by the body due to resistance to the action of insulin. Only a small portion of type II patients need insulin to control hyperglycemia; usually the ingestion of oral hypoglycemic drugs and an appropriate diet are enough to stabilize blood glucose levels [1]. The earliest known record of diabetes is mentioned in a third Dynasty Egyptian papyrus (1500 B.C.) [2]. However, it was not until the early nineteenth century that 1

This is a preprint of an article published in ‘‘An Overview on the Development of a Bio-artiﬁcial Pancreas as a Treament of Insulin Dependent Diabetes Mellitus’’, Silva A. I. et al. Medicinal

Research Reviews, Vol. 26, No. 2, 181–222, 2006 Copyright# 2005 Wiley Periodicals, Inc. (http:// www3.interscience.wiley.com./cgi-bin/jtoc/ 34185/)

Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2008 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

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Tab. 9.1

Main events in the history of diabetes.

Year

Event

1500 B.C. 1869 1889

First known record of diabetesa Paul Langerhans discovers the cellular structure of the pancreasb Josef von Mering and Oskar Minkowski found diabetes could be induced in dogs by removing their pancreasesc Frederick Banting and Charles Best extract insulin from the pancreases of dogsd Two major types of diabetes, insulin dependent and insulin independent are recognized First pancreas transplant performede First islet allograft transplant for diabetes is performede First living donor pancreas transplant in the worlde Human monocomponent insulin is launched, the world’s ﬁrst insulin preparation identical to human insulin. It is extracted from the pancreas of pigs and converted to human insulinf First islet human allograft transplantation using a corticosteroid-free immunosuppressive regimeng

1921 1959 1966 1974 1979 1982

2000 a

Ebers Papyrus. Berlin Pathological Institute, Germany. c University of Strassburg, Germany. d University of Toronto, Canada. e Fairview-University Medical Center, University of Minnesota, USA. f Novo Nordisk A/S. g University of Alberta, Edmonton, Alberta, Canada. b

Source: Adapted from Refs. [2] and [3].

the cause of diabetes mellitus became clearer. Table 9.1 presents some important events in the history of the treatment of diabetes. Currently the more common treatment for insulin-dependent diabetes mellitus (IDDM) is lifelong insulin therapy, which requires daily injections and frequent monitoring of blood sugar concentration. An alternative to insulin therapy is pancreas or islet transplantation. The ﬁrst pancreas transplantation reported was limited by poor graft survival and high morbidity. In 1978, new surgical techniques together with new immunosuppressive regimens became available, greatly improving the outcome of pancreas transplantation [3]. Transplantation of isolated pancreatic islets into the liver by intraportal infusion offers normalization of glucose metabolism similar to that after whole pancreas transplantation, requiring less risky invasive surgery [3]. In the year 2001, an islet transplantation group based in Edmonton, Canada, published a revolutionary paper on human islet allotransplantation describing the use of a different immunosuppressive regimen together with improved islet isolation methods. The follow-up of the transplanted patients [4,5] showed such good islet function over time that this method has become the preferred islet transplantation method for IDDM patients. Still, the patients are subjected to lifelong immunosuppressive drugs and the therapy is conditioned by the scarcity of human donors.

9.1 Bioartificial Pancreas

9.1.2 The Bioartiﬁcial Organ Concept

A large number of people all over the world suffer from hormone deﬁciencies. Currently, many of these hormone deﬁciency diseases can be controlled by the patients with regular intakes of the lacking hormone. However, oral replacement therapy may only delay the onset of complications of the disease, since large variations still occur in the levels of metabolites. Currently, the only deﬁnite cure for these diseases is partial or whole organ transplantation, although the risks involved in transplantation are high. The surgical risk can be reduced drastically, if instead of a whole organ, a speciﬁc cell transplant can be performed. Transplantation of allogeneic organs or cellular tissue from one individual to another or one species to another as in xenotransplantation will always stimulate strong responses from the recipient’s immune system, which means that the recipients will have to take immunomodulating drugs for the rest of their lives in order to avoid organ or cell rejection. In the long run, these drugs have signiﬁcant toxicity that may give rise to other complications, for example, an increased risk of infection and cancer. Over the last 30 years, researchers have been trying to develop devices to prevent the need for immunosuppressing drug regimes. These so-called bioartiﬁcial organs generally contain cells or cell clusters within a synthetic biocompatible semipermeable membrane, which separates the foreign tissue from the host’s immune system [6]. They can be inserted into the blood supply or implanted somewhere in the body with diffusion potential and are meant to fully mimic the behavior and function of a healthy organ. These devices are eliminating the need for immunosuppressing drugs by protecting the cells from the onslaught of immune cells or antibodies, which are not able to cross the synthetic biocompatible semipermeable membrane due to their higher molecular weight. Ideally, the protected cells or cells clusters are kept viable and functional by having access to nutrients and oxygen and stimulatory agents, which can easily diffuse across the membrane and produce the lacking hormones or factors on a minute-to-minute basis or when needed and/or for long periods of time. The bioartiﬁcial organ implant procedure is generally simpler and less risky than whole or partial organ transplantation. Biocompatible synthetic membranes developed for these devices have also other applications, like dialysis, in the development of a bioartiﬁcial kidney device [7,8]. Many parameters can be varied in the bioartiﬁcial organ conception, namely, the biocompatible polymer used and its intrinsic characteristics, the cell immobilization or suspension media, the existence or not of coimmobilized molecules or cells, the number of devices used and the implantation site, just to mention a few. From the number of adjustable parameters, the complexity of developing a functional prototype of such a device is easily perceived. This report attempts to present some of the more recent developments in the conception of a bioartiﬁcial pancreas (BAP) device meant to mimic the insulin secretion of the natural organ.

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9.2 Bioartiﬁcial Pancreas

Several different models of bioartiﬁcial organs have already been studied like the liver, pancreas, and kidney [7–12]. The membrane characteristics and geometry vary, depending on the speciﬁc function of the organ the bioartiﬁcial device should mimic. In any model there are some requirements the device must obey to be accepted as an alternative therapy: biocompatibility and good diffusional properties of the membrane, easy retrievability or biodegradability of the device, and maintenance of cell viability and functionality for long periods of time (when cells are immobilized in the device). A BAP containing viable and functional islets of Langerhans protected by a biocompatible semipermeable membrane is, in theory, a much better solution than currently available IDDM therapies because it would provide a tighter blood glucose control without the need for immunosuppressive drugs. It also makes possible the use of islets from nonhuman donors thereby reducing the problem of scarcity of human pancreata for isolation of islets for transplantation. The physiological feedback inherent in pancreatic islets allows for normoglycemia to be maintained at all times, even during high glucose stimulatory events; therefore, the transport of stimulatory substances to b cells and in response the release of insulin from them is fast and efﬁcient. Consequently, a successful BAP model must provide fast stimulation and response time to secretagogues, which translate into a fast and efﬁcient diffusion potential and must not allow insulin to stay in the vicinity of the cells, thus damaging the physiological feedback system. The material used in a successful BAP model must also be fully biocompatible and should not trigger the immune response system by activating inﬂammatory agents or stimulating ﬁbrosis around the device. Researchers all over the world have studied different membranes and membrane geometries, as well as different immobilization and coimmobilization methods. Devices are usually classiﬁed in two groups: intravascular and extravascular devices. Intravascular devices are implanted as an artery-to-vein (AV) shunt in the recipient’s body and are usually made of tubular hollow ﬁber membranes [13,14]. Extravascular devices can be further classiﬁed into two different categories, macroand microcapsular devices. Macrocapsular devices can have three main geometries: ﬂat sheet [15,16], sealed hollow ﬁbers [17,18], and macrospheres [19]. These devices have not only been mainly implanted in the peritoneal cavity, but have also been implanted in other regions, for example, the subcutaneous tissue site. Microcapsular devices consist of islets suspended in a polymeric gel material (e.g., alginate), gelled and encapsulated in a biocompatible synthetic membrane, to produce microspheres [20]. They are usually implanted in the peritoneal cavity, the kidney capsule, or subcutaneous site. The working principle of all of them is the same: to allow allogeneic or xenogeneic islet transplantation to succeed, by separating the islets from the recipient’s own body cells with an artiﬁcial biocompatible semipermeable membrane.

9.2 Bioartificial Pancreas

9.2.1 Immunoprotection and Biocompatibility of Implanted Devices

Immunoprotection in bioartiﬁcial devices concern the separation of transplanted tissue from the host by a membrane or ﬁlm forming a barrier against the passage of immunologically active solutes and cells, thus preventing immune rejection of the transplant. The pore size of the semipermeable membranes must ideally prevent the passage of cells as well as the diffusion of proteins involved in the humoral component of immune rejection, while allowing nutrients and metabolites to diffuse [21]. Since the cut-off point of synthetic semipermeable membranes is an average value, there are often a small number of larger pores that allow the transport of larger molecules than suggested by the nominal cut-off deﬁnition. Therefore, complex issues of rate of transport, threshold concentration of critical proteins, adsorption and denaturation in contact with polymeric materials, and interactions between proteins involved in immune reactions in an environment where their relative concentrations may be far from normal may all have some impact in immunoprotection [21]. The most important aspects of any graft success are the nature of the membrane and its biocompatibility with the body system. Biocompatibility involves a wide range of requirements. A biocompatible material is deﬁned as any substance (other than a drug), synthetic or natural, that can be used as a system or part of a system that treats, augments, or replaces any tissue, organ, or function of the body without causing a reaction against it [22]. With intravascular implants, clot formation is the most serious issue because it could threaten the patient if thrombi are shed into the bloodstream. With extravascular implants, the most serious threat is the development of ﬁbrotic tissue encapsulation over the membrane or encapsulant surface. One of the major reasons for BAP graft failure is the development of ﬁbrotic and inﬂammatory responses to the chemical surface composition of the device. In addition to the effects of chemical surface composition, the implant shape, morphology, and detailed microscopic topography at the material–tissue interface can inﬂuence a foreign body response. For example, interaction of macrophages with the surface is a crucial aspect of biocompatibility because these cells can induce proliferation and activation of other cell types [1]. The foreign body reaction to implanted materials, resulting in ﬁbrotic overgrowth, is a complex process of sequential events: an acute inﬂammatory response leading to chronic inﬂammatory responses and development of granulation tissue that ultimately leads to the formation of a ﬁbrotic tissue capsule composed largely of collagen, macrophages, ﬁbroblasts, and few, if any, capillaries. The presence of thick ﬁbrotic tissue surrounding the membrane capsule prevents proper nutrient and oxygen diffusion, quickly leading to massive loss of islet viability and functionality. The intensity of this ﬁbrotic response can vary greatly from individual to individual, even within the same species. Some authors [23] concluded that the extrapolation to humans of ﬁndings in rodents is very difﬁcult and, in the particular case of islet transplantation, even high animal models like pigs can differ signiﬁcantly from that

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of human recipients. Although pigs are metabolically similar to humans, their ﬁbrotic response is potentially different. Aside from the foreign body response induced by an inadequate capsule material, a far more dangerous reaction can occur when allo- or xenogeneic tissue or molecules are exposed or released from the capsular environment into the adjacent tissue and initiate recognition of the foreignness of the tissue causing immune rejection. The blockage of immune molecules is a complicated task. Although it is relatively easy to prevent the passage of cytotoxic cells, macrophages, and other larger cellular immune molecules including high molecular weight antibodies through the semipermeable membrane, a potentially more serious problem is blockage of humoral immune components such as low molecular weight cytokines as well as tissue antigens secreted by both the cells inside the membrane and the cells outside the chamber. Antibody binding to a cellular transplant, by itself, usually does not cause a cytotoxic reaction but it is the binding of the complement components that initiates the cytotoxic events [24]. C1q is the key molecule that must be retained by the immunoisolation membrane. This is the largest of the complement components and the ﬁrst that binds to two or more exposed antibody Fc domains. In the absence of immune cells, binding of antibody alone to a cell surface antigen without complement activation cannot lethally damage the cell and is of no consequence unless the epitope is from a site on the cell (e.g., receptor) important for cell function [1]. In order for this cytotoxic process to occur, both C1q (and the other smaller components of the complement system) and the antibody must cross the membrane. This can be minimized or prevented, theoretically, if membrane pore size is controlled. 9.2.2 Vascular Devices

In intravascular devices, islets are encapsulated within one or several hollow biocompatible membrane tubes or ﬁbers and then implanted into the blood system of the recipient’s body (Figure 9.1) [14,25–30]. One of the most relevant in vivo studies was performed with a device that consisted of two concentric tubular membranes, with one end anastomosed to an artery and the other to a vein. The blood ﬂowed through the lumen of the inner membrane, and the islets were immobilized in the annular space between the two membranes [14]. The tubular design was one of the ﬁrst designs ever experimented with and published about. Earlier devices were transplanted in vivo into dogs [25,26], or ex vivo in rats [27–29] and monkeys [27,30]. However, although ex vivo devices maintained recipient normoglycemia for several months, the in vivo grafts never lasted more than a few days (2 days in dogs). This was due to several different problems, some identiﬁed at the time, like the presence of blood coagulation in the ﬁbers (thrombosis), others remain the subject of investigations. The use of intravascular devices presents three great advantages when compared to extravascular ones: First, theislets embedded inside the chamber are near to the ﬂowing blood in the hollow ﬁber and have the most direct access to nutrients and oxygen; secondly, the constantly ﬂowing blood stream allows immediate recognition of

9.2 Bioartificial Pancreas

Fig. 9.1 Schematic representation of the device used by Ref. [14]. Source: Adapted from Monaco et al. [14], with permission from Lippincott, Williams & Wilkins, copyright 1991.

molecular changes in stimulatory agents like glucose and consequently immediate release of insulin; thirdly, a higher diffusion potential is facilitated by the blood ﬂow allowing the glucose and insulin to be carried away faster. Thus, the islets are properly fed, maintain their viability and functionality, and the delay in the response to a glucose peak is reduced. However, all these advantages depend on the size of the hollow ﬁber and the size of the surrounding chamber containing the islets. Diffusion problems may be encountered related to thewidthof thechamber: toonarrow wouldmean insufﬁcient tissue is grafted, and too wide would mean islets suffer from hypoxia and insufﬁcient nutrition, followed by central necrosis and loss of functionality and viability. Nonetheless, this device design is probably the one that mimics more closely the physiological environment of the islets. Moreover, cell re-seeding or complete device removal is easier to perform in case of functional failure. Other major problems are associated with this kind of device [1]: some risk is involved in the surgical implantation procedure, but a high risk is presented by the development of thrombosis, especially near the anastomosed ends. Also, the blood ﬂow is slowed down by the curvature of the device, resulting in turbulence and damage to thrombocytes followed by clot formation. 9.2.2.1 Biocompatible Materials in Vascular Devices Several different polymeric hollow ﬁber membrane materials have been proposed as biocompatible materials for islet encapsulation. Zekorn et al.[31] studied performance and biocompatibility of 50 kDa molecular weight cut-off (MWCO)

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membranes made of different membrane materials: Thomapor (polyamide), Cuprophane HDF (cellulose), and Amicon XM-50 (polyvinyl chloride acrylic copolymer). The membrane devices with seeded rat islets were implanted in the right lobe of the liver of streptozotocin-induced (STZ-induced) diabetic Lewis/Han rats. From these experiments the authors determined that islets enclosed in the selected membranes could not survive in a sufﬁcient number for more than a few days. There was evidence that this was the result of protein coating of the membrane during the culture period, not a problem of islet adaptation to the hollow ﬁbers culture conditions. Instead of being released, insulin was nonspeciﬁcally adsorbed to the membrane surface. Only when all available adsorption sites were ﬁlled did insulin begin to be released into the lumen [31]. Protein precoating of the Amicon XM-50 and PEEK (polyetheretherketone hollow ﬁbers) membranes was then performed by preincubation of the membranes in medium supplemented with 10 % FCS (Fetal Calf Serum)[32]. It improved islet function and cell viability. However, the improvement of diffusional properties of the membranes by protein coating might not have been the only cause for device-increased performance. A very important aspect of membrane effectiveness in a bioartiﬁcial organ is the ability to prevent toxic agents like lymphocytes and cytokines from passing through or adhering to the membrane. Diffusion of toxic cytokines into the inner space causes the destruction of the islets, and cellular adherence to the outer surface of the membrane causes ﬁbrosis and loss of cell viability due to impaired diffusion of nutrients and oxygen. Some tests performed with the 50 kDa MWCO PEEK hollow ﬁbers showed that encapsulation in this particular membrane effectively protected the islets from the toxic effect of Interleukine-1 (IL-1), one of the cytokines thought to play a role in the immune destruction of islets [33,34]. Since the previous work showed improved insulin diffusion with protein precoated membranes, the PEEK was also preincubated with medium supplemented with 10 % FCS. After 2 days of in vitro culturing encapsulated islets in IL-1 supplemented medium, the islets were still able to respond to a glucose stimulus. They secreted insulin in amounts similar to those of control islets (free ﬂoating, without IL-1 supplement) and higher insulin levels compared with those of free islets cultured under the same conditions in IL-1 supplemented medium. Amicon XM-50 membranes were still the ones preferentially used in vascular and extravascular device experiments over the last decade [14,31,32,35–40], especially in the most signiﬁcant set of experiments in vivo with an intravascular prototype, conducted by Monaco and his co-workers [14]. Experiments were performed with dogs, using a prototype device consisting of a single-coiled large-bore Amicon XM-50 membrane, contained in a disk-shaped acrylic housing. The device was AV-anastomosed. To prevent aggregation, which leads to loss of function, the islets were suspended in the annular membrane compartment, which was ﬁlled with sodium alginate (SA) hydrogel. The authors demonstrated the long-term patency of this nonthrombogenic membrane device by implanting unseeded devices and performing the follow up for up to 20 months. The membranes could withstand arterial pressure for long periods of time without losing physical integrity. Nonetheless, 10 out of 22 devices implanted eventually occluded.

