Tissue-Engineered Vascular Grafts [1st ed.] 9783030053352, 9783030053369

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Reference Series in Biomedical Engineering Tissue Engineering and Regeneration Series Editor: Heinz Redl

Beat H. Walpoth · Helga Bergmeister Gary L. Bowlin · Deling Kong · Joris I. Rotmans Peter Zilla  Editors

Tissue-Engineered Vascular Grafts

Reference Series in Biomedical Engineering Tissue Engineering and Regeneration Series Editor Heinz Redl Ludwig Boltzmann Institute for Experimental and Clinical Traumatology/AUVA Research Center Austrian Cluster for Tissue Regeneration Wien, Austria

This series Tissue Engineering and Regeneration consists of comprehensive reference texts encompassing the biological basis of tissue regeneration, basic principles of tissue engineering, and the current state-of-the-art in tissue engineering of specific tissues and organs. Each volume combines established fundamentals and the latest developments, thus forming an invaluable collection for both experienced researchers as well as practitioners from other areas of expertise. The spectrum of topics ranges from the use of cells for tissue regeneration and tissue engineering, growth factors and biological molecules affecting tissue development and regeneration, to the specific roles of biophysical factors in tissue development and regeneration. Tissue engineering lies at the crossroads of medicine, life sciences, and engineering. The field has developed extensively over the last two decades, addressing the requirements of tissue and organ replacement as well as regeneration in a variety of congenital, traumatic, disease, and aging-related conditions, including some of the most critical unmet challenges in modern medicine. Both our increased understanding of the biological basis of tissue engineering as well as significant technological advances mean that engineering design principles can now be used for the de novo construction of functional tissue replacements that meet the requirements of research and clinical applications. More information about this series at http://www.springer.com/series/13441

Beat H. Walpoth • Helga Bergmeister Gary L. Bowlin • Deling Kong Joris I. Rotmans • Peter Zilla Editors

Tissue-Engineered Vascular Grafts With 86 Figures and 21 Tables

Editors Beat H. Walpoth Emeritus, Cardiovascular Surgery and Research University Hospital and University of Geneva Geneva, Switzerland

Helga Bergmeister Center for Biomedical Research Medical University Vienna Vienna, Austria

Gary L. Bowlin Department of Biomedical Engineering University of Memphis Memphis, TN, USA

Deling Kong Key Laboratory of Bioactive Materials of the Ministry of Education College of Life Sciences, Nankai University Tianjin, China

Joris I. Rotmans Department of Internal Medicine Leiden University Medical Center Leiden, The Netherlands

Peter Zilla Chris Barnard Division of Cardiothoracic Surgery, Groote Schuur Hospital University of Cape Town Cape Town, South Africa

ISBN 978-3-030-05335-2 ISBN 978-3-030-05336-9 (eBook) ISBN 978-3-030-05337-6 (print and electronic bundle) https://doi.org/10.1007/978-3-030-05336-9 © Springer Nature Switzerland AG 2020 This work is subject to copyright. All rights are reserved by the Publisher, whether the whole or part of the material is concerned, specifically the rights of translation, reprinting, reuse of illustrations, recitation, broadcasting, reproduction on microfilms or in any other physical way, and transmission or information storage and retrieval, electronic adaptation, computer software, or by similar or dissimilar methodology now known or hereafter developed. The use of general descriptive names, registered names, trademarks, service marks, etc. in this publication does not imply, even in the absence of a specific statement, that such names are exempt from the relevant protective laws and regulations and therefore free for general use. The publisher, the authors, and the editors are safe to assume that the advice and information in this book are believed to be true and accurate at the date of publication. Neither the publisher nor the authors or the editors give a warranty, expressed or implied, with respect to the material contained herein or for any errors or omissions that may have been made. The publisher remains neutral with regard to jurisdictional claims in published maps and institutional affiliations. This Springer imprint is published by the registered company Springer Nature Switzerland AG. The registered company address is: Gewerbestrasse 11, 6330 Cham, Switzerland

This book is dedicated to all the researchers who have been working on the improvement of vascular grafts and tissue engineering, the clinicians who treat patients in need of such grafts, and finally the patients who help us to improve the results of tissue engineered vascular grafts.

Preface

Cardiovascular diseases are the leading cause of mortality for non-transmissible diseases in developed countries despite many years of preventive actions including medical treatment, dietary recommendations, and patient education. Research and innovations have been carried out actively over the last 100 years since Carrel’s Nobel Prize to provide blood supply to organs thanks to vascular anastomoses. Indeed, blood circulation is vital for organ survival as it supplies oxygen and nutrients while carrying away waste products, and the body has a fantastic cardiovascular network of arteries, capillaries, and veins for this purpose. Ageing, disease, trauma, and congenital defects limit or prevent this action and minimal invasive or open surgery should be considered for revascularization of the ischemic organs. In most cases replacement with autologous veins or arteries is preferred but often a vascular graft should be used. Synthetic, shelf-ready vascular grafts are available since 50 years with reasonable results in the large caliber (>6 mm) but with poor patency in the small caliber, limiting their use in coronary artery bypass grafts (CABG), distal peripheral revascularization, or access surgery for hemodialysis. Considering the above, alternative approaches in the form of tissue engineering as pioneered by Langer and Vacanti have also been applied successfully for tissue engineered vascular grafts (TEVG). These approaches range from synthetic, degradable scaffolds implanted directly in the host (in vivo TEVG) or after in vitro seeding/ maturation with the appropriate cell types showing astonishing results with the creation of a neo-artery as pioneered in children by T. Shinoka. Alternatively, scaffolds can be obtained by decellularization of allo- or xenogenic arteries or veins as well as in vivo fibroblast growth around a mandril. Other approaches include tissue engineered Self-Assembly (TESA) or 3D printing with a positive proof-of-concept in humans. Finally, the creation of a human scaffold by bio-reactor maturation of a construct using allogenic human cells, which is then decellularized in order to obtain a human acellular vessel (HAV), has shown good clinical results in access and revascularization surgery. All the above-mentioned tissue engineering approaches were translated successfully in humans, but larger controlled randomized trials (CRT) and long-term results must confirm the positive patient outcomes obtained so far. This Springer Reference e-book on Vascular Tissue Engineering is describing the state-of-the-art in the field of TEVG and will, with on-going updates and editions, vii

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Preface

stay at the forefront of innovation, research, teaching and clinical translation in the field of vascular replacement and access surgery for improved quality of life and survival of our patients.

Acknowledgments

I would like to thank my co-editors for their support, suggestions, and reviews as well as their contributions to this reference book on Vascular Tissue Engineering. I would also like to thank all the authors for their state-of-the-art chapters towards the improvement of TEVG with various experimental and clinical approaches which will contribute to the future well-being of our patients. I also would like to thank the Tissue Engineering and Regenerative Medicine International Society (TERMIS) for supporting this tissue engineering series as well as our Thematic Group of Vascular Tissue Engineering. The organizers of the International Symposium on Vascular Tissue Engineering (ISVTE) have also to be congratulated for their bi-annual global meeting on Vascular Tissue Engineering held in combination with the International Society of Applied Cardiovascular Biology (ISACB) during our last meeting. Last, but not least, my thanks go to Heinz Redl and his team for initiating the Springer Reference Series on Tissue Engineering as well as the team from Springer for their help and professional work. Beat Walpoth

List of Reviewers Elena Aikawa, Toshiharu Shinoka, Anthony S. Weiss

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Contents

Part I Translation and Testing of Tissue Engineered Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1

Clinical Applications and Limitations of Vascular Grafts . . . . . . . . . . . Timothy Pennel and Peter Zilla

3

Mechanical Testing of Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . . . Martin Stoiber, Christian Grasl, Francesco Moscato, and Heinrich Schima

35

Preclinical In Vivo Assessment of Tissue Engineered Vascular Grafts and Selection of Appropriate Animal Models . . . . . . . . . . . . . . . . . . . . Helga Bergmeister and Bruno K. Podesser

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Vascular Tissue Engineering: Pathological Considerations, Mechanisms, and Translational Implications . . . . . . . . . . . . . . . . . . . . . Frederick J. Schoen, Anna Mallone, Leda Klouda, and Carlijn V. C. Bouten

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Part II

Approaches of Tissue Engineered Vascular Grafts . . . . . . .

Synthetic Materials: Processing and Surface Modifications for Vascular Tissue Engineering . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . William E. King III, Benjamin A. Minden-Birkenmaier, and Gary L. Bowlin

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In Vivo Tissue-Engineered Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . Beat H. Walpoth, Sarra de Valence, Jean-Christophe Tille, Damiano Mugnai, Tornike Sologashvili, Wojciech Mrówczyński, Mustafa Cikirikcioglu, Erman Pektok, Suzanne Osorio, Francesco Innocente, Marie-Luce Bochaton-Piallat, Benjamin Nottelet, Afksendyios Kalangos, and Robert Gurny

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Decellularized Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Sotiria Toumpaniari, Andres Hilfiker, Axel Haverich, and Sotirios Korossis

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Autologous Mandril-Based Vascular Grafts . . . . . . . . . . . . . . . . . . . . . Wouter J. Geelhoed, Lorenzo Moroni, and Joris I. Rotmans

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Contents

.....

295

..........

321

.......

339

Small-Diameter Engineered Arteries: The Gel Approach . . . . . . . . . . . Brett C. Isenberg, Chrysanthi Williams, Zeeshan H. Syedain, and Robert T. Tranquillo

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Advances in Cell Seeding of Tissue Engineered Vascular Grafts Justin S. Weinbaum, Darren G. Haskett, Talya F. Mandelkern, and David A. Vorp Vascular Tissue Engineering: The Role of 3D Bioprinting Yu Shrike Zhang and Ali Khademhosseini

Textile-Reinforced Scaffolds for Vascular Tissue Engineering Alicia Fernández-Colino and Stefan Jockenhoevel

Part III Functionalization and Pathology of Tissue Engineered Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Elastin in Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Richard Wang, Bente J. de Kort, Anthal I. P. M. Smits, and Anthony S. Weiss The Incorporation and Release of Bioactive Molecules in Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Muhammad Shafiq, Hongyu Yan, Adam C. Midgley, Kai Wang, Qiang Zhao, and Deling Kong Tissue Engineering to Study and Treat Cardiovascular Calcification . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Mark C. Blaser, Samantha K. Atkins, and Elena Aikawa Vascularization of 3D Engineered Tissues . . . . . . . . . . . . . . . . . . . . . . . Young Min Ju, Anthony Atala, and James J. Yoo Part IV Clinical Application of Tissue Engineered Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Off-the-Shelf Tissue-Engineered Vascular Conduits: Clinical Translation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Emanuela S. Fioretta, Lisa von Boehmer, Melanie Generali, Simon P. Hoerstrup, and Maximilian Y. Emmert

377 379

411

429 469

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Tissue-Engineered Vascular Grafts for Children . . . . . . . . . . . . . . . . . . Toshihiro Shoji, Christopher Breuer, and Toshiharu Shinoka

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Bioengineered Human Acellular Vessels . . . . . . . . . . . . . . . . . . . . . . . . Juan Wang, Jonathan Wu, Jeffrey H. Lawson, and Laura E. Niklason

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Index . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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About the Editors

Beat H. Walpoth M.D. Emeritus, Cardiovascular Surgery and Research University Hospital and University of Geneva Geneva, Switzerland Beat H. Walpoth, M.D., P.D., FAHA, Emeritus, is a trained cardiovascular surgeon and past Director of Cardiovascular Research in the Department of Surgery at Geneva University Hospital, Switzerland. He obtained his medical degree in 1972 from the University of Zurich. Postgraduate training included 2 years at the Peter Bent Brigham, Harvard University, Boston (1973–75), and cardiac transplantation at Stanford University (1982–84). Teaching appointments held at Boston, Zurich, Bern, Geneva, and Verona Universities. Dr. Walpoth is a recipient of several national and international awards, including the ESAO Wichtig Award in 2008 and 2012 for his group’s research on Tissue Engineered Vascular Grafts (TEVG). He has over 200 publications, of which more than 100 are firstauthor papers, in peer-reviewed journals. Dr. Walpoth is Past President of the European Society for Artificial Organs (ESAO) as well as the International Symposium on VTE (ISVTE) and has created the TERMIS Thematic Group on TEVG. His current main areas of interest include TEVG, cell therapy, transplantation, angiogenesis, blood flow measurements, as well as bio-artificial cardiovascular support. His future projects are to support the initiated changes from artificial to bio-artificial, i.e., tissue engineered organs, as well as continuing international teaching and networking in the field for better outcome of clinical applications.

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About the Editors

Helga Bergmeister Center for Biomedical Research Medical University Vienna Vienna, Austria Helga Bergmeister obtained a degree in human medicine from the Medical University of Vienna (M.D.). Prior to that, she was educated in veterinary medicine at the University of Veterinary Medicine Vienna (Bachelor of Veterinary Medicine) and completed a postgraduate doctoral program (Doctor of Veterinary Medicine). Helga Bergmeister obtained education as an experimental surgeon in the Experimental Department of the II. Clinics of Surgery, Faculty of Medicine Vienna. Since 1991 she has been employed at the Center for Biomedical Research, Medical University Vienna. Although Helga Bergmeister is involved in the whole spectrum of experimental studies at the University, the main focus of her work is dedicated to cardiovascular procedures and implants. Since many years she has been performing regulatory approvals as a principal investigator or study director. She is a member of the institutional animal ethics committee. Helga Bergmeister is a teacher in the Medical School and in the Doctoral program and in postgraduate Courses of the Medical University of Vienna. Her research interests are focused on the creation of small diameter vascular grafts made of natural and newly designed synthetic materials. Gary L. Bowlin Department of Biomedical Engineering University of Memphis Memphis, TN, USA Dr. Bowlin is a Professor and Herbert Herff Chair of Excellence at The University of Memphis in the Department of Biomedical Engineering. Dr. Bowlin received his Ph.D. from the University of Akron in Biomedical Engineering and subsequently completed an American Heart Association sponsored postdoctoral fellowship in the Department of Surgery at Akron City Hospital. In 1997, Dr. Bowlin started his first faculty appointment at Virginia Commonwealth University and rose to the rank of Professor and held the Louis and Ruth Harris Exceptional Scholar Professorship. In August 2013, he

About the Editors

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relocated to The University of Memphis to continue his research and entrepreneurial endeavors and training of the next generation of Biomedical Engineers. Dr. Bowlin’s collaborative research has and continues to focus on the application of electrospun templates for tissue engineering and tissue regenerative applications, all in the pursuit of saving lives and improving the quality of life. Dr. Bowlin’s laboratory has published extensively in these areas with over 140 peer-reviewed, highly-cited manuscripts. He is a Fellow of the National Academy of Inventors and the American Institute for Medical and Biological Engineering. He is also the Inaugural and current President of the International Society for Biomedical Polymers and Polymeric Biomaterials. Deling Kong Key Laboratory of Bioactive Materials of the Ministry of Education College of Life Sciences Nankai University Tianjin, China Deling Kong is a Professor of Biochemistry and Polymer Chemistry. He is currently the Chief, Office of Science and Technology of Nankai University. Dr. Kong received his Doctorate in Polymer Chemistry and Physics from Nankai University in 1997. He underwent his first postdoctoral training at the Center of Bioengineering, Rostock, Germany, from 1998 to 2000 and the second postdoctoral training in Brigham and Women’s Hospital, Harvard Medical School, from 2000 to 2003. He came back to Nankai University in 2003. He was Director of the Key Laboratory of Bioactive Materials, Ministry of Education, from 2004 to 2017 and Dean of School of Life Science, Nankai University, from 2017 to 2018. His research interests mainly focus on vascular grafts, bioactive hydrogel, and stem cell for treatment of ischemic diseases. His group has contributed more than 30 patent applications and published over 300 peerreviewed papers. He is a recipient of several awards including the National Science and Technology

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Progress Award (2nd class), Tianjin Award of Natural Science (1st class), and Tianjin Award of Science and Technology Progress (2nd-class). He received the Outstanding Youth Fund in 2007. He was selected Fellow, Biomaterials science and engineering (FBSE) in 2016. Joris I. Rotmans Department of Internal Medicine Leiden University Medical Center Leiden, The Netherlands Joris Rotmans is an internist-nephrologist and Associate Professor in the Department of Nephrology of Leiden University Medical Center (LUMC) in the Netherlands. He obtained his master’s degree in Medicine (cum laude) from the Free University in Amsterdam. He received his Ph.D. in 2005 from the University of Amsterdam on new therapeutic strategies for vascular access for hemodialysis whereupon he started his residency in Internal Medicine. In 2008–2009, he did postdoctoral research on vascular tissue engineering at the Australian Institute of Bioengineering and Nanotechnology in Brisbane, Australia. Since 2010, he combines clinical work as internist-nephrologist with vascular and renal research in the Department of Nephrology at LUMC. His main focus of research is vascular access for hemodialysis. He was the principle investigator of the DialysisXS consortium in which a novel method to generate in vivo engineered blood vessels was developed. He is co-founder of VACIS BV, a spin-off company that aims to bring in situ engineered blood vessels to the clinic. In 2014, he received a prestigious VIDI grant from NWO that allowed him to expand his research group and to continue his research on vascular access for hemodialysis. He is the principle investigator of the LIPMAT trial, a multicenter, randomized clinical trial in the Netherlands in which the efficacy of liposomal prednisolone to enhance AVF maturation is evaluated. Dr. Rotmans is Board Member of the Vascular Access Society and Chair of the Thematic Working Group on Vascular Tissue Engineering of TERMIS.

About the Editors

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Peter Zilla Chris Barnard Division of Cardiothoracic Surgery Groote Schuur Hospital University of Cape Town Cape Town, South Africa Peter Zilla is an academic Cardiovascular and Thoracic Surgeon who is Chief of Adult and Paediatric Cardiothoracic Surgery at Cape Town’s Groote Schuur and Red Cross Children’s Hospitals, Director of the Cardiovascular Research Unit, Co-Director of the Cape Heart Institute, and CEO of Strait Access Technologies at the University of Cape Town, South Africa. He holds an M.D. degree from the University of Vienna, a Dr.Med. degree from the University of Zurich, a Ph.D. from the University of Cape Town, and a P.D. (habil) from the University of Vienna. In his capacity as Chief of the Christiaan Barnard Department, where the first human heart transplantation was performed in 1967, he has built up a training program for cardiothoracic surgeons from the African continent. His main research foci have consistently been in the fields of tissue engineering and prosthetic cardiovascular implants. Pioneering tissue engineering since 1983, he initiated an international multicenter study with in-vitro endothelialisation of peripheral bypass grafts culminating in more than 400 patient-implants. His second focus, improving heart valve prostheses for the young patients of developing countries, has led to tissue treatments that extend the longevity of heart valve prostheses manifold. His engagement in the field of rheumatic heart disease has additionally awarded him international recognition as a leader in this field when he united all major cardiothoracic surgical societies worldwide under one umbrella (Cardiac Surgery Intersociety Alliance/CSIA), thereby creating a worldwide platform for the establishment of local cardiac surgical capacity in regions of the world that are endemic for RHD but have no access to open heart surgery. Addressing this challenge also from another side, he co-founded a University of Cape Town Start-Up Company in 2008 under the name “Strait Access Technologies” (SAT), securing a leading global position in innovative concepts for trans-catheter heart

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About the Editors

valve technologies for the patients and resources of lowto middle-income countries. He is author of almost 200 peer-reviewed full papers and inventor of 41 filed or issued US/PCT patents, having been cited almost 10,000 times with an h-index of 51. He is the editor of 6 books and has authored numerous book chapters. He obtained significant international academic and industry grants and was the recipient of several international awards.

Contributors

Elena Aikawa Center for Interdisciplinary Cardiovascular Sciences, Division of Cardiovascular Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Boston, MA, USA Center for Excellence in Vascular Biology, Division of Cardiovascular Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Boston, MA, USA Anthony Atala Wake Forest Institute for Regenerative Medicine Wake Forest School of Medicine, Medical Center Boulevard, Winston-Salem, NC, USA Samantha K. Atkins Center for Interdisciplinary Cardiovascular Sciences, Division of Cardiovascular Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Boston, MA, USA Helga Bergmeister Center for Biomedical Research, Medical University Vienna, Vienna, Austria Ludwig Boltzmann Institute for Cardiovascular Research, Vienna, Austria Mark C. Blaser Center for Interdisciplinary Cardiovascular Sciences, Division of Cardiovascular Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Boston, MA, USA Marie-Luce Bochaton-Piallat Department of Pathology and Immunology, University of Geneva, Geneva, Switzerland Carlijn V. C. Bouten Department of Biomedical Engineering and Institute for Complex Molecular Systems, Eindhoven University of Technology, Eindhoven, The Netherlands Gary L. Bowlin Department of Biomedical Engineering, University of Memphis, Memphis, TN, USA Christopher Breuer The Tissue Engineering Program and Center for Regenerative Medicine, Nationwide Children’s Hospital, Columbus, OH, USA Mustafa Cikirikcioglu Service of Cardiovascular Surgery, University Hospital of Geneva, Geneva, Switzerland xix

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Contributors

Bente J. de Kort Department of Biomedical Engineering, Eindhoven University of Technology, Eindhoven, The Netherlands Sarra de Valence Ecole de Pharmacie Geneve-Lausanne – EPGL, University of Geneva, Geneva, Switzerland Maximilian Y. Emmert Institute for Regenerative Medicine (IREM), University of Zurich, Zurich, Switzerland Wyss Translational Center Zurich, University and ETH Zurich, Zurich, Switzerland Department of Cardiovascular Surgery, Charité Universitätsmedizin Berlin, Berlin, Germany Department of Cardiothoracic and Vascular Surgery, German Heart Center Berlin, Berlin, Germany Alicia Fernández-Colino Department of Biohybrid & Medical Textiles (BioTex), AME-Institute of Applied Medical Engineering, Helmholtz Institute, RWTH Aachen University, Aachen, Germany Emanuela S. Fioretta Institute for Regenerative Medicine (IREM), University of Zurich, Zurich, Switzerland Wouter J. Geelhoed Department of Internal Medicine, Leiden University Medical Center, Leiden, The Netherlands Eindhoven Laboratory of Experimental Vascular Medicine, Leiden University Medical Center, Leiden, The Netherlands Melanie Generali Institute for Regenerative Medicine (IREM), University of Zurich, Zurich, Switzerland Christian Grasl Center for Medical Physics and Biomedical Engineering, Medical University of Vienna, Vienna, Austria Ludwig Boltzmann Institute for Cardiovascular Research, Vienna, Austria Robert Gurny Emeritus, Ecole de Pharmacie Geneve-Lausanne – EPGL, Geneva, Switzerland Darren G. Haskett Department of Surgery, University of Pittsburgh, Pittsburgh, PA, USA Axel Haverich Leibniz Research Laboratories for Biotechnology and Artificial Organs (LEBAO), Department of Thoracic and Cardiovascular Surgery, Hannover Medical School, Hannover, Germany Department of Cardiothoracic, Transplantation and Vascular Surgery (HTTG) Hannover Medical School, Hannover, Germany Lower Saxony Centre of Biotechnology Implant Research and Development (NIFE), Hannover Medical School, Hannover, Germany

Contributors

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Andres Hilfiker Leibniz Research Laboratories for Biotechnology and Artificial Organs (LEBAO), Department of Thoracic and Cardiovascular Surgery, Hannover Medical School, Hannover, Germany Department of Cardiothoracic, Transplantation and Vascular Surgery (HTTG) Hannover Medical School, Hannover, Germany Simon P. Hoerstrup Institute for Regenerative Medicine (IREM), University of Zurich, Zurich, Switzerland Wyss Translational Center Zurich, University and ETH Zurich, Zurich, Switzerland Francesco Innocente Istituto Policlinico San Donato, Cardiac Surgery, Milan, Italy Brett C. Isenberg Biological Microsystems, Draper, Cambridge, MA, USA Stefan Jockenhoevel Department of Biohybrid & Medical Textiles (BioTex), AME-Institute of Applied Medical Engineering, Helmholtz Institute, RWTH Aachen University, Aachen, Germany AMIBM-Aachen-Maastricht-Institute for Biobased Materials, Maastricht University, Geleen, Netherlands Young Min Ju Wake Forest Institute for Regenerative Medicine Wake Forest School of Medicine, Medical Center Boulevard, Winston-Salem, NC, USA Afksendyios Kalangos Cardiac Surgery, Koc University Hospital, Istanbul, Turkey Ali Khademhosseini Center for Minimally Invasive Therapeutics (C-MIT), University of California-Los Angeles, Los Angeles, CA, USA Department of Radiology, David Geffen School of Medicine, University of California-Los Angeles, Los Angeles, CA, USA Department of Bioengineering, Department of Chemical and Biomolecular Engineering, Henry Samueli School of Engineering and Applied Sciences, University of California, Los Angeles, Los Angeles, CA, USA California NanoSystems Institute (CNSI), University of California-Los Angeles, Los Angeles, CA, USA William E. King III Department of Biomedical Engineering, University of Memphis, Memphis, TN, USA Leda Klouda Department of Biomedical Engineering and Institute for Complex Molecular Systems, Eindhoven University of Technology, Eindhoven, The Netherlands Deling Kong Key Laboratory of Bioactive Materials of the Ministry of Education, College of Life Sciences, Nankai University, Tianjin, China

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Contributors

Sotirios Korossis Centre of Biological Engineering, Wolfson School of Mechanical, Manufacturing and Electrical Engineering, Loughborough University, Loughborough, UK Department of Cardiothoracic, Transplantation and Vascular Surgery (HTTG) Hannover Medical School, Hannover, Germany Lower Saxony Centre of Biotechnology Implant Research and Development (NIFE), Hannover Medical School, Hannover, Germany Jeffrey H. Lawson Humacyte, Inc., Morrisville, NC, USA Department of Surgery, Associate Professor of Pathology, Duke University Medical Center, Durham, NC, USA Anna Mallone Institute for Regenerative Medicine (IREM), University of Zurich, Zurich, Switzerland Talya F. Mandelkern Department of Bioengineering, University of Pittsburgh, Pittsburgh, PA, USA Adam C. Midgley Key Laboratory of Bioactive Materials of the Ministry of Education, College of Life Sciences, Nankai University, Tianjin, China Benjamin A. Minden-Birkenmaier Department of Biomedical Engineering, University of Memphis, Memphis, TN, USA Lorenzo Moroni MERLN Institute for Technology Inspired Regenerative Medicine, Complex Tissue Regeneration, Maastricht University, Maastricht, The Netherlands Francesco Moscato Center for Medical Physics and Biomedical Engineering, Medical University of Vienna, Vienna, Austria Wojciech Mrówczyński Department of Pediatric Cardiac Surgery, Poznan University of Medical Sciences, Poznań, Poland Damiano Mugnai Service of Cardiovascular Surgery, University Hospital of Geneva, Geneva, Switzerland Laura E. Niklason Department of Anesthesia, Yale University, New Haven, CT, USA Department of Biomedical Engineering, Yale University, New Haven, CT, USA Benjamin Nottelet UFR Pharmacie – Université Montpellier, Montpellier, France Suzanne Osorio Department of Pediatric Cardiology, The Children’s Hospital, Aurora, CO, USA Erman Pektok Department of Cardiovascular Surgery, Bursa Uludağ University, Bursa, Turkey Timothy Pennel Chris Barnard Division of Cardiothoracic Surgery, Groote Schuur Hospital, University of Cape Town, Cape Town, South Africa

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Bruno K. Podesser Center for Biomedical Research, Medical University Vienna, Vienna, Austria Ludwig Boltzmann Institute for Cardiovascular Research, Vienna, Austria Joris I. Rotmans Department of Internal Medicine, Leiden University Medical Center, Leiden, The Netherlands Eindhoven Laboratory of Experimental Vascular Medicine, Leiden University Medical Center, Leiden, The Netherlands Heinrich Schima Center for Medical Physics and Biomedical Engineering, Medical University of Vienna, Vienna, Austria Ludwig Boltzmann Institute for Cardiovascular Research, Vienna, Austria Department for Cardiac Surgery, Medical University of Vienna, Vienna, Austria Frederick J. Schoen Department of Pathology, Brigham and Women’s Hospital, Harvard Medical School, Boston, MA, USA Muhammad Shafiq Key Laboratory of Bioactive Materials of the Ministry of Education, College of Life Sciences, Nankai University, Tianjin, China Department of Chemistry, Pakistan Institute of Engineering & Applied Sciences (PIEAS), Islamabad, Pakistan Toshiharu Shinoka The Tissue Engineering Program and Center for Regenerative Medicine, Nationwide Children’s Hospital, Columbus, OH, USA Department of Cardiothoracic Surgery, The Heart Center, Nationwide Children’s Hospital, Columbus, OH, USA Department of Surgery, The Ohio State University, Columbus, OH, USA Toshihiro Shoji The Tissue Engineering Program and Center for Regenerative Medicine, Nationwide Children’s Hospital, Columbus, OH, USA Anthal I. P. M. Smits Department of Biomedical Engineering, Eindhoven University of Technology, Eindhoven, The Netherlands Institute for Complex Molecular Systems (ICMS), Eindhoven University of Technology, Eindhoven, The Netherlands Tornike Sologashvili Service of Cardiovascular Surgery, University Hospital of Geneva, Geneva, Switzerland Martin Stoiber Center for Medical Physics and Biomedical Engineering, Medical University of Vienna, Vienna, Austria Ludwig Boltzmann Institute for Cardiovascular Research, Vienna, Austria Zeeshan H. Syedain Department of Biomedical Engineering, University of Minnesota, Minneapolis, MN, USA Jean-Christophe Tille Service de Pathologie Clinique, University Hospital of Geneva, Geneva, Switzerland

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Contributors

Sotiria Toumpaniari Centre of Biological Engineering, Wolfson School of Mechanical, Manufacturing and Electrical Engineering, Loughborough University, Loughborough, UK Robert T. Tranquillo Department of Chemical Engineering and Materials Science, Minneapolis, MN, USA Department of Biomedical Engineering, University of Minnesota, Minneapolis, MN, USA Lisa von Boehmer Institute for Regenerative Medicine (IREM), University of Zurich, Zurich, Switzerland David A. Vorp Department of Bioengineering, University of Pittsburgh, Pittsburgh, PA, USA Department of Surgery, University of Pittsburgh, Pittsburgh, PA, USA Department of Cardiothoracic Surgery, University of Pittsburgh, Pittsburgh, PA, USA Department of Chemical and Petroleum Engineering, University of Pittsburgh, Pittsburgh, PA, USA McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA, USA Beat H. Walpoth Emeritus, Cardiovascular Surgery and Research, University Hospital and University of Geneva, Geneva, Switzerland Juan Wang Department of Anesthesia, Yale University, New Haven, CT, USA Kai Wang Key Laboratory of Bioactive Materials of the Ministry of Education, College of Life Sciences, Nankai University, Tianjin, China Richard Wang School of Life and Environmental Sciences, The University of Sydney, Sydney, NSW, Australia Charles Perkins Centre, The University of Sydney, Sydney, NSW, Australia Justin S. Weinbaum Department of Bioengineering, University of Pittsburgh, Pittsburgh, PA, USA Department of Pathology, University of Pittsburgh, Pittsburgh, PA, USA McGowan Institute for Regenerative Medicine, University of Pittsburgh, Pittsburgh, PA, USA Anthony S. Weiss School of Life and Environmental Sciences, The University of Sydney, Sydney, NSW, Australia Charles Perkins Centre, The University of Sydney, Sydney, NSW, Australia Bosch Institute, The University of Sydney, Sydney, NSW, Australia Chrysanthi Williams Access Biomedical Solutions, Eden Prairie, MN, USA

Contributors

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Jonathan Wu Department of Biomedical Engineering, Yale University, New Haven, CT, USA Hongyu Yan Key Laboratory of Bioactive Materials of the Ministry of Education, College of Life Sciences, Nankai University, Tianjin, China James J. Yoo Wake Forest Institute for Regenerative Medicine Wake Forest School of Medicine, Medical Center Boulevard, Winston-Salem, NC, USA Yu Shrike Zhang Division of Engineering in Medicine, Department of Medicine, Brigham and Women’s Hospital, Harvard Medical School, Cambridge, MA, USA Qiang Zhao Key Laboratory of Bioactive Materials of the Ministry of Education, College of Life Sciences, Nankai University, Tianjin, China Peter Zilla Chris Barnard Division of Cardiothoracic Surgery, Groote Schuur Hospital, University of Cape Town, Cape Town, South Africa

Part I Translation and Testing of Tissue Engineered Vascular Grafts

Clinical Applications and Limitations of Vascular Grafts Timothy Pennel and Peter Zilla

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 The Clinical Need and Application of Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Congenital Disease . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Atherosclerotic Disease . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3 Peripheral Arterial Disease . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4 Coronary Artery Disease . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1 Bioprosthetic Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 Synthetic Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3 Tissue-Engineered Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Graft Modifications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1 Hemostasis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.2 Hydrophobicity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.3 Antithrombogenicity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.4 Antimicrobial . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.5 Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 Graft Limitations and the Determinants of Failure . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.1 Compliance . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.2 Caliber . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.3 Porosity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

4 5 5 6 6 7 8 9 10 12 15 15 15 16 17 18 18 18 21 23 26 27

Abstract

Prosthetic grafts have been implanted in humans for the treatment of vascular disease for over six decades and have contributed to increased survival and quality of life. Despite the extensive research following this early success, there T. Pennel (*) · P. Zilla Chris Barnard Division of Cardiothoracic Surgery, Groote Schuur Hospital, University of Cape Town, Cape Town, South Africa e-mail: [email protected] © Springer Nature Switzerland AG 2020 B. H. Walpoth et al. (eds.), Tissue-Engineered Vascular Grafts, Reference Series in Biomedical Engineering, https://doi.org/10.1007/978-3-030-05336-9_1

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has been little development or improvement in graft design for clinical application. Multiple modalities have been investigated, from new polymer designs to the completely tissue-engineered blood vessels, but few have translated beyond the preclinical phase. The implementation of a prosthetic graft is dependent on the underlying vascular pathology, the type and size of the target vessel, as well as the inherent limitations of each conduit available. Although it is clear that there is a need for a more robust prosthetic vascular graft, it is essential to have a clear understanding of the clinical applications of these conduit options in order to overcome the limitations of future graft designs. Ultimately, experimental work and graft development requires a multidisciplinary approach between the surgeon and laboratory scientist to ensure translational application.

