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

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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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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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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, Antońio 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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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 Tećnico Av. Rovisco Pais 1049-001 Lisboa Portugal


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Contributors ((Peinemann Vol. 1))

Antońio Norton de Matos University of Porto ICBAS and HGSA-Santo Antońio 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 Tećnico 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 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.


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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


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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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.


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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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.


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.


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


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


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


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.


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


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


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 >20

–— –— 60 000

–— –— >0.6 >0.6

1.5 –— –—

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


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.


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


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


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).


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


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


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 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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4 Membranes in the Biopharmaceutical Industry

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.


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


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).


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.


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 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


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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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.


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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


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


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.


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 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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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


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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170


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).


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


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6 Membranes in Controlled Release

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


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.


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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182


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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184


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


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 ˜ ań-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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186


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


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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delivery, and bionanotechnology. AIChE Journal, 49, 2990–3006. Peppas, N. A. and Meadows, D. L. (1983) Macromolecular structure and solute diffusion in membranes: an overview of recent theories. Journal of Membrane Science, 16, 361–377. Peppas, N. A. and Lustig, S. R. (1985) The role of crosslinks, entanglements and relaxations of the macromolecular carrier in the diffusional release of biologically active materials: conceptual and scaling relationships. Annals of the New York Academy of Sciences, 446, 26–41. Narasimhan, B., Mallapragada, S. K. and Peppas, N. A. (1999) Release kinetics: data interpretation, In Mathiowitz, E. ed. Encyclopedia of Controlled Drug Delivery, John Wiley & Sons, Inc., New York, NY, 921–935. Siepmann, J. and Peppas, N. A. (2001) Modeling of drug release from delivery systems based on hydroxypropyl methylcellulose (HPMC). Advanced Drug Delivery Reviews, 48, 139–157. amEnde, M. T., Hariharan, D. and Peppas, N. A. (1995) Factors influencing drug and protein transport and release from ionic hydrogels. Reactive Polymers, 25, 127–137. Brannon-Peppas, L. and Peppas, N. A. (1988) Structural analysis of charged polymeric networks. Polymer Bulletin, 20, 285–289. Collins, M. C. and Ramirez, W. F. (1979) Mass transport through polymeric membranes. Journal of Physical Chemistry, 83, 2294–2301. Gudeman, L. F. and Peppas, N. A. (1995) pH-Sensitive membranes from poly(vinyl alcohol)/poly(acrylic acid) interpenetrating networks. Journal of Membrane Science, 107, 239–248. Renkin, E. M. (1954) Filtration, diffusion, and molecular sieving through porous cellulose membranes. The Journal of General Physiology, 38, 225–243. Pappenheimer, J. R. (1953) Passage of molecules through capillary walls. Physiologial Reviews, 33, 387 Yasuda, H. and Lamaze, C. E. (1971) Salt rejection by polymer membranes

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in reverse osmosis. I. Nonionic polymers. Journal of Polymer Science, 9, 1537–1551. Yasuda, H., Lamaze, C. E. and Schindler, A. (1971) Salt rejection by polymemembranes in reverse osmosis. II. Ionic polymers. Journal of Polymer Science, 9, 1579–1590. Anderson, J. L. and Quinn, J. A. (1974) Restricted transport in small pores, a model for steric exclusion and hindered particle motion. Biophysical Journal, 14, 130–150. Peppas, N. A. and Reinhart, C. T. (1983) Solute diffusion in swollen membranes. I. A new theory. Journal of Membrane Science, 15, 275–287. Reinhart, C. T.Peppas, N. A. (1984) Solute diffusion in swollen membranes. II. Influence of crosslinking on diffusive properties. Journal of Membrane Science, 18, 227–239. Sassi, A. P., Planch, H. W. and Prausnitz, J. M. (1996) Characterization of size-exclusion effects in highly swollen hydrogels: correlation and prediction. Journal of Applied Polymer Science, 59, 1337–1346. Khare, A. R. and Peppas, N. A. (1995) Swelling/deswelling of anionic copolymer gels. Biomaterials, 16, 559–567. Chou, L. Y., Blanch, H. W. and Prausnitz, J. M. (1992) Buffer effects on aqueous swelling kinetics of polyelectrolyte gels. Journal of Applied Polymer Science, 45, 1411–1423. Kuo, J. H., Amidon, G. L. and Lee, P. I. (1988) pH-Dependent swelling and solute diffusion characteristics of poly(hydroxyethyl methacrylate-comethacrylic acid) hydrogels. Pharmaceutical Research, 5, 592–597. Hariharan, D. L. and Peppas, N. A. (1993) Swelling of ionic and neutral polymer networks in ionic solutions. Journal of Membrane Science, 18, 1–12. Peppas, N. A. and Merrill, E. W. (1976) PVA hydrogels: reinforcement of radiation-crosslinked networks by crystallization. Journal of Polymer Science, Part A: Polymer Chemistry, 14, 441–457.

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(1989) Solute and penetrant diffusion in swellable polymers. IX. The mechanism of drug release from pH-sensitive swelling-controlled systems. Journal of Controlled Release, 8, 267–274. Lowman, A. M. and Peppas, N. A. (1999) Hydrogels, In Mathiowitz, E. ed. Encyclopedia of Controlled Drug Delivery, John Wiley & Sons, Inc., New York, 397–418. Hoffman, A. S. (1997) Intelligent polymers, In Park, K. ed. Controlled Drug Delivery Challenges and Strategies, American Chemical Society, Washington, DC, 485–498. Peppas, N. A. (1995) Controlling protein diffusion in hydrogels, In Lee, V.H.L., Hashida, M. and Mizushima, Y. eds. Trends and Future Perpsectives in Peptide and Protein Drug Delivery, Harwood, Chur, 23–37. Saito, H., Wu, X., Harris, J. M. and Hoffman, A. S. (1997) Graft copolymers of poly(ethylene glycol) (PEG) and chitosan. Macromolecular Rapid Communications, 18, 547–550.

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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.


192

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

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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

195

196


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.


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.

197

198

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.


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

199

200


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.


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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202


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


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].

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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 (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


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].


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].


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


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,


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 Fundaçaõ para a Cieˆncia a Tecnologia SFRH/BD/10594/2002. References 1 Colton, C. K. and Avgoustiniatos, E.

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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


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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