9.2 Bioartificial Pancreas

These in vivo results represented the best results ever obtained for allo- or xenogeneic implants: The longest graft survival times of 284 and 80 days, respectively, without administration of insulin were presented in this article. The most important cause for graft failure in these experiments by Monaco and co-workers was the apparent loss of islet function and/or viability. Fibrotic tissue formation was not observed on the blood stream side of the semipermeable membrane. Infection was an important cause for graft failure, but the main problem seemed to be an inadequate insulin secretion and a delayed response time to glucose stimulation. This poor performance may have been due to an insufﬁcient number of islets implanted, which was conﬁrmed by results from a double-device implantation experiment. Maki et al. [40] devised three other prototypes of larger pore-sized membrane material (MWCO of 70 kDa) and geometry similar to those already described. Totally pancreatectomized dogs were used as models of diabetes, and xenogeneic porcine islets were seeded into these vascularized BAP models, without administration of immunosuppressive drugs. The prototypes differed from each other in length (30 and 45 cm) and volume (7.1–15 mL). The overall reduction of exogenous insulin requirements for the prototypes varied between 54 and 74 %. In seven of nine dogs, administration of exogenous insulin between 4 and 16 U/day was needed to keep the glucose levels below 11.1 mmol/L. Two animals maintained glucose control without additional insulin therapy for more than 8 months. A device removed and dissected on day 271 showed acute vascular occlusion of the device, which the authors attributed to hypercoagulability, but there was no evidence of lymphocyte or inﬂammatory cell inﬁltration. The improvement in long-term survival of discordant xenogeneic encapsulated islet grafts and in glycemic control with a single implanted device was more signiﬁcant when larger prototypes were used, but not necessarily seeded with a larger number of islets. Searching for membranes adequate for intravascular use, with better diffusional properties and which induce less host sensitization and consequent immune responses, is the main current research subject [41–44]. Recently, hydroxymetilated polysulfone (PSU) was tested in vitro and used in the development of a prototype for a vascular device [45]. Nonmodiﬁed PSU had already been used for xenotransplantation of porcine islets into CD1 mice [17]. Graft survival time was very limited, 13 days. Cytotoxic anti-pig islet cell antibodies were present in the recipients after 8 days of coil implantation. The potential of PSU for BAP was demonstrated. The smooth surface of the membrane allowed little or no ﬁbrosis to appear. Lembert et al. [45] used a smaller cut-off, 50 kDa, and performed surface modiﬁcations to improve the permeability characteristics of the biomaterial [45]. PSU is hydrophobic and adsorbs large amounts of proteins, particularly insulin. Rendering it hydrophilic would improve protein diffusion. To reduce diffusion distance between the islets and the bloodstream, the islets were encapsulated in narrow capillaries, which were then coiled to form the inner surface of an artiﬁcial vessel (Figure 9.2). The in vitro testing showed that, over a 3-day period, this material allowed the lag time for insulin to be reduced to 5 min and the maximal release time to be reduced to 30–45 min, which is similar to that of free-ﬂoating islets. The reduced kinetics of

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Fig. 9.2 Vascular prosthesis: (a) Diffusion chamber (b) Compartment with coiled capillary; Immobilization matrix, * islet, –—Immunoprotective membrane, !Blood ﬂow. Source: Modiﬁed from Lembert et al. [115].

glucose-induced insulin occurred only when PSU was chemically modiﬁed by substitution with hydroxymethyl groups, at a degree of substitution near two. The prototype is still waiting to be tested in allogeneic islet transplantation in pancreatectomized pigs. Some authors also tried to ascertain the biocompatibility of other materials by in vitro [46] and in vivo [15,37,47] experiments using extravascular devices. 9.2.3 Extravascular Devices

Extravascular BAP models present some advantages like a smaller surgical implantation risk, easier retrievability (for macrocapsules), possibility of device re-seeding, and size ﬂexibility. However, they also have some major drawbacks, the most important of which is the diffusional hindrance to the transport of nutrients, hormones such as insulin, oxygen, and metabolites. Other problems also limit the success of these BAPs, like the ﬁbrotic response and cellular adherence to the outer wall of the membrane, loss of physical integrity of the capsule, especially in the case of microcapsules, which in turn make total device retrievability difﬁcult. 9.2.3.1 Implantation Sites Extravascular devices can be classiﬁed by the size of the device into two categories: macrocapsular and microcapsular devices. Several implantation sites have been proposed and studied for the different types of capsules: epiploic chambers in the great omentum, area around the splenic vein, external wall of the stomach, back wall of abdominal muscles and peritoneal cavity [23], renal cavity, kidney capsule, subcutaneous sites [48], and so on.

9.2 Bioartificial Pancreas

Icard and co-workers [23] studied the best implantation site for an extravascular hollow ﬁber device in pig recipients. XM-50 hollow ﬁber macrocapsules with 100-mm wall thickness were ﬁlled with human islets and implanted in several sites of STZinduced diabetic pigs. The preferential implantation site (lower ﬁbrotic overgrowth) that emerged from this study was the peritoneal cavity. This became the implantation site chosen by most other authors to perform in vivo studies with extravascular hollow ﬁber BAP even when using other animal models. The subcutaneous implantation site was ﬁrst used by Lacy et al. as an attempt to minimize ﬁbrotic response and cellular adherence by the host and also to determine if this was a better nutrient diffusion site compared with the intraperitoneal site [36]. Others [48,49] evaluated device performance in the subcutaneous site, and it has been used especially when prevascularization is performed, prior to seeding of the device. Prevascularization is promoted in an attempt to minimize diffusional limitations and mimic the pancreatic beta cells, where the cells are directly in contact with the capillaries present within the islets. Compared to intraperitoneal implants subcutaneous implants involve even less surgical risk, are easier to seed and re-seed, or remove if failure occurs. 9.2.3.2 Macrocapsules Macrocapsules can have, essentially, two different geometries: tubular and planar sheet. The tubular membranes are usually made up of a hollow ﬁber membrane, with the islets immobilized inside with both ends sealed. These systems sometimes have groups of individual tubes put together, called modules, to reduce device size and increase the number of islets seeded into it, maintaining the transfer area. Some authors who performed studies with intravascular devices also attempted to develop hollow ﬁbers extravascular devices, using the same membrane material [39,50]. The overall graft survival time under similar experimental conditions but with extravascular instead of intravascular devices was shorter, as expected, since there are additional nutrient, oxygen, and metabolite diffusion problems that condition the islets survival when inside the capsule. Flat sheet membranes are essentially made up as large circular or rectangular sealed chambers with the islets immobilized inside. Due to their size and geometry these devices are generally implanted in the peritoneal cavity. This conﬁguration is especially used for in vivo experiments with vascularized devices, which are preferentially implanted in the subcutaneous space. Because it is difﬁcult to maintain membrane planarity, membrane breakage frequently occurs due to the large size of the device [51], rapidly leading to device failure. This geometry is, however, very useful when performing islet cells viability studies [52] or when trying to assess biocompatibility, diffusional, and immunogenic properties of a particular membrane [15,46]. Insulin and glucose transport have been mathematically modeled for this device geometry [53,54]. Biocompatible Materials in Macrocapsular Devices The biocompatibility of several different membrane materials and their ability to work as immunoprotective barriers in xenotransplantation of islets were established with in vivo experiments [15,37,47].

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Prevost et al. [47] evaluated the AN69 Hydrogel biocompatibility by transplanting and comparing the efﬁcacy of free-versus-encapsulated islets. AN69 membranes of 80 kDa MWCO, common for renal dialysis, are composed of 69 % acrylonitrile and 31% sodium methallyl sulfonate. Four different groups of normal or STZ-induced diabetic Lewis rats were implanted with either 6000 free Lewis rat islets or 9000 Lewis rat islets seeded in an AN69 hollow ﬁber membrane. The diabetic animals transplanted with free islets remained normoglycemic for only 3 weeks, whereas those transplanted with encapsulated islets remained normoglycemic for 54 and 70 days, respectively, until the removal of the membrane devices, after which the animals became hyperglycemic again. The biocompatibility of AN69 proved to be satisfactory, since host reaction to the syngeneic implant was reduced to a thin layer of ﬁbroblasts. The implantation area in these animals remained intact: No tissue necrosis or cellular inﬂammation was observed. Nonetheless, neither glycemia nor diuresis completely returned to normal level, which implied the need for optimization of the diffusion potential. Lanza et al. [37] had performed extravascular xenoimplants with capillary hollow ﬁbers (XM-50) in STZ-induced diabetic rats, to determine the protection against rejection, in the absence of immunosuppressive or anti-inﬂammatory drugs. To minimize ﬁbrosis formation, membranes with smooth external skin were used, which did not result in a ﬁbrotic reaction, after more than 150 days postimplantation. However, these results can hardly be extrapolated since the immune system differs greatly between different species [23]. After conducting studies using intravascular devices, Amicon XM-50 was also tested in the macrocapsular conﬁguration. Allogeneic in vivo transplantation studies in dogs, made diabetic by total pancreatectomy, were conducted to determine overall graft survival, glucose control and adequate insulin release, as well as immune responses induced by the device [38]. After transplantation into the peritoneal cavity of 155–248 chambers, each embedded with 2–4 105 islet equivalents, three of the six dogs regained complete endocrine pancreatic function for 51 to more than 82 days (experiments were still ongoing at the time of publication). These results showed that this device was capable of responding to glycemic stress and could control blood glucose concentrations in large diabetic patients, albeit for relatively short periods of time (compared with 284 days in [14]). The difference in graft survival time was thought to relate to the change from an intravascular to an extravascular BAP, which affects nutrients, oxygen, and insulin availability and diffusion rates. Still, the authors proved it was possible to achieve exogenous insulin independence for prolonged periods of time without the need of any immunosuppressive or anti-inﬂammatory regimen. A planar diffusion chamber was made of Nuclepore membranes with a pore size of 0.1 and up to 0.6 mm to assess the membrane biocompatibility and the inﬂuence of pore size on permeability properties [15]. Diabetic Wistar rats were implanted with transgenic mouse insulinoma cell clusters (MIN6 cells) trapped in a mixed matrix (1 % (w/v) agarose, containing 0.005 % (w/v) HEMA(2-hydroxy ethyl methacrylate) copolymer and 0.15 mg/mL collagen gel) inside the Nuclepore membrane chambers, which were sandwiched and edge-shielded with silicone.

9.2 Bioartificial Pancreas

Only 0.1 mm pore-sized Nuclepore membranes completely reversed diabetes and normalized glycemic levels for up to 20 weeks, suggesting that a device with these characteristics might provide a good option to improve diabetes treatment. Later, these Nuclepore membrane devices with the same pore-size range were further used to perform in vivo xenotransplantation of encapsulated MIN6 cells into STZ-induced diabetic Wistar rats [55]. Both devices with 0.1 or 0.2 mm pore-sized membranes, transplanted intraperitoneally, reversed the diabetic state of the recipients for at least 3 months. Outside this membrane pore-size range, diabetes was only partially reversed and recipient hyperglycemia returned to preimplantation levels in less than 3 weeks. The immunoisolating system with the semipermeable membrane/mixed matrix reduced the IgG and complement permeation into the membrane. This immunoprotective effect was best when 0.1 mm pore-sized membrane was used. These reports explain the classic compromise needed for workable devices in any BAP: Small pore sizes prevent nutrients and insulin from diffusing in and out of the inner space of the membrane device, but larger pore sizes allow immunoglobulins (IgGs) and other cytotoxic factors to enter this space and destroy the islets. Diffusional and permeability properties of membrane materials can also change during the lifetime of a BAP device implanted into an animal. In order to clarify one possible mechanism, protein coating of the membrane was performed. The ﬂat BAP with the AN69 membrane sheet (area of 700 mm2, MWCO of 65 kDa circular sandwich prototype) was coated with proteins before in vitro tissue culture in an attempt to mimic an in vivo peritoneal environment [46]. Its permeability to glucose and insulin was then assessed in culture, and the results compared with those obtained previously with similar membrane devices implanted in rats and recovered 1 or 7 days afterward. The implanted noncoated AN69 membrane presented no tissue necrosis or cellular inﬂammation after 1 or 7 days and maintained a well-preserved porous structure (no cellular adhesion or ﬁbrin deposits on the surface). The amount of proteins adsorbed into the implanted and retrieved membranes were much higher than the amount adsorbed into the in vitro protein-coated of membrane over the same time period. Clearly, different mechanisms were involved in the adsorption of proteins into the membrane in vivo and in vitro, depending, respectively, on the protein composition of the peritoneal ﬂuid and the culture medium. Even so, a signiﬁcant reduction in glucose and insulin diffusion occurred in both precoated and noncoated implanted membranes. Both membranes showed a 50 % reduction of glucose and insulin permeability, indicating that protein coating is a likely reason for the permeability change in vivo. This is in disagreement with the study by Zekorn et al. [31,32], which showed an increase in insulin permeability when XM-50 and PEEK hollow ﬁber membranes were subjected to protein coating. The membranes used in each study had quite different physicochemical properties, which is probably the basis for this discrepancy in results.

275

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Biocompatibility of Inner Immobilization Matrix Aggregation of islets within the membrane capsule results in impairment or loss of islet function within 1–2 weeks due to hypoxia and central necrosis [36]. To prevent aggregation and preserve islet function, the combined use of gel entrapment matrices together with immunoprotective barriers has been studied. A mesh-reinforced polyvinyl alcohol hydrogel tubular (MRPT) device was used for in vitro studies of biocompatibility of material and long-term maintenance of islet function [56]. Agarose-dispersed islets seeded in the MRPT device showed an even dispersion of the intact islets with an acceptable time response to glucose stimulation. Other substances, including collagen, hyaluronic acid, and SA were studied to prevent islet aggregation [57]. Islets were immobilized in collagen, type IA, 0.5 % hyaluronic acid, 1 % SA gel, and 5 or 7 % agarose, before being seeded into the MRPT. Of all substances tested, only the alginate gel allowed islets to stay intact and evenly dispersed for 3 weeks. The differences in the gelling procedure for each substance could have inﬂuenced this outcome, especially the consistence of the gel and gelling time. Agarose was gelled by immediate cooling, minimizing the chance for islets to aggregate. Polymeric mesh structures were also used to assess the efﬁcacy of coimmobilization of molecules or cells as an immunoprotection strategy. Pollok et al., for instance, used a polyglycolic acid (PGA) polymer mesh structure as a scaffold and performed in vivo allo- and xenotransplantation studies with islets coencapsulated with autologous chondrocyte within these scaffolds [58]. Chondrocytes were used to determine if they could act as an immunoprotection barrier. Islets from Lewis rats were seeded in a PGA sheet scaffold (1 1 0.06 cm) and then encapsulated in a monolayer of chondrocytes. After 3 days in static culture, the capsules were implanted subcutaneously into nude mice. The encapsulated islets remained viable and functioned in vivo for up to 2 weeks and recently for up to 5 weeks. The glucose/insulin feedback mechanism remained intact [59,60]. This procedure would need to be repeated in a diabetic animal model. Inﬂuence of Membrane Surface Structure on Biocompatibility Lacy et al. [36] studied the inﬂuence of the outer surface (roughness) of the encapsulating membrane with respect to the ﬁbrotic and immune response using subcutaneous and intraperitoneal xenoimplantation of encapsulated islets. Individual rat islets were ﬁrst encapsulated in poly-L-lysine (PLL) coated alginate microspheres. To prevent islet aggregation islets were embedded in a SA hydrogel. The encapsulated islets were then immobilized inside two different types of hollow ﬁbers: one ﬁber type had a totally fenestrated outer wall and the other had a smooth outer wall. This immobilization process did not interfere with the transport of insulin or nutrients, as was proven by experimental controls. Biocompatibility of both ﬁbers was similar and acceptable (only a thin layer of collagen and ﬁbroblast were present surrounding each membrane). Better immunoprotection and control of glycemic levels were obtained with the smooth ﬁber containing a lower number of islets in the subcutaneous position while the fenestrated ﬁber yielded the best results in the intraperitoneal position. The authors attributed these differences to a greater