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Introduction

The reconstruction of diseased arteries with synthetic grafts represents one of the critical medical advances of the twentieth century (Abbott et al. 1993; Nerem 2000). Despite the early success (Voorhees et al. 1952; Blakemore and Voorhees 1954), there has been little progress in this field over the last six decades, most notably in the development of small-diameter conduits (Zilla et al. 2007). Two significant obstacles remain for contemporary synthetic grafts: the high affinity for acute thrombosis and the chronic development of anastomotic pannus overgrowth, referred to as intimal hyperplasia. Decades of research have been dedicated to creating an inert material that does not elicit coagulation or a chronic inflammatory response. The alternative strategy of tissue engineering aims to integrate the construct within the body, whereby the organization of cells and polymers is used to produce artificial organs. There are four levels of complexity in the development of synthetically engineered tissues. The most simplistic (Level 1) is a two-dimensional layer in the form of epithelial or endothelial replacement. The presence of tubular structures such as blood vessels represent Level 2, while hollow non-tubular structures (the bladder) represent level three. The development of a solid organ is indicative of the highest level of complexity (level four). However, after a half-century of research, there has been little progress in the alternative to autologous conduits (Level 2), and Level 1 regeneration is limited to experimental application (Kapadia et al. 2008). Complete repopulation of a synthetic graft surface with a functional endothelial layer involves highly complex and expensive endothelial cell transplant. Despite its promising results, the application has not translated into widespread clinical practice (Deutsch et al. 2009). It is clear that the lack of clinical translation is the result of several factors including an underappreciation of the complexity of tissue engineering, a lack of communication between clinicians and basic scientists with continuation of false concepts, inadequate preclinical models, and overly stringent regulations. This chapter will focus not only on the current clinical status of vascular grafts but their needs and determinants of failure.

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The Clinical Need and Application of Vascular Grafts

Vascular grafts are required when the cardiovascular system is unable to transport oxygenated blood to the surrounding tissue. The decrease in oxygen delivery (DO2) results in ischemia, hypoplasia, and/or necrosis dependent on the phase of development or the degree of malperfusion. This mismatch of “supply and demand” may have a congenital or acquired etiology and requires disease-specific approaches.

2.1

Congenital Disease

Vascular pathologies in the pediatric population are predominantly congenital, which rarely require revascularization of the systemic circulation (Saad and May 1991; Bell et al. 2003; Kaye et al. 2008). More commonly, vascular conduits in children are used for the treatment of cyanotic heart disease with systemic to pulmonary (Blalock-Taussig shunt, BT shunt), venous to pulmonary (extra-cardiac total cavo-pulmonary connection, TCPC/Fontan circulation), or right ventricular to pulmonary (Sano) shunts. With the exception of the TCPC shunt, which is a medium- to large-diameter graft, these conduits tend to be short, are often temporary, and exist in a high-flow, low-resistance environment with satisfactory patency (Ohuchi et al. 1996; Peries et al. 2005). BT shunts have a reported reintervention rate of 10% usually as a result of postoperative infection (O’Connor et al. 2011). Internal cross-sectional diameters range from 3.5 to 5 mm, and occlusion is associated with poor outcome (Fig. 1). A short interval of anticoagulation is required to prevent thrombosis following which, oral aspirin is sufficient to prevent acute occlusion (Motz et al. 1999). Despite operative mortality decreasing from 15% to 2% over the last seven decades, the total number of BT shunts has also decreased significantly as a result of newer corrective techniques (Williams et al. 2007).

Fig. 1 Explanted Blalock-Taussig shunt explanted after 3 years during corrective surgery. (a) H&E stain ePTFE mid-graft cross section (4
magnification). (b) Immunohistochemistry is representative of the documented graft coverage. CD31, Dapi, actin. (c) SEM of graft surface

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Alternatively, TCPC conduits are considered more “permanent” (until the Fontan circulation fails) and will be required to grow when implanted in a toddler. These conduits are placed in the venous circulation and together with their larger diameters require a lower burst pressure. These conduits have been successfully implemented in first-in-man clinical trials for tissue-engineered grafts (Shinoka et al. 2005). The current innovative tissue engineering work may allow implantation into the higher-pressure pulmonary systems and ultimately the systemic arterial circulation. In contrast to the pediatric population, acquired vascular disease in adults is significantly more prevalent as a result of atherosclerosis, vascular trauma, connective tissue diseases, and infection. Small-diameter conduits are used to bypass stenosed vessels in the coronary and peripheral artery system to provide distal perfusion. Unlike the relatively satisfactory performance of small-diameter pediatric shunts, the high-resistance, low-flow environment of atherosclerotic disease is the principal reason for occlusion and graft failure in this group.

2.2

Atherosclerotic Disease

Atherosclerosis is the most common cause of acquired vascular disease and accounts for around half the deaths in the developed world (Kullo and Rooke 2016), largely as a result of coronary artery disease. The pathogenesis is secondary to deposition of plasma cholesterol, in particular, low-density lipoprotein within the vessel wall (Libby 2002), The plaque progression is much more sophisticated than a collection of mural lipids and involves a complex interaction of circulating monocytes and vessel inflammation (Ross 1999; Moore and Tabas 2011). All arteries are susceptible to atherosclerosis although flow dynamics contribute to disease progression, which is often concentrated at areas of flow disruption such as arterial bifurcations. Furthermore, large- and medium-sized vessels are seldom occluded by atherosclerotic plaques due to the caliber in relation to plaque size, but they may be the source of emboli or aneurysm formation secondary to mural weakening. It is therefore important to differentiate these factors when discussing the application of small-diameter conduits rather than blurring all forms of atherosclerosis into a single pathology.

2.3

Peripheral Arterial Disease

Peripheral arterial disease (PAD) is a diffuse arteriopathy that encompasses all medium to small arteries other than coronary and intracranial vessels although commonly refers to disease distal to the iliac arteries (Kullo and Rooke 2016). The strictest definition of the PAD is an anatomical vessel occlusion; however for epidemiological reference, the symptom of intermittent claudication is required to define the disease. An estimated 200 million people suffer from PAD globally, but the prevalence outside of the developed world is largely unreported (Fowkes et al. 2013). Over

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27 million Americans have PAD, but only 8.5–10.5 million are symptomatic (Kullo and Rooke 2016), and thus around 1–3% of females and 2–3% of males over the age of 60 have PAD (Criqui et al. 1985). Despite the lack of data for non-Caucasian populations, ethnic differences have been noted in patients of the same region (Bennett et al. 2009). In randomly selected black African 50-year-old patients from a Southern African outpatient clinic, one-third were found to have PAD, of which only 9% had risk factors (Kumar et al. 2007). Although PAD is a marker of widespread arteriopathy and associated with coronary artery disease, death is rarely related to this lower limb occlusive condition. As such, management strategies rely on symptom control through alleviating pain and improving quality of life. The femoropopliteal segment is the most intervened division in PAD, of which the mainstay has been open autologous saphenous vein bypass for limb salvage (Szilagyi et al. 1973). In 7–40% of the operations, the saphenous vein is not usable and requires a synthetic alternative (Teebken and Haverich 2002). Patency of these synthetic grafts has been shown to be less than satisfactory (40% at 5 years), with poor outcomes in distal trifurcation reconstruction (Norgren et al. 2007). Endovascular therapy has now established itself as a treatment option in aortoiliac disease but remains controversial in fem-pop disease and is not indicated for more distal intervention. The Trans-Atlantic Inter-Society Consensus for the surgical management of peripheral arterial disease has recommended endovascular treatment as the treatment of choice for type A and B lesions; however it is considered high risk for type C lesions and contraindicated in type D lesions (Norgren et al. 2007). Both bare metal stents and stent grafts have shown to be viable alternatives to bypass grafting (Schillinger et al. 2006). Despite the controversy within the literature and the heterogeneity of studies, the availability of these devices for infrainguinal interventions has resulted in a 97% increase in percutaneous interventions in the last two decades (White and Gray 2007). Of note, however, is that the current interventions do not include stents below 5 mm in diameter, and there has also been an observed trend that 5 mm grafts have decreased patency compared to 6 and 7 mm devices (McQuade et al. 2010). Thus the challenge of synthetic conduit development is whether it is placed as a surgically implanted graft or as a covered stent as there is low evidence to suggest the use of smaller-diameter conduits.

2.4

Coronary Artery Disease

Ischemic coronary artery disease secondary to atherosclerosis can remain asymptomatic for many years. The supply demand mismatch presents as chronic stable angina, which only produces symptoms on exertion. Alternatively, an acute coronary syndrome due to plaque rupture presents as unstable angina or myocardial infarction, which may be precipitated at rest or occur in a crescendo pattern. The treatment of atherosclerotic coronary artery disease requires controlling the risk factors, providing medical therapy and revascularization. Revascularization may include wirebased balloon angioplasty, intra-coronary stents, and coronary artery bypass grafting (CABG). There remains a debate regarding the most appropriate intervention, but

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current recommendations include CABG for triple-vessel coronary artery disease, double-vessel coronary artery disease with proximal LAD stenosis, and survivors of sudden cardiac death that is presumed ischemia related (Patel et al. 2017). The saphenous vein remains the most frequently used conduit for lateral and posterior wall revascularization, with the left internal mammary (IMA) reserved for the left anterior descending (LAD) artery. However the evidence is trending toward total arterial revascularization (Spadaccio et al. 2019) with or without off-pump “notouch” aortic techniques, as well as modified venous harvesting techniques. No-touch versus endoscopic vein harvesting remains the topic of debate in the cardiac surgical literature as there are significant benefits with respect to local morbidity, yet there is a concern with conduit quality (Dewantoro et al. 2018). A positive Allen’s test (preventing the use of the radial artery) as well as contraindications to either IMA harvesting, redo surgery, or previous venous stripping may render standard conduits unavailable, but this is most certainly not as high as the one-third of cases often referenced in the opening paragraph of most smalldiameter manuscripts. A review of the citations for this often ends in a cul-de-sac of citations that dates back to the 1960s and almost certainly stems from PAD literature rather than CABG literature. The authors of this chapter were unable to establish the origins of this misquoted fact and caution against its citation. Lastly, synthetic grafts have implanted experimentally in patients with inadequate autologous conduits, but their dismal patency (4-year patency of 14%) in this position precludes their use (Sapsford et al. 1981; Tomizawa et al. 1998). In summary, it is true that cardiovascular disease has a significant associated mortality, but the coronary surgeons are not in dire need of a synthetic conduit to treat these patients. The preferential use of an “off-the-shelf synthetic graft” could only be considered in the future, should they show equivalent performance. The use of synthetic conduits would save time as well as limit the comorbidities of harvesting (minor digit weakness, infected harvest sites, and venous stasis); however the current literature suggests that the focus of synthetic conduits should not include coronary artery disease until PAD has been resolved.

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Vascular Grafts

Vascular conduits can be broadly classed into bioprosthetic and synthetic grafts. Bioprosthetic grafts include autologous, homologous, and heterologous tissue that may be implanted untreated or fixed through cross-link stabilization with or without synthetic re-enforcement or decellularization. Synthetic grafts on the other hand may be fabricated from synthetic material or through tissue engeneeing as completely tissue-engineered blood vessel (TEBV) or combinations thereof. Surgeons endeavor to implant autologous tissue where possible since these are associated with superior outcomes; however lack of adequate diameter or donor conduit site disruption may force the use of non-autologous alternatives. Ultimately, durability and patency of the surgical intervention are essential to increase probability of survival and improve quality of life.

Clinical Applications and Limitations of Vascular Grafts

3.1

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Bioprosthetic Grafts

3.1.1 Autologous Grafts (Autografts) The role of autografts as conduits is not strictly artificial substitutes, but since an untreated autologous native vessel is used in an “artificial” position, they are included for completeness. They remain the first-choice graft when available, but harvest site morbidity compromises their use. Autografts are limited to small- and medium-sized vessels, usually less than 10 mm. Although dependent on the target vessel, arterial conduits usually have the best long-term patency, particularly for coronary artery intervention (Montalescot et al. 2013). The internal mammary and the radial artery have the longest patency rates, whereas the greater saphenous vein is used in preference to the gastroepiploic artery due to ease of access. Despite these data, venous grafts are still the most commonly used conduit, with the exception of the left internal mammary artery in coronary artery surgery. Regardless of these shortcomings, veins are superior to contemporary synthetic alternatives as they are relatively resistant to infection and enjoy relatively minimal thromboembolic complications in the short term. Conversely, arterial grafts are seldom used in PAD and the greater saphenous vein is definitevley the conduit of choice for peripheral revascularization. It can be “cleansed” of its valves as an in situ graft or reversed to allow for anterograde flow. Although there is a theoretical concern regarding the disruption of the vasa vasorum during graft removal, there is equally a concern regarding endothelial disruption during the removal of vein valves, and thus the reversed vein is thought to be equivalent or possibly superior. Alternative vein conduits such as the basilic vein and the small saphenous vein are often neglected sources of autologous conduits that could be considered before a synthetic alternative. 3.1.2 Homologous Grafts (Homografts) Homografts have been used since the beginning of cardiovascular surgery and are the conduit of choice prior to the use of synthetic grafts. They have been implanted as untreated, frozen, refrigerated in physiological solution, freeze-dried, or preserved in glycerin, alcohol, as well as formalin fixed, but is now almost exclusively used as a valved conduit in the management of aortic root disease, either in the systemic position or in the pulmonary position in a Ross procedure (Carrel 2009). Human umbilical vein grafts (UVg) were initially introduced as a small-diameter conduit in the 1970s, but a widespread concern of incidence of Creutzfeldt-Jakob disease and aneurysmal formation led to the withdrawal of FDA approval on May 25, 2005. This largely hinged on the manufacturer’s unwillingness to screen donors for prionrelated disease (Dardik 2006). 3.1.3 Heterologous Grafts (Xenografts) Cross-species (heterologous) applications for vascular conduits have been prepared as bovine carotid, ureteric, and submucosa of small intestines and mesenteric veins, but have not shown a significant advantage to synthetic alternatives. There are however potential advantages when implanted into an infected space

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(Gharamti and Kanafani 2018). Xenografts are subject to degeneration, thrombosis, and rupture most likely due to an accelerated antigenic tissue response. This antigenicity may be reduced by cross-link modifications, but these grafts remain prone to aneurysmal dilation and still require fabric reinforcement. The bovine jugular graft (Contegra®) is the only commercially available valved xenograft and is widely used in congenital cardiac surgery (Iyer 2012).

3.2

Synthetic Grafts

The initial failures of nonporous materials such as metal, glass, and ivory (Steven and Friedman 2008) resulted in the first successful introduction of porous cloth tubes as vascular conduits (Voorhees et al. 1952). Since then clinicians and engineers have pursued synthetic materials that are strong, pliable, and biostable. This would not only require them to be resistant to thrombosis, inflammation, and infection but also allow for incorporation into the body by healing with a confluent spontaneous endothelium. Most importantly, the grafts should remain patent and intact.

3.2.1 Clinically Implanted Materials Despite sporadic experimental reports of alternative grafts, synthetic conduits are limited to two crystalline, hydrophobic materials: polyethylene terephthalate (PET/polyester) and polytetrafluoroethylene (PTFE). Although each of these vascular conduits has their niche within vascular bypass grafting (polyester – large diameter, PTFE – medium to small diameter), systematic evaluation and metaanalysis of randomized controlled trials comparing polyester and ePTFE showed no evidence of an advantage of one material over the other (Pevec et al. 1992; Takagi et al. 2010). The surgical preference for ePTFE as a small-diameter conduit may rather relate to its ease of handling and the lack of need for pre-clotting rather than any basis of solid scientific evidence (Xue and Greisler 2003). A few studies with small numbers have shown a nonsignificant trend toward improved patency of ePTFE in the distal limb, but this does not justify its overwhelming preferential use in small-diameter grafts (Post et al. 2001). Due to the lack of superiority of either of these materials, selection of an appropriate prosthetic graft should be based on the cost and its handling characteristics (Lau and Cheng 2001). 3.2.2 Polyethylene Terephthalate (PET) Terylene®, Dacron® This semicrystalline, thermoplastic polymer resin is formed from ethylene glycol and terephthalic acid. It was developed in Great Britain in 1941 by two chemists J.R. Whinfield and JT Dickinson and was first patented as Dacron® by DuPont in 1950 (King et al. 1981; Steven and Friedman 2008). Although the majority of texts refer to it by its commercial name, this chapter will refer to PET as polyester. It is bundled and spun as synthetic fibers (10–20 μm in diameter) into a multifilament yarn (24–108 filaments per yarn) and can be woven (over-and-under) or looped into knitted textile vascular grafts giving it a tensile strength of 170 MPa and a tensile modulus of 14,000 MPa (Pennel et al. 2016). These tubes can then be further

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crimped or supported by external spirals to improve kink resistance and radial strength. The material’s high melting point makes it ideal in the textile industry, and the interlocking nature of this graft construction lends itself to large bore diameter and bifurcated grafts. By employing velour techniques, one can extend the loop of yarn onto the surface of the graft, allowing for increased tissue ingrowth of the abluminal surface. Although polyester is proven to be a stable polymer for more than 10 years in vivo, ester bonds may be oxidized and are subject to hydrolysis by activated phagocytes with deterioration in physical properties (Alimi et al. 1994). While the woven grafts have small pores (voids), knitted grafts formed by looping fibers together have a greater distance between the fibers, which make them more compliant and can theoretically promote greater tissue ingrowth (Kannan et al. 2005). The interest in conduit permeability arose from a hypothesis that a lack of fluid transport contributed to rapid calcification. The consequence of the increased permeability in knitted grafts resulted in transmural extravasation of blood. Thus, the polyester graft is often impregnated with a protein (albumin, collagen, or gelatin), which can be further cross-linked (heat, aldehyde) to delay degradation of the sealant from 2 to 8 weeks and alter the water permeability as required (Singh et al. 2015). A further concern for knitted polyester is its tendency to dilate when implanted in the high-pressure arterial system (Alimi et al. 1994) which is even more pronounced in the hypertensive patient (Nunn et al. 1979). Despite these findings, dilatation seldom results in clinically relevant complications (Blumenberg et al. 1991).

3.2.3 Expanded Polytetrafluoroethylene (ePTFE) Gore-Tex®, Teflon® PTFE is a non-biodegradable, inert, and durable fluorocarbon polymer that is made microporous by extrusion and sintering into expanded PTFE (Kannan et al. 2005). It was first marketed under the Teflon trademark by DuPont in 1945 after its accidental discovery by Dr. Roy Plunkett (Yao and Eskandari 2012). The expanded version (ePTFE) was developed by researchers at W.L. Gore & Associates, Inc. (Newark, NJ) and is currently the most commonly used graft for lower-extremity and arteriovenous bypass grafting. It has a moderate stiffness of 0.5GPa (Pennel et al. 2016), with a tensile strength of 14.0 MPa as well as high flexibility and permeability to various biomaterials and gases. ePTFE also poses an electronegative luminal surface that is antithrombotic, and its biostable nature makes it less prone to deterioration in biological environments compared to polyester. Although there is less platelet adhesion or activation distal to the graft compared to that of polyester, this has not translated to clinical superiority for ePTFE (Xue and Greisler 2003). PTFE was unsuccessfully trialed as a textile prosthesis in animals before the expanded form (ePTFE) was implanted as a vascular graft in 1972, which has largely remained unchanged until today (Yao and Eskandari 2012). The use of ePTFE in human circulation was first described in 1972 (Soyer et al. 1972). It is the most commonly used small- to medium-diameter conduit for femoropopliteal implants and arteriovenous ePTFE grafts and comprises as much as 83% of all hemodialysis shunts in the United States.

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ePTFE had been shown to dilate in the arterial system, and thus clinically implantable ePTFE grafts are all composed with an impenetrable wrap re-enforcement. Importantly, it is the wrap reinforcement and not the quoted internodal distance (IND) that determines the porosity. Newer manufacturing technologies are currently in development which may reduce the need for wrap reinforcement and focus research on high-porosity designs (Kapadia et al. 2008).

3.2.4 Experimental Materials Polyurethane is a synthetic material that has been used clinically in the form of intravenous catheters and gained popularity as a potential synthetic vascular reconstructive material in the mid-1980s (Kannan et al. 2005). It comprises of a large family of elastic polymers containing a urethane [-NH-(CO)-O-] group, which have some of the highest elasticity among existing polymers (Pennel et al. 2016). Typically, these are copolymers consisting of three monomers: a crystalline segment derived from diisocyanate to provide rigidity; a soft, amorphous segment to provide flexibility; and a chain extender. This allows the material to have custom-made physical properties for medical devices and great potential for vascular conduits. Although PUs are already extensively used in medical devises, they have not received wide attention as a vascular conduit due to its structural failure. Uncontrolled degradation, dilatation, localized aneurysmal formation, and graft rupture as well as high thrombogenicity have limited their use (Zdrahala 1996; Soldani et al. 2010). Despite these previous failures, polyurethane-based vascular grafts have remained a focus of research for many years. Polyurethanes have been used in a long-term follow-up clinical trial in AV shunts compared to ePTFE, with similar primary patency (54.7% vs. 51.8%) at 2 years (Kiyama et al. 2003). Currently, the only polyurethane graft that is available for use as a vascular device is the Vectra graft (Thoratec Corp, Pleasanton, Calif.) that is implanted for hemodialysis access (Xue and Greisler 2003). The advantage of polyurethanes is their large range of physical properties. The improved compliance and elasticity as well as the customizable microstructure and porosity have more relevance than strength in small-diameter conduits (Lanza et al. 2011). Most importantly, is the ability to mold PU into desired bespoke forms, but their application is limited to experimental work and proof of concept, is the ability to mold PUs. Physical forms can be created that allow for controlled interconnected porosity and permeability (Bezuidenhout et al. 2002; Grenier et al. 2007). As a result, they are the ideal material that can be used to investigate alternative forms of vascular graft healing.

3.3

Tissue-Engineered Grafts

As noted previously, there has been limited translation of tissue-engineered vascular grafts into clinical practice. The future of vascular implants may involve in vivo regenerative approaches or in vitro approaches which may or may not require a scaffold. The complexity can also range from surface modification with cellular

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transplantation (Level 1) to that of a completely tissue-engineered blood vessel (Level 2).

3.3.1 Endothelial Cell Transplantation The concept of endothelial cell transplantation aims to limit the systemic circulatory response to the underlying graft but disguising it with a confluent endothelial monolayer. Despite positive results in experimental research, the widespread use of this technique has proven to be deceptively difficult to upscale into a clinically relevant therapy. 3.3.2 Single-Stage Seeding Dr. Malcolm Herring first described the isolation and transplantation of endothelial cells in the 1970s (Herring et al. 1978; Thornton et al. 1983). Steel wool was used to scrape endothelial cells from saphenous vein segments and immediately transplanted onto polymeric graft surfaces with an autologous blood medium. The cellular yield of this method was inadequate to cover the graft surfaces, but it was hypothesized that the endothelial cell and clot graft modification would serve as a matrix for cellular migration and proliferation (Kempczinski et al. 1985). Subsequent techniques for recovering endothelial cells from veins, arteries, and omental fat have been refined through enzymatic loosening with collagenase and trypsin, with centrifugal concentration (Pawlowski et al. 2004). Despite the improvement in endothelial cell efficiency from 75% to 100% (Deutsch et al. 1999), most cells are lost immediately on implant due to blood flow shear stress. 3.3.3 Two-Stage Seeding The breakthrough in endothelial cell culture in the early 1980s made the two-stage approach to cell seeding possible (Laube et al. 2000). The concept involves the isolation of endothelial cells for single-stage seeding, followed by ex vivo proliferation before seeding onto the vascular prosthesis. The technique allows for large quantities of cells to be cultured to create a confluent surface coverage. This gravitational approach has given rise to sodding, which forcibly deposits cells on grafts through electrostatic and hydrostatic means as well as biological glues. There is limited research on the two-stage technique. Despite this, seeded ePTFE grafts have shown 30-day patency of 92% in seeded grafts compared to 53% in control ePTFE. At 10 years, patency of seeded ePTFE is 61%. Patency outcomes are similar to the saphenous vein in the infrainguinal segment (Deutsch et al. 2009). Furthermore, seeded ePTFE grafts used in CABG have a 91% patency at 28 months, although the long-term outcome remains unknown (Laube et al. 2000). Although the two-stage procedure has shown positive clinical results, there are significant limitations. Firstly, the cell culture expansion takes up to 5 weeks which negates its use in emergency situations. The procedure is also labor-intensive, and the expertise required to maintain the program as well as the risk of contamination, has limited it’s widespread applicatinon and commercialization. Cost and lack of reproducibility are two further limitations associated with this technique (Heyligers et al. 2005). Despite these limitations, the novel adaptions to

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seeding using biodegradable scaffolds have shown promise particularly in the pediatric population as a single-staged approach, with biodegradable scaffold. A porous copolymer of lactide acid and ε-caprolactone was reinforced with poly-Llactide acid (PLLA) or polyglycolic acid (PGA). Mononuclear cells from the bone marrow were centrifuged at the time of congenital cardiac surgery and syringed onto a biodegradable conduit and sealed with fibrin glue. As discussed previously, the use of this technique has been successfully implanted as a Fontan completion, and the low shear of the venous system is less likely to wash away the seeded cells (Matsumura et al. 2003). In a study by Shinoka et al. (2005), the 42 children who received extra-cardiac conduits for the construction of a total cavo-pulmonary connection showed no evidence of aneurysmal formation (Shinoka et al. 2005).

3.3.4 Completely Tissue-Engineered Blood Vessel (TEBV) Traditionally TEBV have been thought to require three major components: a scaffold to provide support, an adhesive matrix, and living cellular components. The concept implies that a living graft will be able to remodel, grow, and self-repair. These features are considered absent in current synthetic grafts. The process requires that the graft possesses sufficient collagen and elastin to ensure the strength and elastic nature of a blood vessel, but also be free of any synthetic material, which could initiate a chronic inflammatory process or act as a nidus for infection. Below the concepts of TEBV development are briefly outlined. Weinberg and Bell reported the first tissue-engineered vessel, which was based on a collagen (natural acellular) scaffold (Weinberg and Bell 1986). This was propagated with bovine fibroblast, smooth muscle, and endothelial cells. This graft failed due to insufficient mechanical strength, even when reinforced with a PET mesh, a problem that was encountered by subsequent authors (Hirai and Matsuda 1996). The second influential study by Niklason and Langer used polyglycolic acid as a temporary scaffold to wrap layers of cell sheets (Niklason and Langer 1997). This sequential wrapping of sheet layers was performed in a bioreactor under flow conditions at 165 beats per minute with 5% radial distension. The PGA was subsequently degraded resulting in a TEBV; this graft could not however sustain systemic pressures. L’Heureux et al. (1998) optimized the Niklason and Langer design by improving mechanical strength by wrapping an acellular extracellular matrix. This consisted of the dehydrated products of adult human skin fibroblasts, which were wrapped around a PTFE mandrill. A layer of umbilical smooth muscle cells was subsequently wrapped around the sheet to form a media; this was then placed into a bioreactor, and 1 week later, fibroblasts were added to construct an outer media and adventitia after which surface endothelial seeding was performed. This multicellular sheet TEBV conduit was been trialed in ten patients as a hemodialysis access shunt, with reported 1- and 6-month patency of 78% and 60%, respectively, which meets the criteria for a high-risk patient cohort (76% expected patency across all patient population) (McAllister et al. 2009). A subsequent pilot study of three patients with the same graft minus the endothelial seeding has reported 11-month stability (Wystrychowski et al. 2014).