9.2 Bioartificial Pancreas

bioburden of cells throughout the walls of the fenestrated ﬁbers and to a lower oxygen tension in the subcutaneous than in the intraperitoneal implantation site. Besides, further experiments [35,38] showed host ﬁbrotic response, and BAP failure could be minimized if the outer surface of the membranes was smooth and continuous. In another xenogeneic experiment, macroencapsulated dog islets, immobilized with SA gel inside 50–80 kDa MWCO acrylic tubular membranes, were transplanted into the peritoneal cavity of diabetic BB/Wor rats [37]. Eight of the ten rat recipients, xenotransplanted with 1.2104 to 2.5104 encapsulated islet equivalents, showed normal glucose levels for at least a month, whereas two animals sustained normoglycemia for 2 and 8 months. Tubular membrane chambers with smooth external skin reduced host inﬂammatory pericapsular responses in these rats. The external membrane surfaces were generally free of ﬁbrosis and host cell adherence. However, many capsules retrieved from animals at 1 and 2 months post-transplantation were broken or bend. This implied the need for improvement in membrane length, wall thickness, and/or diameter. The overall data indicated that the implants were capable of responding to glycemic stress by correcting hyperglycemia within 1 h. Tolerance to intravenous injection of glucose (0.5 g/kg body weight) was signiﬁcantly improved but not normalized. More recently, a group of researchers described their work on the development of a prototype called the Islet Sheet [51]. The Islet Sheet is a multilayered construction of alginate: the islets, inserted in a thin ﬂat sheet without polymer reinforcement, are sandwiched on both sides by an acellular immunoprotective alginate layer. The overall thickness of the Islet Sheet has to be the smallest possible (the typical device measures approximately 4 cm 8 cm 250 mm). Like all devices implanted in the peritoneal cavity, it not only relies on passive diffusion for transport of nutrients, oxygen, insulin, and other metabolites, but also has to assure protection of the islets to prevent host immune response; so a compromise must be reached. As is the case with hollow ﬁber devices [35,37], the outer membrane wall has to be smooth and continuous; any irregularity in the texture of the surface can lead to ﬁbroblast attachment and trigger ﬁbrosis. Allogeneic islets encapsulated in the Islet Sheet and implanted in the omentum of a pancreatectomized beagle dog maintained fasting euglycemia for 84 days [51]. Still, a strong ﬁbrotic capsule (mainly collagen) developed around the macrocapsule. This caused the loss of sheet planarity, further stimulating ﬁbrotic response. Neovascularization of Macrocapsules In order to improve islet nutrient supply and oxygenation as well as insulin secretion, a new concept, that of vascularized devices, was developed during the last decade. To build such a device researchers produced an artiﬁcial implantation site, a scaffold, for the islets. The scaffold is usually implanted a few days or weeks before seeding it with the islets to promote vascularization or neovascularization (development of new capillary vessels around and inside the device) and minimize damage to the islets due to prolonged hypoxia, especially immediately after transplantation. One group of researchers determined the performance of an artiﬁcial islet implantation site made of a polytetraﬂuoroethylene (PTFE) solid support

277

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9 Membranes in Bioartificial Pancreas

sheet [61]. Human islets were encapsulated in a PTFE scaffold and transplanted into the epidermal fat pad of diabetic nude mice. After 6 months, recovery of the implanted devices showed that more than 50 % of the cells stained positive for insulin. Close to the membranes were numerous small blood vessels, suggesting the device had induced vascularization. There was no foreign body response against the scaffold membrane. The effect of protein coating as promoter of device vascularization was explored with the initially developed PTFE device [49]. The scaffolds were coated with acidic ﬁbroblast growth factor (a-FGF) and implanted in rat recipients 4 weeks prior to isogeneic islet seeding. Only the solid PTFE supports that were coated with a-FGF were found covered and penetrated by a large number of blood vessels. The successful islet isografts allowed the diabetic Albino Oxford (AO) rat recipients to remain normoglycemic for the duration of the experiment (6 months). Large numbers of viable islets were found in the solid supports retrieved from the intraperitoneal cavity. The PTFE scaffold allowed successful transplantation of islets to the peritoneal cavity, transforming it into a more efﬁcient transplantation site. Other authors also attempted to induce vascularization of BAP devices, using two different membrane constructions and comparing their performance at different implantation sites (renal subcapsule, intraperitoneal cavity, subcutaneous site) [48]. Thin ﬁbrils of polyvinyl alcohol (PVA) or PGA were woven into soft, pliable, mat-like structures to produce ﬂat sheets where free or encapsulated allogeneic islets were seeded. Free islets implanted under the renal capsule performed better. The best results with encapsulated islets were obtained with PGA polymer sheets implanted subcutaneously (four out of ﬁve mice maintained normoglycemia for 3 months after implantation). Successfully grafted devices showed many intact islets and were neovascularized. The authors could not determine which speciﬁc contributions of the PGA material caused this graft success. But the bad results obtained with PVA polymer were probably a consequence of PVA being more loosely woven than PGA, thus causing islet leakage from the membrane. The efﬁcacy of prevascularization of other planar diffusion chambers was assessed by performing xenotransplantation experiments of macroencapsulated rat islets into the subcutaneous site of STZ-induced diabetic Swiss Webster Nude mice [62]. The 20 mL macrocapsules used (Thera Cyte, from Baxter) were made of three different layers of membrane laminates: an inner cell-retentive 0.45 mm membrane, laminated to a 5 mm pore diameter vascularizable PTFE membrane, with an outer polyester mesh providing support. Islets in culture medium were seeded into capsules previously implanted subcutaneously (2 weeks earlier); their performance was compared with implants without prevascularization. Prevascularization did not improve the outcome of the implants; the time needed to achieve normoglycemia was similar, between prevascularized or non-prevascularized macrocapsules. Evidence of AN69 hollow ﬁber membrane neovascularization potential was also reported [18]. Intraperitoneal xenotransplantation of porcine islets into diabetic NOD (Non Obese Diabetic) mice was originally performed, in this study, to

9.2 Bioartificial Pancreas

determine the ability of the devices as immunoprotection barriers. Seeded capsules (40 cm-long ﬁbers, 500 mm i.d., 80 mm wall-thickness and 80 kDa MWCO) reversed the diabetic state in only 18 out of 54 recipient NOD mice for 6 weeks. This showed that the ﬁbers provided at least transient protection from xenograft rejection to porcine islets. Capsules retrieved from successfully implanted mice showed functional islets 26 days after implantation. Two out of four ﬁbers retrieved from mice that remained hyperglycemic contained islets showing a response to glucose. These ﬁbers showed periﬁber inﬂammatory tissue (ﬁbrotic reaction and cellular adherence). Additionally, neovascularization of the periﬁber tissue was observed in more than 20 % of the random sections. This neovascularization process may have improved the supply of nutrients and oxygen to islets. Nonetheless, the efﬁciency of the encapsulated islets transplant was limited by the percentage of successful recipients (only 37 %) and by the duration of graft function (6 weeks). Although it was not proven, the formation of the periﬁber inﬂammatory tissue might have been associated with the promotion of neovascularization, having a beneﬁcial effect on the viability of entrapped islets. Smooth surface regenerated cellulose hollow ﬁbers (Thomapor) have also been reported to spontaneously promote vascularized connective tissue overgrowth, when implanted intraperitoneally in rodents [63]. Nevertheless, only grafts without surrounding connective tissue had a stimulatory response to high levels of glucose. In an attempt to improve oxygen delivery to the islets encapsulated in agarose-PSSa, a study was undertaken where a subcutaneous site was coimplanted with basic ﬁbroblast growth factor (b-FGF) impregnated microspheres, as shown in [48,49]. The b-FGF-impregnated microspheres were implanted into C57BL/6 mice 2 weeks before implantation of the macrocapsular device [64]. The device implanted was made of a mixture of agarose crosslinked with polystyrene sulfonic acid (PSSa) with a tubular shape. This type of hydrogel is more usually found in experiments with microcapsular devices (vide Section Agarose-PSSa Microcapsules). The performance of devices coimplanted with or without b-FGF-impregnated microspheres was assessed. Intraperitoneal glucose tests were performed 1 and 2 months after transplantation. Vascularization assessment around the implantation site was histologically performed with subcutaneous tissue samples. In two recipients, graft survival was evaluated after device retrievement at 1 and 2 months post-transplantation. All the recipients of the subcutaneous xenografts returned to normoglycemia. The average normoglycemia period was 68.4 25.6 days, in contrast with the control groups without b-FGF-impregnated microspheres, which maintained normoglycemia only for 9.5 5.1 days. Glucose tolerance tests at 1 and 2 months with b-FGFþ mice showed signiﬁcant improvement but were still not as good as for control mice. b-FGFimpregnated gelatine microspheres, unlike nonimpregnated ones, induced signiﬁcant subcutaneous neovascularization in diabetic mouse 14 days after injection. Viability assessment in recipients of the b-FGFþ group revealed the islets within the graft were viable and intact and no ﬁbrotic overgrowth or pericapsular cellular inﬁltrate were present around the surface of the graft. In the b-FGF– and control groups, however, remnant islets in the removed graft were found to be necrotic, although no

279

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9 Membranes in Bioartificial Pancreas

signiﬁcant ﬁbrotic overgrowth was observed, indicating islet malnutrition as the probable cause of necrosis. When recipients returned to normoglycemic state, they were subjected to an intraperitoneal injection glucose tolerance test. Their insulin secretion levels were found to be improved but reduced and returning more slowly to basal levels than in normal animals, which pointed to plasma insulin responses being insufﬁcient. The cylindrical device presents many technical advantages: it is injectable, retrievable, durable, difﬁcult to mechanically disrupt, and suitable for subcutaneous implantation. The additional use of b-FGF-impregnated gelatine hydrogel or gelatine microspheres strongly induced neovascularization and tissue granulation around the implantation or injection site. However, biocompatibility of this agarose-PSSa mixed gel-based BAP device was attenuated over time, which eventually led to graft failure. Also, if a tighter homeostatic control of glucose metabolism could be achieved by this xenotransplantation technique, it could be possible to achieve long-term reversal of IDDM in human patients. 9.2.3.3 Microcapsular Devices Aside from macrocapsular devices, researchers have also developed a microspherebased BAP. In this case, the cells are immobilized inside microspheres of a biocompatible material usually alginate or agarose gel, coated or not with a semipermeable membrane, and implanted in the patient’s body. This kind of device was the most extensively used in the development of a BAP mainly because of the simplicity of its construction and its ﬂexibility for modifying key components. This construction ﬂexibility allowed researchers to play with key parameters like wall thickness, pore size, or membrane compactness to overcome signiﬁcant nutrient and oxygen diffusion limitations seen in all extravascular implantation sites. Nonetheless, the geometry presented technical problems, of which the major was in many cases the difﬁculty to retrieve the microspheres from the patient’s body, especially if ﬁbrotic reaction or cellular adherence occurred. Several problems were addressed initially and others were evidenced when microcapsules were used in vivo, like the development of a strong ﬁbrotic response, difﬁculties in microcapsule construction, and ideal polymer composition. Poly-L-Lysine (PLL) and Alginate-PLL-Alginate (APA) Microcapsules Sun et al. were the ﬁrst to report the use of microcapsules as a BAP [65] and later optimized prototype characteristics. Brieﬂy, the microcapsules were produced in the following manner: Islets were suspended in a SA solution, and droplets of the suspension were produced by syringe pump extrusion with a droplet-forming apparatus and dropped into a calcium chloride solution, for bead gelling to occur. The alginate beads were then incubated in a PLL solution, washed, and incubated with sodium citrate to liquefy the alginate gel inside the capsule. Several alterations to this method were made in follow-up experiments, namely the droplet generation process, alginate purity, and its concentration and incubation times. Consequently, capsule properties changed which, as will be shown, had a crucial effect on graft survival. A special

9.2 Bioartificial Pancreas

modiﬁcation with respect to bead size allowed obtaining microcapsules smaller in diameter than the previous ones (250–350 mm), most of them containing only one islet (Table 9.2). Many studies were performed where chemical composition of capsules [66,67], purity [68,69], bead size [70,71] and shape, integrity, and mechanical properties [72–74] of APA microcapsules were examined and related to the ability to induce or not an immune response, when implanted in an animal recipient. Other modiﬁcations included the use of fetal rat islets [75] because their isolation technique is relatively simple, requiring no Ficoll separation or laborious handpicking. The results proved that these islets were well suited for this kind of device construction. Sun and co-workers also used Wistar rats as donors and Balb/c mice as recipients (concordant xenografts) and a similar microencapsulation method [20] but with smaller capsules, which more than tripled the duration of normoglycemia following transplantation when compared to the already good results obtained with fetal rat islets. Since smaller size beads allowed easier access to oxygen and nutrients, the islet viability was increased, which, in turn, promoted prolonged graft survival. The eventual graft failure that was observed for some recipients may have occurred for several reasons: insufﬁcient number of islets transplanted, presence of cellular overgrowth, or capsules clumping together (which can lead to islet central necrosis). All of these aspects were absent in recipients that remained normoglycemic and emphasized the need to clarify the inﬂuence of capsule characteristics on graft survival including mechanical structure, size, and alginate composition. Zhou et al. tried this microcapsular prototype in a large animal model with pig islets into diabetic monkey recipients. Porcine islets were microencapsulated in alginate-PLL microspheres and transplanted into spontaneously diabetic Cynomologus monkeys. The islets were isolated and puriﬁed by means of two different processes, which yielded two islet preparations with different degrees of purity, 75– 90 % and more than 95 %, respectively. The microcapsules used were similar to those of [20]. The general condition of all recipients improved substantially [76]. Sun et al. reported the ﬁrst ever long-term discordant xenograft function without any immunosuppression in a large animal model, resulting in normal glycemic control [77]. Several months after transplantation, the reappearance of hyperglycemia could be reversed after a second implant with a similar number of encapsulated islets. In two recipients that never became insulin independent, graft failure was attributed to damage caused to islets by contaminating exocrine tissue (these recipients were transplanted with lower purity islet preparations). The long-term survival of the microencapsulated islet grafts in these experiments could be attributed to two major factors: the strength of the capsular membrane and the purity of the porcine islet tissue. The purity of the alginate solution was also very important with respect to the biocompatibility of the capsules. This was investigated later by other authors, using immunogenic assays. Effect of Alginate Chemical Composition Alginate gels exist in varied forms and compositions and usually their exact and complete composition is not known, which

281

M.c.s had smooth surfaces and Ø between 400 and 600 mm

STZ-induced diabetic Balb/c mice, i.p. transplant, seven with free islets and nine with m.i.s

STZ-induced diabetic Balb/c mice, i.p. transplant, four with free and 10 with m.i.s

800–1000 free or encapsulated Wistar Rat fetal pancreatic islets/recipient

900–1000 Wistar rat free islets or m.i.s/recipient

M.c.s with Ø between 250 and 350 mm

N/A

STZ-induced Wistar Lewis Rats, i.p. transplant, ﬁve with m.i.s

Wistar rat islets (number of islets N/A)

Microcapsule parameters

Recipient/ implantation site

Experiments with APA microcapsular devices.

Number of islets transplanted

Tab. 9.2

Until hyperglycemia returned

Until hyperglycemia returned

APA m.c.s

2–3 wks, in vivo

PLL: 37 kDa MW

APA m.c.s: PLL with 17–22 kDa MW

15 wks, in vitro

Experiment duration time

APA m.c.s

Encapsulation method/material

End of experiment, recurrence of hyperglycemia

Recurrence of hyperglycemia, in vivo End of experiment, recurrence of hyperglycemia

Loss of functionality, in vitro

Device removal reasons

M.i.s morphologically and functionally intact throughout in vitro culture studies (15 wks) Five recipients of m.i.s maintained normoglycemia for 3 wks Reversal of mice diabetic state for between 54 and 171 days, without immunosuppression Capsules removed from two recipients (144 days postimplantation) were essentially free of ﬁbrosis and cell overgrowth and had minimal or no structural damage Xenografts functioned for periods ranging from 172 to 308 days Recovery of approximately 80 % of the capsules implanted About 70 % of recovered capsules free of cell overgrowth and with viable islets

Major outcomes/results

[20]

[75]

[65]

Ref.

Mongrel dog islets (number of islets N/A)

One, two, or three transplants of 3104 – 7104 porcine free islets or m.i.s/ recipient

One or two injections of 3104 – 7104 porcine m.i.s/ recipient

Spontaneously diabetic dogs with IDDM transplanted into the abdominal cavity, three with free islets, seven with m.i.s, four reimplanted with m.i.s

Four spontaneously diabetic cynomolgus monkeys, i.p. transplant with porcine m.i.s Spontaneously diabetic cynomolgus monkeys i.p. transplant, two with free islets and nine with m.i.s

Same as Ref. [65], with high-G alginate

Same as Ref. [20]

Same as Ref. [20]

High-G alginate-PLL m.c.s

APA m.c.s

APA m.c.s

Graft survival occurred when recipient remained euglycemic w/o any exogenous insulin

Until hyperglycemia returned

Until hyperglycemia returned

Until hyperglycemia returned

Recurrence of hyperglycemia

Recurrence of hyperglycemia

Not referred

Normoglycemia restored from 120 to 803 days, in seven of nine recipients No antiporcine islet antibodies detected in the recipients, even after return to hyperglycemia Recovered m.c.s free of cell overgrowth, physically intact and not clumped Insulin secretion of recovered islets was similar to that of freshly isolated ones Graft survival occurred in all animals transplanted with m.i.s for 63–172 days (CysA administered to all recipients) Lower exogenous insulin requirements to maintain euglycemia, following graft failure before transplantation, suggesting that islet survival occurred several days beyond those reported

Complete restoration of normoglycemia for 106–390 days, w/o exogenous insulin or immunosuppression

[78]

[77]

[76]

Recipient/ implantation site

In vitro assay

Normoglycemic AO rats, i.p. implant

Five to six Balb/c mice and ﬁve to six Wistar rats, i.p. transplant

In vitro assay

None

ICCs from porcine fetuses (number of ICCs N/A)

(Continued )

Number of islets transplanted

Tab. 9.2

Average alginate bead Ø: 0.7 mm; one or two ICCs/capsule

Same as in Ref. [69]

M.c.s Ø between 500 and 1800 mm

Microcapsule parameters

APA high-G and high-M, empty with fetal porcine ICCs or with Dynabeads (Dynal, Oslo, Norway) covered with BSA

PLL Bariumalginate m.c.s with Manugel and Keltone LV

APA m.c.s, external alginate layer made of Manugel or Keltone LV

Encapsulation method/material

Different types of m.c.s implanted twice, with a 2-month interval

1 month

In vitro assay

Experiment duration time

End of the experiment

End of the experiment

In vitro assay

Device removal reasons

Keltone LV: lowest overall percentage of protruding islets; islet-containing defective capsules were lowest with 800-mm Ø m.c.s. Manugel: percentage of incomplete encapsulation and defective m.c.s lowest for 500-mm Ø (apparently due to swelling properties of Keltone LV gel) High-M m.c.s were found more biocompatible than high-G ones Minor modiﬁcations in the encapsulation procedure lead to high variability of m.c.s biocompatibility and graft survival Mice developed antibodies against the high-M but not the high-G material Rats did not develop any antibody responses to either of the m.c.s

Major outcomes/results

[80]

[67]

[66]

Ref.

1000–2000 Fisher rat or 2000–5000 mongrel dog m.i.s/mice recipient 300000 mongrel dog m.i.s/beagle dog recipient

29 STZ-induced diabetic Balb/c mice and one beagle dog, islets i.p. implant

Microspheres of 250–300 mm in diameter, 1–3 islet/m.c. Prepuriﬁed high-G and high-M alginate m.c.s (no PLL used) In vivo, graft failure determined of the end of the experiment

Islets cultured for 4 days in vitro

Persistence or return to hyperglycemic state

Immune response proved not only to be connected to alginate type but also to recipient species Twenty-three out of 29 mice transplanted with rat m.i.s sustained normal plasma glucose levels for 56–419 days. Twenty-two out of 40 mice transplanted with dog m.i.s sustained normal plasma glucose levels for 21–630 days M.c.s recovered from mice contained rat islets; a few exhibited central damage, indicating insufﬁcient amounts of nutrients and oxygen in the peritoneal cavity Mongrel dog m.i.s retrieved from mice showed well-granulated endocrine cells (a few necrotic islets) Fibrosis and cell adhesion were observed when islets protruded from the m.c.s Beagle dog remained euglycemic (after gradual withdrawal of exogenous insulin), for over 400 days

[68]

Recipient/ implantation site

STZ-induced diabetic AO/G-rats, i.p. transplant, empty or isletcontaining m.c.s

AO/G rats, i.p. transplant, three with empty m.c.s, seven with small m.c.s, ﬁve with large m.c.s

AO/G or Lewis rats (number of islets N/A)

Lewis rat islets (number of islets N/A)

(Continued )

Number of islets transplanted

Tab. 9.2

M.c.s with two different Ø ranges: 400–500 (small) and 700–800 mm (large)

M.c.s with 600–700 mm Ø and 100 kDa MWCO

Microcapsule parameters

APA m.c.s

Puriﬁed and nonpuriﬁed SA (40% guluronic acid, 60% mannuronic acid contents)

Encapsulation method/material

4 wks after transplantation or 2 wks after recurrence of hyperglycemia

Empty capsules removed 1-month postimplantation; graft failure determined or the end of the experiment

Experiment duration time

Persistent or recurrent hyperglycemia

Persistence or recurrence of hyperglycemia

Device removal reasons

Within 1 month of postimplant, all empty nonpuriﬁed alginate m.c.s were overgrown with ﬁbrotic tissue and surrounded by collagen deposits Majority of puriﬁed alginate m.c.s freely ﬂoated in abdominal cavity, w/o adhesion to abdominal organs Ten out of 13 rats with allogeneic m.i.s implants (puriﬁed alginate) remained normoglycemic for 6–20 wks No relation was observed between the duration of graft function and the number of capsules overgrown by ﬁbrotic tissue Nonovergrown capsules had nonviable islet remnants, after graft failure One in seven recipients became normoglycemic with small m.c.s (>5 % vital b cells in native pancreas) for 8 wks

Major outcomes/results

[70]

[69]

Ref.

2000–3000 Sprague–Dawley rat m.i.s/recipient

STZ-induced diabetic Balb/c mice, i.p. transplant, 20 with m.i.s

Two populations of m.c.s: standard (Ø 350 mm, method of Ref. [82]) and compact (Ø 300 mm modiﬁcation to method of Ref. [82])

Adequacy of encapsulation: % of inadequate m.c.s in total number of islet-containing m.c.s

APA m.c.s

Until blood glucose levels rose above 200 mg/dL for m.i.s

1, 3, and 6-month after implantation, for empty m.c.s

Blood glucose levels rose above 200 mg/dL

Majority of small m.c.s adhered to the abdominal organs (very low m.c.s retrieval rate); extensive ﬁbrosis and cellular inﬁltration present; islets in m.c.s were necrotic and contained no vital b-cells All ﬁve recipients of large m.c.s became normoglycemic for 7–16 wks Majority of large m.c.s were free ﬂoating in the peritoneal cavity and nonadherent to abdominal organs Recipients of compact m.c.s maitained blood glucose levels below 180 mg/dL for 108 days; standard m.c.s maintained these same levels for only 76 days. Empty APA m.c.s were recovered 6 months postimplantation: 81% compact ones intact, 38% standard ones intact Fibrotic neogrowth occurred only when m.c.s were broken or damaged (islets exposed)

[73]

Recipient/ implantation site

Normal Lewis rats, i.p. transplant, four with alginate beads, 14 with APA m.c.s

STZ-induced diabetic AO/G rats, i.p. implant, ﬁve with free islets and nine with m.i.s

Adult porcine or bovine calve islets (number of islets N/A)

2800–3000 free Lewis rat islets or 2500–2900 Lewis rat m.i.s/ recipient

(Continued )

Number of islets transplanted

Tab. 9.2

Alginate beads or APA m.c.s

APA m.c.s

Uniform smooth spherical m.c.s, Ø 800–900 mm; 1–3 islets/m.c.

Encapsulation method/material

Alginate beads: Ø 300–600 mm, 1–2 IE/ microsphere m.c.s: Ø 800–1200 mm, 66 kDa MWCO (measured)

Microcapsule parameters

5–6 wks

2 wks, 1, 2, 3, 7, and 9 months

Experiment duration time

Relapse into hyperglycemia

Removal of device at the end of the experiment

Device removal reasons

Islets encapsulated in uncoated alginate beads nonviable for 2 wks or 1 month after implantation in normal rats Islets immobilized in m.c.s survived transplantation and had similar viability at preimplantation and removal days Nonadherent and intact m.c.s retrieved, with external surfaces mostly free from ﬁbrosis or host cellular adherence. M.i.s removed at 9 months had an average twofold increased insulin secretion with a static glucose challenge Similarly good results were obtained for the xenotransplantation of encapsulated bovine islets into rats Four out of nine recipients of m.i.s remained normoglycemic for 10–14 wks (three of them showed b-cell regeneration in native pancreases)

Major outcomes/results

[83]

[74]

Ref.

Healthy B6 or Balb/c mice or Sprague– Dawley rats

Same as in Ref. [73] APA m.c.s

4 wks

Removal of device at the end of the experiment

M.c.s retrieved from animals that returned to hyperglycemic state were overgrown with ﬁbrous tissue and contained no viable cells Islet recovery ratio was signiﬁcantly different for allogeneic versus xenogeneic and syngeneic versus xenogeneic transplants: 60 % versus 0 % and 1.6 % versus 4 %; neogrowth was very different in both cases: 1.6 % versus 4 % and 2.8 % versus 4 % B6 mice and Sprague–Dawley rats are incompatible with APA m.c.s, but Balb/c mice are a good model for transplantation of these m.c.s

[84]

Abrevitations: m.i.s, microencapsulated islets; m.c.s, microcapsules; wks, weeks; STZ, streptozotocin; Ø, diameter; i.p., intraperitoneal; N/A, not available; IDDM, insulindependent diabetes mellitus; CysA, cyclosporin; w/o, without; ICCs, islet-like cell clusters; IE, islet equivalents; NPCC, Neonatal porcine pancreatic cell clusters.

C57BL/6 (B6) mice or Balb/c mice or Sprague– Dawley rats, 2000–3000 m.i.s/recipient

Removal of device at the end of the experiment

290

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means they may contain substances toxic to the islets or to the recipient. Pure alginates are linear polysaccharides composed of homopolymeric regions of 1,4linked b-D-mannuronic (M) and 1,4-linked a-L-guluronic (G) acid, interspaced with regions of mixed sequence [68]. But commercial alginates are usually crude products extracted from algae and are contaminated with different inﬂammation provoking components, of which only polyphenols and endotoxins have been identiﬁed [69]. The alginate composition was studied by Soon-Shiong et al. [78]. This was the ﬁrst report describing long-term successful reversal of spontaneous diabetes in a large animal model transplanting allogeneic islets with this kind of device. The report focused on the enhancement of microcapsule biocompatibility and mechanical integrity by using high-G-alginate PLL microcapsules and introducing modiﬁcations to the microencapsulation method. Longer periods of euglycemia were veriﬁed for high-G PLL capsules. In an attempt to determine if the development of ﬁbrosis around the microcapsules was due to the chemical composition of the alginate used in the immobilization process, De Vos et al. studied the inﬂuence of the size of the microspheres and their guluronic and mannuronic acid alginate content (G:M ratio) in microcapsule physical resilience and ability to completely envelop the islets, in vitro [66]. Encapsulation was performed as usual, followed by an additional external alginate layer, of Manugel or Keltone LV; Manugel has higher G:M ratio than Keltone LV. The percentage of incomplete encapsulation varied with capsule size and type of alginate used. The authors concluded from this study that alginates with high G:M ratio were better suited for encapsulation. Quantitatively, incomplete encapsulation was the most important manifestation and was lower with Manugel than with Keltone LV. Since Manugel has a higher G:M ratio than Keltone LV, it is associated with less swelling in the absence of divalent ions, which diminishes the possibility of protrusion (Figure 9.3). De Vos et al. continued to investigate the effect of alginate composition (this time with barium-alginate microcapsules) on the induction of ﬁbrotic response induced by different chemical composition alginates [67]. The ﬁbrotic reaction was more severe against Manugel alginate (capsules over-

Fig. 9.3 Schematic presentation of the process of swelling and shrinkage during the encapsulation procedure. Swelling and shrinkage contribute to the occurrence of incomplete encapsulation. Source: Reproduced from De Vos et al. [66], with permission from Lippincott, Williams & Wilkins, copyright 1996.)

9.2 Bioartificial Pancreas

grown and adherent to abdominal organs) than against Keltone LV alginate (capsules free ﬂoating and without adherent cells). No ﬁbrotic response was observed with both kinds of alginates in the absence of the PLL membrane, indicating the problem related to the binding of PLL to guluronic acid present in alginate, which was in agreement with other authors (see the section ‘‘Assessment of ﬁbrotic response’’, Ref. [79]). Kulseng et al. [80] tried to determine the inﬂuence of the alginate composition on the ability to induce antibody responses or act as an adjuvant to antibody responses to antigens leaked from the microcapsule. The authors compared high-G and high-M alginates performance. Two different in vivo sets of experiments were performed: the ﬁrst, to determine the antigenicity of empty capsules; the second, to determine the responses of encapsulated porcine islet-like cell clusters (ICCs), when transplanted into mice and rats. They showed that high-M alginate had a higher immunestimulating activity than high-G alginate, both in terms of antibody and cytokine production, thus supporting results by [81]. To determine the generation of antibodies against ICCs, rats received 50 free or encapsulated ICCs, in high-G alginate-PLL capsules. Antibodies against ICCs were detected 6 weeks after implantation of free or encapsulated islets, and there was an increase in antibody levels 2 weeks after the second injection; after 8 weeks the antibodies started to decrease. This result clearly showed that ICCs, either free or encapsulated, always evoke a humoral immune response, which decreases with time. Surprisingly the highest antibody response occurred in animals transplanted with encapsulated islets. However, as was suggested before, this happened because proteins released from the encapsulated ICCs can pass through the membrane and bind themselves to the outer capsule surface, triggering complement activation and local inﬂammation, thus negating the protective effect of the microcapsular membrane. By implanting into rats BSA covered Dynabeads (commercial beads), with or without alginate coating, it was further shown that the alginate coating completely prevented antibody response, where the absence of it lead to the generation of antibodies against BSA.

Effect of Alginate Puriﬁcation Cochrum et al. [68] also conducted in vitro and in vivo xenotransplantation experiments with microcapsules, but this time high mannuronic and high guluronic acid alginate were puriﬁed (proteins, fucose, and polyphenols were removed) before being used to produce PLL microspheres [68]. This study [68] reported prolonged reversal of diabetes in mice and dog by transplantation of encapsulated xenogeneic islets. Furthermore, it showed that highly puriﬁed mannuronic and guluronic calcium-alginates were stable in vivo and could provide immunoisolation of islets, especially in conjunction with a small diameter biocompatible membrane coating. The absence of ﬁbrosis on implanted empty spheres and absence of host mononuclear cell inﬁltration (except where islets were exposed) was supportive of the efﬁcacy of alginate puriﬁcation in removing immune response inducing contaminants. However, when the survival of the allotransplants

291

292

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in rodents (56–419 days) was compared with that shown by Sun and co-workers [20] (172–308 days) without puriﬁcation, no signiﬁcant improvement on graft survival was observed, thereby questioning the need for this alginate puriﬁcation. The reduced size of the microcapsules used in this work led to protrusion of islets from the capsules. This fact rendered the proper evaluation of the effect under study impossible. Where total recovery of microspheres was possible, the PLL coatings were generally free of ﬁbrosis and host cells; however, if islet tissue was left exposed on the microsphere surface, host mononuclear cell inﬁltration occurred, as was later conﬁrmed by [66]. The inﬂuence of capsule size on graft survival is mentioned below. De Vos et al. also compared the performance of puriﬁed and nonpuriﬁed SA [69]. The capsules had a larger diameter and pore size than those of [68]. In order to test the effect of the immune response against allogeneic islet transplants, isografts were performed. The mean graft survival and percentage of capsules retrieved was very similar with both iso- and allografts, indicating that the limited graft function was not caused by insufﬁcient immunoprotection of the allogeneic tissue. The biocompatibility of the grafts were inﬂuenced both by the chemical composition of the alginate and by the individual imperfections of the microcapsules (similar to the case of extravascular hollow ﬁber devices). Graft failure could have had several causes, including cytokine release (induced by small ﬁbrotic reactions toward exposed islet tissue), insufﬁcient number of islets (leading to graft exhaustion), and problems of nutrient diffusion. The main cause of limited duration of graft function, in this case, seems to be the presence of the capsule itself combined with the choice of implantation site. The lifespan of the islets is very much conditioned by the implantation site. Functional microencapsulated islets remain free ﬂoating in the peritoneal cavity. They depend completely on diffusional phenomena for supply of nutrients, growth factors, and other substances essential for the homeostasis of islet cells. This fact necessarily led to the idea of minimizing, as much as possible, the size of microcapsules, but without compromising the integrity of islet encapsulation procedures. However, for longterm survival of the graft, further functional studies on the diffusion and its limitations are needed.