Clinical Applications and Limitations of Vascular Grafts

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The final design concept involves a non-scaffold approach, which is analogous to a pearl formation within an oyster. In this novel design, Campbell et al. used an animal’s peritoneal cavity as a bioreactor for a mesenchymal-based TEBV (Campbell et al. 1999). This method required no scaffold other than a silastic tube mandrill that was placed in the peritoneal cavity as a nidus. This initiated an inflammatory reaction, which resulted in successive layers of myofibroblasts and collagen matrix and a monolayer of mesothelial cells covering the tube. The mesothelium-covered myofibroblast capsule was then inverted to serve as an endothelial surface. Over 4 months the vessel underwent transformation that resembled elastic-lamellar and high-volume myofilaments. Subsequently in a porcine model, grafts were shown to have sufficient mechanical strength with vascular phenotype differentiation (Rothuizen et al. 2016).

4

Graft Modifications

Regulatory demands and limited progress in preclinical innovation have resulted in optimizing pre-existing materials being explored as an alternative for improving patency and accelerating access into clinical practice. These surface modifications aim to improve the blood-device interface without influencing the structural properties.

4.1

Hemostasis

The most rudimentary of all surface modifications is intraoperative pre-clotting of synthetic vascular grafts. This effect may have an additional biological component when applied with growth factors and adhesion proteins within the clot. Despite multiple descriptions of pre-clotting, including submersion and intermitted clamping with or without pre-wetting, a standardized method was published in the 1970s (Yates et al. 1978). This four-step protocol which aimed to combat transmural exsanguination has largely been abandoned since the introduction of protein impregnation and the reduced use of knitted grafts. More recently gelatin impregnation has been used to prevent transmural blood leakage.

4.2

Hydrophobicity

Inert hydrophobic surfaces (as in ePTFE) were thought to prevent thrombosis; however in situ evaluation has shown a propensity to initiate coagulation through protein adsorption and platelet adhesion (Brash 1977). However, this does not translate to cell adherence (Bull et al. 1995), as ePTFE grafts do not endothelialize beyond a few millimeters from the anastomosis (Berger et al. 1972). Surface modification to increase hydrophilicity through alcohol wetting and plasma film

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deposition are thought to improve patency as well as endothelial cell adhesion (Pektok et al. 2009). Although alcohol treatment may enhance endothelialization, the mechanism likely relates to the wicking of blood through the pores and subsequent displacement of air (Trudell et al. 1978; Stronck et al. 1992). The subsequent milieu of proteins and growth factors as in pre-clotting is more likely the source of increased cellularity than a physical interaction (Lehle et al. 2003). Conversely, plasma surface modification (PSM) is thought to have a direct effect on the wettability, hardness, inertness, and biocompatibility of the material (Chu et al. 2002), with subsequent cellular adhesion (Ramires et al. 2000). Plasma is created when gases are charged though an energy source, exciting molecular, ionic, and radical species. This plasma polymerization has the ability to change the properties of the underlying surface though cross-linking.

4.3

Antithrombogenicity

All current prosthetic vascular materials are inherently thrombogenic and may fail due to thrombus formation within the graft or as a result of micro-emboli shearing from the surface to create a downstream occlusion. Although platelet adhesion has a significant role in the patency of small-diameter grafts, it alone does not determine failure. PET grafts have a well-documented increase platelet deposition, but this does not result in inferior patency (Callow et al. 1982). Despite these findings, the theoretical basis has prompted further research in antithrombogenic performance, but this has not yet translated into improved performance (Kapfer et al. 2006). Surface alteration, prolonged clotting times both in vitro and in vivo in graphitecoated surface, and surface manipulation with biologically active components which have become a standard were first described in the 1960s by Gott et al. (1963). Aspirin, thrombomodulin, hirudin, and prostaglandin E1 have been used experimentally (San Román et al. 1996; Yoneyama et al. 2000), but none have had a clinical impact like heparin surface modification. Extracorporeal life support systems such as cardiopulmonary bypass have heparin bonded tubing for short-term use but are limited by the thromobgenicity of the oxygenator. More recently extracorporeal membrane oxygenators (ECMO) are licenced for 30 days due to their reduced thrombogencic polymethyl pentene oxygenators and closed circuit centrifugal pumps. Both systems require additional systemic anticoagulation, but in the case of ECMO, there are a number of case reports and series of systemic heparin-free runs (Ryu and Chang 2018). Although covalent bonding of heparin to graft surfaces does not completely prevent platelet adhesion, there are sufficient studies that show improved patency, with reduced downstream micro-emboli (Ritter et al. 1997). Alternative applications of heparin bonding make use of polyethylene glycol (Devine et al. 2001), as well as heparan sulfate substitutes which may induce wound healing and improve graft patency (Sprugel et al. 1987). However, heparin-induced thrombocytopenia remains a significant concern, which would be particularly difficult to manage with an implanted device (Bosiers et al. 2006).

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17

Antimicrobial

Infections are a serious complication of vascular grafts, and prosthetic material is always at risk of bacterial contamination at the time of implant or subsequently through bacteremia. In particular staphylococcal species are capable of graft attachment and maturation of an impenetrable barrier. In particular staphylococcal species are capable of graft attachment and maturation of an impenetrable barrier known as biofilms, which shield the bacteria from antimicrobials and allow for the bacteria subsequently detach causing systemic spread (Otto 2008). The incidence of graft infection ranges from 0.5% to 4%, and the subsequent morbidity and mortality are high despite aggressive surgical intervention. Limb loss in infected infrainguinal vascular bypass is as high as 79% (Kikta et al. 1987), and mortality ranges from 25 to 75% dependent on the site of infection, the underlying comorbidities of the patient, and the responsible organism (Calligaro and Veith 1991). Conservative treatment alone results in high mortality approaching 100% at 2 years (Smeds et al. 2016). There have however been sporadic reports of high cure rates in nonsurgical management (Erb et al. 2014). Infections are predominantly gram positive with Staphylococcus aureus that usually occurs in the first 4 months with less virulent trains of S. epidermidis appearing later (Schmacht et al. 2005). Ultimately, graft infections generally follow with systemic sepsis and are only treated with complete excision of the infected graft and infected perigraft tissue and require staged in situ of simultaneous extra-anatomical implant. The first report of graft surface modification with the intention of treating infection by antimicrobial soaking was described in the 1970s (Richardson et al. 1970). Two decades later, this concept of protecting grafts with antimicrobial modification was introduced in the clinic (Hernández-Richter et al. 2003). Silver has shown to have in vitro antimicrobial activity, but subsequent graft silver treatment of vascular grafts has not shown to be resistant to infection in clinical studies. This may be a result of the in vivo conversion from silver acetate to silver chloride, which reduces the effectiveness of the silver ion's antimicrobial effect. A further concern is the potential toxicity through graft leaching (Hernández-Richter et al. 2003). Despite the concerns of altered healing, there is no evidence of this in preclinical data; however rifampicin-treated grafts have shown peri-anastomotic necrosis (Ueberrueck et al. 2005; Schneider et al. 2008). Pretreatment of grafts with sealants such as gelatin serves as a vehicle for other antimicrobial agents such as cephalothin, rifampicin, and triclosan. Grafts can be easily pre-soaked in a solution for 15–25 min, as longer duration has not shown to increase the concentrations. Rifampicin has activity against staphylococci as well as gram-negative pathogens and bonds well, due to its strong affinity for gelatinimpregnated polyester grafts (Malassiney et al. 1996). Rifampicin-soaked polyester grafts have shown a marked transient resistance to MRSA colonization, but this has not translated to long-term sustained results. In mice inoculated with sensitive Staphylococcus aureus, rifampicin-soaked grafts were free of infection, compared to control grafts which were all infected.

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In situ replacement of grafts with or without surface treatments cannot be recommended for previously infected grafts despite the theoretical advantage of antimicrobial treatment (Schmacht et al. 2005).

4.5

Summary

Untreated autologous tissue remains the gold standard conduit for small-diameter grafting, but the limited availability and morbidity associated with harvest often preclude its use. Contemporary grafts are limited to ePTFE and polyester, which have remained the materials of choice for the last five decades. Although these two polymers are not ideal, current alternatives have not shown significantly superior results to justify change. Cellular modifications have shown an improved patency, but cost and complexity prevent widespread application. This is even more evident in the TEBV, which remains a proof of concept rather than a clinically viable option. New techniques for decellularized tissue stabilization, as well as new degradable polymers, hold promise for the future, but regenerative potential may be limited by designs that hinder spontaneous endothelialization (Pennel et al. 2013, 2016).

5

Graft Limitations and the Determinants of Failure

Vascular conduits have to remain both patent and chemically and structurally stable within the circulation to reduce risk of graft failure. Large-diameter and to a lesser degree medium caliber grafts have demonstrated this; however compliance mismatch remains a challenge. Smaller-diameter grafts have much lower patency rates particularly in the high-resistance, low-flow environments. This section outlines the mechanisms of failure of synthetic grafts and specifically the failure of autologous vein grafts, which are often referred to as the benchmark for tissue-engineered vascular grafts.

5.1

Compliance

Compliance is defined by the alteration of cross-sectional diameter of a vessel between diastolic and systolic pressure states and is usually expressed as %/100 mm Hg: CD ¼

ðDs  Dd Þ r 1 ¼ ðPs  Pd Þ
Dd t Ew

Compliance (CD) relates to the diameter (D) and pressure (P) changes, the radius (r), wall thickness (t), and bulk modulus of the wall (EW), where subscripts s and d denote systole and diastole, respectively. Blood vessels and in particular veins have nonlinear elastic properties, and therefore this equation only serves as a concept rather than practical application in

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the assessment of conduit-host artery mismatch. Furthermore, experimental grafts are subject to various measurement modalities, which may confound comparison between materials and configurations (Salacinski et al. 2001). The influence of compliance mismatch on altered healing and subsequent graft failure in small-diameter grafts was first proposed in the mid-1970s (Baird and Abbott 1976). This in turn created a niche for mathematicians and graft developers to collaborate in order to identify the relationship between compliance and flow dynamics in cardiovascular disease and graft patency through finite element algorithms. Physical forces created in the interaction between a conduit and its adjacent in situ vessel trigger tissue remodeling and anastomotic thickening, which is most pronounced at the distal anastomosis. The incompatible mechanical strain and shear wall stress with subsequent turbulent flow have influence on the endothelial cells and trigger intimal hyperplasia. This theoretical model can be confirmed by stiffening an arterial graft through glutaraldehyde cross-linking, which has higher occlusions compared to non-fixed arteries (Abbott et al. 1987). Two forms of compliance mismatch have been described: Tubular mismatch refers to the energy loss of pulsatile flow between a compliant and non-compliant graft which is thought to result in endothelial cell damage and aneurysm formation. Since all grafts longer than 3 centimeters are devoid of the endothelium, this largely refers to compliance mismatch of native arteries that are altered by atherosclerosis and calcification (Salacinski et al. 2001). Anastomotic mismatch is more complex as the low compliant graft elicits a hyper-compliant, peri-anastomotic area extending a few millimeters from the anastomosis. This paradoxically doubles the perianastomotic compliance mismatch compared to the reference value, which is thought to be the primary driver for intimal hyperplasia (Tiwari et al. 2003). Graft patency depends on the combined effect of tubular mismatch and the hypercompliant zone created in anastomotic mismatch. Both autologous and synthetic grafts display tubular compliance mismatch. Whereas the discrepancy between artery and prosthetic grafts are more substantial, veins also exhibit reduced compliance range due to lower elastin levels and less smooth muscle compared to arteries. This gives rise to venous compliance properties similar to that of a rigid tube even in the low-pressure environment of the venous circulation (Sarkar et al. 2006). When exposed to the systemic arterial circulation, veins de-endothelialize with subsequent media disruption resulting in “arterialization” from overdistension. This involves platelet deposition, vascular smooth muscle apoptosis, monocyte infiltration, and fibroblastic activity, ultimately ending in myointimal thickening and SMC proliferation (Stooker et al. 2001). A significant increase in wall stiffness occurs as a result of wall thickening and reduction of luminal size, all of which further reduce the compliance of the venous conduit.

5.1.1 Large-Diameter Grafts The clinical implications and limitations of synthetic vascular grafts seldom include large-diameter conduits reserved for the aortic position. In general the outcomes for these grafts have been deemed acceptable with a research focus toward antimicrobial

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surface treatment particularly in the management of pre-infected sites, rather than mechanical failure or patency (Gharamti and Kanafani 2018). Large-diameter grafts in the thoracic and abdominal aorta are almost exclusively polyester (Dacron®) and predominantly woven. The discrepancy in compliance between the recipient aorta and the synthetic graft has remained a theoretical concern. Compliance mismatch has a significant systemic effect, with long-term impact on reintervention and mortality despite good patency. This is in contrast to small-diameter grafts where compliance is strongly implicated in loss of patency due to local effects yet has no upstream impact on the cardiovascular system. Elastic mechanical properties are also an important feature of conduits. Conduits function as both a hydraulic buffer and a passive diastolic pump. This loss of pulsatile blood flow transmission has specific effects on the myocardium, the aortic valve complex, and coronary flow. Polyester grafts and ePTFE have an 8-fold to 18-fold lower compliance value compared to arteries (1.18  1.2, 1.1  0.3, 8.0  5.9%.mmHgx102, respectively) (Tai et al. 2000; Tremblay et al. 2009). This loss of compliance and dynamic loading of the implanted graft has a significant decrease on the Windkessel effect of an elastic vessel (Ferrari et al. 2019). This loss in downstream pulse wave has two distinct consequences for the left ventricle where both result in increased workload. Firstly, there is a loss of energy in diastole with a widening of the pulse pressure and requires a 25% increase in systolic pressure to maintain an adequate mean arterial pressure. Furthermore, the lack of the hydraulic buffer translates to increased afterload through “dynamic stenosis” all resulting in left ventricular hypertrophy (LVH) through loss of ventriculoarterial coupling (Appleyard and Sauvage 1986). This progressive LVH has been observed more frequently in patients with longer bypasses in the thoracic to abdominal aorta compared to shorter abdominal aortic grafts (Mitsui et al. 1986). A 40% increase risk in cardiovascular and all-cause mortality has been found with one standard deviation of pulse wave velocity (Vlachopoulos et al. 2010). This may be explained by left ventricular remodeling, diastolic dysfunction, as well as decrease coronary flow during diastole. There may be direct effect on the aortic valve in grafts adjacent to the aortic annulus in reimplantation and root remodeling procedures. The aortic valve is stressed in diastole, and the increased impedance of vascular grafts would not create direct strain on the coapting leaflets. Despite this, progressive aortic valve regurgitation is well documented. In valve-sparing aortic root surgery, valvular dysfunction has been attributed to the periannular tubular stiff polyester graft and disruption of the normal sinus physiology, rather than increased impedance. Although the reported 15-year freedom from intervention in centers of excellence is quoted at 97  5.5 = 3% and freedom of severe AR at 89.12%, up to 40% of patients are reported to have mild to moderate insufficiency in other cohorts. It is difficult to separate the continuation of the degenerative process in the leaflets from cusp injury against the polyester tube or sinus disruption through surgical techniques. Long-term analysis of compliant Valsalva™ (Terumo. Tokyo, Japan) grafts with preservation of the sinuses of Valsalva has shown improved outcome (Beckerman et al. 2018).

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Dilatation of the knitted polyester grafts has been well documented, and this is important to differentiate when considering apparent graft “growth” in tissueengineered or biodegradable grafts. Knitted grafts have been implicated in continual graft dilatation and have been reported to dilate up to 26  6% in the first 2 weeks of implant with an annual dilatation of 3.7  1.3% per year. The preference for woven grafts is largely based on the recognition of progressive knitted polyester dilatation, but there is little clinical impact to justify this choice (Takami et al. 2012). Adjacent aortic false aneurysm and rupture may occur in 2–6% of patients with polyester grafts, which is associated with foreign body reaction, the interaction between graft suture and native aorta, and stress created by compliance mismatch. It is important to note that prosthetic grafts that have been anastomosed to an inherently diseased aorta are at risk of failure from the underlying progression of aortopathy rather than factors directly related to the synthetic graft. Furthermore, it is a challenge to excise all of the diseased tissue in surgery. Surgical strategies can reduce the risk of rupture from high to low by replacing vulnerable regions such as the ascending aorta, large-diameter aneurysm, and those pending rupture; however the risk of progression is not negated following surgery. Large-diameter grafts, like their small-diameter counterparts, remain devoid of an endothelial surface throughout their “lifespan,” and they are less susceptible to occlusion due to their diameter. However, these grafts are still responsible for surface thrombi with downstream micro-emboli. However, unlike small-diameter grafts, conduits in the aorta tend to cause distant or systemic failure rather than local occlusion due to their high flow. These systemic effects are now becoming more recognized as having a clinical impact that may require further medical intervention. In conclusion, compliance mismatch has an important role in graft failure and occurs at multiple levels. It has been demonstrated that the compliance of biological conduits is significantly greater than those of current non-compliant (rigid) prosthetic graft materials (Veith et al. 1986). Clarification is required to define the degree of compliance necessary to maintain patency and to determine the length of time the vessel remains compliant.

5.2

Caliber

Surgical preference is often to oversize a conduit to allow for a technically simpler anastomosis through end-to-side or side-to-side beveling. Smaller-diameter grafts are more susceptible to occlusion; however oversized grafts have been shown to have an increase rate of occlusion compared to better size-matched grafts due to shear disturbances and propagation of intimal hyperplasia (Zilla et al. 2008). Overly distensible conduits, more specifically veins in the systemic arterial circulation, will undergo further dilatation which amplifies the mismatch. The cross-sectional quotient (Qc) is used to describe the mismatch between conduit and host vessels. Saphenous vein grafts in a baboon model have a twofold to threefold increase in mural diameter when anastomosed the coronary circulation compared to a better size-matched femoral circulation (Zilla et al. 2012).

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Zilla et al. have previously noted that a Qc of >0.5 would create enough shear to prevent intimal hyperplasia formation; however most clinical small-graft situations have a Qc in the range of 0.25–0.35. The Qc of an internal mammary artery is 1.0, whereas the average saphenous vein is 0.25, another contributing factor to the superior patency of this arterial conduit (Franz et al. 2010). Although the study states that the Qc is not the only factor contributing to inferior outcomes of saphenous veins as coronary conduits, it is clear that the low shear, relative stasis, and flow separation in an oversized conduit can create an environment for an orchestrated inflammatory cascade. The recruitment of inflammatory cells stimulates vascular smooth muscle in the media to proliferate and migrate, resulting in intimal proliferation and graft failure. Furthermore, the caliber mismatch/hyperplasia correlation in venous conduits subjected to arterial hemodynamic force is more complex than in synthetic grafts. This is due to the fact that autologous grafts more susceptible to intimal proliferation throughout their length, whereas current sythetic alternatives only heal at the perianastomotic region (Powell and Gosling 1998). The normal pressure in the venous system is 10 mmHg, which dramatically increases to over 100 mmHg with an immediate increase in flow, shear stress, and pulsatile wall tension and increase in caliber (Barra et al. 1986). Due to the progressive dilatation of venous grafts and worsening of the Qc, the concept of external support to prevent this distension has been created.

5.2.1 Constrictive and Nonrestrictive External Supports As is the case with many novel concepts in device development and tissue engineering, similar work has often been previously published in multiple forms. In 1963 Parsonnet et al. first described the concept of an external support sheath for endarterectomized arteries and securing difficult anastomoses but did not pursue the biological implications. Fifteen years later, the use of a support system to reduce intimal hyperplasia was first described (Karayannacos et al. 1978). These were loose-fitting, no-restrictive “stents” that were thought to function by sequestering immune and inflammatory cells on the abluminal surface of the graft and promoted adventitial angiogenesis (Desai et al. 2010). Oversized meshes have been less successful with 100% occlusion being reported in one experimental series (Murphy et al. 2007). More recently David Taggart has reported on a clinical series employing the restrictive concept with the VEST external stent (Vascular Graft Solutions, Tel Aviv, Israel), which functions by suppressing the proliferative responses induced by high wall tension by preventing stretch. The trial showed improved patency compared to the control at 4.5 years (Taggart et al. 2018). These recent clinical trials rest on the backbone of extensive preclinical work investigating the optimum amount of size reduction so as not to over-constrict, resulting in luminal folds (Zilla et al. 2011). A 27% reduction in cross section appears to be the optimum allowable restriction to prevent luminal irregularities and eddy flow but still limits intimal hyperplasia (Human et al. 2009). Although the mechanism of action remains unclear, it is thought that the surround sheath or mesh limits wall tension and shear forces, which trigger intimal

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23

hyperplasia. This downregulation of potent mitogenic substances prevents vascular smooth muscle proliferation through hemodynamic modulation (Zilla et al. 2008).

5.3

Porosity

Solid tubes were first implanted as vascular grafts at the turn of the twentieth century. The use of paraffinated glass as well as aluminum was early evidence that completely impervious materials fail as vascular conduits. Voorhees initially demonstrated a successful conduit in the form of a porous material, but it took a decade before porosity was recognized as critical in vascular graft performance when Wesolowski described direct correlation between graft patency and its porosity (Wesolowski et al. 1962). Porosity and permeability are often used interchangeably in the literature; however they have distinctly different definitions. Porosity is the measure of void (empty) spaces within a material, and thus the void fraction refers to the volume of voids divided by the total volume of the material (Guidoin et al. 1987). Permeability is the ability of a material to be permeated (penetrated) or defined as the admittance of liquids or gases that is usually measured at 120 mmHg and is expressed as ml.cm2. min1 (Zilla et al. 2007). The less well-known terms of “permeance” or “permittivity” are more precise in describing this feature. However, surgeons and manufacturers have traditionally referred to this feature as permeability and will be referenced as such in this text (Bezuidenhout and Zilla 2004). Therefore, porosity refers to the size of empty space, whereas permeability expresses the ease at which fluid flows through a material and is an inference of interconnectivity. It is thus important to note that it is often assumed that porosity is interconnected but may result in cul-de-sacs, where a “highly porous” graft may in fact be impermeable. Clinically implanted grafts have standard physical features, which have been determined by earlier research and subsequent FDA approval. These properties have become established for each of the graft types; it should be borne in mind that all porosity is not equal.

5.3.1 Polyester (PET) The porosity of polyester grafts is measured by their permeability (permeance), where 200–1000 ml.cm2.min1 is regarded as low “porosity” and 1500–4000 ml. cm2.min1 as high “porosity” (Zilla et al. 2007). Wesolowski suggested that an ideal graft should have a fluid permeability of less than 10 ml.cm2.min1 to resist hemorrhage with a “biological” permeability of 50,000 ml.cm2.min1 to assist tissue ingrowth. Although these measures are contradictory, it is important to note that contemporary grafts fall short of both of these recommendations. In vitro measurements are with referenced to water or a crystalloid solution and will not necessarily translate to blood in a circulatory system. Furthermore, hemodynamic stresses further alter the physical properties of the grafts, which redistribute the intravascular pressures across the yarns and fibers which can also be distorted

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through handling, suturing, and kinking. Ultimately these processed can alter the permeability of the conduit. Although polyester textiles have not been specifically designed to allow for mid-graft healing, sporadic reports have documental mid-graft endothelial growth in knitted grafts. The larger voids in knitted grafts may allow for increased transmural tissue growth at the detriment to resistance to dilatation. It must be noted however that tissue ingrowth is not exclusively dependent on porosity as polyester has been known to incite perigraft inflammation and despite the high void index does not appear to allow tissue ingrowth for equivalent porosity as ePTFE. It may be necessary to take into account both the physical and chemical features when describing the healing characteristics of vascular grafts.

5.3.2 ePTFE In contrast to fabric polyester grafts, the “porosity” of ePTFE is defined on the IND or fibrillar length (Pennel et al. 2016). The extrusion of the PTFE creates solid nodes that are linked to each other by a dense meshwork of fine fibrils. The void space between these fibrils is 5 μm, which many researchers consider to be nonporous irrespective of the internodal distance. The standard IND of clinical grafts is 30 μm, but the measured IND varies considerably irrespective of the manufactures’ reported dimensions. Although there is no strict definition of porosity in ePTFE, 30 μm is generally regarded as low porosity and 60 μm as high porosity, where 45 μm has been suggested as the cut point (Zilla et al. 2007). These definitions however disregard the impact of the impact of internodal void space between fibrils, which is subject to distortion at high pressure. This void space increases when thin-walled grafts (50 μm wall thickness) are exposed to systemic pressure, which is independent of the IND, allowing for transmural tissue ingrowth (Pennel 2014). Furthermore, graft wall thickness is an independent factor of tissue ingrowth. Capillaries are unable to extend beyond 200 fibril layers grafts, which are 10% of the graft wall thickness of standard grafts (IND 30 μm, wall thickness 350 μm). ePTFE grafts may be further modified by external reinforcement, a term which itself is poorly described and as a result, equally poorly understood. Almost all clinically implanted ePTFE grafts contain a thin impervious wrap reinforcement (Fig. 2) The use of this wrap was popularized following reports of aneurysmal dilatation of unwrapped ePTFE grafts (Campbell et al. 1976). As a result ePTFE grafts without an outer wrap are not clinically available due to their perceived physical fragility. This thin external layer renders all grafts impervious to tissue ingrowth irrespective of the quoted IND. Thus, it should always be taken into account when reviewing healing of these grafts that the intention to wrap was not only to increase the strength of the graft but control the fibrous tissue ingrowth. Ring reinforcement, which is less commonly used, prevents kinking of the graft and may be used in combination with a wrap in the clinical setting. In the absence of a wrap, ring reinforcement would have little effect on ingrowth. It is therefore essential to be sure of the methods of reinforcement used in an experiment before generating a correlation between ingrowth and porosity.

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Fig. 2 SEM
100 magnification of (a) luminal surface and (b) abluminal wrap reinforcement, Gore-Tex® 30 μm IND ePTFE. Luminal (c) and abluminal (d) surface at
500 magnification. Scale bar white = 50 μm, black = 100 μm

There are however proponents of an alternative, even opposing, hypotheses, supporting reduced porosity (Kimura et al. 1976). Since graft patency is not only related to endothelialization but is also affected by intimal hyperplasia, the reduction of tissue ingrowth can theoretically prevent intimal hyperplasia as long as the surface can remain non-thrombogenic. When observing the patency of vascular grafts in animal models, Campbell et al. found that limiting the fibril length improved patency in the carotid and femoral artery of dogs and reported the optimal IND of 22 μm (outperforming 34 μm) (Campbell et al. 1975). Alterations to the lumen of the graft to create a smoother surface with silicone of wrapped ePTFE have also shown a reduction in intimal hyperplasia in both baboon and canine models (Lumsden et al. 1996).

5.3.3 PU The development of stable polyurethanes has allowed for a variety of methods to establish graft porosity, which can be classified into two groups: fibrillar and foamy PUs. Fibrillar PU grafts are textiles, structurally similar to polyester grafts, and with a few exceptions, these grafts are largely impenetrable to tissue ingrowth (Hess et al. 1986). They are constructed by standard textile weaving and knitting or alternatively by electrostatic spinning (electrospinning). As a result, PU porosity is defined by fiber thickness and the winding/spinning angle. Although the parameters of the polymer as well as the construction methods can be manipulated, the resultant void created is

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coincidental rather than intentional and does not directly address graft interconnectivity that is essential for transmural tissue ingrowth. Foamy designs, such as foam floatation, dip-coating, gas expansion, and laser perforation, similar to fibrillar designs do not specifically address interconnectivity either. Phase inversion porogen extraction does allow for interconnectivity; however the use of irregular porogens such as salt also relies too heavily on randomness of contact to ensure interconnectivity. Uniformly sized and tightly prepacked porogens can be extracted by phase inversion that yields well-defined, equally spaced interconnections (Bezuidenhout et al. 2002). Foamy porosity of less than 15 μm does not allow for tissue ingrowth; however from 20 to 40 μm, capillaries start penetrating the graft wall (Xue and Greisler 2003). Patency improved from 8 to 76% ( p < 0.001) when increasing the porosity of foamy PU from 30–70 μm to 70–130 μm but also showed more extensive endothelialization (Okoshi et al. 1996). Okoshi et al. investigated polyurethane grafts with varying pore sizes (10 and 60 μm) and found both grafts had the same permeability of 41  8 and 39  8 ml. cm2.min1 at 100 mmHg (Okoshi et al. 1992; Okoshi et al. 1993) yet are considered low- and high-porosity grafts, respectively. The likely explanation for this equal permeability lies in the lack of interconnectivity of the pores, which in turn will not facilitate transmural tissue ingrowth. Despite this, increased endothelialization is quoted as being superior in the more porous grafts, which may be a reflection of improved transanastomotic ingrowth rather than capillary traversing the graft wall. It is apparent that porosity is important at two levels: Not only is sufficient void space and interconnectivity required for capillary ingrowth and subsequent transmural endothelialization, but it also appears that some form of transmural subcellular communication is required for patency. As the first investigators learned when attempting to anastomose impervious materials, solid tubes do not remain patient in the circulatory system. When the outer surface of a highly porous graft is sealed, patency decreases to 0% (Pennel et al. 2018). It is hypothesized that the transportation of cytokines and growth factors between the perigraft and lumen is essential for graft healing and patency (Zhang et al. 2007). Alternative methods of wall porosity have been described but have not yet translated into clinically implantable devices. The process of laser-induced microporosity was initially developed in an attempt to improve myocardial perfusion, known as laser angioplasty. An ultraviolet excimer laser prevents surrounding thermal injury and peripore tissue swelling (Wollenek et al. 1986). Subsequently, a 308 nm excimer laser was shown to create micropores in postmortem femoral arteries (Laufer et al. 1989), and in vivo animal models have shown improved mid-graft healing laser-perforated vascular grafts. Similar approaches have remained in the laboratory but may have potential to create porosity in tissue-engineered grafts (Grabenwöger et al. 1999).