Effect of Microcapsules Size/Mechanical Stability Chicheportiche and Reach studied, in vitro, the effect of PLL microcapsules size in insulin secretion of microencapsulated islets [71]. They used 47 kDa MW PLL polymer chains to produce two populations of microcapsules: 350 mm (small) and 650 mm (large) average diameters, respectively. Static glucose challenges with evaluation of insulin concentrations were performed. Insulin secretion increased signiﬁcantly with small microcapsules, but insulin secretions of both groups of microcapsules were much lower than that of free islets. These results clearly showed the signiﬁcance of capsule diameter in the diffusional process and, consequently, on the encapsulated islets’ survival. Islets encapsulated by small microcapsules respond faster to stimulation than larger microcapsules, possibly due to the time taken for nutrients and

9.2 Bioartificial Pancreas

metabolites to reach the islet within the capsule, and/or accumulation of insulin in the microcapsule. The driving force for diffusion is the difference in concentration across the membrane and the surrounding milieu, therefore the exchange of stimulatory agents and the ensuing secretory compounds will result in a longer time lag for larger microcapsules. De Vos et al. also studied the inﬂuence of the size of the microcapsules. They hypothesized that a portion of microcapsules in any transplantation procedure would always be overgrown with ﬁbrotic tissue due to the unavoidable presence of protruding islets. They used Sephadex beads (S-beads) to simulate islet encapsulation and studied their distribution and placement in PLL microcapsules, using a common encapsulation protocol. S-beads immobilized in Ca-alginate beads (no PLL outer capsule) were mostly located in the periphery of the latter, although some were located centrally; a small number of S-beads even protruded from the Ca-alginate beads. APA microcapsules with immobilized S-beads showed a higher number of these beads of protruding from the capsule, than those with Ca-alginate beads [66]. In another set of experiments [70], they used microcapsules of two different size ranges and assessed the effect of islet protrusion from the microcapsules in graft survival. With small sized microcapsules only one in seven recipients became normoglycemic and remained so for the ﬁrst 8 weeks after transplantation. Thus, reducing the microcapsule diameter was associated with not only a substantial reduction in volume but also a substantial increase in the percentage of inadequately encapsulated islets. Inadequate encapsulation is a major factor inducing ﬁbrotic overgrowth and subsequent necrosis of the islet cells, and may, therefore, lead to graft failure. The inﬂammatory response developed seemed to be proportional to the number of inadequate microcapsules. The percentage of inadequate capsules was higher for small-sized microcapsule (24 %) than for larger sized microcapsule (6 %) preparations. In contrast, all ﬁve recipients became normoglycemic for 7–16 weeks after transplantation with islets encapsulated in large capsules. The majority of capsules were found free ﬂoating in the peritoneal cavity. Occasionally, some large capsules were found to be adhering to the omentum. However, no islets were found in these attached large capsules. The islets in nonadherent capsules were found to contain large necrotic zones with only a few viable b cells left. The main factor that seemed to lead to graft failure with large microcapsules was islet necrosis, brought about by insufﬁcient nutrition and oxygenation. These results seemed to be in disagreement with other reports [20,75–77] using small microcapsules (250–350 mm in diameter). Size of the microcapsules, however, is not the only factor inﬂuencing graft survival. Alginate purity, PLL MW cutoff, and especially the mechanical stability of the microcapsules are very important, and may inﬂuence the success of the encapsulation procedure. Mechanical stability of APA microcapsules was addressed by Hsu et al. [73], who evaluated microcapsules membrane integrity and its inﬂuence on immunoprotection. Calcium concentration in the alginate gel was increased to prepare predictably more compact APA microcapsules, and their performance was compared to those

293

294

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of a standard APA microcapsule preparation method [82]. Although APA microcapsules protected the encapsulated islets from direct antibody attack, neither could prevent donor antigens from diffusing out of the capsule and inducing host immune responses: cytokines involved in tissue rejection are small enough in size, can enter the microcapsules, and damage the islets. They can also induce ﬁbroblast proliferation, causing the disintegration of the microcapsule. It was very likely that the membrane was rendered more resistant to biodegradation because of its more compact membrane, with less water molecules and stronger and closer ionic interaction between interacting polymers, emphasizing that better in vivo long-term glycemia control results were obtained with compact APA microcapsules. Lanza et al., who had performed several hollow ﬁber studies, also studied the mechanical stability of standard APA and newly developed biodegradable microcapsules, varying alginate concentrations [74]. Calcium alginate microcapsules coated with a 10 kDa MW PLL plus an external alginate layer were produced (800– 1200 mm in ﬁnal diameter). The alginate concentration was varied (0.75–1.5 %) to produce different sets of microspheres, with which mechanical stability and permeability studies were performed. Good in vivo mechanical stability of the microcapsules was observed only for alginate concentrations of 1.5 %. No capsule breakdown was observed when 1.5 % alginate microcapsules were transplanted, and virtually all of the microcapsules were recovered from the animals and had remained intact for the whole duration of the study (1 year). IgG was detected on the external surface of the capsules, but no appreciable amounts of IgG were found inside. The resulting microcapsules allowed transplantation of discordant bovine and porcine tissue into rats and excluded IgGs for over a year, while maintaining islet viability similar to preimplantation levels. One can conclude from these experiments that membrane and alginate gel characteristics synergistically interact with each other and together determine the durability, integrity, and biocompatibility of APA microcapsules.

Fibrotic Response to PLL-Microcapsules Islets encapsulated in APA-microcapsules have also been used as a BAP for evaluation of the induction of a ﬁbrotic response [83,84]. In the work of Fritschy et al., Lewis rat islets were encapsulated in APA microspheres and implanted in STZ-induced diabetic AO rats. The capsules prepared were of 800– 900 mm diameter [83]. Although microencapsulated islet allografts were able to restore normoglycemia, there was reduced insulin secretion. The microcapsules used were relatively large, thereby delaying insulin and glucose diffusion kinetics. Insufﬁcient islet mass transplanted could also have contributed to graft failure, but the most important factor was the occurrence of ﬁbrotic overgrowth, further hindering diffusion of nutrients and metabolites. The intensity of ﬁbrotic response can vary greatly from individual to individual, even within the same species. Phylogenetic differences of islet donors and recipients strongly inﬂuence the ﬁbrotic overgrowth of microcapsules [84]. Hsu et al. used islets from three different rat and mice donor species. They were microencapsulated in APA

9.2 Bioartificial Pancreas

microcapsules and transplanted intraperitoneally into healthy concordant and discordant mice or rats. Three parameters were assessed: islet recovery ratio (IRR), neogrowth, and glucose-stimulated insulin secretion (GSIS). With C57BL/6 mice as recipients, neogrowth in syngeneic islet transplantation was signiﬁcantly less than that of allo- or xenotransplantation. When these mice were used as donors, however, the neogrowth score was less severe than that of syngeneic transplantation and the IRR was signiﬁcantly higher for allogeneic transplantation. No islets were recovered from xenografts. These results showed that phylogenetic disparity affects graft survival, but the authors were unable to clarify the immunological processes involved. Barium-Alginate Microcapsules Instead of using alginate coated with PLL membrane, several researchers [85–87] crosslinked alginate microdroplets with barium ions and developed beads that are usually referred to as Ba-alginate microcapsules. Their performance as immunoprotection barriers has been studied from then on (Table 9.3). In vitro Biocompatibility Assessment In vitro studies were performed to test the properties of Ba-alginate microcapsules with pig islets [85]. Encapsulation of islets in Ba-alginate was shown to be a simple and quick method of immunoprotection, allowing islet functionality to be preserved, although insulin output was greatly diminished when compared to that of free islets. A similar in vitro study compared the performance of Ba-alginate capsules versus thin layer (10 mm) alginate-coated islets [86]. The results also emphasized the importance of having low microcapsule thickness. Still the effectiveness of such thin capsules as immunoprotective barriers was not always good, especially due to lower mechanical stability and deﬁcient islet encapsulation. Effect of Microcapsule Size Concerned with the microcapsule diffusional limitations, Petruzzo et al. developed a different encapsulation process, allowing the production of Ba-alginate capsules of 200 mm average size, within the size range of 50–500 mm, and pore size between 60 and 70 kDa [88]. The capsules retrieved from in vivo rat experiments showed well-preserved islets and absence of ﬁbrosis, if they were still from normoglycemic animals. Those removed from animals that regressed to a hyperglycemic state were free of ﬁbrosis but contained unhealthy islets. The authors attributed this to damage to the islets during the isolation and encapsulation procedures, because porcine islets are extremely sensitive and prone to disintegration. Assessment of Fibrotic Response Evaluation of ﬁbrotic response to Ba-alginate capsules was extensively performed in syngeneic [89], allogeneic [87,89,90], and xenogeneic [79,91] transplantation experiments. Allogeneic transplantation of Wistar rat islets into preimmunized and non-preimunnized STZ-induced Lewis rats, to evaluate recipient immunological response, was performed [90]. There was tissue reaction by the host, which seemed to be

295

In vitro assays

In vitro assays

10 STZ-diabetic rats (species not mentioned) i.p. transplant

50 rat or porcine m.i.s (species names not mentioned)

50 Lewis rat m.i.s for static and 100 Lewis rat m.i.s for perfusion assays

150–200 porcine free islets or m.i.s in vitro

5000–7500 m.i.s in vivo (species not mentioned)

Recipient/ implantation site

M.c.s Ø: 50–500 mm, average diameter 200 mm; 60–70 kDa pore size

N/A

N/A

Microspheres parameters

Experiments with Ba-alginate beads.

Number of islets or islet cells

Tab. 9.3

Ba-alginate beads

Ba-alginate beads, with or without a 10 mm-thickness alginate coating

Ba-alginate beads

Encapsulation method/material

In vitro: 135 min glucose perfusion experiments In vivo: 2, 4, 12, and 24 wks

120 min static glucose stimulation Perfusion glucose stimulation for 60 min

120 min static glucose stimulation

7 or 14 days in culture

Experiment duration time

Until graft failure (two consecutive blood glucose levels higher than 10 mm)

–—

–—

Device removal reasons

M.c.s recovered from normoglycemic animals: well-preserved islets; no ﬁbrotic overgrowth; m.c.s recovered from hyperglycemic animals: free of ﬁbrotic overgrowth, but contained unhealthy islets

Rat m.i.s responded to glucose static challenges at 7 and 14 days with a ﬁvefold increase in insulin secretion, substantially lower than the insulin output of free-ﬂoating islets Porcine m.i.s had poorer insulin secretion response to the glucose challenge, which worsened after 14 days M.i.s and alginate-coated islets had similar insulin response to static glucose challenges. M.i.s had a 2-min delay compared to controls, in perfusion experiments; alginate-coated islets had no measurable delay No observable difference of insulin secretion of free islets versus m.i.s in perifusion experiments, in vitro In vivo xenografts functioned for 17–180 days

Major outcomes/results

[88]

[86]

[85]

Ref.

STZ-induced diabetic Lewis rats i.p. transplant, seven immunized and nine nonimmunized

STZ-induced diabetic Lewis rats i.p. transplant, seven syngeneic, 10 allogeneic, four with empty capsules

6000 Wistar rat free islets and/or 3500 Wistar rat m.i.s

6000 Lewis and Wistar rat free islets or 3500 Lewis and Wistar rat m.i.s M.c.s Ø: 600–700 mm

m.c.s average Ø: 600 mm, 1 or 2 islets/m.c.

Ba-alginate beads

Ba-alginate beads

15 wks

1 month

Graft failure or the end of the experiment

Relapse into hyperglycemia or the end of the experiment

Four out of nine nonimmunized recipients of m.i.s remained normoglycemic for 1 month; the remaining ﬁve stayed normoglycemic for 28 days, or less Five out of seven preimmunized recipients were normoglycemic on day 28; nonrelapsed into hyperglycemia on days 7 and 10. Intact tissue was present in devices explanted from normoglycemic animals In failing animals, islets were destroyed or badly damaged and ﬁbrotic overgrowth was increased No differences between immunized and nonimmunized recipients Reversal of hyperglycemia in six of seven animals with syngeneic transplants for 15 wks Two out of 10 recipients of allogeneic transplants stayed normoglycemic for more than 100 days; remaining eight animals stayed normoglycemic for 0–56 days Negligible foreign body reaction to empty m.c.s Slightly higher reaction to syngeneic m.i.s than empty m.c.s Allogeneic normoglycemic transplanted animals: m.i.s free from ﬁbrotic overgrowth, with viable islets; relapsing hyperglycemic animals: m.i.s with severe ﬁbrosis, nonviable islets present

[89]

[90]

Recipient/ implantation site

Four groups of rats: 25 Balb/c mice treated or untreated with HOE 077, transplanted with empty m.c.s 10 Balb/c mice treated or untreated with HOE 077 transplanted with m.i.s, in the peritoneal cavity or the liver

STZ-induced diabetic NOD mice (seven with free islets, 12 with m.i.s) and Balb/c mice (three with free islets, 11 with m.i.s), i.p. transplant

Porcine islets (number of islets N/A)

9000–10 000 NOD or B6AF1 mice free IEs or m.i.s

(Continued )

Number of islets or islet cells

Tab. 9.3

49, 105, and 210 days, for m.i.s

1 month for empty m.c.s

Ba-alginate beads, highly puriﬁed high-G or high-M alginate

M.c.s Ø: 900–1100 mm

in vitro: static 24 h assays, on 2 consecutive days

In vivo: 4 wks

Ba-alginate beads

M.c.s Ø: 250–450 mm,

Experiment duration time

30–50 m.c.s implanted in the liver

Encapsulation method/material

Microspheres parameters

The end of the experiment, graft failure

The end of the experiment

Device removal reasons

Ba-alginate m.i.s allogratfs survived for >340 days and xenografts for >350 days Retrieved m.c.s free of capsular overgrowth and with viable islets Retrieved m.i.s had a 5–10-fold increase of glucose-stimulated insulin secretion (perfusion)

23% encapsulated islet viability, after 4 wks, in recipients treated with HOE 077, thinner ﬁbrotic overgrowth No difference detected between the HOE 077 treated and untreated groups after i.p. transplant Administration of HOE 077 signiﬁcantly reduced ﬁbrosis after intrahepatic transplant Percentage of retrieved empty m.c.s lower for high-G than high-M Ba-alginate

15 % encapsulated islet viability in untreated recipients, 4 wks after intrahepatic transplantation; high ﬁbrotic overgrowth

Major outcomes/results

[87]

[91]

Ref.

10 000 IEs of free or encapsulated NPCCs

STZ-induced diabetic B6AF1 mice, i.p. transplant, 32 with m.i.s

Same as in [87]

Ba-alginate beads, highly puriﬁed high-G or high-M alginate

2, 6, and 20 wks

The end of the experiment

Two-six out of 32 animals transplanted with microencapsulated NPCCs had normalized blood glucose levels for 20 wks M.c.s removed at 2, 6, and 20 wks still had increased insulin secretory responses to static glucose challenges High m.c. recovery: 91, 85, and 93% at 2, 6, and 20 wks, respectively 80 % of m.c.s free of overgrowth at 2 and 20 wks

[79]

300

9 Membranes in Bioartificial Pancreas

triggered by alloantigens that had been released out of the bead, suggesting the need for a reduction in pore size. Also, in failing animals most of the islets were destroyed or severely damaged and the ﬁbrotic and inﬂammatory reaction toward beads was markedly increased, when compared to transplanted animals still normoglycemic. No differences in inﬂammatory response were observed for immunized and nonimmunized animals (immunization will be further discussed in Section 9.2.4). The same group of researchers proceeded with syngeneic or allogeneic islets transplants, for longer periods [89]. Syngeneic transplanted microbeads were mostly recovered with little cellular or ﬁbrotic overgrowth and had mostly viable cells. The reaction to syngeneic implants was only slightly bigger to that of empty microcapsules. When allogeneic transplant was performed, only the animals that sustained normoglycemia for more than 100 days had microcapsules free from ﬁbrotic overgrowth and containing viable cells; those retrieved from recurring diabetic animals had severe ﬁbrotic overgrowth, nonviable islet remnants, and decreased number of microcapsules. Zhang et al. veriﬁed that a strong ﬁbrotic response occurred in vivo to Ba-alginate material and that the use of anti-ﬁbrotic agent HOE 077 (developed as an anti-ﬁbrotic agent for human liver ﬁbrosis therapy) could signiﬁcantly reduce it if encapsulated islets were xenotransplanted into adequate sites [91]. Empty capsules retrieved 4 weeks after implantation had a marked pericapsular inﬁltrate (PCI) surrounding the capsules. Administration of HOE 077 after intrahepatic transplantation resulted in a signiﬁcant reduction of PCI formation. However, treatment with this substance after intraperitoneal transplantation showed no inhibitory effect on ﬁbrotic overgrowth. HOE 077 exists in a much higher concentration in the liver and only inhibits inactivation of cells present on this organ, which explains this outcome. This proved that the administration of an anti-ﬁbrotic drug could be a different but feasible approach to improve islet viability and graft survival of microencapsulated islets. Two highly puriﬁed alginates (high-G or high-M) were used to encapsulate mice islets in Ba-alginate microcapsules and transplanted into the peritoneal cavity of NOD and Balb/c mice. The ability of such a microcapsule as an immunoprotective barrier against allorejection, was compared to crude SA beads coated with a PLL membrane, namely in the prevention of a ﬁbrotic response [87]. Overall, the Ba-alginate capsules made of highly puriﬁed alginate proved to be stable and highly biocompatible and protect against allorejection and autoimmunity in murine models, even without a PLL barrier. If such a barrier provides the same kind of immune protection in larger animal models is still unknown. Considering these positive results, these highly puriﬁed alginate capsules were used to determine their immunoprotective ability for neonatal pancreatic cell clusters (NPCC), when these are transplanted into STZ-induced diabetic B6AF1 mice [79]. Aside from graft survival or rejection times, the proliferation and maturation of NPCCs in the Ba-alginate capsules were also evaluated. In animals with successful grafts, maturation of the NPCCs signiﬁcantly increased with time. Thus, in agreement with a superior number of b-cells, the total insulin content per islet equivalent also increased in the NPCCs: 4-fold, 12-fold, and 10-fold, at 2, 6, and 20 weeks, respectively.