6

Conclusions

Surface thrombogenicity and compliance mismatch between the conduit and the host target vessel remain the most significant limitations of contemporary vascular grafts. Although the development of the TEBV is the most attractive solution as it addresses

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both concepts, it remains a theoretical model as preclinical success has not translated into clinical adaption due to the extreme cost. There is little doubt that TEBV will be the future of vascular replacement, but the slow progress has forced the development of interim solutions. Endothelial seeding has shown the most robust clinical data; however this too is a complex and expensive route to consider. It is therefore important that graft designs focus on in vivo healing, improved elastic properties, as well as degradable biomaterials to promote graft integration. A multidisciplinary approach is essential to ensure that past failures are not repeated by devising preclinical devices that have no translational application.

References Abbott WM, Megerman J, Hasson JE et al (1987) Effect of compliance mismatch on vascular graft patency. J Vasc Surg 5:376–382 Abbott WM, Callow A, Moore W et al (1993) Evaluation and performance standards for arterial prostheses. J Vasc Surg 17:746–756. https://doi.org/10.1016/0741-5214(93)90120-B Alimi Y, Juhan C, Morati N et al (1994) Dilation of woven and knitted aortic prosthetic grafts: CT scan evaluation. Ann Vasc Surg 8:238–242 Appleyard RF, Sauvage LR (1986) Haemodynamic consequences of arterial replacement with a synthetic graft. Cardiovasc Res 20:26–35 Baird RN, Abbott WM (1976) Pulsatile blood-flow in arterial grafts. Lancet 2:948–950 Barra JA, Volant A, Leroy JP et al (1986) Constrictive perivenous mesh prosthesis for preservation of vein integrity. Experimental results and application for coronary bypass grafting. J Thorac Cardiovasc Surg 92:330–336 Beckerman Z, Leshnower BG, McPherson L et al (2018) The evidence in a Bentall procedure with Valsalva graft: is this standard of care? J Visc Surg 4:98–98. https://doi.org/10.21037/ jovs.2018.04.05 Bell P, Mantor C, Jacocks MA (2003) Congenital abdominal aortic aneurysm: a case report. J Vasc Surg 38:190–193. https://doi.org/10.1016/S0741-5214(03)00146-0 Bennett PC, Silverman S, Gill PS, Lip GYH (2009) Ethnicity and peripheral artery disease. QJM 102:3–16. https://doi.org/10.1093/qjmed/hcn140 Berger K, Sauvage LR, Rao AM, Wood SJ (1972) Healing of arterial prostheses in man: its incompleteness. Ann Surg 175:118 Bezuidenhout D, Zilla P (2004) Vascular grafts. In: Encyclopedia of biomaterials and biomedical engineering, pp 1715–1725. https://doi.org/10.1081/E-EBBE Bezuidenhout D, Davies N, Zilla P (2002) Effect of well defined dodecahedral porosity on inflammation and angiogenesis. ASAIO J 48:465–471 Blakemore AH, Voorhees AB (1954) The use of tubes constructed from vinyon N cloth in bridging arterial defects; experimental and clinical. Ann Surg 140:324–334 Blumenberg RM, Gelfand ML, Barton EA et al (1991) Clinical significance of aortic graft dilation. J Vasc Surg 14:175–180 Bosiers M, Deloose K, Verbist J et al (2006) Heparin-bonded expanded polytetrafluoroethylene vascular graft for femoropopliteal and femorocrural bypass grafting: 1-year results. J Vasc Surg 43:313–318, discussion 318–9. https://doi.org/10.1016/j.jvs.2005.10.037 Brash JL (1977) Hydrophobic polymer surfaces and their interactions with blood. Ann N Y Acad Sci 283:356–371. https://doi.org/10.1111/j.1749-6632.1977.tb41781.x Bull DA, Hunter GC, Holubec H et al (1995) Cellular origin and rate of endothelial cell coverage of PTFE grafts. J Surg Res 58:58–68. https://doi.org/10.1006/jsre.1995.1010 Calligaro KD, Veith FJ (1991) Diagnosis and management of infected prosthetic aortic grafts. Surgery 110:805–813

28

T. Pennel and P. Zilla

Callow AD, Connolly R, O’Donnell TF et al (1982) Platelet – arterial synthetic graft interaction and its modification. Arch Surg 117:1447–1455 Campbell CD, Goldfarb D, Roe R (1975) A small arterial substitute: expanded microporous polytetrafluoroethylene: patency versus porosity. Ann Surg 182:138 Campbell CD, Brooks DH, Webster MW et al (1976) Aneurysm formation in expanded polytetrafluoroethylene prostheses. Surgery 79:491–493 Campbell JH, Efendy JL, Campbell GR (1999) Novel vascular graft grown within recipient’s own peritoneal cavity. Circ Res 85:1173–1178. https://doi.org/10.1161/01.RES.85.12.1173 Carrel T (2009) Aortic valve and/or aortic root replacement using an aortic homograft. Multimedia Manual Card Thorac Surg. https://doi.org/10.1510/mmcts.2009.003905 Chu PK, Chen JY, Wang LP, Huang N (2002) Plasma-surface modification of biomaterials. Materials Science and Engineering: R: Reports, 36(5–6):143–206. https://doi.org/10.1016/ S0927-796X(02)00004-9 Criqui MH, Fronek A, Barrett-Connor E et al (1985) The prevalence of peripheral arterial disease in a defined population. Circulation 71:510–515 Dardik H (2006) A 30-year odyssey with the umbilical vein graft. J Am Coll Surg 203:582–583. https://doi.org/10.1016/j.jamcollsurg.2006.07.003 Desai M, Mirzay-Razzaz J, von Delft D et al (2010) Inhibition of neointimal formation and hyperplasia in vein grafts by external stent/sheath. Vasc Med 15:287–297. https://doi.org/ 10.1177/1358863X10366479 Deutsch M, Meinhart J, Fischlein T et al (1999) Clinical autologous in vitro endothelialization of infrainguinal ePTFE grafts in 100 patients: a 9-year experience. Surgery 126:847–855. https:// doi.org/10.1016/S0039-6060(99)70025-5 Deutsch M, Meinhart J, Zilla P et al (2009) Long-term experience in autologous in vitro endothelialization of infrainguinal ePTFE grafts. J Vasc Surg 49:352–362. https://doi.org/10.1016/j. jvs.2008.08.101 Devine C, Hons B, McCollum C (2001) Heparin-bonded Dacron or polytetrafluoroethylene for femoropopliteal bypass grafting: a multicenter trial. J Vasc Surg 33:533–539 Dewantoro D, Nenna A, Satriano U et al (2018) Advantages and disadvantages of total arterial coronary artery bypass graft as compared to venous coronary artery bypass graft. Vessel Plus 2:20–29. https://doi.org/10.20517/2574-1209.2018.50 E Niklason L, S Langer R (1997) Advances in tissue engineering of blood vessels and other tissues. Transpl Immunol 5:303–306. https://doi.org/10.1016/S0966-3274(97)80013-5 Erb S, Sidler JA, Elzi L et al (2014) Surgical and antimicrobial treatment of prosthetic vascular graft infections at different surgical sites: a retrospective study of treatment outcomes. PLoS One 9: e112947. https://doi.org/10.1371/journal.pone.0112947 Ferrari G, Balasubramanian P, Tubaldi E et al (2019) Experiments on dynamic behaviour of a Dacron aortic graft in a mock circulatory loop. J Biomech 86:132–140. https://doi.org/10.1016/ j.jbiomech.2019.01.053 Fowkes FGR, Rudan D, Rudan I et al (2013) Comparison of global estimates of prevalence and risk factors for peripheral artery disease in 2000 and 2010: a systematic review and analysis. Lancet 382:1329–1340. https://doi.org/10.1016/S0140-6736(13)61249-0 Franz T, Human P, Dobner S et al (2010) Tailored sizes of constrictive external vein meshes for coronary artery bypass surgery. Biomaterials 31:9301–9309. https://doi.org/10.1016/j. biomaterials.2010.08.054 Gharamti A, Kanafani ZA (2018) Vascular graft infections: an update. Infect Dis Clin N Am 32:789–809. https://doi.org/10.1016/j.idc.2018.06.003 Gott VL, WHIFFEN JD, Dutton RC (1963) Heparin bonding on colloidal graphite surfaces. Science 142:1297–1298 Grabenwöger M, Fitzal F, Sider J et al (1999) Endothelialization of biosynthetic vascular prostheses after laser perforation. Ann Thorac Surg 66:S110–S114 Grenier S, Sandig M, Mequanint K (2007) Polyurethane biomaterials for fabricating 3D porous scaffolds and supporting vascular cells. J Biomed Mater Res A 82:802–809. https://doi.org/ 10.1002/jbm.a.31194

Clinical Applications and Limitations of Vascular Grafts

29

Guidoin R, King M, Marceau D et al (1987) Textile arterial prostheses: is water permeability equivalent to porosity? J Biomed Mater Res 21:65–87. https://doi.org/10.1002/ jbm.820210111 Hernández-Richter T, Schardey HM, Wittmann F et al (2003) Rifampin and Triclosan but not silver is effective in preventing bacterial infection of vascular dacron graft material. Eur J Vasc Endovasc Surg 26:550–557. https://doi.org/10.1016/S1078-5884(03)00344-7 Herring M, Gardner A, Glover J (1978) A single-staged technique for seeding vascular grafts with autogenous endothelium. Surgery 84:498–504 Hess F, Jerusalem C, Braun B, Grande P (1986) The inner prosthetic surface structure and re-endothelialization: an experimental study in the rat using two types of microvascular prostheses for aortic implantation. Microsurgery 7:29–37 Heyligers JMM, Arts CHP, Verhagen HJM et al (2005) Improving small-diameter vascular grafts: from the application of an endothelial cell lining to the construction of a tissue-engineered blood vessel. Ann Vasc Surg 19:448–456. https://doi.org/10.1007/s10016-005-0026-0 Hirai J, Matsuda T (1996) Venous reconstruction using hybrid vascular tissue composed of vascular cells and collagen: tissue regeneration process. Cell Transplant 5:93–105 Human P, Franz T, Scherman J et al (2009) Dimensional analysis of human saphenous vein grafts: implications for external mesh support. J Thorac Cardiovasc Surg 137:1101–1108. https://doi. org/10.1016/j.jtcvs.2008.10.040 Iyer KS (2012) The Contegra bovine jugular valved conduit: living up to expectations? Ann Pediatr Cardiol 5:34–35 Kannan RY, Salacinski HJ, Butler PE et al (2005) Current status of prosthetic bypass grafts: a review. J Biomed Mater Res 74B:570–581. https://doi.org/10.1002/jbm.b.30247 Kapadia MR, Popowich DA, Kibbe MR (2008) Modified prosthetic vascular conduits. Circulation 117:1873–1882. https://doi.org/10.1161/CIRCULATIONAHA.107.714170 Kapfer X, Meichelboeck W, Groegler F-M (2006) Comparison of carbon-impregnated and standard ePTFE prostheses in extra-anatomical anterior tibial artery bypass: a prospective randomized multicenter study. Eur J Vasc Endovasc Surg 32:155–168. https://doi.org/10.1016/j. ejvs.2005.12.015 Karayannacos PE, Hostetler JR, Bond MG et al (1978) Late failure in vein grafts: mediating factors in subendothelial fibromuscular hyperplasia. Ann Surg 187:183–188 Kaye AJ, Slemp AE, Chang B et al (2008) Complex vascular reconstruction of abdominal aorta and its branches in the pediatric population. J Pediatr Surg 43:1082–1088. https://doi.org/10.1016/j. jpedsurg.2008.02.035 Kempczinski RF, Rosenman JE, Pearce WH et al (1985) Endothelial cell seeding of a new PTFE vascular prosthesis. J Vasc Surg 2:424–429 Kikta MJ, Goodson SF, Bishara RA et al (1987) Mortality and limb loss with infected infrainguinal bypass grafts. J Vasc Surg 5:566–571 Kimura T, Matsumoto M, Sugiura M, Saigusa M (1976) Studies on expanded polytetrafluoroethylene as the vascular prosthesis (the fourth report): portal vein replacement, its pore size and neointima. Jpn J Artif Organs 5:85–89 King M, Blais P, Guidoin R et al (1981) Polyethylene terephthalate (Dacron) vascular prosthesesmaterial and fabric construction aspects. Biocompatibility of clinical implants materials, vol 2. CRC Press, Boca Raton, pp 177–207 Kiyama H, Imazeki T, Kurihara S, Yoneshima H (2003) Long-term follow-up of polyurethane vascular grafts for hemoaccess bridge fistulas. Ann Vasc Surg 17:516–521. https://doi.org/ 10.1007/s10016-003-0061-7 Kullo IJ, Rooke TW (2016) Peripheral artery disease. N Engl J Med 374:861–871. https://doi.org/ 10.1056/NEJMcp1507631 Kumar A, Mash B, Rupesinghe G (2007) Peripheral arterial disease – high prevalence in rural black south Africans. S Afr Med J 97:285–288 Lanza R, Langer R, Vacanti JP (2011) Principles of tissue engineering. Academic Press, Amsterdam Lau H, Cheng SW (2001) Is the preferential use of ePTFE grafts in femorofemoral bypass justified? Ann Vasc Surg 15:383–387. https://doi.org/10.1007/s100160010085

30

T. Pennel and P. Zilla

Laube HR, Duwe J, Rutsch W, Konertz W (2000) Clinical experience with autologous endothelial cell-seeded polytetrafluoroethylene coronary artery bypass grafts. J Thorac Cardiovasc Surg 120:134–141. https://doi.org/10.1067/mtc.2000.106327 Laufer G, Wollenek G, Rüeckle B et al (1989) Characteristics of 308 nm excimer laser activated arterial tissue photoemission under ablative and non-ablative conditions. Lasers Surg Med 9:556–571 Lehle K, Buttstaedt J, Birnbaum DE (2003) Expression of adhesion molecules and cytokines in vitro by endothelial cells seeded on various polymer surfaces coated with titaniumcarboxonitride. J Biomed Mater Res A 65:393–401. https://doi.org/10.1002/jbm.a.10492 L’Heureux N, P^aquet S, Labbé R et al (1998) A completely biological tissue-engineered human blood vessel. FASEB J 12:47–56 Libby P (2002) Inflammation and atherosclerosis. Circulation 105:1135–1143. https://doi.org/ 10.1161/hc0902.104353 Lumsden AB, Chen C, Coyle KA et al (1996) Nonporous silicone polymer coating of expanded polytetrafluoroethylene grafts reduces graft neointimal hyperplasia in dog and baboon models. J Vasc Surg 24:825 Malassiney P, Goëau-Brissonnière O, Coggia M, Pechère JC (1996) Rifampicin loading of vascular grafts. J Antimicrob Chemother 37:1121–1129 Matsumura G, Hibino N, Ikada Y et al (2003) Successful application of tissue engineered vascular autografts: clinical experience. Biomaterials 24:2303–2308 McAllister TN, Maruszewski M, Garrido SA et al (2009) Effectiveness of haemodialysis access with an autologous tissue-engineered vascular graft: a multicentre cohort study. Lancet 373:1440–1446. https://doi.org/10.1016/S0140-6736(09)60248-8 McQuade K, Gable D, Pearl G et al (2010) Four-year randomized prospective comparison of percutaneous ePTFE/nitinol self-expanding stent graft versus prosthetic femoral-popliteal bypass in the treatment of superficial femoral artery occlusive disease. J Vasc Surg 52:584–591.e7. https://doi.org/10.1016/j.jvs.2010.03.071 Mitsui T, Maeta H, Fukuda I et al (1986) Left ventricular hypertrophy due to aortic bypass grafting with a long prosthesis. J Card Surg 27:201–206 Montalescot G, Sechtem U, Achenbach S et al (2013) 2013 ESC guidelines on the management of stable coronary artery disease. Eur Heart J 34:2949–3003. https://doi.org/10.1093/eurheartj/ eht296 Moore KJ, Tabas I (2011) Macrophages in the pathogenesis of atherosclerosis. Cell 145:341–355. https://doi.org/10.1016/j.cell.2011.04.005 Motz R, Wessel A, Ruschewski W, Bürsch J (1999) Reduced frequency of occlusion of aortopulmonary shunts in infants receiving aspirin. Cardiol Young 9:474–477 Murphy GJ, Newby AC, Jeremy JY et al (2007) A randomized trial of an external Dacron sheath for the prevention of vein graft disease: the extent study. J Thorac Cardiovasc Surg 134:504–505. https://doi.org/10.1016/j.jtcvs.2007.01.092 Nerem RM (2000) Tissue engineering a blood vessel substitute: the role of biomechanics. Yonsei Med J 41:735–739 Norgren L, Hiatt WR, Dormandy JA et al (2007) Inter-society consensus for the management of peripheral arterial disease (TASC II). J Vasc Surg 45:S5–S67. https://doi.org/10.1016/j. jvs.2006.12.037 Nunn DB, Freeman MH, Hudgins PC (1979) Postoperative alterations in size of Dacron aortic grafts: an ultrasonic evaluation. Ann Surg 189:741–745 O’Connor MJ, Ravishankar C, Ballweg JA et al (2011) Early systemic-to-pulmonary artery shunt intervention in neonates with congenital heart disease. J Thorac Cardiovasc Surg 142:106–112. https://doi.org/10.1016/j.jtcvs.2010.10.033 Ohuchi H, Okabe H, Nagata N, Koseni K, Kaneko Y, Itoh K (1996) Long-term patency after the Blalock-Taussig operation–comparison between classic and modified shunts. [Zasshi] [Journal]. Nihon Kyōbu Geka Gakkai, 44(8):1108–1113

Clinical Applications and Limitations of Vascular Grafts

31

Okoshi T, Chen H, Soldani G et al (1992) Microporous small diameter PVDF-TrFE vascular grafts fabricated by a spray phase inversion technique. ASAIO J 38:M201–M206 Okoshi T, Soldani G, Goddard M, Galletti P (1993) Very small-diameter polyurethane vascular prostheses with rapid endothelialization for coronary artery bypass grafting. J Thorac Cardiovasc Surg 105:791 Okoshi T, Soldani G, Goddard M, Galletti PM (1996) Penetrating micropores increase patency and achieve extensive endothelialization in small diameter polymer skin coated vascular grafts. ASAIO J. https://doi.org/10.1097/00002480-199609000-00019 Otto M (2008) Staphylococcal biofilms. Curr Top Microbiol Immunol 322:207–228 Owens CD (2010) Adaptive changes in autogenous vein grafts for arterial reconstruction: clinical implications. J Vasc Surg 51:736–746. https://doi.org/10.1016/j.jvs.2009.07.102 Patel MR, Calhoon JH, Dehmer GJ et al (2017) ACC/AATS/AHA/ASE/ASNC/SCAI/SCCT/STS 2017 Appropriate Use Criteria for Coronary Revascularization in Patients With Stable Ischemic Heart Disease: A Report of the American College of Cardiology Appropriate Use Criteria Task Force, American Association for Thoracic Surgery, American Heart Association, American Society of Echocardiography, American Society of Nuclear Cardiology, Society for Cardiovascular Angiography and Interventions, Society of Cardiovascular Computed Tomography, and Society of Thoracic Surgeons. J Am Coll Cardiol 69:2212–2241. https://doi.org/10.1016/j. jacc.2017.02.001 Pawlowski KJ, Rittgers SE, Schmidt SP, Bowlin GL (2004) Endothelial cell seeding of polymeric vascular grafts. Front Biosci 9:1412–1421 Pektok E, Cikirikcioglu M, Tille J-C et al (2009) Alcohol pretreatment of small-diameter expanded polytetrafluoroethylene grafts: quantitative analysis of graft healing characteristics in the rat abdominal aorta interposition model. Artif Organs 33:532–537. https://doi.org/10.1111/j.15251594.2009.00731.x Pennel TC (2014) Isolating and optimising transmural endothelialization as an independent mode of spontaneous vascular graft healing. University of Cape Town Pennel T, Zilla P, Bezuidenhout D (2013) Differentiating transmural from transanastomotic prosthetic graft endothelialization through an isolation loop-graft model. J Vasc Surg 58:1053–1061. https://doi.org/10.1016/j.jvs.2012.11.093 Pennel T, Zilla P, Bezuidenhout D (2016) Biomaterials in vascular graft surgery Pennel T, Bezuidenhout D, Koehne J et al (2018) Transmural capillary ingrowth is essential for confluent vascular graft healing. Acta Biomater 65:237–247. https://doi.org/10.1016/j. actbio.2017.10.038 Peries A, Al-Hay AAA, Shinebourne EA (2005) Outcome of the construction of a Blalock-Taussig shunt in adolescents and adults. Cardiol Young 15:368–372. https://doi.org/10.1017/ S1047951105000788 Pevec WC, Darling RC, L’Italien GJ, Abbott WM (1992) Femoropopliteal reconstruction with knitted, nonvelour dacron versus expanded polytetrafluoroethylene. J Vasc Surg 16:60–65. https://doi.org/10.1016/0741-5214(92)90418-8 Post S, Kraus T, Müller-Reinartz U et al (2001) Dacron vs. polytetrafluoroethylene grafts for femoropopliteal bypass: a prospective randomised multicentre trial. Eur J Vasc Endovasc Surg 22:226–231. https://doi.org/10.1053/ejvs.2001.1424 Powell JT, Gosling M (1998) Molecular and cellular changes in vein grafts: influence of pulsatile stretch. Curr Opin Cardiol 13:453–458 Ramires PA, Mirenghi L, Romano AR et al (2000) Plasma-treated PET surfaces improve the biocompatibility of human endothelial cells. J Biomed Mater Res 51:535–539 Richardson RL, Pate JW, Wolf RY et al (1970) The outcome of antibiotic-soaked arterial grafts in guinea pig wounds contaminated with E. coli or S. aureus. J Thorac Cardiovasc Surg 59:635–637 Ritter EF, Kim YB, Reischl HP et al (1997) Heparin coating of vascular prostheses reduces thromboemboli. Surgery 122:888–892

32

T. Pennel and P. Zilla

Ross R (1999) Atherosclerosis – an inflammatory disease. N Engl J Med 340:115–126. https://doi. org/10.1056/NEJM199901143400207 Rothuizen TC, Damanik FFR, Lavrijsen T et al (2016) Development and evaluation of in vivo tissue engineered blood vessels in a porcine model. Biomaterials 75:82–90. https://doi.org/10.1016/j. biomaterials.2015.10.023 Ryu KM, Chang SW (2018) Heparin-free extracorporeal membrane oxygenation in a patient with severe pulmonary contusions and bronchial disruption. Clin Exp Emerg Med 5:204–207. https:// doi.org/10.15441/ceem.17.252 Saad SA, surgery AMJOP (1991). (n.d.). Abdominal aortic aneurysm in a neonate. Elsevier. https:// doi.org/10.1016/0022-3468(91)91055-4 Salacinski HJ, Goldner S, Giudiceandrea A et al (2001) The mechanical behavior of vascular grafts: a review. J Biomater Appl 15:241–278. https://doi.org/10.1106/NA5T-J57A-JTDD-FD04 San Román J, Buján J, Bellón JM et al (1996) Experimental study of the antithrombogenic behavior of Dacron vascular grafts coated with hydrophilic acrylic copolymers bearing salicylic acid residues. J Biomed Mater Res 32:19–27. https://doi.org/10.1002/(SICI)1097-4636(199609) 32:13.0.CO;2-S Sapsford RN, Oakley GD, Talbot S (1981) Early and late patency of expanded polytetrafluoroethylene vascular grafts in aorta-coronary bypass. J Thorac Cardiovasc Surg 81:860–864 Sarkar S, Salacinski HJ, Hamilton G, Seifalian AM (2006) The mechanical properties of infrainguinal vascular bypass grafts: their role in influencing patency. Eur J Vasc Endovasc Surg 31:627–636. https://doi.org/10.1016/j.ejvs.2006.01.006 Schillinger M, Sabeti S, Loewe C et al (2006) Balloon angioplasty versus implantation of nitinol stents in the superficial femoral artery. N Engl J Med 354:1879–1888. https://doi.org/10.1056/ NEJMoa051303 Schmacht D, Armstrong P, Johnson B, Pierre K, Back M, Honeyman A, et al. (2005) Graft infectivity of rifampin and silver-bonded polyester grafts to MRSA contamination. Vascular and Endovascular Surgery, 39(5):411–420. https://doi.org/10.1177/153857440503900505 Schneider F, O’Connor S, Becquemin JP (2008) Efficacy of collagen silver-coated polyester and rifampin-soaked vascular grafts to resist infection from MRSA and Escherichia coli in a dog model. Ann Vasc Surg 22:815–821. https://doi.org/10.1016/j.avsg.2008.06.011 Shinoka T, Matsumura G, Hibino N et al (2005) Midterm clinical result of tissue-engineered vascular autografts seeded with autologous bone marrow cells. J Thorac Cardiovasc Surg 129:1330–1338. https://doi.org/10.1016/j.jtcvs.2004.12.047 Singh C, Wong C, Wang X (2015) Medical textiles as vascular implants and their success to mimic natural arteries. J Funct Biomater 6:500–525. https://doi.org/10.3390/jfb6030500 Smeds MR, Duncan AA, Harlander-Locke MP et al (2016) Treatment and outcomes of aortic endograft infection. J Vasc Surg 63:332–340. https://doi.org/10.1016/j.jvs.2015.08.113 Soldani G, Losi P, Bernabei M et al (2010) Long term performance of small-diameter vascular grafts made of a poly(ether)urethane-polydimethylsiloxane semi-interpenetrating polymeric network. Biomaterials 31:2592–2605. https://doi.org/10.1016/j.biomaterials.2009.12.017 Soyer T, Lempinen M, Cooper P et al (1972) A new venous prosthesis. Surgery 72:864–872 Spadaccio C, Rainer A, Barbato R, Trombetta M, Chello M, Meyns B (2019) The long-term follow-up of large-diameter Dacron (R) vascular grafts in surgical practice: a review. The Journal of Cardiovascular Surgery, 60(4):501–513. https://doi.org/10.23736/S00219509.16.08061-7 Sprugel KHK, McPherson JMJ, Clowes AWA, Ross RR (1987) Effects of growth factors in vivo. I. Cell ingrowth into porous subcutaneous chambers. Am J Pathol 129:601–613 Steven G, Friedman MD (2008) A history of vascular surgery. Wiley, New York Stooker W, Niessen HWM, Baidoshvili A et al (2001) Perivenous support reduces early changes in human vein grafts: studies in whole blood perfused human vein segments. J Thorac Cardiovasc Surg 121:290–297. https://doi.org/10.1067/mtc.2001.111656 Stronck JW, van der Lei B, Wildevuur CR (1992) Improved healing of small-caliber polytetrafluoroethylene vascular prostheses by increased hydrophilicity and by enlarged fibril length. An experimental study in rats. J Thorac Cardiovasc Surg 103:146–152

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Szilagyi DE, Elliott JP, Hageman JH et al (1973) Biologic fate of autogenous vein implants as arterial substitutes: clinical, angiographic and histopathologic observations in femoro-popliteal operations for atherosclerosis. Ann Surg 178:232–246 Taggart DP, Webb CM, Desouza A et al (2018) Long-term performance of an external stent for saphenous vein grafts: the VEST IV trial. J Cardiothorac Surg 13:117. https://doi.org/10.1186/ s13019-018-0803-9 Tai NR, Salacinski HJ, Edwards A et al (2000) Compliance properties of conduits used in vascular reconstruction. Br J Surg 87:1516–1524. https://doi.org/10.1046/j.1365-2168.2000.01566.x Takagi H, Goto S, Matsui M et al (2010) A contemporary meta-analysis of Dacron versus polytetrafluoroethylene grafts for femoropopliteal bypass grafting. J Vasc Surg 52:232–236. https://doi.org/10.1016/j.jvs.2010.02.010 Takami Y, Tajima K, Kato W et al (2012) Long-term size follow-up of knitted Dacron grafts (Gelseal ) used in the ascending aorta. Interact Cardiovasc Thorac Surg 14:529–531. https://doi. org/10.1093/icvts/ivr086 Teebken OE, Haverich A (2002) Tissue engineering of small diameter vascular grafts. Eur J Vasc Endovasc Surg 23:475–485. https://doi.org/10.1053/ejvs.2002.1654 Thornton SC, Mueller SN, Levine EM (1983) Human endothelial cells: use of heparin in cloning and long-term serial cultivation. Science 222:623–625 Tiwari A, Cheng KS, Salacinski H et al (2003) Improving the patency of vascular bypass grafts: the role of suture materials and surgical techniques on reducing anastomotic compliance mismatch. Eur J Vasc Endovasc Surg 25:287–295. https://doi.org/10.1053/ejvs.2002.1810 Tomizawa Y, Takanashi Y, Noishiki Y et al (1998) Evaluation of small caliber vascular prostheses implanted in small children: activated angiogenesis and accelerated calcification. ASAIO J 44: M496–M500 Tremblay D, Zigras T, Cartier R et al (2009) A comparison of mechanical properties of materials used in aortic arch reconstruction. Ann Thorac Surg 88:1484–1491. https://doi.org/10.1016/j. athoracsur.2009.07.023 Trudell LA, Boudreau L, Van De Water JM et al (1978) Alcohol-treated PTFE vascular grafts. ASAIO J 24:320 Ueberrueck T, Meyer L, Zippel R et al (2005) Healing characteristics of a new silver-coated, gelatine impregnated vascular prosthesis in the porcine model. Zentralbl Chir 130:71–76. https://doi.org/10.1055/s-2004-836244 Veith FJ, Gupta SK, Ascer E et al (1986) Six-year prospective multicenter randomized comparison of autologous saphenous vein and expanded polytetrafluoroethylene grafts in infrainguinal arterial reconstructions. J Vasc Surg 3:104–114. https://doi.org/10.1016/0741-5214(86)90073-X Vlachopoulos C, Aznaouridis K, Stefanadis C (2010) Prediction of cardiovascular events and all-cause mortality with arterial stiffness: a systematic review and meta-analysis. J Am Coll Cardiol 55:1318–1327. https://doi.org/10.1016/j.jacc.2009.10.061 Voorhees AB, Jaretzki A, Blakemore AH (1952) The use of tubes constructed from vinyon “N” cloth in bridging arterial defects. Ann Surg 135:332–336 Weinberg C, Bell E (1986) A blood vessel model constructed from collagen and cultured vascular cells. Science 231:397–400. https://doi.org/10.1126/science.2934816 Wesolowski SA, FRIES CC, Karlsen KE et al (1962) Porosity: primary determinant of ultimate fate of synthetic vascular grafts. Plast Reconstr Surg 29:131–132. https://doi.org/10.1097/ 00006534-196201000-00083 White CJ, Gray WA (2007) Endovascular therapies for peripheral arterial disease: an evidence-based review. Circulation 116:2203–2215. https://doi.org/10.1161/CIRCULATIONAHA.106.621391 Williams JA, Bansal AK, Kim BJ et al (2007) Two thousand Blalock-Taussig shunts: a six-decade experience. Ann Thorac Surg 84:2070–2075, discussion 2070–5. https://doi.org/10.1016/j. athoracsur.2007.06.067 Wollenek G, Laufer G, Fasol R et al (1986) Laser-induced vascular lesions by cw-NdYAG or pulsed UV lasers during angioplastic procedures. Thorac Cardiovasc Surg 34:63–65 Wystrychowski W, McAllister TN, Zagalski K et al (2014) First human use of an allogeneic tissueengineered vascular graft for hemodialysis access. J Vasc Surg 60:1353–1357. https://doi.org/ 10.1016/j.jvs.2013.08.018