9.2 Bioartificial Pancreas

Recipient IgG was found throughout the NPCC-containing capsules at 6 and 20 weeks after transplantation, which conﬁrms that IgG diffuses through the gel without exerting any adverse effect on the encapsulated NPCCs. The authors could not fully explain why these simple capsules worked so well for xenografts, but they advanced two probable reasons for it: the avoidance of PLL, which might work as an adjuvant of xenograft reaction and the use of highly puriﬁed alginate, with very low endotoxin levels. This experiment showed that it is possible to successfully transplant encapsulated NPCCs into diabetic mice and that they can differentiate into b-cells in this environment, thus providing a large number of insulin-producing cells for transplantation. Comparison of APA With Ba-Alginate Although these materials have been extensively used to develop microcapsular BAPs, their comparison has been studied using macrocapsular devices [19] to avoid interferences due to a nonsmooth outer membrane surface, the existence of protruding islets, or the clumping in vivo of the microcapsules. Trivedi et al. performed and compared two different islets macroencapsulation protocols (protocol 1: islets are suspended in alginate, drops are gelliﬁed and coated with PLL and ﬁnally covered with alginate, as already described; protocol 2: islets are suspended in alginate and the droplets crosslinked with barium ions). Both kinds of macrocapsules (35 islets/each; 2–3 mm in diameter) were studied in vitro and in vivo by the syngeneic transplantation of encapsulated Lewis rat islets into the peritoneal cavity. Static glucose challenges performed with Ba-alginate macrocapsules, before graft transplantation, showed these had signiﬁcantly lower total insulin secretion in basal media, compared to free islets, and no signiﬁcant response on exposure to a high glucose challenge. STZ-induced diabetic Lewis rats syngeneically transplanted with protocol 1 or protocol 2 macrobeads became normoglycemic signiﬁcantly faster than animals with free islets transplanted under the kidney capsule. Regarding protocol 1 graft performance: after a meal challenge, no signiﬁcant difference was observed in the plasma glucose proﬁles between alginate-PLL macrobeads and kidney capsule recipients. The plasma insulin levels for the kidney capsule group displayed a sharp increase after a meal, unlike the macrobead group, for which plasma insulin remained at basal level. Grafts using protocol 2 macrobeads showed divergent behavior: after an overnight fast, animals with Ba-alginate macrobeads had signiﬁcantly lower plasma glucose levels than kidney-capsule grafted animals. Also, their plasma insulin levels increased signiﬁcantly compared to basal insulin levels in response to a meal at 100 and 120 min. During the intravenous infusion tests, both rat groups, protocol 1 and protocol 2 macrobeads, had no measurable increases in plasma insulin, whereas in control rats (normal rats or diabetic rats with free-islet grafts) there were clear insulin responses. Islet retrievability after 5 weeks of transplantation was larger for Ba-alginate (80–90 % intact) than for APA macrocapsules (50% broken) indicating lesser

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mechanical stability of the materials in the latter. Retrieved capsules from protocol 1 had mostly viable islets but with some evidence of central necrosis, unlike those of protocol 2 for which islets had a healthy appearance. Several reasons were advanced to explain the poor insulin secretion dynamics observed. It may, in part, be due to the implantation site, as was shown by other researchers, although the authors thought that diffusion limitations were responsible for the absence of acute insulin responses. This directly relates to the large size of the macrobeads, which unquestionably induced time lags in glucose and insulin diffusion. The residual negative charge on alginate may also play a role in binding the insulin to the gel, which would slow the dynamic response to glucose stimulation. Agarose Beads Without Immunoprotecting Membrane Since a synergistic effect seemed to occur between the agarose gel, PLL membrane, and the immune response reaction [67,79], instead of conjugating alginate bead formation and PLL membrane or other polymeric membrane coating, some authors tried to use simple agarose beads to microencapsulate islets and tested their performance in vivo. The immunoprotective efﬁciency of agarose capsules was tested by Iwata et al. [92]. C57BL/6 mice islets were microencapsulated with agarose and implanted in STZBalb/c and NOD mice. Balb/c mice remained lifelong normoglycemic and four of ﬁve NOD mice had functional graft survival for more than 80 days. It was therefore determined that a capsule made of this material only can effectively immunoprotect islets and prolong graft functioning in these recipient mice, without any immunosuppression. Comparative experimental microcapsular geometry studies with agarose gel constructed BAP were also performed [93]. Three geometries were studied: microbead, rod, and disc. The devices were in vitro and in vivo analyzed. In vitro examination showed the islets survived well and released basal insulin regardless of the geometry of the device used. In vivo, ﬁve of the eight rods and ﬁve of the six disc recipients could not function for longer than 15 days, although low dosages of 15deoxyspergualin, an immunosuppressive drug, were used. This occurred even though no difference in ﬁbrotic response could be observed between the three kinds of devices. The best graft survival was obtained with microcapsules, which was attributed to their smaller diameters and lower diffusion times. These same capsules were used for allotransplant of microencapsulated islets in dogs [94]. Implantation of microencapsulated islets completely supplanted exogenous insulin therapy in three out of the ﬁve recipients for 28, 42, and 49 days. Good fasting glucose control was achieved but response to an intravenous glucose tolerance test remained abnormal. The histologic appearance of retrieved allografts revealed viable islets and no evidence of adherence or inﬁltration of immunocytes or inﬂammatory cells. Thus, although with limited success, they showed it is possible to use these microcapsules to reverse IDDM in a larger animal model. Agarose hydrogel microcapsules have also been used to relate graft survival time with the amount of islets encapsulated. Ao et al. [95] compared the performance of microcapsule devices, with different amounts of islets seeded, to the performance

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of free islets (xenogeneic mongrel dog islets transplant into alloxan-induced diabetic nude mice). Long-term graft survival occurred when 500 encapsulated islet equivalents (IE) or more were implanted; best results were obtained with 1000 IE (all mice transplanted had more than 119 days of graft survival). Two thousand free islets implanted in the kidney capsule, or more than 4000 islets intraperitoneally, were needed to reverse diabetes. These results have, however, a limited scope since diabetic nude mice are immuno-incompetent animals. The above-mentioned results are modest, when compared, for example, to those of [68], which used similar experimental parameters but encapsulated islets in APA microcapsules. These authors had better graft survival times, both with dog islet allografts and xenografts, and the microcapsules were usually free of ﬁbrotic overgrowth (except if islets were exposed or protruded from the capsule) (see Table 9.2). Comparing both results, the number of islets transplanted seems to be crucial to obtain long-term graft survival and ﬁbrotic overgrowth appears to occur not due to the presence or absence of the PLL membrane or to the alginate gel used but rather due to an efﬁcient coating of the islets, thus preventing direct exposure of islet tissue to the host immune cells. Agarose-PSSa Microcapsules A mixture of agarose crosslinked with polystyrene sulfonic acid (PSSa) coated with a polyanionic carboxymethyl cellulose (CMC) membrane material was also used for the microencapsulation of pancreatic islets or islet cells. The ﬁrst time this kind of microcapsule was used, a hamster-to-mouse xenotransplantation was performed for material biocompatibility assessment [96]. The longest survival time achieved was 90 days, and three out of ﬁve NOD recipients showed prolonged normoglycemia. Biocompatibility experiments proceeded with xenotransplantation of cells from a b-cell line, MIN6, to the subcutaneous space of STZ-induced diabetic Lewis rats [97]; the MIN6 cell line seems to retain the characteristics of glucose metabolism and GSIS similar to those of normal pancreatic islets cells. Microspheres were prepared with agarose and PSSa as before to yield droplets of 300–500 mm in diameter, and these were covered with hexadimethrine bromide to form a polyanionic complex, preventing PSSa from leaking, and further coated with CMC. Two thousand microcapsules were formed, seeded with MIN6 cells and implanted in the subcutaneous site for a week, recovered and, afterwards, subjected to in vitro static glucose challenges. Retrieved agarose-PSSa microbeads were intact, spherical, and smooth. Cells were found viable inside the capsules, and adhesions or ﬁbrotic overgrowth were rarely found in the subcutaneous site. The in vitro static glucose challenges showed that the encapsulated cells had a 2.7-fold increase of insulin secretion relative to basal secretion levels, which conﬁrmed cell functionality. The studies evidenced the microcapsules biocompatibility and immunoprotective properties. Microspheres alone were rarely used for implantation in the subcutaneous site. However, tubular capsules together with b-FGF impregnated microspheres, both made of PSSa, were tested in [48], as was discussed in Section Neovascularization of

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Macrocapsules. Average graft survival time was 68.4 25.6 days, also due to neovascularization induced by the presence of the b-FGF impregnated beads. These results conﬁrmed and improved the ones from Ref. [97]. The biocompatibility of the material produced by this crosslinking method was thus conﬁrmed, since no ﬁbrotic overgrowth or pericapsular cellular inﬁltrate were observed either with b-FGF+ and b-FGF– groups, even though islets retrieved from the second group were found necrotic. Nevertheless, graft survival time was reduced, when compared to similar xenograft experiments with Ba-alginate beads [89]. Regarding the success in long-term reversal of diabetes with this type of BAP, PSSa capsules were used in Hamster-to-rat xenotransplantation studies, this time to determine the amount of islets needed to reverse diabetes in this model [98]. Long-term graft survival was only achieved in recipients with 3000 encapsulated islets, the mean period of normoglycemia being 124 days. Encapsulated islets, recovered from recipients when they returned to hyperglycemia, had mostly retained their morphologic appearance but lost functionality. A ﬁbrous layer surrounded the surface of the microbeads and was probably the primary reason for graft failure. In spite of this, the authors believed these microcapsules could be used to reverse diabetes in higher mammals, with a more complex immune system. The ability of the agarose-PSSa microspheres to prevent porcine islet degradation had been determined in vitro [94]. The encapsulated porcine islets were kept morphologically intact, viable, and functional during 7 days in culture, whereas the nonencapsulated ones were rapidly disintegrated. An extensive study was then conducted to examine the use of agarose-PSSa microcapsules in xenogeneic transplantation of islets in higher animal models and determine if their immunoprotective ability is as good as that of smaller animals [99]. In the larger animal discordant xenotransplant, transplantation of pig islets into pancreatectomized diabetic beagle dogs, two out of ﬁve animals became normoglycemic without exogenous insulin for 50 and 119 days [99]. Agarose-PSSa microcapsules could protect the xenogeneic islets from humoral attack by the host’s immune system. The limited success was attributed especially to the difﬁculty of isolating and handling pig islets. Pancreatic pig-to-dog xenografts with microencapsulated islets were not attempted again. These results are slightly better than those of Ref. [14], without the problem of device occlusion, which occurred with intravascular devices. Even better results than those presented in Ref. [99] were obtained by Ref. [40], which achieved xenograft survival for more than 8 months. However administration of a low dosage of exogenous insulin was necessary, which might have prevented an earlier recurrence of hyperglycemia. Islets or Islet Cells Covered With a Siliceous Membrane Over the last 7 years, silanization of islets (reaction of silanes with exposed hydroxides on the cell surface) has been proposed as a method of microencapsulation. The potential for long-term survival of viable and functional islet grafts was evaluated [100–104].

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The permeability of two sol–gel synthesized silicate membranes was assessed [102]: Si(OCH3)4 (tetramethoxysilane, TMOS) and (CH3O)3Si(CH2)3NH2 (3-aminopropyltrimethoxysilane, APTrMOS). Microcapsules were prepared in a similar manner as APA capsules, with Ca- or Ba-alginate, although the siliceous substances were used instead of PLL, and their permeabilities to glucose, myoglobin, ovalbumin, bovine serum albulin (BSA), and g-globulin were determined. BSA and g-globulin could adsorb onto the silicate derived from TMOS alone. Only the membrane derived from APTrMOS allowed g-globulin (the biggest molecule used in this study) to permeate, even with high concentrations of this siliceous substance. Permeability assays were performed with Ca-and Ba-alginate capsules coated with a mixture of silanes (best ratio APTrMOS/TMOS ¼ 2.40). For this ratio of silanes, coated Ba-alginate beads had good permeability to g-globulin, but coated Ca-alginate beads could reject g-globulin. This membrane was found to have a MWCO of approximately 60 kDa. These changes in permeability could be explained by the change of electric charge of substances used during capsule formation and electrostatic interaction of Ba and Ca ions with amino and carboxyl groups. The in vitro insulin secretion indexes of encapsulated rat islets were shown to be dependent on the molar ratios of the silanes [103]. For APTrMOS/TMOS ¼ 0.6, onefourth of the ratio for optimal permeability, the ability of islets to secrete insulin upon a glucose stimulus was hampered, while for ratios of 1.2 and 2.4 stimulation was similar to that of free islets. In vivo xenotransplantation studies (encapsulated Wistar rat islets peritoneally inserted into STZ-induced diabetic DDY mice; 1000 islets/recipient) demonstrated marked cellular overgrowth just after the 21st days of transplant. That led to islet necrosis and recurrence of hyperglycemia. Some mice (6/10) were able to maintain normoglycemia levels up to 2–3 months after transplantation. The cause of overgrowth, although not due to the chemistry of the membrane, had not yet been deﬁned [104]. Desai et al. analyzed, in vitro, the behavior of islet cells when in contact with the microfabricated immunoprotecting membrane, using two microfabricated half capsules bound together with silicon elastomer adhesive or when direct deposition of a thin ﬁlm of siliceous material on the cell’s membrane surface was performed [100]. Preliminary biocompatibility studies had already been performed with silicon membranes, which suggested that silicon substrates were well tolerated and nontoxic in vitro and in vivo [24]. Using a microfabrication technology (combination of UV-lithography, silicon thinﬁlm deposition and selective etching of sacriﬁcial oxide layers) a homogeneous siliceous membrane was formed. This technology allows thickness and pore size uniform distribution to be controlled, while maintaining original cell morphology. The microfabricated biocapsule had 78 nm pore-sized membranes. Rat islets were placed in the microfabricated capsules made of boron-doped silicon and polysilicon (Figure 9.4), cultured, and their viability and functionality assessed. The insulin secretion response obtained with encapsulated and free islets was similar, which validated the pore density and length chosen when constructing the microcapsules, and showed these biocapsules provide a stable functional environment for the islets.

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Fig. 9.4 Diagram of basic microfabricated immunoisolation biocapsule concept. Source: Reproduced from Desai et al. [24] with permission from Elsevier, copyright 2000.

Insulin secretion levels still decreased after 1 month of culture, though. After 4 weeks islets removed from the capsules showed very little central and no peripheral necrosis and were attached to the inorganic substrate. In vitro and in vivo assays were then performed with micromachined biocapsules of 18 or 66 nm pore-sized membranes and insulinoma cells (RIN and bTCC6F7) [24]. The in vitro assays showed both kinds of insulinoma cells had higher stimulatory indices (stimulatory/basal insulin secretion) with 66 nm than with 18 nm membranes, even though the latter had higher immunoprotection abilities. In vivo, microfabricated capsules with RIN and bTCC6F7 cells remained mechanically intact until removal (8 days postimplantation). Both types of transplanted microencapsulated cells remained viable and had higher stimulatory indices with 66 nm than 18 nm pore-sized membranes in vivo, indicating that the pore size greatly affected the secretory response of the cells. However, in vitro experiments with recovered capsules, showed a sharp reduction in the stimulatory indices of both cell lines with 66 nm biocapsules, but not with 18 nm biocapsules, when compared to those of nonimplanted microencapsulated islets. This suggests a good immunoisolation effect of the 18 nm pore-sized biocapsules, which was lacking with the 66 nm biocapsules. Although no long-term in vivo experiments were performed so far with these micromachined biocapsules, they remain a very attractive alternative, since microfabrication technology allows the precise control of pore size, conﬁguration, and distribution, and thus protect encapsulated islets from the host immunologic response. Recently, another research group produced the microcapsules by direct deposition of a siliceous material, after suspending the islets in a ﬁlter holder and ﬂuxing them with air saturated by CH3SiH(OEt)2 and Si(OEt)4 (50/50 molar ratio). The siliceous membrane formed had at least 0.1-mm thickness (BioSil membrane). Capsule performance was evaluated in vitro and in vivo [101].

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This encapsulation procedure allowed the islets to preserve their original dimensions and viability (about 90 %). The membrane formed was continuous and homogeneously distributed over the islets’ surface. Also, perifusion experiments indicated that islet function was not reduced by the encapsulation method. The siliceous envelope did not reduce the insulin diffusion rate when compared to free islets. Islets from inbred Lewis rats were encapsulated with the siliceous membrane and transplanted into Lewis or Sprague–Dawley (SD) rats, under the left kidney capsule, and normoglycemia was restored. The implant promoted a faster return to normal glucose levels and a prolonged effect in glycemic control than in animals transplanted with free islets. Longer graft survival time was observed for encapsulated islets, both with syngeneic (6 out of 10 rats remained normoglycemic for at least 70 days) and allogeneic transplants (six out of 10 rats remained normoglycemic for at least 55 days). Marked improvement was observed with encapsulated allografts, both in terms of prolonged graft survival and faster decrease in glycemia levels. This clearly demonstrated that the siliceous membrane is well tolerated by the living organism, allows proper feeding of islets, provides immunological protection, and allows a fast response to hyperglycemia episodes, thus improving the long-term effect of the islets grafts. In view of these favorable results, the testing of these microcapsules in a larger animal model with more complex immunological defenses and more important diffusional limitations is needed. Multicomponent Microcapsules Containing Poly-Methylene-co-Guanidine At the Vanderbilt University, Nashville, USA, Dr Taylor Wang and co-workers developed a method for SA-CS/PMCG microcapsules fabrication in which a combination of polyanions, SA, cellulose sulfate (CS), and the polycation poly(methylene-coguanidine) (PMCG) were used, together with a speciﬁc encapsulation technology, resulting in polyelectrolyte complexation. The authors claim that capsule parameters, like size, wall thickness, mechanical strength, permeability, and surface characteristics, can be controlled independently [105]. Animal trials were conducted after characterization of microcapsules (morphology, strength, molecular exclusion limits, and biocompatibility) [106] with a rat-to-mouse intraperitoneal transplantation model. SD rat islets (1000/recipient) were encapsulated in 800 mm microcapsules (100 mm wall thickness; 230 kDa of exclusion limit) and transplanted into STZ-induced diabetic C57/ BL6 mice and into spontaneously diabetic mice. The authors followed up transplantation in both systems, for about 6 and 3 months, respectively, before recurrence of hyperglycemia. Microcapsules retrieved after 1 year from C57/BL6 transplanted animals were shown individualized and absent of ﬁbrosis, and the encapsulated islets still retained signiﬁcant insulin response to glucose stimulus. In contrast, microcapsules inserted in NOD mice were referred to as clumped and surrounded by ﬁbrosis; the failure possibly occurred due to immune or autoimmune attack.

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9.2.4 Inﬂuence of Recipients Sensitization to Donor Antigens in Graft Survival

Immunoprotecting barriers are far from being absolute. BAP devices may sensitize diabetic recipient patients to donor antigen, because seeded islet tissue can still trigger an immune response [39]. Xenogeneic humoral response can occur with extravascular immunoisolation devices, even with a membrane that prevents passage of large molecular weight substances. Due to the release of donor proteins by the islets, which leak through the membrane, the host indirect immune response system is activated. The number of islets implanted seems to condition this humoral response, by causing an overload of the system with immune complexes. The islet tissue mass needed to reverse diabetes may be signiﬁcantly less when using other types of BAP. Microcapsules, narrower-bore membrane chambers, or vascular devices may evoke a smaller xenogeneic humoral response. In order to verify the hypotheses, Lanza et al. tried a different BAP design [50]. Injectable membrane small capsules made from biodegradable materials (MWCO < 150 kDa) seeded with bovine calve islets were introduced in STZinduced diabetic rats and mongrel dogs. Several weeks after postimplantation, the injected capsules were recovered from the host animals. They still contained multiple viable and functional islets, mostly free of ﬁbrotic overgrowth, with sufﬁcient insulin response to high glucose levels (one to threefold and four to sixfold increase above basal levels, in rats and dogs, respectively). The authors successfully proved that discordant islet xenograft survival can be achieved in rodents and dogs, without immunosuppression, with this encapsulation technology, thus establishing its clinical utility. 9.2.5 Islet Oxygenation Studies

A crucial aspect in the prevention of islet loss of viability and functionality is to have proper feeding and oxygenation conditions [107–112]. Schrezenmeir et al. performed extensive studies of islet oxygenation [109]. They used both piscine islet organs, called Brockmann bodies (BBs) from Osphronemus gorami and Lewis rat islets, together with a special recessed-type microelectrode to determine oxygenation conditions around and inside the islets in culture, correlating islet functionality with oxygen distribution throughout the islets. The islets were encapsulated in smooth surface 12 kDa MWCO Thomapor hollow ﬁbers (regenerated cellulose, 2.2 and 1.5 mm outer and inner diameter, respectively) [63]. Free and encapsulated rat islets were cultured in vitro for 34 days, and their functionality determined for 5 and 34 days. The secretion of insulin by free islets under stimulatory conditions at the beginning and end of the culturing period was signiﬁcantly higher than the basal values, and increased signiﬁcantly during the observation interval. The encapsulated islets also had increased insulin release under stimulatory conditions, at the beginning, but considerable loss of viability was observed at the end of the culture period. A typical ﬁnding in these cases was central

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necrosis. This clearly demonstrated the burden introduced by additional unstirred water layers through the membrane. In vivo tests of graft survival and functionality proceeded with the implantation of insert peritoneally encapsulated BBs and rat islets in adult male Wistar rats. Histological ﬁndings conﬁrmed that islet function was best preserved if the immunoseparating membrane remained free from surrounding connective tissue. Finally, in vitro tests with individualized BBs were performed to determine oxygen distribution proﬁles. BBs were cultured in vitro, in unstirred aerated medium (pO2 tension in medium was 140 mmHg), and the oxygen tensions were measured. Oxygen tension inside the islet organs declined with organ depth from the pressure of the surrounding medium following a sigmoidal curve. The decline began about 300 mm outside the organ and attained values close to 0 mmHg in the vicinity of the organ center, at 300–400 mm from the organ’s surface. The oxygen tension fell to 30 % of the oxygen saturated surroundings at 64 mm inside the organ tissue with basal glucose concentrations and at half this distance with stimulatory glucose concentrations. Also, the depths for which the oxygen tension fell to 10 % of initial oxygen value were 176 24 mm with basal glucose concentrations and 128 21 mm with stimulatory ones. Given these results, three main aspects were made clear: 1. The diameter of unstirred water layers should be reduced to a minimum, and the outer limit of the device should not exceed 300 mm; preventing ﬁbrous tissue formation or inducing vascularization was later proven to produce better results [64]; 2. Energy-consuming tissue layers should be reduced to unavoidable dimensions; 3. Euglycemia improves the chances of survival and function of the graft, by preventing unnecessary energy oxygen expenditure (since O2 is needed for the degradation of glucose). To clarify the relation between oxygen supply to the islets and their viability and functionality and also to assess the critical tissue volumes and critical pO2 tensions for intact islet function, other studies followed [113]. Three groups of nonencapsulated BBs were established, identical in mass of viable cells: smaller organs (diameter 680 mm), larger organs (diameter 1120 mm), and larger organs segmented to produce approximately three equal parts (each organ produced three segments of diameter 805 mm). The organs were kept in culture for several days. On days 4/5 and 18/19, insulin secretion was tested for all the BBs at basal and stimulation glucose levels. No signiﬁcant difference was observed 4 days after isolation. Two weeks later, however, the values of insulin secretion for large organs were substantially lower than those of small or segmented organs, both under basal and stimulated conditions. This conﬁrmed the previous hypotheses that increased energy expenditure and oxygen consumption for insulin secretion increase the decline of oxygen tension to the center of the BB.

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Comparing the results of smaller islet organs with those of larger islet organs showed pO2 tension had a tendency to be higher in the smaller ones at the same deﬁned distances, but the difference is not very signiﬁcant. A shift of sigmoidal curve to minor pO2 values occurred when higher glucose concentrations were used. In larger organs central necrosis inevitably occurs, since pO2 values reach a point where function is hindered. The calculated critical value for these organs is between 8 and 16 mmHg. This is lower than but close to the values reported for rat islets (12 or 20 mmHg [107,108]), which indicates that BBs are less sensitive to low pO2 than rat islets or that the different adopted methodologies in each case may have contributed to the discrepancy of results. Aside from static culture of BBs for a prolonged period of time, freshly isolated BBs were cultured also, for 14–16 h only, for pO2 levels measurement. The pO2 proﬁles had a sigmoidal shape and a decline in oxygen tension beginning in the medium surrounding the BB, as before. Oxygen tensions at 50 mm outside the organ were consistently and signiﬁcantly lower in stimulating than in basal level of glucose concentrations. Based on the data assessed, the authors assumed that, under normoglycemic conditions and high oxygen tension, the outer rim of piscine islets that are sufﬁciently supplied with oxygen is approximately 200 mm in depth. This distance will be reduced if glucose concentration is elevated and oxygen tension in the surrounding milieu is lower than the one used in this study. It also seems to be smaller in mammalian islets. Schrezenmeir et al. analyzed the effect of coimmobilizing the islets with hemoglobin to facilitate oxygen diffusion and obtain better islet survival, preventing hypoxia and islet central necrosis [110]. Capillary 20 kDa MWCO Cuprophan Type 11/600 membranes were used to encapsulate the islets, suspended in sodium alginate gel. The devices were implanted in Wistar rats and several in vitro and in vivo experiments were performed, with and without the alginate matrix and/or hemoglobin. The best results in vitro were obtained with the islets suspended and hemoglobin molecules immobilized in the alginate matrix. By the time the assay ended (34 days), about 1/3 of the islets in the membrane were still 100 % viable (cells with intact chromatine structure). This membrane-gel matrix system was then tested in vivo. Very poor results were obtained without hemoglobin, and the improvement with coimmobilized hemoglobin molecules was not substantial (by the end of the assay only approximately 2 % of the cells were intact). Even so, the authors concluded that the permanent presence of hemoglobin might improve function and viability of the immunoisolated islets. However, they could not distinguish to which hemoglobin properties this was due, oxygen donation or nitric oxide/oxygen radical scavenging. A recent study performed by this group, with piscine islets encapsulated in microspheres, allowed to clarify the function of hemoglobin in the improvement of oxygen distribution to the islets and their functionality [112]. The results obtained, which will be discussed in further detail when microsphere devices are mentioned,

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still did not completely clarify the mechanism responsible for longer graft survival, when islets are coimmobilized with hemoglobin. 9.2.5.1 Oxygenation of Ba-Alginate Microencapsulated Islets Schrezenmeir et al., who had already performed some studies with piscine islets, free or encapsulated in macrocapsular extravascular devices, to determine the oxygen distribution in the islet and its inﬂuence in graft survival, continued their studies using Ba-alginate microcapsules. They were the ﬁrst to study the inﬂuence of convection and microencapsulation on oxygen distribution in islet organs encapsulated in Ba-alginate microcapsules. The oxygen measurement was performed by means of polarography using oxygensensitive microelectrodes [111]. When free BBs were cultured with and without convection, signiﬁcant differences were observed in pO2 values. With convection, the values outside and inside of BBs were considerably higher. Without convection, the oxygen-depleted zone outside the BBs was larger, the diameter of the oxygen-consuming rim was smaller, and the oxygenation values in the BBs center were reduced to 6 % of the initial value. Better results were obtained, in every case, with convection. Microencapsulation eradicates the effect of direct convection on the BBs. The authors tried to clarify if encapsulation exerts a similar effect as lack of convection and the way in which it inﬂuences oxygen distribution. The oxygen proﬁles of nonencapsulated BBs without convection corresponded well to those of microencapsulated BBs; all electrode positions inside and outside the BBs showed signiﬁcantly lower pO2 values when compared to nonencapsulated BBs. According to these results, the encapsulation has primarily an effect on the extension of the oxygen-depleted zone outside the organ. Thinning down the alginate layer, and thus achieving better oxygenation, could reduce the depth of this zone. Compared to single mammalian islets, piscine islet organs have a much higher cell mass, so a smaller decrease of pO2 in mammalian islets could be expected. However, a similar cell mass to that of BBs can be attained when mammalian islets are packed in a BAP. To avoid these problems, the authors suggested a packing mechanism using alginate immobilization for increased spacing between islets or singlecell suspensions and choosing appropriate transplantation sites, with increased convection. They also showed, in an additional study, that islets from different animal models can have different oxygen requirements and, therefore, different hypoxia-tolerance capabilities [114]. For example, rat islet cells cultured in 38 mmHg tolerate hypoxia better than porcine islets, and the neonatal form is more hypoxia tolerant than the adult, which means the choice of the islet donor species can inﬂuence the outcome of xenotransplantation. As a continuation to another study already performed with hollow ﬁbers, a study was performed to clarify whether hemoglobin may improve oxygen supply within the microcapsules by increasing its oxygen solubility [112]. In the previous study the presence of coimmobilized hemoglobin deﬁnitely improved encapsulated islet organs survival but the mechanism responsible for this was not made clear.

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In this study, the islets were cultured free or encapsulated, with or without hemoglobin, and functional tests were performed, in vitro. Naked BBs had considerably higher pO2 outside and inside the islet tissue than BBs encapsulated without any hemoglobin and these had higher pO2 at any electrode position than islet organs encapsulated with hemoglobin. However, on the second and third days of culture, BBs of all groups showed normal function with a signiﬁcant increase of insulin release as a response to a glucose stimulus, when compared to basal insulin secretion levels. After 4 weeks both basal and stimulated insulin secretion had decreased in all groups, particularly the stimulated secretion. Viability after 4 weeks of culture was similar for naked BBs and BBs coencapsulated with hemoglobin (approximately 65 %) but lower for BBs encapsulated without hemoglobin (approximately 55 %). In the previously mentioned report, it had already been shown that microencapsulation reduced oxygen availability and tension in the islet tissue. In this study this was conﬁrmed. Also, the addition of hemoglobin to the capsule matrix signiﬁcantly improved the viability following long-term culture. The oxygenconsuming rim tended to be higher in BBs encapsulated with hemoglobin compared to BBs encapsulated without hemoglobin. The oxygen plateau reﬂecting central necrosis tended to be lower with hemoglobin in the matrix. The improved survival of islet tissue in BBs encapsulated with hemoglobin does not seem to be a result of increased oxygen supply, though. Other factors must be considered, since the introduction of hemoglobin into the capsule tended to reduce the oxygen tension at the organ surface and signiﬁcantly decrease oxygen tension at the organ center. Thus, the mechanism by which the beneﬁcial effect of hemoglobin on islet cell survival is mediated remains unclear, but it is deﬁnitely not due to improved oxygen availability. 9.3 Final Comments

From this review of the developments toward a BAP device, major aspects are evidenced: i. None of the devices developed so far can bring ﬁbrotic response to acceptable negligible levels; ii. Diffusional limitations imposed by the encapsulating material and the capsule size are crucial for graft success. The ﬁrst aspect remains as one of the main reasons for graft failure and limited graft survival. Regarding the second aspect, although some authors achieved good results with microcapsules, graft survival time and response time to a glucose stimulus still have to be improved. The use of new biomaterials and/or of coimmobilization of other molecules, cells or cell aggregates with the islets, are two possible strategies to overcome this problem. Also, device design other than the ‘‘classical’’ macro- and microcapsular design has been attempted with some degree of success and may eventually lead to better results in the future.

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9.4 Acknowledgment

Ana Isabel Silva wants to acknowledge the PhD grant attributed to her by the Portuguese Government through FCT Fundac¸a˜o para a Cieˆncia a Tecnologia SFRH/BD/10594/2002. References 1 Colton, C. K. and Avgoustiniatos, E.

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Kobayashi, T., Urizono, Y., Yamada, T., Shibaji, T., Kanamura, T., Ogawa, S., Iwata, H., Nakano, H. (2000) Efﬁcacy of agarose/polystyrene sulfonic acid microencapsulation for islet xenotransplantation. Transplantation Proceedings, 32, 1071– 1072. Desai, T. A., Chu, W. H., Tu, J. K., Beattie, G. M., Hayek, A., Ferrari, M. (1998) Microfabricated immunoisolating biocapsules. Biotechnology and Bioengineering, 57, 118–120. Boninsegna, S., Bosetti, P., Carturan, G., Dellagiacoma, G., Dal Monte, R., Rossi, M. (2003) Encapsulation of individual pancreatic islets by sol–gel SiO2 : a novel procedure for perspective cellular grafts. Journal of Biotechnology, 100, 277–286. Sakai, S., Ono, T., Ijima, H., Kawakami, K. (2001) Newly developed aminopropyl-silicate immunoisolation membrane for a microcapsule-shaped bioartiﬁcial pancreas. Annals of the New York Academy of Sciences, 944, 277–283. Sakai, S., Ono, T., Ijima, H., Kawakami, K. (2002) Alginate/ aminopropyl-silicate/alginate membrane immunoisolatability and insulin secretion of encapsulated islets. Biotechnology Progress, 18, 401–403. Sakai, S., Ono, T., Ijima, H., Kawakami, K. (2002) In vitro and in vivo evaluation of alginate/sol–gel synthesized aminopropyl-silicate/ alginate membrane for bioartiﬁcial pancreas. Biomaterials, 23, 4177– 4183. Lacı´k, I., Brissova´, M., Anilkumar, A. V., Powers, A. C., Wang, T. (1998) New capsule with tailored properties for the encapsulation of living cells. Journal of Biomedical Materials Research, 39, 52–60. Wang, T., Lacı´k, I., Brissova´, M., Anilkumar, A. V., Prokop, A., Hunkeler, D., Green, R., Shahrokhi, K., Powers, A. C. (1997) An encapsulation system for the

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immunoisolation of pancreatic islets. Nature Biotechnology, 15, 358–362. Dionne, K. E., Colton, C. K., Yarmush, M. L. (1989) Effect of oxygen on isolated pancreatic tissue. ASAIO Transactions, 35, 739–741. Dionne, K. E., Colton, C. K., Yarmush, M. L. (1993) Effect of hypoxia on insulin secretion by isolated rat and canine islets of Langerhans. Diabetes, 42, 12–21. Schrezenmeir, J., Laue, C., Sternheim, E. T., Wolbert K.,Darquy S.,Chicheportiche D., Kirchgessner J., Reach G. (1992) Long-term function of single-cell preparations of piscine principal islets in hollow ﬁbers. Transplantation Proceedings, 24, 2941– 2945. Schrezenmeir, J., Velten, F., Beyer, J. (1994) Immobilized hemoglobin improves islet function and viability in the bioartiﬁcial pancreas in vitro and in vivo. Transplantation Proceeding, 26, 792–800. Schrezenmeir, J., Kirchgessner, J., Gero, L., Kunz, L. A., Beyer, J., Mueller-Klieser, W. (1994) Effect of microencapsulation on oxygen

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distribution in islets organs. Transplantations, 57, 1308–1314. Schrezenmeir, J., Hyder, A., Vreden, M., Laue, C., Mueller-Klieser, W. (2001) Oxygen proﬁle of microencapsulated islets: effect of immobilised hemoglobin in the alginate matrix. Transplantation Proceedings, 33, 3511–3516. Schrezenmeir, J., Gero, L., Solhdju, M., Kirchgessner, J., Laue, C., Beyer, J., Stier, H., Muller-Klieser, W. (1994) Relation between secretory function and oxygen supply in isolated islet organs. Transplantation Proceedings, 26, 809–813. Hyder, A., Laue, C., Schrezenmeir, J. (1998) Variable responses of islet cells of different ages and species to hypoxia. Transplantation Proceedings, 30, 578–580. Lembert, N., Petersen, P., Wesche, J., Zschocke, P., Enderle, A., Doser, M., Planck, H., Becker, H. D., Ammon, H. P. (2001) In vitro test of new biomaterials for the development of a bioartiﬁcial pancreas. Annals of the New York Academy of Sciences, 944, 271–276.