34

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Xue L, Greisler HP (2003) Biomaterials in the development and future of vascular grafts. J Vasc Surg 37:472–480. https://doi.org/10.1067/mva.2003.88 Yao JST, Eskandari MK (2012) Accidental discovery: the polytetrafluoroethylene graft. Surgery 151:126–128. https://doi.org/10.1016/j.surg.2011.09.036 Yates SG, Barros D’Sa AA, Berger K et al (1978) The preclotting of porous arterial prostheses. Ann Surg 188:611 Yoneyama T, Ito M, Sugihara K et al (2000) Small diameter vascular prosthesis with a nonthrombogenic phospholipid polymer surface: preliminary study of a new concept for functioning in the absence of pseudo- or neointima formation. Artif Organs 24:23–28 Zdrahala RJ (1996) Small caliber vascular grafts. Part II: Polyurethanes Revisited Zhang Z, Briana S, Douville Y et al (2007) Transmural communication at a subcellular level may play a critical role in the fallout based-endothelialization of dacron vascular prostheses in canine. J Biomed Mater Res A 81A:877–887. https://doi.org/10.1002/jbm.a.31124 Zilla P, Bezuidenhout D, Human P (2007) Prosthetic vascular grafts: wrong models, wrong questions and no healing. Biomaterials 28:5009–5027. https://doi.org/10.1016/j.biomaterials.2007.07.017 Zilla P, Human P, Wolf M et al (2008) Constrictive external nitinol meshes inhibit vein graft intimal hyperplasia in nonhuman primates. J Thorac Cardiovasc Surg 136:717–725. https://doi.org/ 10.1016/j.jtcvs.2008.02.068 Zilla P, Moodley L, Wolf MF et al (2011) Knitted nitinol represents a new generation of constrictive external vein graft meshes. J Vasc Surg 54:1439–1450. https://doi.org/10.1016/j. jvs.2011.05.023 Zilla P, Moodley L, Scherman J et al (2012) Remodeling leads to distinctly more intimal hyperplasia in coronary than in infrainguinal vein grafts. J Vasc Surg 55:1734–1741. https://doi.org/ 10.1016/j.jvs.2011.11.057

Mechanical Testing of Vascular Grafts Martin Stoiber, Christian Grasl, Francesco Moscato, and Heinrich Schima

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 1.1 Material Requirements and Mechanical Properties of Natural Vessels . . . . . . . . . . . . . . . 2 Classification of Mechanical Tests . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.1 Measurement of Basic Material Properties . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 Predicting the Vascular Grafts Behavior in Physiologic Working Range . . . . . . . . . . . . . 2.3 Attempts for Complete Mechanical Characterization of Vascular Grafts . . . . . . . . . . . . . 2.4 Standards for Mechanical Characterization of Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . 3 Test Samples . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.1 Specimen Geometry . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.2 Storage . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.3 Environmental Conditions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3.4 Preconditioning . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Fundamental Material Parameters . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.1 Stress and Strain . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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M. Stoiber (*) · C. Grasl Center for Medical Physics and Biomedical Engineering, Medical University of Vienna, Vienna, Austria Ludwig Boltzmann Institute for Cardiovascular Research, Vienna, Austria e-mail: [email protected]; [email protected] F. Moscato Center for Medical Physics and Biomedical Engineering, Medical University of Vienna, Vienna, Austria e-mail: [email protected] H. Schima Center for Medical Physics and Biomedical Engineering, Medical University of Vienna, Vienna, Austria Ludwig Boltzmann Institute for Cardiovascular Research, Vienna, Austria Department for Cardiac Surgery, Medical University of Vienna, Vienna, Austria e-mail: [email protected] © Springer Nature Switzerland AG 2020 B. H. Walpoth et al. (eds.), Tissue-Engineered Vascular Grafts, Reference Series in Biomedical Engineering, https://doi.org/10.1007/978-3-030-05336-9_3

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4.2 Tensile Strength, Tensile Strain, Strain at Break . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.3 Elastic Modulus . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4.4 Compliance . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 Methods for Mechanical Characterization of Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.1 Uniaxial Tensile Test . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.2 Uniaxial Hoop Tensile Test . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.3 Planar Biaxial Test . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.4 Tubular Inflation Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.5 Membrane Bulge Test . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.6 Kink Diameter Test . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.7 Strength after Repeated Puncture . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.8 Tear Propagation Strength . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5.9 Suture Retention Strength . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6 Dynamic Measurements on Vascular Grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.1 Creep/Recovery Tests . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.2 Stress Relaxation Tests . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.3 Cyclic Test . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.4 Bioreactor Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.5 Fatigue Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.6 Accelerated Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.7 Dynamic Mechanical Analysis . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 Additional Test Methods . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.1 Nondestructive Tests . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.2 Nano/Micromechanics Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8 Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

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Abstract

Beside other requirements, the mechanical behavior plays an important role in the success of vascular grafts. For the development, quality control, and assessment of vascular grafts, precise and predictive mechanical characterization is essential. Today’s technology offers a wide range of mechanical tests for various applications. The selection of the proper test method out of these numerous options is a challenging task, because different scientific questions and issues require different testing methods. In this chapter, an overview and review of important methods with numerous relevant references for the mechanical characterization of vascular grafts is given. In addition, a classification of the testing methods according to the measurement purpose is presented. Special attention is paid to dynamic measurement techniques, which are more close to the target application. Challenges and possibilities of these measurement methods are demonstrated and discussed. This shall form a basis to assess one’s own needs and to help at the selection process to find the most suitable testing methods.

1

Introduction

For the success of vascular grafts, their mechanical behavior plays an important role. Comparable biomechanical properties to native host vessels are essential for successful functioning and long term patency (Collins et al. 2012; Natasha et al. 2015). In this book chapter, an overview and review of important methods for mechanical

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characterization of vascular grafts is given. Challenges and possibilities of these mechanical measurements are demonstrated and discussed. Mechanical characterization is of course only one part of the necessary characterization methods for biomaterials and tissue engineered constructs.

1.1

Material Requirements and Mechanical Properties of Natural Vessels

Blood vessels can be seen as inflated tube-like structures, where the arterial pressure is translated into tension-dominated mechanical stress in circumferential and axial direction (Bäck et al. 2013). In vivo the artery is in a prestretched state under an internal pressure load (Holzapfel et al. 2000). Blood vessels are soft tissues, which behave like viscoelastic solids, having both viscous and elastic behavior. In a soft biological tissue, as for every solid with a microstructure, there is a correlation between its internal structure and its macroscopic mechanical properties. The passive mechanics of the vascular wall involve elastin and collagen. Elastin fibers are easily distended and have a linear stress-strain relation. The function of elastin is to provide elasticity at lower stress levels. In contrast, collagen fibers are extremely stiff and break at low strains. Collagen fibers in vascular tissues are wavy and loosely arranged; therefore, they take the load only at a certain distension. At low strain only elastin fibers are taking the load, followed by gradually recruitment of the stiffer collagen fibers at higher strain (VanBavel et al. 2003). This is the reason, for the nonlinear stress-strain behavior in the pressure-diameter curves of blood vessels. The arterial wall is composed of three distinct layers, the intima, the media, and the adventitia. Due to different fiber orientations in these layers, there is a complex, anisotropic mechanical behavior (Peterson and Bronzino 2008). To better understand this behavior, constitutive laws for modeling the mechanical properties of arterial tissue are developed (Gasser et al. 2006). The detailed knowledge of the mechanical properties of blood vessels and their substitutes is important, because mechanical properties influence the vessels physiology, can be reason for vascular diseases, and have effects on blood flow. The measurement of the mechanical properties of arterial walls can be challenging due to these complex phenomena, which have to be taken into account. Due to anisotropy, different physical properties in the radial, circumferential, and longitudinal directions are present. Viscoelasticity is the reason that the stiffness of the vessel depends on the rate at which it is deformed. The nonlinear stress-strain relationship is a consequence of the fact that blood vessels are composed of several materials, each with different elastic properties. Moreover, residual stresses are present, which are the forces that remain within the vessel’s wall when all external loads have been removed.

2

Classification of Mechanical Tests

The target information that the mechanical test should deliver has a big influence in the choice of the used testing method. Different mechanical tests can give results for different purposes. Of course there is also the factor of costs, time, complexity, and

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a

j

b

d

c

e

k

f

g

h

i

l

Fig. 1 Important loading modes for mechanical characterization of vascular grafts include: (a) uniaxial tension, (b) uniaxial compression, (c) circumferential hoop tension, (d) in-plane biaxial tension, (e) biaxial tension by pressurizing a vessel, (f) torsion, (g) kinking load, (h) suture tear load, (i) tear propagation load, (j) diametric compression, (k) bending load, (l) membrane bulge load. (Modified from Roeder 2013)

availability which has an influence in the selection a suitable testing method. The principle “keep it as simple as possible and as complex as necessary” is appropriate in that case. There also exist different modes of loading a test sample, which are chosen for different measurement purposes. Common modes of loading materials (Roeder 2013) are represented in Fig. 1 as important loading modes for mechanical characterization of vascular grafts. These loading modes include (a) uniaxial tension, (b) uniaxial compression, (c) circumferential hoop tension, (d) in-plane biaxial tension, (e) biaxial tension by pressurizing a vessel, (f) torsion, (g) kinking load, (h) suture tear load, (i) tear propagation load, (j) diametric compression, (k) bending load, and (l) membrane bulge load. It has to be mentioned that material properties measured in different loading modes cannot be considered as equivalent. A direct comparison is not possible and it is important to precisely describe the type of loading to the samples when reporting the measured mechanical properties. To give the reader an orientation, the authors have established a coarse classification of the mechanical tests according to different objectives of the measurements. This classification is shown in Fig. 2, where some of these testing methods are highlighted, as they are recommended by standards. The borders in this classification are of course not sharply delimited and there is no claim for completeness.

2.1

Measurement of Basic Material Properties

In the early product design stage, where a material is compared to an already known material, or to check the constant production quality, tests which give basic material

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Classification of mechanical tests for vascular grafts Pure material properties

Component testing

Complete characterization

For material comparison, qualitiy control and fulfilment of specifications

To predict the graft‘s behavior in at physiologic loading

For constitutive, microstructural and degradation models and as basis for Finite Element Simulations

Uniaxial tensile test

Uniaxial hoop tensile test

Stress relaxation test

Repeated puncture test

Cyclic test

Dynamic Mechanical Analysis

Tear propagation strength

Creep/recovery test

Tubular inflation test /burst pressure

Nano/Micromechanics test

Suture retention test Membrane bulge test Planar biaxial measurements Kink diameter test Fatique testing

Accelerated tests Bioreactor testing Recommended by International Standard ISO 7198:2016, "Cardiovascular implants - Tubular vascular prostheses," International Organization for Standardization Genève, Switzerland, 2016 Recommended by ASTM Standard F2150-13, "Standard Guide for Characterization and Testing of Biomaterial Scaffolds Used in Tissue-Engineered Medical Products," ASTM International, West Conshohocken, PA, 2013.

Fig. 2 Classification of mechanical test according to the different objectives of the measurements. By standards recommended test methods are highlighted

parameters are important. Here the uniaxial tensile test, suture retention test, and burst pressure measurements are common test methods. To fulfill given specifications or for comparison to literature data, often tensile stress is calculated. In vascular grafts however, the sample size is often limited, which reduces the options in mechanical test and hinders the measurement of pure material properties. It has to be also kept in mind that influencing factors like different pretreatment, boundary conditions, sample geometry, and the load transmission to the sample can reduce comparability. In literature, often varying data for the same tissues due to applications of different evaluation methods are reported (Awad et al. 2018).

2.2

Predicting the Vascular Grafts Behavior in Physiologic Working Range

The finished vascular graft is often not a single material, it has a sophisticated structure, a geometry which is not ideal for performing standardized test, and

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complex boundary conditions act on it. Therefore, different test methods which allow a better prediction of the graft’s performance might be useful. Such test methods simulate dynamic loading in the physiological range, like cyclic pressure inflation tests, dynamic hoop tensile tests, or biaxial measurements. Tests considering the viscoelasticity and long term mechanical behavior of the material are also important to get a deeper understanding how the graft would behave after implantation. Of course this requires more complex measurement equipment and a time consuming test procedures and evaluation of the measured results. Mechanical properties can change after implantation, blood and cells may enter the graft’s wall, and material internally changes and degradation can occur over the time. Also the graft’s final environment with growing surrounding tissue and fluid pressures effects the materials behavior. Therefore, in vivo tests are indispensable to completely asses the vascular graft’s performance. In vivo measurement techniques to measure the elasticity are essential at this stage. Techniques using ultrasound, magnetic resonance tomography, pulse wave velocity, x-ray, and contactless optical strain measurements are used (Chirinos 2012; Mugnai et al. 2013).

2.3

Attempts for Complete Mechanical Characterization of Vascular Grafts

Complete constitutive models allow a better understanding of the fundamental behavior of arteries and vessel structures (Holzapfel et al. 2000). A lot of experimental data is necessary to establish viscoelastic and hyper-elastic vessel behavior. Sophisticated models are even used to predict degradation behavior or include fiber orientation and fiber recruitment. For such a detailed characterization, test methods of different vessel layers in multiaxial directions are performed (Holzapfel and Ogden 2010). The models reliability is depending on the quality and completeness of the experimental data. Such data can then also be used as boundary condition for finite element simulations (Holzapfel et al. 2002).

2.4

Standards for Mechanical Characterization of Vascular Grafts

Standards like ISO 7198, Cardiovascular implants — Tubular vascular prostheses (International Standard ISO 7198:2016 2016) and ASTM F2150 Standard Guide for Characterization and Testing of Biomaterial Scaffolds Used in Tissue-Engineered Medical Products (ASTM Standard F2150-13 2013) report a wide range of mechanical testing methods. Test methods are suggested, for the characterization, quality control, and 100% inspections of vascular grafts. The type of necessary test methods can be subdivided in biological, coated, synthetic nontextile, and synthetic textile structures. Mechanical testing of vascular grafts is based on tests for plastics (ISO/DIS 6721-1:2018 2018; ASTM Standard D3039-00 2000; ASTM Standard

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D2290-00 2000; ASTM Standard D2290-01 2001; ASTM Standard D638-02a 2002; ASTM Standard D882-02 2002) and textiles (ASTM Standard D1388-96 1996) and therefore normative references to these procedures are common in these standards. However, such regulations do not define test methods in complete detail, and they can only give certain suggestions. It is in the responsibility of the researcher or producer to choose proper testing methods and to establish adequate measurement protocols.

3

Test Samples

3.1

Specimen Geometry

The specimen geometry is critical for accurate measurements of material properties. Loading the test sample can create a nonuniform stress or strain distribution, especially in vascular structures, where the specimen size is often limited. The size and type of load transmission can have strong effects on the measured results (Waldman and Lee 2005). A certain length is required to allow a uniform stress distribution and to reduce the effect of specimen clamping. Local strain measurements can reduce this error (Sun et al. 2005), because they exclude the deformation and movement at the clamps. For stress calculations, it is essential to have an exactly measured sample geometry, which can be challenging in soft tissues. Using thickness gauges may deform the material and falsify the measured values. In native structures also the thickness can vary within the sample, or tissue which takes no load is attached to it. This has to be considered and it is recommended to use contactless measurements techniques (de Gelidi et al. 2017). An alternative could be to use the measured force only, without calculation of the engineering stress. For direct comparison of different materials within the same measurement setup is this a simple way to reduce inaccuracy (Stoiber et al. 2015).

3.2

Storage

When testing samples of native source or samples including cells, the time span until the measurement takes place is critical. Keeping the samples at defined storage conditions is important, because they can influence the mechanical properties, like young’s modulus and tensile stress. Effects of different storage protocols have been extensively studied (Stemper et al. 2007; Virues Delgadillo et al. 2010; Hemmasizadeh et al. 2013; O’Leary et al. 2014; Caro-Bretelle et al. 2015). To have low change in mechanical properties, measurements of fresh samples should be favored. If storage is necessary, the samples should be kept in the fridge not longer than 24 h. For longer storage periods, freezing of the samples is recommended (Macrae et al. 2016).

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Environmental Conditions

For application-oriented mechanical measurements, it is essential to provide environmental conditions, which are close to the final application. The samples shall be measured in wet condition, with proper fluid, like physiological saline solution (Sacks 2000). Temperature of 37
C should be maintained during the test (Bernal et al. 2011) and for cell-containing tissues, the chemical environment plays also an important role (O’Brien 2011). In short term measurements, it may be appropriate to keep the samples superficially wet. In measurements where the temperature has to be regulated and for tests with a longer timeframe, immersion of the samples is necessary or the use of a humidity chamber is required to maintain a constant, wet surrounding (Tonge et al. 2013). To establish a setup with such environmental conditions can be challenging and problematic for optical deformation measurements. If marker points on the samples are required, e.g., speckles out of ink, the surface has to be dried, but such changes in water content of the tissue can influence the mechanical properties (Lee et al. 2013). Alternatively, the movement of the machine’s cross-head is often used for displacement acquisition and strain calculation. In this case the clamping influence on the soft samples cannot be eliminated and can cause some error. However, it has been reported that this error is comparably low and acceptable in uniaxial tensile test for larger arteries (Tian et al. 2014). A good planning of the setup is therefore essential to keep proper surrounding conditions for the sample without influencing the measurement.

3.4

Preconditioning

Initial loading curves of certain materials can be considerable different to subsequent loading cycles (Hosseini et al. 2014). To reach a steady state and to have a comparable strain history for all samples, a certain number of loading cycles, the preconditioning cycles are applied (Macrae et al. 2016). The internal structure can change with the cycling; therefore, at least 3–10 loading cycles should be conducted on the samples prior testing (Fung 1993). These internal changes of the material are strain and strain rate dependent, which requires preconditioning comparable to the measurement protocol. The transition from the preconditioning cycles to the actual measurement is also critical, and it has to be considered that after a short unloading period significant partial recovery can occur (Hosseini et al. 2014).

4

Fundamental Material Parameters

Basics of material properties in biomaterials are comprehensively published in literature (Fung 1993; Brown 2002; Rectors et al. 2003; Czichos et al. 2006; Macrae et al. 2016). It is important to understand the underlying mechanical principles for the choice of the proper mechanical test and to allow an interpretation of the results. In the following section, an overview of basic mechanical parameters, important for vascular grafts are presented.

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Stress and Strain

The most common parameters to characterize the mechanical properties of biomaterials are stress and strain (Roeder 2013). These values result from monitoring the force required to pull a material apart and the material’s displacement, as a result of this applied force. In basic tests, usually a constant deformation rate is applied. The required force is depending on the specimen’s material properties and its orthogonal cross-section and the tensile length of the specimen. To determine pure material parameters, force and displacement are normalized against geometry, by calculating stress and strain. Stress is the force divided by the orthogonal cross section and strain is defined as the percentage change in length divided by the initial length. In an ideal case, geometrically defined specimen are used, which provide uniform stress distribution over the cross section and provide results regardless of the geometry. In native materials, this can however be challenging and the choice and exact measurement of the sample geometry is critical, because of this direct influence on the stress-strain calculation.

4.2

Tensile Strength, Tensile Strain, Strain at Break

Out of the stress-strain curve, several values can be determined, which help to assess and compare different materials. The tensile strength or ultimate tensile strength is the stress as the highest point in the stress strain curve. This value gives information on the stability of a material. In ring-shaped specimen, the tensile strength can also be used to estimate the burst pressure (Laterreur et al. 2014). Tensile strength however gives no information on the deformation history, or if the material is tough or brittle and should be taken more as theoretical value, as the material is fairly damaged at this point. For example, in small diameter blood vessels, tensile stresses are reported which would correspond to burst pressures of several thousand mmHg (Nerem 2000). These values are also present in native blood vessels and necessary to achieve a certain safety factor, but the load is nonphysiological.Tensile strain is the strain at ultimate tensile strength and gives information about the toughness of a material. Strain at break, when compared with the tensile strain, gives information about how a material fails. If there is less difference between tensile strain and strain at break, a more brittle fracture is present. In a material with ductile fracture, a larger interval between both values is present (Tanzi and Farè 2017).

4.3

Elastic Modulus

The elastic modulus describes the relationship between stress and strain and is measured in N/m2. In common materials, the elastic modulus or Young’s modulus is reported as the slope of the linear elastic region of the stress strain curve. In materials with a nonlinear stress-strain behavior, this calculation can be problematic.

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Especially in blood vessels and their substitutes, the working range is in the nonlinear region (Fung 1993). In that case, the exact determination point of the elastic modulus has to be stated in the specification. In literature it is distinguished between different elastic moduli at different zones in the stress-strain curve (Sokolis 2007; Duprey et al. 2010). In some cases the incremental modulus, defined as the differentiate function of the stress–strain relationship, is reported (Thubrikar et al. 2001). Out of inflation measurements the elastic modulus can be reported as pressure-strain elastic modulus (Greenwald and Berry 2000), as a measure of its effectiveness as an elastic reservoir. To evaluate the elastic behavior as it would be present in vivo, the elastic modulus should be determined within the physiologic load range (Duprey et al. 2010).

4.4

Compliance

Compliance is the ability of a prosthesis to elastically expand and contract in the circumferential direction in response to a pulsatile pressure. To express compliance as a measure of elasticity, the percent change in diameter at a change in pressure is stated. Compliance is not a pure material property but it represents a more practical value, to allow a better prediction of the vessels in vivo behavior. Compliance can be measured either in pressure inflation tests, with recording the vessels diameter depending on the inner pressure, or it can be measured in ring tensile test. In ring tensile tests, the measured force has to be transferred into a corresponding inner pressure. This requires the consideration of LaPlace’s law, where the influence of the diameter change is included (Laterreur et al. 2014). According to ISO 7198 standard (International Standard ISO 7198:2016 2016), the compliance can be calculated as following: Equation 1 Calculation of the circumferential compliance in a tubular structure  C

Rp2 Rp1  % Rp1 ¼  104 100mmHg P2  P1

ð1Þ

where C is the circumferential compliance expressed as a percentage of the radius change per 100 mmHg Rp1 is the internal radius at pressure P1, in mm Rp2 is the internal radius at pressure P2, in mm P1 is the lower pressure value, in mmHg P2 is the higher pressure value, in mmHg

5

Methods for Mechanical Characterization of Vascular Grafts

The following section is presenting mechanical characterization methods with focus on quasi-static tests. Most of these presented methods can also be used for dynamic tests, but dynamic testing is presented in a separate section. It is important to

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distinguish between quasi-static and dynamic measurements, because they provide different types of information. In quasi-static mechanical tests, the loads or displacements are applied at comparably slow rates. Therefore, these measurements are considered as time independent and dynamic effect are assumed as small enough to be negligible. In dynamic measurements the factor time is also being considered in the mechanical behavior, which is especially important for dynamically loaded blood vessels and their substitutes.

5.1

Uniaxial Tensile Test

The uniaxial tensile test is the most common test to measure mechanical properties. Specimens with defined geometry are used, with dumbbell or rectangular shape (Macrae et al. 2016). The sample is mounted between two clamps which are positioned at a defined distance. The clamps are moved apart, usually with a constant speed and the applied load and displacement are recorded. The techniques for these measurements are well described in standards for plastics (ASTM Standard D63802a 2002). Stress and strain can be calculated to normalize against geometry and to get the pure material parameters. Therefore, the geometry, the boundary conditions, and applied load must be exactly known to allow this calculation of material parameters. These tests are usually performed until rupture, the stress-strain curve, tensile stress, and strain at break are common measuring results. Material stiffness can be taken from the curve’s steepness. For many materials, a linear behavior is seen at the beginning of the curve where stress and deformation are linearly related. Considering linear behavior, Hooke’s law can be applied and the elastic modulus is calculated by the slope of the stress-strain curve. In blood vessels and also vascular grafts which mimic their behavior, the physiologic range is however in the nonlinear region. It is therefore necessary to state the exact region in the stress strain curve where the elastic modulus is calculated. It is also important to know that the test speed has an influence on the measured results and comparison is difficult for samples measured at different test speeds. For soft materials, the influence of the gripping is significant. To reduce this influence displacement is measured at a certain gauge length, either mechanically or optically between markings on the specimen. The dumbbell shaped specimen are preferential, but in small vascular tissue the size is limited and requires different sample geometry and sample attachment. When rectangular specimen are cut, the required uniform deformation at the central region cannot be guaranteed. To reduce the influence of the gripping, the grips can be equipped with sand paper, or the specimens are glued on the specimen holder or some suture is used (Ng et al. 2005).

5.2

Uniaxial Hoop Tensile Test

To reduce the gripping problem for small diameter vessels, tensile tests on ring specimen can be an alternative. This test is described in the ISO 7198 standard for tubular vascular prostheses (International Standard ISO 7198:2016 2016) and the standard test method for apparent hoop tensile strength of plastic (ASTM Standard

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D2290-00 2000). The test specimens are gathered by cutting ring samples of the tubular graft via circular cuts. The exact cutting of the ring length and the exact measurement of the wall thickness are essential. The ring is mounted between two parallel pins and stretched in circumferential direction. In standards, a tensile speed of 50–200 mm/min is specified (International Standard ISO 7198:2016 2016). The diameter of the sample can be calculated when the exact distance of the pins is known. The strain is calculated as percent increase related to the initial diameter. This assumes that the strain at the whole ring is uniform, which is only true when bending effects and friction are negligible. This measurement allows also the calculation of the theoretical burst pressure and apparent compliance, which is the percent change in diameter at a change in pressure. This requires the application of LaPlace’s law, which takes into account that a change in diameter causes a change in the load of the vessel’s wall. Of course, the hoop tensile tests can be applied only in circumferential direction and therefore are not applicable to measure the graft’s anisotropic behavior.

5.3

Planar Biaxial Test

For testing the mechanical properties more similar to in vivo conditions, planar biaxial measurements can be conducted. The load is applied independently on each axis, by controlling the force/stretch ratio. The forces and displacement at each axis are measured. Variations of this force/ratio allow a more complete characterization of the biaxial properties. The planar samples are cut in squared or cruciform shape and gripped along their edges. Like in uniaxial tests, the gripping is essential because it has strong influence on the results. Clamps equipped with sandpaper, roughed surface are used, or gripping is improved by using a glue. Alternatively, the samples are attached with thin hooks, which reduces the influence of the boundary. When performing this biaxial tests until rupture the use of clamps and cruciform samples are recommended, because hooks would induce rupture near the gripping points. However, the required amount of tissue is higher.Local deformations are measured by optical methods like digital image correlation, marker tracking, or electronic speckle pattern interferometry. With this deformation and corresponding force measurements, stress and strain values are calculated. However, to fully characterize the three-dimensional behavior of anisotropic material, these planar biaxial tests are not sufficient.

5.4

Tubular Inflation Testing

More closely to in vivo loading are pressure inflation tests. Internal fluid pressure and a force in axial direction are applied to mimic physiologic conditions. The tubular sample is attached at its ends to cannulas, usually secured by a wrapped suture or fixated by glue. Pressure, force, and the vessel’s geometry are recorded. The vessel’s geometry is measured with contactless methods like marker tracking, digital image correlation, or optical measurement of the outer diameter (Soletti et al. 2011).

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Additionally, torsional load can be applied by connecting a rotational motor and torque transducer to one end of the sample. This allows the determination of the shear modulus, fatigue due to three dimensional movements and investigation of torsional buckling (Vorp et al. 1996; Garcia et al. 2013). A challenge for pressure inflation tests is porous grafts, where leakage has to be prevented or compensated (Soletti et al. 2011). Burst pressure measurements are particular dependent on the boundary conditions. Considerable influences of the rate of pressure and the interface between the graft and the pressurizing fluid were reported (Sarkar et al. 2006).

5.5

Membrane Bulge Test

In membrane bulge tests, flat samples are firmly clamped to the system and loaded by fluid or air pressure to form a hemispherical dome. This allows a three dimensional loading close to in vivo conditions and the measurement of anisotropic elastic properties. This method provides low influence of the sample clamping and is therefore advantageous in determining the strength or rupture information. In other biaxial tests, this is not possible because samples are influenced by and fail near the specimen gripping (Marra et al. 2006). Deformation is measured by contactless methods where the radius and curvature of the dome is analyzed or by applying markings or speckles to record local strains. Calculation of local stress is performed by comprising change of the marker distances (Hsu et al. 1995), and use of numerical methods (Marra et al. 2006) and finite element simulations (Romo et al. 2014).

5.6

Kink Diameter Test

Vascular grafts have to resist not only loads by blood pressure, there act also mechanical loads from the surrounding. Depending on the location in the body, the vessels are bent and twisted (Dobrin et al. 2001; Biemans et al. 2002). In kinking tests, the smallest radius of curvature that the graft can withstand without kinking or lumen loss of more than 50% occurs (Bode et al. 2015; Bensch et al. 2016; BrandtWunderlich et al. 2016; Kanapathy et al. 2016). In these tests, the graft should be pressurized at 100 mmHg and the sample is placed on a radius template. The template radius is decreased until slight narrowing or kinking is determined. Alternatively, a loop of the test sample can be formed and the ends are pulled in opposite direction. Thus, the loop is reduced until a kink is observed. The diameter of this loop is reported.

5.7

Strength after Repeated Puncture

This test method determines the strength of a prosthesis after repeated puncture. Samples are repeatedly punctured and then tested for pressurized burst strength or

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tensile strength. These tests are common for measurement of dialysis grafts. However, also for other vascular grafts this test might be important to evaluate how the material behaves at a puncture. Such a locally induced damage can be critical in certain materials, which is also evaluated in tear propagation strength tests.

5.8

Tear Propagation Strength

For thin plastic sheets, the tear propagation strength is a common criterion (ASTM Standard D1938-02 2002), which represents how a crack in the material affects its tearing strength. For tissue material, only few work is published according to such measurements (Purslow 1983; Koop and Lewis 2003; Claramunt et al. 2013) and also no regulations to this are given in the standards for vascular grafts. However, in highly extensible materials the deformation energy of the specimen is significantly greater than the tearing energy (ASTM Standard D1938-02 2002), which underlines the importance of such measurement. Measurements on tear propagation become imminent also when studying formation of aortic dissections (Carson and Roach 1990; Haslach et al. 2018). Tear propagation in vascular grafts is partially covered by suture retention tests.

5.9

Suture Retention Strength

For the determination of suture retention strength, the force which is necessary to pull a suture through a material and causes the wall to fail is measured (International Standard ISO 7198:2016 2016). Rectangular sheets, segments, or tubular vascular samples are mostly used (Nieponice et al. 2008; Konig et al. 2009; Browning et al. 2012; Bergmeister et al. 2013; Wang et al. 2014; Chaparro et al. 2016; Pensalfini et al. 2018). One end is going to be clamped and in the other end, a suture is inserted. A typical suture as in clinical applications should be used. At a distance of 2 mm from the end of the stretched prosthesis the suture is pierced through one wall of the prosthesis to form a half loop. The graft is placed in a tensile testing machine and the suture is pulled, as given in standards, at a rate of 50–200 mm/min (International Standard ISO 7198:2016 2016).

6

Dynamic Measurements on Vascular Grafts

In vivo, vascular grafts are exposed to a surrounding which has a very dynamic behavior. The functionality of vascular grafts at these fluctuating boundary conditions, like the different blood pressures and changes in pressure curve, variations in heart rate, changes in temperature, and movements of the body, has to be guaranteed. Additionally, a high number of loading cycles, a long implantation time for many years, and an aggressive surrounding will cause erosion, material fatigue, as well as desired changes in the material like cell ingrowth, degradation, and remodeling.

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Biologic material and vascular grafts which consists of different substructures, fibers, cells, and fluid have material properties with a considerable viscoelastic behavior and are therefore sensitive to dynamic and time-dependent change of boundary conditions (Valdez-Jasso et al. 2011; Huang and Niklason 2014). Viscoelasticity is a combination of pure elastic material behavior, which is recoverable and a pure permanent viscous deformation. Fung (1993) describes the phenomenon of viscoelasticity which can be observed in different loading situations: When a constant strain is applied to a material, the corresponding stresses in the material decrease with time, which is called stress relaxation. When a constant stress is applied to the material, it will continuously deform, which is called creep behavior. At cyclic loading conditions, differences in the stress-strain relationship at loading and unloading of the sample exist and a hysteresis curve is observed (see Fig. 3). All these behaviors are features of a mechanical property called “viscoelasticity.” Beside of viscoelastic effects, under cyclic loading also fatigue can occur, which is the failure of material at lower stress levels than assumed under quasi-static conditions. To capture these effects and to better understand and predict the material behavior, dynamic measurements including “time” are essential. Measurements with dynamic characteristics can be realized with previously described measurement setups and sample geometries. Samples are usually loaded at physiologic conditions, but also measurements at higher impact are performed to simulate effects of injury and to get information about the safety margin of the material. For a full understanding of viscoelastic behavior, measurements with variations in frequency, strain, temperature, and over a long time period are necessary. Such extensive measurements are important

Viscoelasticity in stress-strain curve

Viscoelastic model

spring represents elasticity stress

energy absorbed

load

energy returned

damper represents viscosity

unload

strain Fig. 3 Viscoelasticity is a combination of pure elastic behavior (like a spring) and pure viscous behavior (like a damper). Due to viscoelasticity, hysteresis between loading and unloading is seen in a stress-strain curve

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to develop constitutive models, which have the goal to create general laws describing the mechanical behavior of materials and structures (Avril and Evans 2017; Chuong and Fung 1983; Maurel et al. 1998; Holzapfel et al. 2000, 2002; Holzapfel 2006; Gasser et al. 2006; Valdez-Jasso et al. 2011; Hosseini et al. 2013).

6.1

Creep/Recovery Tests

In creep tests, the time-dependent deformation in response to a constant stress is measured. Viscoelastic behavior is investigated by applying a stress-step to the test sample which is held for a certain time period. The time period can vary from some minutes to investigate short term creeping, up to hours or weeks for long term creeping test. Measurements can be performed either on ring-specimen (Soffer et al. 2009) or on flat samples in uniaxial (Theron et al. 2010) and biaxial measurements (Zou and Zhang 2011). The creeping curve can be classified in three zones (Wang et al. 1995). At the start, primary creep acts at a rapid rate and slows with time. Secondary creep has a relatively uniform rate, whereas the tertiary creep has an accelerated creep rate and comes active when the material begins to rupture. Creeping is calculated as the ratio between the difference of strain in investigated time period and the strain at starting time. If the applied load is removed a viscoelastic recovery of the material can be observed, which is the recovery phase. The rate of deformation is the creep rate, which is the slope of the line in the creep-strain vs. time curve.

6.2

Stress Relaxation Tests

Stress relaxation tests give information of the time-dependent modulus of materials. The stress relaxation behavior influences the residual stress in the material and describes how the material reduces stress. The specimen is strained to a fixed level for a defined time period, and the decay of stress is monitored. At unloading, the strain recovery can be investigated. During the stress relaxation, the modulus of the material typically decays from an initial value E0, to a final stable value E1 (Obaid et al. 2017). The speed of this process is characterized in terms of a relaxation time constant τ. Which is the time needed for the modulus to decrease to 1/e of the interval between E0 and E1. It is also common to perform stepwise stress relaxation tests (Gauvin et al. 2010; Loy et al. 2018). Incremental strain steps are applied and relaxation profiles are recorded. This allows the determination of the initial modulus and the equilibrium modulus of a material.

6.3

Cyclic Test

To study and predict the elastic behavior of vascular grafts, measurements in a region of interest corresponding to physiological loads are important. Figure 4 shows

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Comparison of dynamic compliance in physiologic range 60 58 56

Load [mN]

54 52 50 48 46 44 42 40 0

20

40

60

Displacement [µm] Rat aorta

electrospun - coarse mesh

ePTFE graft

electrospun - fine mesh

Fig. 4 Result of cyclic tests on small diameter vascular prostheses compared to the native rat aorta – authors unpublished data

authors unpublished data on cyclic tests of small diameter vascular grafts compared to the native rat aorta. Also viscoelastic effects can be captured, which is not possible in quasi-static tests, but affects their results (Fung 1993). Measurements with a periodic cycle of stress and strain are applied or tubular inflation test is performed for the assessment of dynamic compliance, as stated in ISO 7198 Standard (International Standard ISO 7198:2016 2016). The dynamic compliance is measured at applying dynamic pressure on the inside of a tubular test sample (International Standard ISO 7198:2016 2016). The sample is kept at constant tension or at a fixed length. Diameter changes are measured either directly, at multiple sites along the test specimen or by measuring the intraluminal volume and length changes. For biological material however, volume methods which provide an average compliance are inappropriate due to inherent biological variability (International Standard ISO 7198:2016 2016). In principle, prostheses should be tested under conditions that approximate the in vivo environment. An additional advantage of dynamic compliance measurements is the application of a certain number of loading cycles, to study material changes over time.

6.4

Bioreactor Testing

In Tissue Engineered Vascular Grafts, culturing techniques have been established, which mimic in vivo loading conditions (McClure et al. 2012; Diamantouros et al. 2013). Such setups allow at the same time the application of mechanical tests

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to the sample, which provides data over a long time period and shows dynamic response to different loading conditions (Konig et al. 2009). Due to control of fluid flow, the influence of shear stress additionally to transmural pressure and axial loading can be investigated (van Haaften et al. 2018). The combination of cell culturing and mechanical testing gives information on dynamic processes like cellular response, remodeling, or degradation of the vascular graft material (Gleason et al. 2004; Konig et al. 2009).

6.5

Fatigue Testing

Fatigue is the fail of a material after a certain time subjected to constant or repeated load below the fractural stress. The reason of failure can be microcracks in the material or on its surface. In such tests, the number of cycles to failure is determined (L’Heureux et al. 2007; Konig et al. 2009). However, for the assessment of the design safety of a product at different stress levels, a high number of such measurements are required. For vascular grafts, fatigue resistance is also determined in tests under cycling physiological loading for up to several months, to create a representative number of loading cycles (L’Heureux et al. 2007). Such tests can however be only a proof that the material can withstand the applied cyclic loads without dilatation, and a prediction of the further behavior is hardly possible.

6.6

Accelerated Testing

Vascular grafts have to withstand a long-term use with approximately 40-million loading cycles per year (James and Sire 2010). In heart valves, a material is defined as durable, if it withstands more than 200-million loading cycles (International Standard ISO 5840-3 2013; Vaesken et al. 2014). To guarantee the proper function for the whole period of use, the knowledge of how long the implant can withstand a certain stress without failing is important. To simulate aging of the material due to a high number of loading cycles and long time use, accelerated tests are performed, which create significant results within an acceptable time range. To produce such an acceleration, some test variables have to be intensified, for example, temperature raised, the environment changed, or the loading frequency increased (Brown 2002). Accelerated aging techniques are based on the assumption that the chemical reactions involved in the deterioration of materials follow the Arrhenius reaction rate function. This function states that a 10
C increase in temperature of a homogeneous process results in an approximately two times higher rate of a chemical reaction (ASTM Standard F1980-02 2002). Accelerated test conditions are also used to study premature degradation (Vaesken et al. 2014; Takmakov et al. 2017). However, it must be noticed that the necessary changes themselves could induce effects that would not occur in the reality, for example, different internal load at higher test frequency due to viscoelastic behavior (Vaesken et al. 2014).

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Dynamic Mechanical Analysis

Dynamic Mechanical Analysis (DMA) is a rheological method to characterize the viscoelastic behavior of samples at varying stresses, frequencies, and temperature. An advantage compared to conventional stress relaxation and creep methods is that the physiological loading conditions can be more closely replicated (Burton et al. 2017). DMA measurements can be performed for various loading types. Basics of the DMA are well described in David (1999) and Menard and Menard (2017). In brief, during the measurement a sinusoidal force is applied to the sample, which results to a sinusoidal deformation. In materials with a viscoelastic behavior, both elastic (according to an ideal spring) as well as viscous (corresponding an ideal damper) properties are present (ASTM Standard D5026-01 2001). Because of this viscoelastic behavior, there is a shift between the deformation, which is the material response and the force excitation (see Fig. 5). This deviation is defined as the phase shift δ. Out of these measurements the storage modulus E0 (refers to the reversible, elastic portion) and the loss modulus E00 (refers to the irreversible, viscous component) can be determined. The loss factor tan δ results the ratio of loss and storage modulus (tanδ = E00 / E0 ). Generally, the storage modulus represents the rigidity of a material while the loss modulus is a measure of the dynamic energy that is converted into heat. Tan δ characterizes the mechanical damping or internal friction of a viscoelastic system (ISO/DIS 6721-1:2018 2018). A simplified demonstration of loss- and storage modulus on a bouncing ball is given in Fig. 6. DMA Measurements are common for characterization of plastics (ISO/DIS 6721-1:2018 2018), but they are also reported used for cartilage (Fulcher et al. 2009), heart

Elastic Behavior

strain ε stress σ

ε

σ

time

Viscoelastic Behavior

σ

E' time phase angle δ

E''

δ E* ε

E'...storage modulus E''...loss modulus E*...complex modulus

Fig. 5 Viscoelastic behavior at sinusoidal deformation

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Fig. 6 A bouncing ball as demonstration of loss- and storage modulus

Loss Modulus

Storage Modulus

E''

E'

valves (Baxter et al. 2017), vascular structures (Hinnen et al. 2007), and vascular grafts (Nezarati et al. 2014).

7

Additional Test Methods

7.1

Nondestructive Tests

Most of the common mechanical tests require a certain destruction of the measured product. Either when taking a test sample, the final part is getting unusable, or during the measurement the material is damaged. For evaluation of the products quality and to guarantee proper mechanical properties of a final product, nondestructive measurements are required. For the investigation of the mechanical properties in course of the implantation time, also measurements in the living environment without influencing the prosthesis are necessary. Compliance of vascular grafts can be measured in vivo using high-resolution ultrasound imaging (Mugnai et al. 2013). The vessel’s diameter is measured at different locations and simultaneously the blood pressure is recorded. The application of elastography measurements is also an option to asses shear modulus or stiffness of soft tissues in vivo. Basically, a stress in the tissue is generated and the response of the tissue is measured by imaging techniques (Sarvazyan et al. 2011). The measurement method can be performed using diverse physical principles including magnetic resonance imaging (MRI) (Woodrum et al. 2006; Mariappan et al. 2010), ultrasound imaging(De Korte and Van Der Steen 2002; Mahmood et al. 2016; Li and Cao 2017), X-ray imaging, optical (Shah et al. 2017) and acoustic signals. These techniques have however limited resolution (Disney et al. 2018). Another possibility for nondestructive in vitro measurement are instrumented indentation tests, where the indentation depth is measured as a function of indentation force. Different material parameters can be derived from the measured indentation curves, including hardness parameters, elastic properties, and tensile strength (Schindler 2005).

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Nano/Micromechanics Testing

To optimize the mechanical function, knowledge in the mechanisms of microstructural deformations and local strains at different loading conditions is important (Disney et al. 2018). Fiber sliding, deformation, uncrimping, and alignment are some mechanisms which are investigated in inflation and tensile measurements, combined with optical methods (Yu et al. 2018). Measurements on microstructural organization using multiphoton microscopy are also used to establish constitutive models of local mechanical environment within biological tissues and to investigate the relationship between biological responses and mechanical stimuli (Wan et al. 2012).

8

Conclusions

For the development, quality control, and assessment of vascular grafts of vascular grafts, precise and predictive mechanical characterization is essential. Today’s technology offers a wide range of mechanical tests for various applications. Perhaps the most challenging task is however the selection of the proper test method. Standards offer a good basis for the choice of the methods, but at certain development stages, maybe different tests are required. At the development stage, basic material parameters are important but the behavior of a material in the final application should be also predicted and the success guaranteed. To mimic native blood vessels, structural complexity and therefore mechanical complexity is required, which is hard to predict by measurements of common material parameters. On the other hand simple and easy understandable comparative values are desired, which give clear information where eventual improvement is necessary. Measurements determining the viscoelastic behavior of the material are essential to predict in vivo material performance, but of course the parameters are more complex and require deep understanding of the underlying physical principles. To investigate fundamental material behavior, even more sophisticated tests are necessary, which allow the formation of constitutive material models. Unfortunately, a clear statement about the ideal testing methods cannot be given. The presented overview, with numerous relevant references and the established classification according to the measurement purpose, shall be a basis to assess one’s own needs and should help at the selection process to find the most suitable testing methods.

References ASTM Standard D1388-96 (1996) Standard test method for stiffness of fabrics. ASTM International, West Conshohocken ASTM Standard D2290-00 (2000) Standard test method for apparent hoop tensile strength of plastic or reinforced plastic pipe by split disk method. ASTM International, West Conshohocken. https://doi.org/10.1520/D2290-08.2

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ASTM Standard D3039-00 (2000) Standard test method for tensile properties of polymer matrix composite materials. ASTM International, West Conshohocken. https://doi.org/10.1520/ D3039_D3039M-08 ASTM Standard D2290-01 (2001) Standard test methods for tensile, compressive, and flexural creep and creep-rupture of plastics. ASTM International, West Conshohocken. https://doi.org/ 10.1520/D2990-09.2 ASTM Standard D5026-01 (2001) Standard test method for plastics: dynamic mechanical properties: in tension. ASTM International, West Conshohocken. https://doi.org/10.1520/D502606R14.1 ASTM Standard D1938-02 (2002) Tear-propagation resistance of plastic film and thin sheeting by a single-tear method. ASTM International, West Conshohocken. https://doi.org/10.1520/ D1922-15.2 ASTM Standard D638-02a (2002) Standard test method for tensile properties of plastics. ASTM International, West Conshohocken. https://doi.org/10.1520/D0638-14.1 ASTM Standard D882-02 (2002) Standard test method for tensile properties of thin plastic sheeting. ASTM International, West Conshohocken. https://doi.org/10.1520/D0882-12 ASTM Standard F1980-02 (2002) Standard guide for accelerated aging of sterile medical device packages. ASTM International, West Conshohocken. https://doi.org/10.1520/D7102-10. Copyright ASTM Standard F2150-13 (2013) Standard guide for characterization and testing of biomaterial scaffolds used in tissue-engineered medical products. ASTM International, West Conshohocken. https://doi.org/10.1520/F2150-07.Copyright Avril S, Evans S (2017) Material Parameter Identification and Inverse Problems in Soft Tissue Biomechanics [Internet]. Springer International Publishing, pp 37–67. Available from: https:// link.springer.com/book/10.1007%2F978-3-319-45071-1 Awad NK, Niu H, Ali U et al (2018) Electrospun fibrous scaffolds for small-diameter blood vessels: a review. Membranes (Basel) 8:1–26. https://doi.org/10.3390/membranes8010015 Bäck M, Gasser TC, Michel J-B, Caligiuri G (2013) Biomechanical factors in the biology of aortic wall and aortic valve diseases. https://doi.org/10.1093/cvr/cvt040 Baxter J, Buchan KG, Espino DM (2017) Viscoelastic properties of mitral valve leaflets: an analysis of regional variation and frequency-dependency. Proc Inst Mech Eng Part H J Eng Med 231:938–944. https://doi.org/10.1177/0954411917719741 Bensch M, Müller M, Bode M, Glasmacher B (2016) Automation of a test bench for accessing the bendability of electrospun vascular grafts. Curr Dir Biomed Eng 2:3–7. https://doi.org/10.1515/ cdbme-2016-0068 Bergmeister H, Schreiber C, Grasl C et al (2013) Healing characteristics of electrospun polyurethane grafts with various porosities. Acta Biomater 9:6032–6040. https://doi.org/10.1016/j. actbio.2012.12.009 Bernal M, Urban MW, Rosario D et al (2011) Measurement of biaxial mechanical properties of soft tubes and arteries using piezoelectric elements and sonometry. Phys Med Biol 56:3371–3386. https://doi.org/10.1088/0031-9155/56/11/012 Biemans JMA, Van Den Berg JC, Moll FL (2002) Beware the kinking bypass – a hidden cause of graft reocclusion. EJVES Extra 3:88–90. https://doi.org/10.1053/ejvx.2002.0145 Bode M, Mueller M, Zernetsch H, Glasmacher B (2015) Electrospun vascular grafts with antikinking properties. Curr Dir Biomed Eng 1:524–528. https://doi.org/10.1515/cdbme-2015-0125 Brandt-Wunderlich C, Schwerdt C, Behrens P et al (2016) A method to determine the kink resistance of stents and stent delivery systems according to international standards. Curr Dir Biomed Eng 2:289–292. https://doi.org/10.1515/cdbme-2016-0064 Brown RP (2002) Handbook of polymer testing: short-term mechanical tests. Rapra Technology, Shrewsbury Browning MB, Dempsey D, Guiza V et al (2012) Multilayer vascular grafts based on collagenmimetic proteins. Acta Biomater 8:1010–1021. https://doi.org/10.1016/j.actbio.2011.11.015

Mechanical Testing of Vascular Grafts

57

Burton HE, Freij JM, Espino DM (2017) Dynamic viscoelasticity and surface properties of porcine left anterior descending coronary arteries. Cardiovasc Eng Technol 8:41–56. https://doi.org/ 10.1007/s13239-016-0288-4 Caro-Bretelle AS, Gountsop PN, Ienny P et al (2015) Effect of sample preservation on stress softening and permanent set of porcine skin. J Biomech 48:3135–3141. https://doi.org/10.1016/ j.jbiomech.2015.07.014 Carson MW, Roach MR (1990) The strength of the aortic media and its role in the propagation of aortic dissection. J Biomech 23:579–588. https://doi.org/10.1016/0021-9290(90)90050-D Chaparro FJ, Matusicky ME, Allen MJ, Lannutti JJ (2016) Biomimetic microstructural reorganization during suture retention strength evaluation of electrospun vascular scaffolds. J Biomed Mater Res Part B Appl Biomater 104:1525–1534. https://doi.org/10.1002/jbm.b.33493 Chirinos JA (2012) Arterial stiffness: basic concepts and measurement techniques. J Cardiovasc Transl Res 5:243–255. https://doi.org/10.1007/s12265-012-9359-6 Chuong CJ, Fung YC (1983) Three-dimensional stress distribution in arteries. J Biomech Eng 105:268. https://doi.org/10.1115/1.3138417 Claramunt R, Álvarez-Ayuso L, García-Páez JM et al (2013) Changes in the mechanical properties of chemically treated bovine pericardium after a short uniaxial cyclic test. Artif Organs 37:183–188. https://doi.org/10.1111/j.1525-1594.2012.01538.x Collins MJ, Li X, Lv W et al (2012) Therapeutic strategies to combat neointimal hyperplasia in vascular grafts. Expert Rev Cardiovasc Ther 10:635–647. https://doi.org/10.1586/erc.12.33 Czichos H, Saito T, Smith LR (2006) Springer handbook of materials measurement methods. Springer, New York David SJ (1999) Dynamic mechanical analysis of polymeric systems of pharmaceutical and biomedical significance. Int J Pharm 179:167–178. https://doi.org/10.1016/S0378-5173(98) 00337-8 de Gelidi S, Tozzi G, Bucchi A (2017) The effect of thickness measurement on numerical arterial models. Mater Sci Eng C 76:1205–1215. https://doi.org/10.1016/j.msec.2017.02.123 De Korte CL, Van Der Steen AFW (2002) Intravascular ultrasound elastography: an overview. Ultrasonics 40:859–865. https://doi.org/10.1016/S0041-624X(02)00227-5 Diamantouros SE, Hurtado-Aguilar LG, Schmitz-Rode T et al (2013) Pulsatile perfusion bioreactor system for durability testing and compliance estimation of tissue engineered vascular grafts. Ann Biomed Eng 41:1979–1989. https://doi.org/10.1007/s10439-013-0823-5 Disney CM, Lee PD, Hoyland JA et al (2018) A review of techniques for visualising soft tissue microstructure deformation and quantifying strain Ex Vivo. J Microsc 00:1–15. https://doi.org/ 10.1111/jmi.12701 Dobrin PB, Hodgett D, Canfield T, Mrkvicka R (2001) Mechanical determinants of graft kinking. Ann Vasc Surg 15:343–349. https://doi.org/10.1007/s10016001 Duprey A, Khanafer K, Schlicht M et al (2010) In vitro characterisation of physiological and maximum elastic Modulus of ascending thoracic aortic aneurysms using uniaxial tensile testing. Eur J Vasc Endovasc Surg 39:700–707. https://doi.org/10.1016/j.ejvs.2010.02.015 Fulcher GR, Hukins DWL, Shepherd DET (2009) Viscoelastic properties of bovine articular cartilage attached to subchondral bone at high frequencies. BMC Musculoskelet Disord 10:1–7. https://doi.org/10.1186/1471-2474-10-61 Fung YC (Yuan-cheng) (1993) Biomechanics: mechanical properties of living tissues. Springer, New York Garcia JR, Lamm SD, Han HC (2013) Twist buckling behavior of arteries. Biomech Model Mechanobiol 12:915–927. https://doi.org/10.1007/s10237-012-0453-0 Gasser TC, Ogden RW, Holzapfel GA (2006) Hyperelastic modelling of arterial layers with distributed collagen fibre orientations. J R Soc Interface 3:15–35. https://doi.org/10.1098/ rsif.2005.0073 Gauvin R, Sc M, Ahsan T et al (2010) A novel single-step self-assembly approach for the fabrication of tissue-engineered. Tissue Eng Part A 16:1737–1747

58

M. Stoiber et al.

Gleason RL, Gray SP, Wilson E, Humphrey JD (2004) A multiaxial computer-controlled organ culture and biomechanical device for mouse carotid arteries. J Biomech Eng 126:787. https:// doi.org/10.1115/1.1824130 Greenwald SE, Berry CL (2000) Improving vascular grafts: the importance of mechanical and haemodynamic properties. J Pathol 190:292–299. https://doi.org/10.1002/(SICI)1096-9896 (200002)190:33.0.CO;2-S Haslach HW, Siddiqui A, Weerasooriya A et al (2018) Fracture mechanics of shear crack propagation and dissection in the healthy bovine descending aortic media. Acta Biomater. https://doi. org/10.1016/j.actbio.2017.12.027 Hemmasizadeh A, Darvish K, Autieri M (2013) Characterization of changes to the mechanical properties of arteries due to cold storage using nanoindentation tests. Ann Biomed Eng 40:1434–1442. https://doi.org/10.1007/s10439-011-0506-z.Characterization Hinnen JW, Rixen DJ, Koning OHJ et al (2007) Development of fibrinous thrombus analogue for in-vitro abdominal aortic aneurysm studies. J Biomech 40:289–295. https://doi.org/10.1016/j. jbiomech.2006.01.010 Holzapfel GA (2006) Determination of material models for arterial walls from uniaxial extension tests and histological structure. J Theor Biol 238:290–302. https://doi.org/10.1016/j. jtbi.2005.05.006 Holzapfel GA, Ogden RW (2010) Constitutive modelling of arteries. Proc R Soc A Math Phys Eng Sci 466:1551–1597. https://doi.org/10.1098/rspa.2010.0058 Holzapfel GA, Gasser TC, Ogden RW (2000) A new constitutive framework for arterial wall mechanics and a comparative study of material models. J Elast 61:1–48. https://doi.org/10.1023/ A:1010835316564 Holzapfel GA, Gasser TC, Stadler M (2002) A structural model for the viscoelastic behavior of arterial walls: continuum formulation and finite element analysis. Eur J Mech A Solids 21:441–463 Hosseini SM, Wilson W, Ito K, van Donkelaar CC (2013) How preconditioning affects the measurement of poro-viscoelastic mechanical properties in biological tissues. Biomech Model Mechanobiol. https://doi.org/10.1007/s10237-013-0511-2 Hosseini SM, Wilson W, Ito K, Van Donkelaar CC (2014) How preconditioning affects the measurement of poro-viscoelastic mechanical properties in biological tissues. Biomech Model Mechanobiol 13:503–513. https://doi.org/10.1007/s10237-013-0511-2 Hsu FPK, Liu AMC, Downs J et al (1995) A triplane video-based experimental system for studying axisymmetrically inflated biomembranes. IEEE Trans Biomed Eng 42:442–450. https://doi.org/ 10.1109/10.376148 Huang AH, Niklason LE (2014) Engineering of arteries in vitro. Cell Mol Life Sci 71:2103–2118. https://doi.org/10.1007/s00018-013-1546-3 International Standard ISO 5840-3 (2013) Cardiovascular implants – cardiac valve prostheses – part 3: heart valve substitutes implanted by transcatheter techniques. International Organization for Standardization, Geneva International Standard ISO 7198:2016 (2016) Cardiovascular implants – tubular vascular prostheses. International Organization Standards, Genève ISO/DIS 6721-1:2018 (2018) Plastics – determination of dynamic mechanical properties – part 1: general principles James BA, Sire RA (2010) Fatigue-life assessment and validation techniques for metallic vascular implants. Biomaterials 31:181–186. https://doi.org/10.1016/j.biomaterials.2009.10.028 Kanapathy M, Kalaskar D, Mosahebi A, Seifalian AM (2016) Development of a tissue-engineered lymphatic graft using nanocomposite polymer for the treatment of secondary lymphedema. Artif Organs 40:E1–E11. https://doi.org/10.1111/aor.12604 Konig G, McAllister TN, Dusserre N et al (2009) Mechanical properties of completely autologous human tissue engineered blood vessels compared to human saphenous vein and mammary artery. Biomaterials 30:1542–1550. https://doi.org/10.1016/j.biomaterials.2008.11.011 Koop BE, Lewis JL (2003) A model of fracture testing of soft viscoelastic tissues. J Biomech 36:605–608. https://doi.org/10.1016/S0021-9290(02)00435-9

Mechanical Testing of Vascular Grafts

59

L’Heureux N, Dusserre N, Marini A et al (2007) Technology insight: the evolution of tissueengineered vascular grafts – from research to clinical practice. Nat Clin Pract Cardiovasc Med 4:389–395. https://doi.org/10.1038/ncpcardio0930 Laterreur V, Ruel J, Auger FA et al (2014) Comparison of the direct burst pressure and the ring tensile test methods for mechanical characterization of tissue-engineered vascular substitutes. J Mech Behav Biomed Mater 34:253–263. https://doi.org/10.1016/j.jmbbm.2014.02.017 Lee J-H, Lee S-S, Chang J-D et al (2013) A novel method for the accurate evaluation of Poisson’s ratio of soft polymer materials. Sci World J 2013:1–7. https://doi.org/10.1155/2013/930798 Li GY, Cao Y (2017) Mechanics of ultrasound elastography. Proc R Soc A Math Phys Eng Sci 473. https://doi.org/10.1098/rspa.2016.0841 Loy C, Pezzoli D, Candiani G, Mantovani D (2018) A cost-effective culture system for the in vitro assembly, maturation, and stimulation of advanced multilayered multiculture tubular tissue models. Biotechnol J 13:1–9. https://doi.org/10.1002/biot.201700359 Macrae RA, Miller K, Doyle BJ (2016) Methods in mechanical testing of arterial tissue: a review. Strain 52:380–399. https://doi.org/10.1111/str.12183 Mahmood B, Ewertsen C, Carlsen J, Nielsen M (2016) Ultrasound vascular elastography as a tool for assessing atherosclerotic plaques – a systematic literature review. Ultrasound Int Open 02: E106–E112. https://doi.org/10.1055/s-0042-115564 Mariappan YK, Glaser KJ, Ehman RL (2010) Magnetic resonance elastography: a review. Clin Anat 23:497–511. https://doi.org/10.1002/ca.21006 Marra SP, Kennedy FE, Kinkaid JN, Fillinger MF (2006) Elastic and rupture properties of porcine aortic tissue measured using inflation testing. Cardiovasc Eng 6:123–131. https://doi.org/ 10.1007/s10558-006-9021-5 Maurel W, Thalmann D, Wu Y et al (1998) Biomechanical models for soft tissue simulation. Springer, Berlin McClure MJ, Simpson DG, Bowlin GL (2012) Tri-layered vascular grafts composed of polycaprolactone, elastin, collagen, and silk: optimization of graft properties. J Mech Behav Biomed Mater 10:48–61. https://doi.org/10.1016/j.jmbbm.2012.02.026 Menard KP, Menard N (2017) Dynamic mechanical analysis. Encycl Anal Chem Online 2–25. https://doi.org/10.1002/9780470027318.a2007.pub3 Mugnai D, Tille JC, Mrówczyński W et al (2013) Experimental noninferiority trial of synthetic small-caliber biodegradable versus stable vascular grafts. J Thorac Cardiovasc Surg 146. https:// doi.org/10.1016/j.jtcvs.2012.09.054 Natasha G, Tan A, Gundogan B et al (2015) Tissue engineering vascular grafts a fortiori: looking back and going forward. Expert Opin Biol Ther 15:231–244. https://doi.org/10.1517/ 14712598.2015.980234 Nerem RM (2000) Tissue engineering a blood vessel substitute: the role of biomechanics. Yonsei Med J 41:735–739 Nezarati RM, Eifert MB, Dempsey DK, Cosgriff-Hernandez E (2014) Electrospun vascular grafts with improved compliance matching to native vessels. J Biomed Mater Res B Appl Biomater:1–11. https://doi.org/10.1002/jbm.b.33201 Ng BH, Chou SM, Krishna V (2005) The influence of gripping techniques on the tensile properties of tendons. Proc Inst Mech Eng Part H J Eng Med 219:349–354. https://doi.org/10.1243/ 095441105X34239 Nieponice A, Soletti L, Guan J et al (2008) Rotational vacuum seeding technique. Glass 29:825–833. https://doi.org/10.1016/j.biomaterials.2007.10.044.Development O’Brien FJ (2011) Biomaterials & scaffolds for tissue engineering. Mater Today 14:88–95. https:// doi.org/10.1016/S1369-7021(11)70058-X O’Leary SA, Doyle BJ, McGloughlin TM (2014) The impact of long term freezing on the mechanical properties of porcine aortic tissue. J Mech Behav Biomed Mater 37:165–173. https://doi.org/10.1016/j.jmbbm.2014.04.015 Obaid N, Kortschot MT, Sain M (2017) Understanding the stress relaxation behavior of polymers reinforced with short elastic fibers. Materials (Basel) 10:1–15. https://doi.org/10.3390/ ma10050472