321

Index a Abel, John 2 ACCUREL process 59, 82 active electrodes 202 active systems 212 active TDD Systems 216 acute phase reaction 37 adhesion 228, 241 advanced glycation end products 34 afﬁnity chemistry 146 air diffusion tests 120 alginate 280, 300 –chemical composition 281 –effect of puriﬁcation 291, 300 –guluronic and mannuronic acid content 290 –inﬂammation induction 290 Althane 9 Althin Medical AB 9 ALZA 175 AM50-Bio 29 AM-SD18M 31 AN 69ST 13, 37 AN69 acrylonitrile and sodium methallyl sulfonate copolymer 13, 27f, 30ff, 274 anaphylactoid reactions 26 angiotensin converting enzyme 27 anion-exchange membrane adsorbers 149 APA (alginate-PLL-alginate) overview 282 APMA 234f apoptosis 31f applications in drug delivery 184 artiﬁcial lungs 49 Arylane 16, 19 arylether 16 Asahi AM-UP-75 37 Asahi Medical 8, 11f, 16, 34

Asahi PAN 27 asymmetric membranes 97, 118 autoclaved assemblies 112

b B. diminuta 105 B. revundimonas diminuta 101 back pressure 116 bacteriophage 120 BAP, (bioartiﬁcial Pancreas) 266 –classiﬁcation 266 –implantation sites 266 Baxter Healthcare Corporation 9 Bemberg 4 benzyl-modiﬁed cellulose 11 biocompatibility 24, 267, 273, 276, 281, 290, 295, 303 –assessment 272, 295, 303 –ﬁbrotic overgrowth 267, 279, 290f –foreign body reaction 267 –macrocapsules 273f –membrane surface structure 276 –purity of alginate solution 281 –requirements 267 –thrombosis 267, 268 bind-elute 147 binding capacity 147 bioartiﬁcial organs 264 –Immunoprotection 266 biodegradable membranes 187f bioelectrochemistry 171 biohybrid organ 242ff, 249, 253 biological membrane 191 biotransformation 162 BK series U/P/F 14 blood 75 –speciﬁc features during application 75 blood plasma treatment 87 –medical indications 87 blood fractionation 69

Membranes for the Life Sciences. Edited by Klaus-Viktor Peinemann and Suzana Pereira Nunes Copyright ß 2007 WILEY-VCH Verlag GmbH & Co. KGaA. All rights reserved ISBN: 978-3-527-31480-5

322

Index blood oxygenation 49 bone 236ff bradykinin 26ff, 37, 38 bubble point equation 112 bubble point 96, 101 bubble-free gassing 158 buffers 103

c cake model 102 calcitonin 203 capacity 94, 104, 139 –trials 107 capillary membrane 84 –velocity proﬁle 84 capillary membrane plasma separation ﬁlter 86 capsules 110 cardiodiapulmonary bypass (CPB) 64 cartridges 110 cassettes 116 categories for hemodialysis membranes 22 caustic stability 98 cell 103 –culture media 107 –culture 103 –debris 103 –harvest 107 –lysate 103 cell activation 31 Cellophane 3 cellulose 5 cellulose acetate (CA) 9 cellulose acetate dialyzers 29 cellulose diacetate 5 cellulose triacetate 10 charge of the drug 200 chemical compatibility 104 chromatography columns 103 clariﬁcation 103 clearances 25 cofactor regeneration 166 collodium tubes 3 colony forming units 100 compatibility 104, 115 complement activation 29 complement factor C3b 29 complement factor D 35 composite membranes 95, 115 constant conductance alternating current (CCAC) 202 contact phase activation 13

controlled release kinetics 176 convective clearance 23 convective treatments 9 Cordis Dow 9 CR-DDS (controlled release drug delivery system) 175f crossﬂow rate 127 crossﬂow (TFF) 116, 137 cross-laid knitted hollow ﬁber mats 63 cross-wound mat device 64 Cuoxam process 4 cuprammonium rayon 8 Cuprophan 4, 8, 27–29, 33 current density 200 cytokine elimination 36 cytokine generation 35 cytokine mRNA production 35

d dead-end 137 DEAE cellulose 5 DEAE-cellulose (Hemophan) 29 DEAE-modiﬁed cellulose 10, 28 degradable polymers 236 degranulation of neutrophils 34 dense membranes 55, 57 dermal ﬁbroblasts 229 dermis 192 description of diffusion 20 device controlled delivery 214 diabetes mellitus 263 –characterization 263 –history 264 –insulin-dependent diabetes mellitus 263 –STZ-induced 270 –treatment 263 diaﬁltration 123, 125, 129 diaﬁltration optimization parameter 130 dialysis patients 1 dialysis 157, 168 dialyzer 2 –constructions 17, 20 –housing 18 –sales 21 DIAPES 16 diffusion –coefﬁcient of the drug 195 –membrane controlled release 176 –tests 112 diffusive ﬂow 132 direct current (DC) 202 dirt holding capacity 102 double ﬁltration ﬁrst 74

Index double ﬁltration set up 74 drug ionic current ﬂow 200 drug partition coefﬁcient 195 dynamic binding 149

e electrodialysis 157, 168 electromigration 199 electroosmosis 198 electroosmotic ﬂux 200 electroporation 204 –electrode material and design 205 –pulse speciﬁcations 205 –safety issues 205 elute puriﬁcation 147 endogenous retroviruses 135 endogenous viral clearance 135 endotoxins 32 ENKA company 4 enzyme retention 163 epidermis 192 epithelial cells 252 erythropoietin (EPO) 13 ESRD patient population 2 EVAL 27, 33 Excebrane membrane 33f extractibles 137 extravascular devices 272ff

f factor D 35 factors H and B 30 fed batch 128 fermenter 102 ﬁber bundle 19 ﬁbrotic response –Ba-alginate microcapsules 295 –PLL microcapsules 294 Fick’s law 176, 177 ﬁlter selection 104 ﬁltration mode 137 ﬂat membrane ﬁlter 86 ﬂow rate 104 ﬂow-through 147, 148 ﬂux excursion 126 FMC FX class series 19 formulation 121 fractionation membrane 81 Franz diffusion cell 197 freeze-thaw 141 Fresenius polysulfone 8, 15, 16, 28, 29, 33, 35 Fresenius 6, 16

FX-class of dialyzers 20

g Gambro 16 gamma irradiation 105, 110 gas transfer 53 –carbondioxide 53 –oxygen 53 gas transfer rates 58 –microporous PP membrane 58 –PMP membrane 58 glyco-protein GMP-140 29 gold nanoparticles 120 good manufacturing practice 135 graft survival 270 –extravascular 274, 279, 281, 293, 300, 302f., 308 –hemoglobin 311, 312 –intravascular 270, 271 –oxygen limitations 309 –siliceous membrane 307 Graham, Thomas 2 guarding layer 97 guided bone generation 237, 239 guided tissue generation 237, 239

h Haas, Georg 1ff. Hagen–Poiseulle’s law 23 heat lung machine 65 Helixone 16 Helixone membranes 6 hemapheresis 70 hematocrit (Hct) 21 hemodialysis 1 Hemoﬂux 17 hemolytic dispositions 85 Hemophan 10, 28, 30, 35f. hepatoblastoma 230, 249 hollow ﬁber 114 –devices 117 –dialyzers 17 Hospal (Gambro) 12 Hospal AN69 27 Hospal/Code 16 hourglass 95 housing 18 human skin structure 192 hydrogel 180–186 hydrophilic 105, 112 hydrophobic 105 hydrophobic interaction chemistry 146 hydrophobic microporous ﬂat sheet membranes 52

323

324

Index

i

m

IDDM (insulin dependent diabetes mellitus) 264 IDEMSA, Spain 8 immune rejection 267, 294 –cytotoxic reaction 268 –humoral immune response 268, 291, 308 inert electrodes 202 inner diameter 24f. insulin secretion 271, 292, 301, 305, 308 –insulin release time 271, 292, 301 integrity test 98, 112, 132 –air ﬂow integrity test 132 integrity 125, 144 interleukin-1 35 intermediate buffer exchange 121 ion-exchange 146 islets covered with a siliceous membrane 304f. –3-aminopropyltrimethoxysilane, APTrMOS 305 –tetramethoxysilane, TMOS 305 islet-like cell clusters (ICCs) 291 islets of Langerhans 263, 276 –depth of oxygened outer rim 310 –direct deposition of a siliceous membrane 306 –gel entrapment matrices 276 –nutrient supply and oxygenation 277, 293 –oxygenation studies 308 –preventing aggregation 276

macrocapsules 273, 301 –acrylonitrile 274 –alginate-PLL-alginate (APA) 280, 282 –AN69 274 –biocompatibility 274 –Amicon XM-50 274 –AN69 274 –poly-L-lysine (PLL) 276 macrovoid 115, 118 magnetophoresis 211 manufacturing of microporous membranes 81 matrix systems 212 melt spin stretch process 55 Membrana GmbH 4, 8, 16 membrane adsorbers 146 membrane effectiveness 270 –citotoxicity testing 270 –diffusional properties 271 –effect of pore size 306 –insulin permeability 275 –insulin release time 271 –pore-sized 275 –thrombogenicity testing 270 membrane is asymmetric 13 membrane or reservoir systems 212 membrane oxygenators 50 membrane processes 155 membrane production 60 –melt spin stretch process 60 –temperature-induced phase separation process 60 membrane selection 124 membrane sieving 123 membrane wall thickness 7 microcapsules 280, 292ff, 300, 302, 305, 307, 311 –agarose beads 302 –agarose-PSSa 303 –Barium-alginate 295f –comparison of APA with Ba-alginate 301 –effect of size 292, 295 –encapsulation process 303, 307 –fabrication 307 –ﬁbrotic overgrowth 300 –ﬁbrotic response 294f –immune rejection 294 –mechanical stability 294 –microfabrication technology 305 –multicomponent, containing polymethylene co-guanidine 307 –oxygenation islets 311

k Kalle company 3 kallikrein-kinin system 26 keratinocytes 232, 233, 234, 236 kidney 244, 245, 249 Kimal 16 kinin system 37 Kolff, Willem 3

l linearly scaleable 124 liquid porosimetric 120 liver 243, 244, 245, 246 log reduction value 120 lontophoresis 198

Index –poly-L-lysine (PLL) 280 –protruding islets 293 –swelling and shrinking phenomena 290 –use of anti-ﬁbrotic agent 300 microﬁltration membranes 93 microporous capillary membranes 62 –blood inside 62 –blood outside 62 microporous membranes 55, 61, 83 –melt spin stretch process 83 –production 55 microporous PMP membrane 59 –dense outer skin 59 microporous PP membrane 56 microprojection/microneedle patch 210 Minntech 16 Moire´ structure 19 molecular pore sizes 183 molecular weight cutoff 22 monoclonal antibody 129 Murine Leukemia Virus 142 mycoplasma 94, 101

n neonatal pancreatic cell clusters (NPCC) 300 nervous system 240 NFF, see normal ﬂow Niels Alwall 3 Nikkiso 16 noncellular blood components 78 –size 78 normal ﬂow (NFF) 118, 137 normal molecular weight cutoff 116 Nucleopore 274

o OROS 175 osmotic pressure 122 oxidative stress 32 oxygen species production 32 oxygenation membranes 55 oxygenator 54, 61 –extraluminal ﬂow 54 –membrane makeup 61 –operational modes 61 OXYPLUS 59

p PAN DX 27, 37 PAN 27 PAN17DX 31

pancreas 243 pancreas, bioartiﬁcial 265 –exocrine tissue contamination 281 –extravascular devices 272 –graft survival 270 –hollow ﬁbers 268 –implantation sites 272 –insulin adsorption 270, 271 –long-term discordant xenograft function 281, 304 –oxygen distribution 308, 311 –vascularization 277 particle challenge tests 120 Parvoviruses 136, 142 passive diffusion 195 passive drug permeability coefﬁcient 195 passive systems 212 passive TDD systems 212 patch 191 PCL (poly-e-caprolactone) 236 PEG-grafted cellulose 11 PEPA 16 PES, see polyethersulfone PH effect 181f, 185 pilot-scale studies 141 plasma fractionation membranes 79f –producers 80 –sieving coefﬁcients 80 plasma kallikrein 38 plasma separation 75 –set up 74 plasma separation membranes 77, 79 Plasmaﬂo 79 Plasmaphan 79 plasmapheresis 69 –procedures 83 pleated cartridges 93, 97 Pmax 109 PMMA, see Polymethyl Methacrylate polyacrylonitrile 5, 29, 33 polyacrylonitile membranes 13 polyacrylonitrile (PAN) 12, 234f, 251 polyamide (PA) 5, 17, 36 polyamide S 16 Polyamix 16 polyethersulfone (PES) 5, 95, 117 polyethylenimine (PEI) 13 Polyﬂux 17 polylactide 239 polymethyl methacrylate (PMMA) 5, 7, 14, 27, 30, 33–36 Polyphen 16 polysulfone (PSu) 15 polysulfone 28, 31, 36, 95, 250, 251, 253

325

326

Index polysynthane 11 polyurethane 17 polyvinylpyrrolidone 5 poly-L-lysine (PLL) 276 –coimmobilizing the islets with hemoglobin 310 –comparison of APA with Ba-alginate 301 –effect of barium crosslinking 301 –effect of size 292 –encapsulation process 280 –geometry 273 –Mechanical stability 293, 301 –polyvinyl alcohol hydrogel tubular (MRPT) 276 –sodium methallyl sulfonate 274 –vascularization 277 pore blockage 102 pore-plugging 107 –model 141 post-harvest concentration 121 potting material 18 preﬁltering 140 process control 134 process development 124, 139 process ﬂux 115 processing plan 128 product recovery strategy 133 protein A column chromatography 147f PVDF (polyvinylidene ﬂuoride) 117

r reactive oxygen species 33 red biotechnology 155 regenerated cellulose 8, 117 retention 94, 104, 105, 115 –assurance 94 –conﬁdence 100 retention assurance 101 retrovirus like particles (RVLPs) 135 retroviruses 136 reversed phase chemisty 146 Rhone Poulenc 12 robust 133 robustness 94, 115 rotating disc plasma separator 87 rotating drum 3f rotating plasma separation device 86

s Salmon calcitonin 203

Saxonia 16 scaffold 231 scalability 115 scale-up 131 –linear-scale 131 scaling 109 sieving coefﬁcients 21f simulation study 141 size exclusion 113 sizing ﬁlters 107 skin 191, 232 skin penetration enhancers 207 skin fouling 202 skin Controlled Delivery 214 small-scale testing 111 spacer yarns 19 spiral wound module 116 steaming 98 steam-in-place 112 sterile ﬁltration 103 sterile microﬁltration 93 sterility 145 –assurance 99f sterilization 99, 110 –autoclave sterilization 110 –ethylene oxide (ETO) 26 –hydrophobic sterilization membranes 102 –method 112 –pre-sterilized 99 –sterilizing assurance 100 –sterilizing grade 100 sterilizing grade 101, 105 sterilizing membranes 105 Stewart, Richard 4 stratum corneum 193 stretching process 61 structural parameters 182 surface modiﬁcation 98 symmetric membranes 118

t tangential ﬂow 118 TAT formation 27 temperature effect 186 temperature-induced phase separation (TIPS) 55 Terumo 11 Terumo polysulfone 34 TFF, see cross ﬂow Thalhimer, William 3 thermally induced phase separation 82 thrombin-anti thrombin III complex (TAT) 27 thrombogenicity 26

Index

ultraﬁltrate (QF) 24 ultraﬁltration 113, 121 –factor (UE) 21 –membranes 114f –process 121, 123, 128 –theory 122 ultrasound-assisted TDD – sonophoresis 208 –high frequency 208 –low frequency 208 –medium frequency 208 undulation 19

vascularization 273, 278, 303 –AN69 hollow ﬁber 278 –ﬁbroblast growth factor 278f, 303 –hydroxymetilated polysulfone 271 –neovascularization of macrocapsules 277 –polyglycolic acid (PGA) ﬂat sheet 278 –polystyrene sulfonic acid (PSSa) 279 –polytetra ﬂuoroethylene (PTFE) scaffold 278 –regenerated cellulose hollow ﬁbes 279 –subcutaneous implantation site 273, 303 virus capture 149 virus ﬁlters 143 –virus clearance ﬁlters 144 virus ﬁltration 113, 137 –ﬁltration process 137 –validation 142 virus removal 117, 136 –virus clearance ﬁlters 117 –virus removal membranes 117 virus spike 142 Visking Company 3 Vitabran E 16 vitamin E 12, 33, 34 vitamin E-modiﬁed cellulosic 11 vitamine C 34 Vmax 107ff, 141

v

w

vascular devices 268ff. –Amicon XM-50 (polyvinyl chloride acrylic capolymer) 269 –Cuprophan HDF (cellulose) 269 –hydroxymetilated polysulfone 271 –polyethertherketone (PEEK) 270 –Thomapor (polyamide) 269

wall thickness 9 water permeability 116 Weingand 3 wetting 112 white biotechnology 155

timolol 197 tissue 267 tissue engineering 227ff, 228, 231, 232, 236, 240, 241 TNFa (tumor nekrose factor alpha) 35 Toray Industries 14, 16 Toraysulfone 16 transdermal drug delivery (TDD) 191 transmembrane pressure (TMP) 21, 23 transplantation 264, 265 –allogeneic or xenogeneic cellular tissue 265 –pancreatic islets 264 transport number 200

u

x Xanthogenate 9

z zeta potential 38

327

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