60

M. Stoiber et al.

Pensalfini M, Meneghello S, Lintas V et al (2018) The suture retention test, revisited and revised. J Mech Behav Biomed Mater 77:711–717. https://doi.org/10.1016/j.jmbbm.2017.08.021 Peterson DR, Bronzino JD (2008) Biomechanics: principles and applications. CRC Press, Boca Raton Purslow PP (1983) Measurement of the fracture toughness of extensible connective tissues. J Mater Sci 18:3591–3598. https://doi.org/10.1007/BF00540731 Rectors T, Madrid MGV, General TS et al (2003) Biomechanics of soft tissue in cardiovascular systems. Springer, Wien Roeder RK (2013) Mechanical characterization of biomaterials. Elsevier, Amsterdam Romo A, Badel P, Duprey A et al (2014) In vitro analysis of localized aneurysm rupture. J Biomech 47:607–616. https://doi.org/10.1016/j.jbiomech.2013.12.012 Sacks MS (2000) Biaxial mechanical evaluation of planar biological materials. In: Cardiovascular soft tissue mechanics. Kluwer Academic Publishers, Dordrecht, pp 199–246 Sarkar S, Hillery C, Seifalian A, Hamilton G (2006) Critical parameter of burst pressure measurement in development of bypass grafts is highly dependent on methodology used. J Vasc Surg 44:846–852. https://doi.org/10.1016/j.jvs.2006.07.023 Sarvazyan A, Hall TJ, Urban W et al (2011) An overview of elastography-an emerging branch of medical imaging. Curr Med Imaging Rev 7:255–282. https://doi.org/10.2174/ 157340511798038684 Schindler HJ (2005) On quasi-non-destructive strength and toughness testing of elastic-plastic materials. Int J Solids Struct 42:717–725. https://doi.org/10.1016/j.ijsolstr.2004.06.029 Shah RG, Pierce MC, Silver FH (2017) Morphomechanics of dermis-A method for non-destructive testing of collagenous tissues. Skin Res Technol 23:399–406. https://doi.org/10.1111/srt.12349 Soffer L, Wang X, Zhang X et al (2009) Silk-based electrospun tubular scaffolds for tissueengineered vascular grafts. J Biomater Sci Polym 19:653–664. https://doi.org/10.1163/ 156856208784089607.Silk-Based Sokolis DP (2007) Passive mechanical properties and structure of the aorta: segmental analysis. Acta Physiol 190:277–289. https://doi.org/10.1111/j.1748-1716.2006.01661.x Soletti L, Hong Y, Guan J et al (2011) A bi-layered elastomeric scaffold for tissue engineering of small-diameter vascular grafts. Acta Biomater 6:110–122. https://doi.org/10.1016/j. actbio.2009.06.026.A Stemper BD, Yoganandan N, Stineman MR et al (2007) Mechanics of fresh, refrigerated, and frozen arterial tissue. J Surg Res 139:236–242. https://doi.org/10.1016/j.jss.2006.09.001 Stoiber M, Messner B, Grasl C et al (2015) A method for mechanical characterization of small blood vessels and vascular grafts. Exp Mech 55:1591–1595. https://doi.org/10.1007/s11340015-0053-x Sun W, Sacks MS, Scott MJ (2005) Effects of boundary conditions on the estimation of the planar biaxial mechanical properties of soft tissues. J Biomech Eng 127:709. https://doi.org/10.1115/ 1.1933931 Takmakov P, Ruda K, Phillips KS et al (2017) Rapid evaluation of the durability of cortical neural implants using accelerated aging with reactive oxygen species. J Neural Eng 12:1–29. https:// doi.org/10.1088/1741-2560/12/2/026003.Rapid Tanzi MC, Farè S (2017) Characterization of polymeric biomaterials. Woodhead Publishing, Cambridge Theron JP, Knoetze JH, Sanderson RD et al (2010) Modification, crosslinking and reactive electrospinning of a thermoplastic medical polyurethane for vascular graft applications. Acta Biomater 6:2434–2447. https://doi.org/10.1016/j.actbio.2010.01.013 Thubrikar MJ, Labrosse M, Robicsek F et al (2001) Mechanical properties of abdominal aortic aneurysm wall. J Med Eng Technol 25:133–142. https://doi.org/10.1080/0309190011005780 Tian L, Henningsen J, Salick MR et al (2014) Stretch calculated from grip distance accurately approximates mid-specimen stretch in large elastic arteries in uniaxial tensile tests. J Mech Behav Biomed Mater 47:107–113. https://doi.org/10.1016/j.jmbbm.2015.03.016

Mechanical Testing of Vascular Grafts

61

Tonge TK, Murienne BJ, Coudrillier B et al (2013) Minimal preconditioning effects observed for inflation tests of planar tissues. J Biomech Eng 135:114502. https://doi.org/10.1115/1.4025105 Vaesken A, Heim F, Chakfe N (2014) Fiber heart valve prosthesis: influence of the fabric construction parameters on the valve fatigue performances. J Mech Behav Biomed Mater 40:69–74. https://doi.org/10.1016/j.jmbbm.2014.08.015 Valdez-Jasso D, Bia D, Zócalo Y et al (2011) Linear and nonlinear viscoelastic modeling of aorta and carotid pressure-area dynamics under in vivo and ex vivo conditions. Ann Biomed Eng 39:1438–1456. https://doi.org/10.1007/s10439-010-0236-7 van Haaften EE, Wissing TB, Rutten M et al (2018) Decoupling the effect of shear stress and stretch on tissue growth & remodeling in a vascular graft. Tissue Eng Part C Methods 24:ten. TEC.2018.0104. https://doi.org/10.1089/ten.TEC.2018.0104 VanBavel E, Siersma P, Spaan J a E (2003) Elasticity of passive blood vessels: a new concept. Am J Physiol Heart Circ Physiol 285:H1986–H2000. https://doi.org/10.1152/ajpheart.00248.2003 Virues Delgadillo JO, Delorme S, El-Ayoubi R et al (2010) Effect of freezing on the passive mechanical properties of arterial samples. J Biomed Sci Eng 03:645–652. https://doi.org/ 10.4236/jbise.2010.37088 Vorp DA, Severyn DA, Steed DL, Webster MW (1996) A device for the application of cyclic twist and extension on perfused vascular segments. Am J Phys 270:H787–H795 Waldman SD, Lee JM (2005) Effect of sample geometry on the apparent biaxial mechanical behaviour of planar connective tissues. Biomaterials 26:7504–7513. https://doi.org/10.1016/j. biomaterials.2005.05.056 Wan W, Dixon JB, Gleason RL (2012) Constitutive modeling of mouse carotid arteries using experimentally measured microstructural parameters. Biophys J 102:2916–2925. https://doi. org/10.1016/j.bpj.2012.04.035 Wang XT, Ker RF, Alexander RM (1995) Fatigue rupture of wallaby tail tendons. J Exp Biol 198:847–852. https://doi.org/10.1016/J.JBIOMECH.2007.05.005 Wang F, Mohammed A, Li C et al (2014) Degradable/non-degradable polymer composites for insitu tissue engineering small diameter vascular prosthesis application. Biomed Mater Eng 24:2127–2133. https://doi.org/10.3233/BME-141023 Woodrum DA, Romano AJ, Lerman A et al (2006) Vascular wall elasticity measurement by magnetic resonance imaging. Magn Reson Med 56:593–600. https://doi.org/10.1002/ mrm.20991 Yu X, Turcotte R, Seta F, Zhang Y (2018) Micromechanics of elastic lamellae: unravelling the role of structural inhomogeneity in multi-scale arterial mechanics. J R Soc Interface 15:20180492. https://doi.org/10.1098/rsif.2018.0492 Zou Y, Zhang Y (2011) The orthotropic viscoelastic behavior of aortic elastin. Biomech Model Mechanobiol 10:613–625. https://doi.org/10.1007/s10237-010-0260-4

Preclinical In Vivo Assessment of Tissue Engineered Vascular Grafts and Selection of Appropriate Animal Models Helga Bergmeister and Bruno K. Podesser

Contents 1 Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2 Preclinical In-Vivo Assessment of Tissue Engineered Vascular Grafts . . . . . . . . . . . . . . . . . . . . . 2.1 Animal Models for Vascular Graft Testing . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.2 General Considerations for Graft Testing and Selection Criteria for Animal Models . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.3 Ethical Issues and Responsible Animal Use . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2.4 Acquisition of Experimental Animals . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 3 Housing of Animals and Care . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 4 Preoperative Treatment . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5 Surgical Procedure . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6 Postoperative Care . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.1 Observation Period . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.2 Assessment of Graft Patency and Diagnostic Imaging of Prostheses Outcomes . . . . . 6.3 Small Animal Models . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.4 Large Animal Models . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.5 Nonhuman Primates . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.6 Common Implant Locations for Conduit Evaluation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6.7 Large Diameter Implant Locations (>6 mm) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 Pulmonary Artery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7.1 Small Diameter Implant Locations (6 mm)

6.7.1 Aorta Less frequently conduits or patches using TE approaches have been applied in the aortic position in large animal models (Shum-Tim et al. 1999; Ichihara et al. 2015) because of the challenging hemodynamic forces. The descending thoracic (Kajbafzadeh et al. 2016) as well the abdominal aorta have been used as implant locations. Both locations have desirable length and diameter in large animals to implant conduits of human dimensions. The descending thoracic aorta in adult sheep is approximately 23–25-cm long and has a mean diameter of 1.7 cm (Joscht et al. 2016; Wells et al. 1998).

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Pulmonary Artery

The pulmonary artery in large animals has often been utilized to test biodegradable conduits for surgical reconstruction between the right ventricle and the pulmonary artery because of its appropriate length and diameter. Especially (Kampmeier et al. 2017) growing lambs have been extensively used to investigate the performance of tissue-engineered conduits, which are intended for the application as degradable ventricular-pulmonary artery connections in children (Hoerstrup et al. 2006; Shinoka et al. 1998). It has been shown that in sheep with 40 kg bodyweight, the pulmonary artery has a length of approximately 5.5 cm and an average diameter of 2.2 cm (Harper et al. 2001; Mrocki et al. 2018). In awake and anesthetized sheep, the mean pulmonary arterial pressure is approximately 18 mm Hg. TEVGS have also been applied in canines as pulmonary conduits (8 mm diameter) successfully (Matsumura et al. 2013).

7.1

Small Diameter Implant Locations ( 6 mm) (Niklason et al. 1999; Shin’oka et al. 2005; McAllister et al. 2009; Dahl et al. 2011b). However, they have not yet been shown to be usable at small calibers (3 to 4 mm) due to perceived risks of thrombosis, especially for decellularized vessels that have an exposed collagen surface (Kaushal et al. 2001; Cho et al. 2005). To overcome this issue, efforts to establish an endothelium on the luminal surface of acellular TEVGs have had some success. Autologous endothelial progenitor cells (EPCs) or ECs were perfused in fibronectin-coated grafts with an outer diameter of 4 mm for 2 days. All the EPCand EC-seeded TEVGs remained patent for 30 days as an end-to-side anastomosis to the common carotid artery in a porcine implant study, whereas the contralateral control vein grafts and non-seeded TEVGs were patent in only 3/8 and 0/3 implants, respectively (Quint et al. 2011). In the case of EC-seeded canine TEVGs in a canine model, evaluating a total of six carotid and coronary artery bypass grafts, primary patency was 83% at 30 days (Dahl et al. 2011b). Although the introduction of endothelial cells onto the acellular TEVGs greatly improved the patency of smalldiameter vascular grafts, complete EC coverage was never achieved preimplantation. Rather, EC coverage varied widely between grafts, ranging from 5% to 60%, with an average of 35% on sections sampled from grafts before implant (Quint et al. 2011). Another challenge is that EPCs or ECs must be autologous and must be

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freshly prepared before implantation, which are both significant hurdles to largescale clinical implementation of this otherwise off-the-shelf technology. To prevent thrombosis, some researchers have immobilized anticoagulants, such as heparin (Liao et al. 2009; Dimitrievska et al. 2015) or peptides (Mahara et al. 2015), to coat the luminal surface of the decellularized vessel (Nakamura et al. 2016). One study treated acellular collagen vascular grafts derived from decellularized subintestinal mucosa with heparin benzalkonium chloride (HBAC) complex and found that thrombus formation on the HBAC-treated grafts after 90 days was significantly lowered in rabbit carotid artery interposition grafts (Huynh et al. 1999). Hydrophilic spacers, such as poly(ethylene glycol) (PEG), have been shown to effectively enhance AT-III binding activity of heparin and maintain the patency of vascular grafts for 3 months in canine models (Nojiri et al. 1990). Other alternatives to heparin include direct thrombin inhibitors such as hirudin (Hashi et al. 2010), or aptamers with DNA sequences specific to thrombin binding motifs (Pagano et al. 2008), and are available or in development. To prevent platelets from adhering to and activating on the vessel surface, researchers have studied nitric oxide (Smith et al. 1996), glycosaminoglycans (Kito and Matsuda 1996), phosphorylcholine (Jordan et al. 2006), and albumin (Hubbell 1993) to serve as coatings on the luminal surface of synthetic grafts. Achieving a stable antithrombogenic chemical coating may be pivotal for the extension of this “off theshelf” TEVG to small-diameter vascular indications.

3.3

Nonaggressive Intimal Hyperplasia

Intimal hyperplasia is one of the most vexing failure mechanisms for arteriovenous fistulas and arterial bypass grafts. Synthetic arteriovenous grafts suffer from a propensity for occlusion due to thrombosis and intimal hyperplasia (40–60% in the first year) (Helling et al. 1992; Zibari et al. 1997; Cinat et al. 1999; Schild et al. 2008). The Niklason Lab has tested human acellular TEVGs as arteriovenous grafts in a baboon model (Prichard et al. 2011). Less venous intimal hyperplasia was observed in histological sections with arteriovenous TEVGs than with commercial ePTFE grafts. This apparent improvement in the baboon model may be due to the fact that the TEVGs are comparatively noninflammatory and compliant and thereby may reduce two common triggers of intimal hyperplasia in arteriovenous ePTFE grafts. TEVGs that resist substantial intimal hyperplasia could provide a new treatment option for patients with end stage renal disease. Both reduced incidence of failure due to intimal hyperplasia and slowed progression of intimal hyperplasiarelated failure could lead to a decrease in surgical intervention of arteriovenous access grafts. Decreased surgical intervention, in turn, could decrease overall endstage renal disease health-care costs and patient morbidity. Encouragingly, decellularized TEVGs also resist dilatation and calcification in multiple large animal models (dog, pig, and baboon) of arteriovenous access, peripheral artery bypass, and coronary artery bypass (Dahl et al. 2011b; Prichard et al. 2011; Quint et al. 2011, 2012).

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561

Preclinical Studies

In canine models of peripheral and coronary bypass, implanted 3–4-m-diameter canine TEVGs luminally seeded with autologous ECs, integrated well with native vasculature at anastomotic sites, and resisted intimal hyperplasia. In porcine models of carotid bypass, implanted 4-mm-diameter porcine TEVGs luminally seeded with either autologous ECs or endothelial progenitor cells, resisted both clotting and intimal hyperplasia. A baboon arteriovenous model demonstrated the functional capabilities of implanted 6 mm-diameter human TEVGs. In all the TEVGs, there was an infiltration of α-smooth muscle actin-positive cells, and ECs were detected on graft lumens. Importantly, histological examination of the grafts after explant showed no indication of peri-graft fibrosis or calcification. In addition, long-term patency was demonstrated for up to 1 year in the canine model.

5

Clinical Applications

5.1

HAVs for Hemodialysis Vascular Access

Following the successful preclinical work in baboons with decellularized, allogeneic human bioengineered vessels, Humacyte entered into its first Phase I/II clinical study using the HAV as a conduit for hemodialysis access. Two single-arm Phase II trials involving 60 end-stage renal disease (ESRD) patients from six centers in Poland and the USA were reported in The Lancet in 2016 (Lawson et al. 2016). These patients were implanted with HAVs for hemodialysis access and were followed up for a minimum of 12 months and up to 24 months (Fig. 3). HAVs were fabricated in vitro by culturing human vascular smooth muscle cells from deceased organ and tissue donors on biodegradable tubular PGA scaffolds (Dahl et al. 2011b). In these clinical trials (Table 1), HAVs had 63% primary patency (functional access patency without Fig. 3 Implantation of a HAV into a clinical patient

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Table 1 HAV patency and interventions for dialysis in patients with end-stage renal disease (Lawson et al. 2016) Dialysis patients with end-stage renal disease (n = 60 patients pooled from the USA and Poland) Patency Primary Primary Patency Assisted Secondary Follow% Follow-up % Follow-up up time (mos) time (mos) Time Time (mos) Time (mos) (mos) 6 63% 6 73 6 12 28% 12 38 12

%

18

18%

18

30

18

81

24

15

24

25

24

80

Interventions

97 89

Type Thrombectomy

Number 116

Angioplasty

55

Revision Removal or ligation Total procedures Interventions per patient-year

16 2 155 1:89

intervention) and 97% secondary patency (functional patency with or without intervention) at 6 months. At 18 months, they had 18% primary and 81% secondary patencies. In comparison, synthetic ePTFE grafts have an approximate 33% primary and 55% secondary patencies at 18 months. The high secondary patency of the implanted vessels in this study was primarily due to the rarity of terminal vessel loss caused by infection or rethrombosis. These high rates of secondary patency could confer both clinical and economic benefits by reducing the need for new access placement, which is more complex and expensive than rescuing an access already in situ. For one engineered vessel that was partially biopsied at 16 weeks, the implant was repopulated with endothelial and mural cells, indicating re-endothelialization and vessel wall maturation over time (Fig. 4). Moreover, the vessels provided adequate blood flow for hemodialysis and did not display excessive post-cannulation bleeding or aneurysm formation. Another implant specimen that was obtained at 55 weeks showed more extensive host cell repopulation and little inflammatory response (Fig. 4). Although the Lancet results revealed no immune sensitization and a low infection rate as compared to historical values for ePTFE (Akoh and Patel 2010), more work remains to improve primary patency rates and reduce the need for interventions (Song et al. 2018). To adequately assess the efficacy and safety of HAVs in a prospective manner, Humacyte is conducting two Phase III prospective, open-label, double-armed, randomized, multicenter, multinational, trials comparing HAV with ePTFE grafts, and with autogenous fistulas, as conduits for hemodialysis. This is one of the first ever Phase III trials for a bioengineered implantable tissue. Each Phase III trial is designed for several hundred patients, and the follow-up period is up to 2 years. If significant clinical improvements in functional patency and/or infection rates are noted when

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Fig. 4 Human acellular vessel remodeling: Immunoperoxidase histology staining of explant at 16 weeks shows abluminal repopulation by (a) CD68-positive cells; (b) smooth muscle actinpositive cells; and (c) CD31-positive cells on the graft lumen. Hematoxylin counterstain shows nuclei. Immunoperoxidase histological staining of explant at 55 weeks shows (d) few CD68 positive brown cells; (e) extensive smooth muscle actin-positive cell repopulation extending almost throughout the graft wall; and (f) CD31-positive staining on the graft lumen. Histological assessment of midgraft segments resected at 44 weeks because of infected perigraft hematoma shows two previous cannulation sites. H&E staining shows (g) a very recent cannulation site with fresh clot in

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acellular-engineered vessels are compared to ePTFE grafts, this would represent a critical step toward making engineered grafts a valuable new option for dialysis patients.

5.2

HAVs for Peripheral Artery Disease

Humacyte is investigating the safety, tolerability, and patency of HAV for peripheral arterial bypass via an ongoing Phase II clinical trial in the USA and a recently completed Phase II clinical trial in Poland (manuscript in preparation). The patient population for these studies is patients with symptomatic peripheral arterial occlusive disease. The patency and intervention results of the clinical trial showed that all HAVs functioned as intended. Primary unassisted patency was 63.3% at 1 year and 58.1% at 2 years; these rates are slightly lower than those reported in literature (Klinkert et al. 2004). Secondary patency was 84.2% at 1 year and 73.7% at 2 years; these rates are consistent with the historic literature for above-knee femoro-popliteal bypass with vein or ePTFE (Klinkert et al. 2004). Average resting ankle to brachial indices (ABIs) increased to 0.96, and claudication distance improved significantly when compared to baseline assessments; this remained consistent throughout the trial. All subjects shared the benefit of symptomatic relief from rest pain and claudication as well as sustained, significant improvement in ABI from baseline. There was no evidence of immune rejection of the HAV detected from clinical explants or from the 6-month serological assessments. There was a very low rate of overall infection in this study (0.06 per patient year or 2 cases in 34.4 patient years), none of which led to the abandonment or explantation of the HAV.

5.3

HAVs for Peripheral Vascular Trauma

An additional potential application of HAV is for use in vascular injury as a result of peripheral vascular trauma, since vascular reconstruction is often needed to save tissues as a result of arterial damage, laceration, and/or thrombosis. Examples of situations which could produce peripheral vascular trauma with associated arterial injuries include combat and fractures as a result of motor vehicle accidents, gunshot wounds, dog bites, and other types of blunt or penetrating trauma (Helfet et al. 1990; Andrikopoulos et al. 1995; Akingba et al. 2013). The currently available vascular reconstruction methods have several limitations for use in peripheral vascular trauma. For the treatment of combat injuries as a result of multiple limb injuries ä Fig. 4 (continued) the cannulation tract; (h) a healing cannulation site with cell repopulation occurring from the luminal surface. (i) CD31 immunoperoxidase stain shows luminal positive cells in this mid-vessel segment. For all panels, the vessel lumen is at the bottom of the panel. SMA smooth muscle actin, H&E hematoxylin&eosin. (Figure reproduced with permission from Elsevier Limited, The Lancet)

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following improvised explosive device (IED) detonations, harvesting autologous vein is often difficult, risky, or impossible (Holcomb 2011). In addition, because bacteria are often abundant in IED wounds, and bacteria can colonize synthetic grafts, ultimately causing abscesses and/or sepsis, the use of synthetic grafts is often not advised in these situations. Similar limitations exist for the use of autologous vein and synthetic grafts for the treatment of fractures with associated vascular damage (Akingba et al. 2013). Thus, in both military and civilian circumstances, there is a need for a vascular conduit that is readily available that is based on human (as opposed to synthetic) tissue, but does not require collection of vessel or cells from the patient. To investigate the utility of HAV in peripheral vascular trauma, Humacyte is conducting a pilot Phase II prospective, multicenter, single-arm, nonrandomized trial in patients with lower limb vascular trauma which threatens the viability of the leg and who require reconstruction of the superficial femoral or popliteal artery. The goal of this trial is to follow patients for 12 months to evaluate the safety, tolerability, and rate of patency and limb salvage following the use of HAV as a vascular bypass or interposition vessel.

6

Limitations, Challenges, and Perspective

6.1

Resistance to Thrombosis

Despite nearly 30 years of research, there is currently no surface coating capable of reliably preventing long-term thrombosis. Many advancements in nonthrombogenic surface treatments have been proven to be effective in vitro, but there is still a major gap that separates in vitro outcomes from those implemented in vivo. This gap is likely the result of the lack of understanding of the specific mechanisms involved in thrombus formation in vivo due to biomaterials (Jackson 2007). While antithrombogenic modifications with biomolecules ensure that the surfaces are non-thrombogenic in the short term, in situ endothelialization is needed to maintain the long-term non-thrombogenic property of implants without relying on the constant presence of anticoagulation or antiplatelet drugs in the blood circulation (Li and Henry 2011).

6.2

Non-immunogenicity

Although decellularization could largely decrease immunogenicity, there is no decellularization strategy that removed all traces of DNA and cellular proteins from tissue implants. With the rise of gene-editing technology that uses clustered regularly interspaced short palindromic repeats (CRISPR), a new push has emerged for “humanizing” animals to decrease, and one day completely remove, xenograft immunogenicity. Recently, the Yang and Church research labs successfully inactivated all genetic copies of porcine endogenous retroviruses. These retroviruses

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are a risk for integration into the human genome and cause safety concerns for xenotransplantation from porcine sources, but the CRISPR-based inactivation may prevent this from happening (Patience et al. 1997; Niu et al. 2017). If a similar approach can be used to inactivate other immunogenic biomolecules or mutate them to be compatible with humans, non-immunogenic TEVG derived from animal cells or tissues could become a reality.

6.3

Elastin Network

Elastin is a recoil protein that acts like a rubber band, stretching with an artery at each pulse, then contracting and pulling the vessel back to its original diameter, thus making elastin a key factor that determines the compliance of a vessel. An elastin network in engineered arteries would prevent the vessel from dilating in response to the continuous pressures exerted by blood flow in vivo. Cross-link formation between elastin fibers creates an insoluble elastin network, which offers greater compliance for vascular grafts (Mitchell and Niklason 2003). Although tissuebased constructs are generally more compliant than synthetic counterparts, higher compliance that further mimics the properties of native vessels may be preferred over a strictly collagen-based tissue. Until now, elastin fibers have been absent from all tissue-engineered grafts before implantation. In efforts to enhance elastin in engineered tissues, the Niklason Lab took advantage of the fact that miR-29 downregulates the expression of elastin (Zhang et al. 2012). When a miR-29a inhibitor was added to medium used for growth of TEVGs, miR-29a inhibition increased the appearance of elastin “islands,” and the elastin produced was more cross-linked than under control conditions. Through antagonizing the actions of miR-29, increased elastin levels can be promoted in cultured tissues, and such an approach may one day have therapeutic value for conditions such as enhanced elastinolysis or elastin deficiencies. As such, reliable deposition of load-bearing, insoluble elastin would allow better control over the compliance of engineered vessels. However, maintaining the formation of an insoluble and truly functional elastin networks in vitro remains a challenge.

6.4

Cell Sources

Currently, collagen accumulation thickness and mechanical characteristics of TEVGs are largely dependent on the age of the donor as well as donor cell quality. With the advent of personalized medicine due to new advances in stem cell technology, new genetic approaches to modify implanted cells offer the possibility that engineered vessels can be free of immunological rejection. Some cell sources, such as human induced pluripotent stem cell (hiPSC)-derived SMCs, have been shown to generate abundant collagen for vessels within 8–9 weeks of culture (Gui et al. 2016). Adult stem cells and hiPSCs thus have the potential to eliminate the need to harvest cells from patients or donors (Dimitrievska and Niklason 2018). A current challenge

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to this approach, however, is that the suture retention and burst pressure of hiPSCderived TEVGs are only up to 70 g and 500 mm Hg (Gui et al. 2016), respectively. These values are lower than those for native human arteries and likely mean that vessels derived from hiPSCs are not yet suitable for long-term human implantation. Thus, the mechanical properties of these TEVGs need to be improved before they can serve as a viable clinical alternative.

6.5

Shorten Manufacturing Time

Currently it takes approximately 3 months to manufacture TEVGs. This total time derives from the fact that it takes 2 weeks to isolate and proliferate the cells, 1 week to set up the bioreactor culture system, 8 weeks to culture the vessel, and finally another week to decellularize the cultured vessels. Shortening the manufacturing time will lower the costs of TEVGs and make widespread adoption more straightforward.

6.6

Regulatory Perspective

HAVs are considered biological grafts and thus must fulfill strict regulatory requirements to be clinically approved in different countries. The high-throughput production of HAVs also should be cost-effective for successful global implementation of this cutting-edge technology. Once these criteria are met, provided that HAVs have superior qualities to various alternatives, HAVs could displace synthetic vascular grafts as a key option in hemodialysis treatment. Furthermore, HAVs also may have promise for a wider range of vascular bypass procedures. There is work required to determine whether decellularization approaches can be used to provide new options to patients needing vessels for coronary artery bypass procedures and other smallcaliber, low blood flow scenarios such as distal extremity arterial bypass (Chang and Niklason 2017). This makes it even more important for these vessels to be assessed over long periods to ensure safety and efficacy. Successful translation of engineered graft technology to dialysis grafts would be an important proof of principle for future clinical applications.

7

Conclusions

The overall goal of regenerative therapies is to repair or replace damaged tissue with new tissues or organs that biologically mimic and provide similar benefits of native tissues or organs. The recent development of HAVs for when patients either lack suitable autologous tissue or cannot receive synthetic grafts shows promise for providing access in hemodialysis and as a conduit in cases of peripheral arterial surgery.

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Acknowledgments Juan Wang is supported by AHA (17POST33661238). This work was supported by R01 HL127386 (Niklason) and R01 HL128406-01A1 (Dardik) and by an unrestricted research gift from Humacyte Inc. LEN is a founder and shareholder in Humacyte, which is a regenerative medicine company. Humacyte produces engineered blood vessels from allogeneic smooth muscle cells for vascular surgery. LEN’s spouse has equity in Humacyte, and LEN serves on Humacyte’s Board of Directors. LEN is an inventor on patents that are licensed to Humacyte and that produce royalties for LEN. LEN has received an unrestricted research gift to support research in her laboratory at Yale.

References Abbott WM, Megerman J, Hasson JE, L'Italien G, Warnock DF (1987) Effect of compliance mismatch on vascular graft patency. J Vasc Surg 5(2):376–382 Akingba AG, Robinson EA, Jester AL, Rapp BM, Tsai A, Motaganahalli RL, Dalsing MC, Murphy MP (2013) Management of vascular trauma from dog bites. J Vasc Surg 58(5):1346–1352 Akoh JA, Patel N (2010) Infection of hemodialysis arteriovenous grafts. J Vasc Access 11(2):155–158 Albers M, Romiti M, Pereira C, Antonini M, Wulkan M (2004) Meta-analysis of allograft bypass grafting to infrapopliteal arteries. Eur J Vasc Endovasc Surg 28(5):462–472 Albers M, Romiti M, Brochado-Neto FC, De Luccia N, Pereira CAB (2006) Meta-analysis of popliteal-to-distal vein bypass grafts for critical ischemia. J Vasc Surg 43(3):498–503. e492 Andrikopoulos V, Antoniou I, Panoussis P (1995) Arterial injuries associated with lower-extremity fractures. Cardiovasc Surg 3(1):15–18 Beamish JA, He P, Kottke-Marchant K, Marchant RE (2010) Molecular regulation of contractile smooth muscle cell phenotype: implications for vascular tissue engineering. Tissue Eng Part B Rev 16(5):467–491 Benrashid E, McCoy CC, Youngwirth LM, Kim J, Manson RJ, Otto JC, Lawson JH (2016) Tissue engineered vascular grafts: origins, development, and current strategies for clinical application. Methods 99:13–19 Bethesda, M. (2011). Annual data report: atlas of chronic kidney disease and end-stage renal disease in the United States Bitondo JM, Daggett WM, Torchiana DF, Akins CW, Hilgenberg AD, Vlahakes GJ, Madsen JC, MacGillivray TE, Agnihotri AK (2002) Endoscopic versus open saphenous vein harvest: a comparison of postoperative wound complications. Ann Thorac Surg 73(2):523–528 Carver W, Nagpal ML, Nachtigal M, Borg TK, Terracio L (1991) Collagen expression in mechanically stimulated cardiac fibroblasts. Circ Res 69(1):116–122 Chang WG, Niklason LE (2017) A short discourse on vascular tissue engineering. NPJ Regen Med 2(1):1–8. Chard R, Johnson D, Nunn G, Cartmill T (1987) Aorta-coronary bypass grafting with polytetrafluoroethylene conduits. Early and late outcome in eight patients. J Thorac Cardiovasc Surg 94(1):132–134 Cho S-W, Lim SH, Kim I-K, Hong YS, Kim S-S, Yoo KJ, Park H-Y, Jang Y, Chang BC, Choi CY (2005) Small-diameter blood vessels engineered with bone marrow–derived cells. Ann Surg 241(3):506 Cinat ME, Hopkins J, Wilson SE (1999) A prospective evaluation of PTFE graft patency and surveillance techniques in hemodialysis access. Ann Vasc Surg 13(2):191–198 Collins P, Webb C, Chong C, Moat N (2008) Radial artery versus saphenous vein patency (RSVP) trial investigators. Radial artery versus saphenous vein patency randomized trial: five-year angiographic follow-up. Circulation 117(22):2859–2864 Dahl SL, Koh J, Prabhakar V, Niklason LE (2003) Decellularized native and engineered arterial scaffolds for transplantation. Cell Transplant 12(6):659–666 Dahl SL, Rhim C, Song YC, Niklason LE (2007) Mechanical properties and compositions of tissue engineered and native arteries. Ann Biomed Eng 35(3):348–355

Bioengineered Human Acellular Vessels

569

Dahl SL, Vaughn ME, Hu JJ, Driessen NJ, Baaijens FP, Humphrey JD, Niklason LE (2008) A microstructurally motivated model of the mechanical behavior of tissue engineered blood vessels. Ann Biomed Eng 36(11):1782–1792 Dahl SL, Blum JL, Niklason LE (2011a) Bioengineered vascular grafts: can we make them off-theshelf? Trends Cardiovasc Med 21(3):83–89 Dahl SL, Kypson AP, Lawson JH, Blum JL, Strader JT, Li Y, Manson RJ, Tente WE, DiBernardo L, Hensley MT (2011b) Readily available tissue-engineered vascular grafts. Sci Transl Med 3(68):68ra69-68ra69 Das N, Bratby MJ, Shrivastava V, Cornall AJ, Darby CR, Boardman P, Anthony S, Uberoi R (2011) Results of a seven-year, single-Centre experience of the long-term outcomes of bovine ureter grafts used as novel conduits for haemodialysis fistulas. Cardiovasc Intervent Radiol 34(5):958–963 Davies A, Magee T, Baird R, Sheffield E, Horrocks M (1992) Vein compliance: a preoperative indicator of vein morphology and of veins at risk of vascular graft stenosis. Br J Surg 79(10):1019–1021 Dimitrievska S, Niklason LE (2018) Historical perspective and future direction of blood vessel developments. Cold Spring Harb Perspect Med 8(2):a025742 Dimitrievska S, Cai C, Weyers A, Balestrini JL, Lin T, Sundaram S, Hatachi G, Spiegel DA, Kyriakides TR, Miao J, Li G, Niklason LE, Linhardt RJ (2015) Click-coated, heparinized, decellularized vascular grafts. Acta Biomater 13:177–187 Dixon BS, Beck GJ, Vazquez MA, Greenberg A, Delmez JA, Allon M, Dember LM, Himmelfarb J, Gassman JJ, Greene T (2009) Effect of dipyridamole plus aspirin on hemodialysis graft patency. N Engl J Med 360(21):2191–2201 Dohmen P, Liu J, Lembcke A, Konertz W (2003) Reoperation in a Jehovah’s witness: 22 years after aortic allograft reconstruction of the right ventricular outflow tract. Tex Heart Inst J 30(2):146 Gao J, Niklason L, Langer R (1998) Surface hydrolysis of poly (glycolic acid) meshes increases the seeding density of vascular smooth muscle cells. J Biomed Mater Res 42(3):417–424 Gui L, Chan SA, Breuer CK, Niklason LE (2009) Novel utilization of serum in tissue decellularization. Tissue Eng Part C Methods 16(2):173–184 Gui L, Dash BC, Luo J, Qin L, Zhao L, Yamamoto K, Hashimoto T, Wu H, Dardik A, Tellides G, Niklason LE, Qyang Y (2016) Implantable tissue-engineered blood vessels from human induced pluripotent stem cells. Biomaterials 102:120–129 Gupta V, Grande-Allen KJ (2006) Effects of static and cyclic loading in regulating extracellular matrix synthesis by cardiovascular cells. Cardiovasc Res 72(3):375–383 Harlander-Locke M, Jimenez JC, Lawrence PF, Gelabert HA, Derubertis BG, Rigberg DA, Farley SM (2014) Bovine carotid artery (Artegraft) as a hemodialysis access conduit in patients who are poor candidates for native arteriovenous fistulae. Vasc Endovasc Surg 48(7–8):497–502 Harris L, O'Brien-Irr M, Ricotta JJ (2001) Long-term assessment of cryopreserved vein bypass grafting success. J Vasc Surg 33(3):528–532 Hashi CK, Derugin N, Janairo RRR, Lee R, Schultz D, Lotz J, Li S (2010) Antithrombogenic modification of small-diameter microfibrous vascular grafts. Arterioscler Thromb Vasc Biol 30(8):1621–1627 Haskal ZJ, Trerotola S, Dolmatch B, Schuman E, Altman S, Mietling S, Berman S, McLennan G, Trimmer C, Ross J (2010) Stent graft versus balloon angioplasty for failing dialysis-access grafts. N Engl J Med 362(6):494–503 Hatzibaloglou A, Velissaris I, Kaitzis D, Grekas D, Avdelidou A, Kiskinis D (2004) ProCol ® vascular bioprosthesis for vascular access: midterm results. J Vasc Access 5(1):16–18 Helfet D, Howey T, Sanders R, Johansen K (1990) Limb salvage versus amputation. Preliminary results of the mangled extremity severity score. Clin Orthop Relat Res 256:80–86 Helling TS, Nelson PW, Shelton L (1992) A prospective evaluation of plasma-TFE and expanded PTFE grafts for routine and early use as vascular access during hemodialysis. Ann Surg 216(5):596 Holcomb J (2011) Devastating dismounted IED injuries in OEF. Increasing amputation and genital injury rates admitted to LRMC 2009–2010. Presentation to the defense health board. (http://www.

570

J. Wang et al.

health.mil/Reference-Center/Presentations/2011/03/08/Devastating-Dismounted-Improvised-Explo sive-Device-Injuries-in-Operation-Enduring-Freedom). Huang AH, Niklason LE (2014) Engineering of arteries in vitro. Cell Mol Life Sci 71(11):2103–2118 Huang AH, Lee YU, Calle EA, Boyle M, Starcher BC, Humphrey JD, Niklason LE (2015) Design and use of a novel bioreactor for regeneration of biaxially stretched tissue-engineered vessels. Tissue Eng Part C Methods 21(8):841–851 Hubbell JA (1993) Pharmacologic modification of materials. Cardiovasc Pathol 2(3):121–127 Huynh T, Abraham G, Murray J, Brockbank K, Hagen P-O, Sullivan S (1999) Remodeling of an acellular collagen graft into a physiologically responsive neovessel. Nat Biotechnol 17(11):1083 Jackson SP (2007) The growing complexity of platelet aggregation. Blood 109(12):5087–5095 Jordan SW, Faucher KM, Caves JM, Apkarian RP, Rele SS, Sun X-L, Hanson SR, Chaikof EL (2006) Fabrication of a phospholipid membrane-mimetic film on the luminal surface of an ePTFE vascular graft. Biomaterials 27(18):3473–3481 Kaushal S, Amiel GE, Guleserian KJ, Shapira OM, Perry T, Sutherland FW, Rabkin E, Moran AM, Schoen FJ, Atala A (2001) Functional small-diameter neovessels created using endothelial progenitor cells expanded ex vivo. Nat Med 7(9):1035 Kennealey PT, Elias N, Hertl M, Ko DS, Saidi RF, Markmann JF, Smoot EE, Schoenfeld DA, Kawai T (2011) A prospective, randomized comparison of bovine carotid artery and expanded polytetrafluoroethylene for permanent hemodialysis vascular access. J Vasc Surg 53(6):1640–1648 Kingston GT, Darby CR, Roberts IS (2009) The pathology of depopulated bovine ureter xenografts utilized for vascular access in haemodialysis patients. Histopathology 55(2):154–160 Kito H, Matsuda T (1996) Biocompatible coatings for luminal and outer surfaces of small-caliber artificial grafts. J Biomed Mater Res 30(3):321–330 Klinkert P, Post P, Breslau P, Van Bockel J (2004) Saphenous vein versus PTFE for above-knee femoropopliteal bypass. A review of the literature. Eur J Vasc Endovasc Surg 27(4):357–362 Konig G, McAllister TN, Dusserre N, Garrido SA, Iyican C, Marini A, Fiorillo A, Avila H, Wystrychowski W, Zagalski K (2009) Mechanical properties of completely autologous human tissue engineered blood vessels compared to human saphenous vein and mammary artery. Biomaterials 30(8):1542–1550 Laschke MW, Menger MD (2016) Bioengineered vascular grafts off the shelf. Lancet 387(10032):1976–1978 Laub GW, Muralidharan S, Clancy R, Eldredge WJ, Chen C, Adkins MS, Fernandez J, Anderson WA, McGrath LB (1992) Cryopreserved allograft veins as alternative coronary artery bypass conduits: early phase results. Ann Thorac Surg 54(5):826–831 Lawson JH, Glickman MH, Ilzecki M, Jakimowicz T, Jaroszynski A, Peden EK, Pilgrim AJ, Prichard HL, Guziewicz M, Przywara S, Szmidt J, Turek J, Witkiewicz W, Zapotoczny N, Zubilewicz T, Niklason LE (2016) Bioengineered human acellular vessels for dialysis access in patients with end-stage renal disease: two phase 2 single-arm trials. Lancet 387(10032):2026–2034 Leung D, Glagov S, Mathews MB (1976) Cyclic stretching stimulates synthesis of matrix components by arterial smooth muscle cells in vitro. Science 191(4226):475–477 L'Heureux N, Dusserre N, Konig G, Victor B, Keire P, Wight TN, Chronos NA, Kyles AE, Gregory CR, Hoyt G (2006) Human tissue-engineered blood vessels for adult arterial revascularization. Nat Med 12(3):361 Li S, Henry JJ (2011) Nonthrombogenic approaches to cardiovascular bioengineering. Annu Rev Biomed Eng 13:451–475 Liao D, Wang X, Lin PH, Yao Q, Chen C (2009) Covalent linkage of heparin provides a stable anticoagulation surface of decellularized porcine arteries. J Cell Mol Med 13(8b):2736–2743 MacLeod A (2009) NCEPOD report on acute kidney injury—must do better. Lancet 374(9699):1405–1406

Bioengineered Human Acellular Vessels

571

Madden R, Lipkowitz G, Benedetto B, Kurbanov A, Miller M, Bow L (2002) Decellularized cadaver vein allografts used for hemodialysis access do not cause allosensitization or preclude kidney transplantation. Am J Kidney Dis 40(6):1240–1243 Mahara A, Somekawa S, Kobayashi N, Hirano Y, Kimura Y, Fujisato T, Yamaoka T (2015) Tissueengineered acellular small diameter long-bypass grafts with neointima-inducing activity. Biomaterials 58:54–62 Manduz S, Katrancioglu N, Ozker E, Dogan K (2008) Early thrombosis in bovine mesenteric vein grafts after infrainguinal reconstruction. Int J Angiol 17(1):37 Manson RJ, Unger JM, Ali A, Gage SM, Lawson JH (2012) Tissue-engineered vascular grafts: autologous off-the-shelf vascular access? Semin Nephrol 32(6):582–591. Elsevier McAllister TN, Maruszewski M, Garrido SA, Wystrychowski W, Dusserre N, Marini A, Zagalski K, Fiorillo A, Avila H, Manglano X (2009) Effectiveness of haemodialysis access with an autologous tissue-engineered vascular graft: a multicentre cohort study. Lancet 373(9673):1440–1446 Mitchell SL, Niklason LE (2003) Requirements for growing tissue-engineered vascular grafts. Cardiovasc Pathol 12(2):59–64 Naghavi M, Abajobir AA, Abbafati C, Abbas KM, Abd-Allah F, Abera SF, Aboyans V, Adetokunboh O, Afshin A, Agrawal A, Ahmadi A, Ahmed MB, Aichour AN, Aichour MTE, Aichour I, Aiyar S, Alahdab F, Al-Aly Z, Alam K, Alam N, Alam T, Alene KA, Al-Eyadhy A, Ali SD, Alizadeh-Navaei R, Alkaabi JM, Alkerwi AA, Alla F, Allebeck P, Allen C, Al-Raddadi R, Alsharif U, Altirkawi KA, Alvis-Guzman N, Amare AT, Amini E, Ammar W, Amoako YA, Anber N, Andersen HH, Andrei CL, Androudi S, Ansari H, Antonio CAT, Anwari P, Ärnlöv J, Arora M, Artaman A, Aryal KK, Asayesh H, Asgedom SW, Atey TM, Avila-Burgos L, Avokpaho EFG, Awasthi A, Babalola TK, Bacha U, Balakrishnan K, Barac A, Barboza MA, Barker-Collo SL, Barquera S, Barregard L, Barrero LH, Baune BT, Bedi N, Beghi E, Béjot Y, Bekele BB, Bell ML, Bennett JR, Bensenor IM, Berhane A, Bernabé E, Betsu BD, Beuran M, Bhatt S, Biadgilign S, Bienhoff K, Bikbov B, Bisanzio D, Bourne RRA, Breitborde NJK, Bulto LNB, Bumgarner BR, Butt ZA, Cahuana-Hurtado L, Cameron E, Campuzano JC, Car J, Cárdenas R, Carrero JJ, Carter A, Casey DC, Castañeda-Orjuela CA, Catalá-López F, Charlson FJ, Chibueze CE, Chimed-Ochir O, Chisumpa VH, Chitheer AA, Christopher DJ, Ciobanu LG, Cirillo M, Cohen AJ, Colombara D, Cooper C, Cowie BC, Criqui MH, Dandona L, Dandona R, Dargan PI, das Neves J, Davitoiu DV, Davletov K, de Courten B, Defo BK, Degenhardt L, Deiparine S, Deribe K, Deribew A, Dey S, Dicker D, Ding EL, Djalalinia S, Do HP, Doku DT, Douwes-Schultz D, Driscoll TR, Dubey M, Duncan BB, Echko M, El-Khatib ZZ, Ellingsen CL, Enayati A, Ermakov SP, Erskine HE, Eskandarieh S, Esteghamati A, Estep K, Farinha CS e S, Faro A, Farzadfar F, Feigin VL, Fereshtehnejad S-M, Fernandes JC, Ferrari AJ, Feyissa TR, Filip I, Finegold S, Fischer F, Fitzmaurice C, Flaxman AD, Foigt N, Frank T, Fraser M, Fullman N, Fürst T, Furtado JM, Gakidou E, Garcia-Basteiro AL, Gebre T, Gebregergs GB, Gebrehiwot TT, Gebremichael DY, Geleijnse JM, Genova-Maleras R, Gesesew HA, Gething PW, Gillum RF, Giref AZ, Giroud M, Giussani G, Godwin WW, Gold AL, Goldberg EM, Gona PN, Gopalani SV, Gouda HN, Goulart AC, Griswold M, Gupta R, Gupta T, Gupta V, Gupta PC, Haagsma JA, Hafezi-Nejad N, Hailu AD, Hailu GB, Hamadeh RR, Hambisa MT, Hamidi S, Hammami M, Hancock J, Handal AJ, Hankey GJ, Hao Y, Harb HL, Hareri HA, Hassanvand MS, Havmoeller R, Hay SI, He F, Hedayati MT, Henry NJ, Heredia-Pi IB, Herteliu C, Hoek HW, Horino M, Horita N, Hosgood HD, Hostiuc S, Hotez PJ, Hoy DG, Huynh C, Iburg KM, Ikeda C, Ileanu BV, Irenso AA, Irvine CMS, Islam SMS, Jacobsen KH, Jahanmehr N, Jakovljevic MB, Javanbakht M, Jayaraman SP, Jeemon P, Jha V, John D, Johnson CO, Johnson SC, Jonas JB, Jürisson M, Kabir Z, Kadel R, Kahsay A, Kamal R, Karch A, Karimi SM, Karimkhani C, Kasaeian A, Kassaw NA, Kassebaum NJ, Katikireddi SV, Kawakami N, Keiyoro PN, Kemmer L, Kesavachandran CN, Khader YS, Khan EA, Khang Y-H, Khoja ATA, Khosravi MH, Khosravi A, Khubchandani J, Kiadaliri AA, Kieling C, Kievlan D, Kim YJ, Kim D, Kimokoti RW, Kinfu Y, Kissoon N, Kivimaki M, Knudsen AK, Kopec JA, Kosen S, Koul PA, Koyanagi A, Kulikoff XR, Kumar GA, Kumar P, Kutz M, Kyu HH, Lal DK, Lalloo R, Lambert

572

J. Wang et al.

TLN, Lan Q, Lansingh VC, Larsson A, Lee PH, Leigh J, Leung J, Levi M, Li Y, Kappe DL, Liang X, Liben ML, Lim SS, Liu PY, Liu A, Liu Y, Lodha R, Logroscino G, Lorkowski S, Lotufo PA, Lozano R, Lucas TCD, Ma S, Macarayan ERK, Maddison ER, El Razek MMA, Majdan M, Majdzadeh R, Majeed A, Malekzadeh R, Malhotra R, Malta DC, Manguerra H, Manyazewal T, Mapoma CC, Marczak LB, Markos D, Martinez-Raga J, Martins-Melo FR, Martopullo I, McAlinden C, McGaughey M, McGrath JJ, Mehata S, Meier T, Meles KG, Memiah P, Memish ZA, Mengesha MM, Mengistu DT, Menota BG, Mensah GA, Meretoja TJ, Meretoja A, Millear A, Miller TR, Minnig S, Mirarefin M, Mirrakhimov EM, Misganaw A, Mishra SR, Mohamed IA, Mohammad KA, Mohammadi A, Mohammed S, Mokdad AH, Mola GLD, Mollenkopf SK, Molokhia M, Monasta L, Montañez JC, Montico M, Mooney MD, Moradi-Lakeh M, Moraga P, Morawska L, Morozoff C, Morrison SD, Mountjoy-Venning C, Mruts KB, Muller K, Murthy GVS, Musa KI, Nachega JB, Naheed A, Naldi L, Nangia V, Nascimento BR, Nasher JT, Natarajan G, Negoi I, Ngunjiri JW, Nguyen CT, Nguyen QL, Nguyen TH, Nguyen G, Nguyen M, Nichols E, Ningrum DNA, Nong VM, Noubiap JJN, Ogbo FA, Oh I-H, Okoro A, Olagunju AT, Olsen HE, Olusanya BO, Olusanya JO, Ong K, Opio JN, Oren E, Ortiz A, Osman M, Ota E, Pa M, Pacella RE, Pakhale S, Pana A, Panda BK, PandaJonas S, Papachristou C, Park E-K, Patten SB, Patton GC, Paudel D, Paulson K, Pereira DM, Perez-Ruiz F, Perico N, Pervaiz A, Petzold M, Phillips MR, Pigott DM, Pinho C, Plass D, Pletcher MA, Polinder S, Postma MJ, Pourmalek F, Purcell C, Qorbani M, Quintanilla BPA, Radfar A, Rafay A, Rahimi-Movaghar V, Rahman MHU, Rahman M, Rai RK, Ranabhat CL, Rankin Z, Rao PC, Rath GK, Rawaf S, Ray SE, Rehm J, Reiner RC, Reitsma MB, Remuzzi G, Rezaei S, Rezai MS, Rokni MB, Ronfani L, Roshandel G, Roth GA, Rothenbacher D, Ruhago GM, Sa R, Saadat S, Sachdev PS, Sadat N, Safdarian M, Safi S, Safiri S, Sagar R, Sahathevan R, Salama J, Salamati P, Salomon JA, Samy AM, Sanabria JR, Sanchez-Niño MD, Santomauro D, Santos IS, Milicevic MMS, Sartorius B, Satpathy M, Schmidt MI, Schneider IJC, SchulhoferWohl S, Schutte AE, Schwebel DC, Schwendicke F, Sepanlou SG, Servan-Mori EE, Shackelford KA, Shahraz S, Shaikh MA, Shamsipour M, Shamsizadeh M, Sharma J, Sharma R, She J, Sheikhbahaei S, Shey M, Shi P, Shields C, Shigematsu M, Shiri R, Shirude S, Shiue I, Shoman H, Shrime MG, Sigfusdottir ID, Silpakit N, Silva JP, Singh JA, Singh A, Skiadaresi E, Sligar A, Smith DL, Smith A, Smith M, Sobaih BHA, Soneji S, Sorensen RJD, Soriano JB, Sreeramareddy CT, Srinivasan V, Stanaway JD, Stathopoulou V, Steel N, Stein DJ, Steiner C, Steinke S, Stokes MA, Strong M, Strub B, Subart M, Sufiyan MB, Sunguya BF, Sur PJ, Swaminathan S, Sykes BL, Tabarés-Seisdedos R, Tadakamadla SK, Takahashi K, Takala JS, Talongwa RT, Tarawneh MR, Tavakkoli M, Taveira N, Tegegne TK, Tehrani-Banihashemi A, Temsah M-H, Terkawi AS, Thakur JS, Thamsuwan O, Thankappan KR, Thomas KE, Thompson AH, Thomson AJ, Thrift AG, Tobe-Gai R, Topor-Madry R, Torre A, Tortajada M, Towbin JA, Tran BX, Troeger C, Truelsen T, Tsoi D, Tuzcu EM, Tyrovolas S, Ukwaja KN, Undurraga EA, Updike R, Uthman OA, Uzochukwu BSC, van Boven JFM, Vasankari T, Venketasubramanian N, Violante FS, Vlassov VV, Vollset SE, Vos T, Wakayo T, Wallin MT, Wang Y-P, Weiderpass E, Weintraub RG, Weiss DJ, Werdecker A, Westerman R, Whetter B, Whiteford HA, Wijeratne T, Wiysonge CS, Woldeyes BG, Wolfe CDA, Woodbrook R, Workicho A, Xavier D, Xiao Q, Xu G, Yaghoubi M, Yakob B, Yano Y, Yaseri M, Yimam HH, Yonemoto N, Yoon S-J, Yotebieng M, Younis MZ, Zaidi Z, Zaki MES, Zegeye EA, Zenebe ZM, Zerfu TA, Zhang AL, Zhang X, Zipkin B, Zodpey S, Lopez AD, Murray CJL (2017) Global, regional, and national age-sex specific mortality for 264 causes of death, 1980–2016: a systematic analysis for the global burden of disease study 2016. Lancet 390(10100):1151–1210 Nakamura N, Kimura T, Kishida A (2016) Overview of the development, applications, and future perspectives of decellularized tissues and organs. ACS Biomater Sci Eng 3(7):1236–1244 Niklason L, Gao J, Abbott W, Hirschi K, Houser S, Marini R, Langer R (1999) Functional arteries grown in vitro. Science 284(5413):489–493 Niu D, Wei H-J, Lin L, George H, Wang T, Lee I-H, Zhao H-Y, Wang Y, Kan Y, Shrock E (2017) Inactivation of porcine endogenous retrovirus in pigs using CRISPR-Cas9. Science 357(6357):1303–1307 Nojiri C, Park KD, Grainger DW, Jacobs HA, Okano T, Koyanagi H, Kim SW (1990) In vivo nonthrombogenicity of heparin immobilized polymer surfaces. ASAIO Trans 36(3):M168–M172

Bioengineered Human Acellular Vessels

573

Pagano B, Martino L, Randazzo A, Giancola C (2008) Stability and binding properties of a modified thrombin binding aptamer. Biophys J 94(2):562–569 Palumbo R, Niscola P, Calabria S, Fierimonte S, Bevilacqua M, Scaramucci L, Tolu B, De Fabritiis P, Bondanini F (2009) Long-term favorable results by arteriovenous graft with Omniflow II prosthesis for hemodialysis. Nephron Clin Pract 113(2):c76–c80 Patience C, Takeuchi Y, Weiss RA (1997) Infection of human cells by an endogenous retrovirus of pigs. Nat Med 3(3):282–286 Peck M, Gebhart D, Dusserre N, McAllister TN, L’heureux N (2012) The evolution of vascular tissue engineering and current state of the art. Cells Tissues Organs 195(1–2):144–158 Prabhakar V, Grinstaff MW, Alarcon J, Knors C, Solan AK, Niklason LE (2003) Engineering porcine arteries: effects of scaffold modification. J Biomed Mater Res A 67(1):303–311 Prichard HL, Manson RJ, DiBernardo L, Niklason LE, Lawson JH, Dahl SL (2011) An early study on the mechanisms that allow tissue-engineered vascular grafts to resist intimal hyperplasia. J Cardiovasc Transl Res 4(5):674–682 Quint C, Kondo Y, Manson RJ, Lawson JH, Dardik A, Niklason LE (2011) Decellularized tissue-engineered blood vessel as an arterial conduit. Proc Natl Acad Sci U S A 108(22):9214–9219 Quint C, Arief M, Muto A, Dardik A, Niklason LE (2012) Allogeneic human tissue-engineered blood vessel. J Vasc Surg 55(3):790–798 Roh JD, Sawh-Martinez R, Brennan MP, Jay SM, Devine L, Rao DA, Yi T, Mirensky TL, Nalbandian A, Udelsman B (2010) Tissue-engineered vascular grafts transform into mature blood vessels via an inflammation-mediated process of vascular remodeling. Proc Natl Acad Sci 107(10):4669–4674 Roy-Chaudhury P, Kelly BS, Miller MA, Reaves A, Armstrong J, Nanayakkara N, Heffelfinger SC (2001) Venous neointimal hyperplasia in polytetrafluoroethylene dialysis grafts. Kidney Int 59(6):2325–2334 Santo VJ, Dargon PT, Azarbal AF, Liem TK, Mitchell EL, Moneta GL, Landry GJ (2014) Open versus endoscopic great saphenous vein harvest for lower extremity revascularization of critical limb ischemia. J Vasc Surg 59(2):427–434 Schecter AD, Giesen PL, Taby O, Rosenfield C-L, Rossikhina M, Fyfe BS, Kohtz DS, Fallon JT, Nemerson Y, Taubman MB (1997) Tissue factor expression in human arterial smooth muscle cells. TF is present in three cellular pools after growth factor stimulation. J Clin Invest 100(9):2276–2285 Schild AF, Perez E, Gillaspie E, Seaver C, Livingstone J, Thibonnier A (2008) Arteriovenous fistulae vs. arteriovenous grafts: a retrospective review of 1,700 consecutive vascular access cases. J Vasc Access 9(4):231–235 Schinstock CA, Albright RC, Williams AW, Dillon JJ, Bergstralh EJ, Jenson BM, McCarthy JT, Nath KA (2011) Outcomes of arteriovenous fistula creation after the fistula first initiative. Clin J Am Soc Nephrol 6(8):1996–2002 Schmidli J, Savolainen H, Heller G, Widmer M, Then-Schlagau U, Baumgartner I, Carrel T (2004) Bovine mesenteric vein graft (ProCol) in critical limb ischaemia with tissue loss and infection. Eur J Vasc Endovasc Surg 27(3):251–253 Schon D, Blume SW, Niebauer K, Hollenbeak CS, De Lissovoy G (2007) Increasing the use of arteriovenous fistula in hemodialysis: economic benefits and economic barriers. Clin J Am Soc Nephrol 2(2):268–276 Sharp M, Phillips D, Roberts I, Hands L (2004) A cautionary case: the SynerGraft vascular prosthesis. Eur J Vasc Endovasc Surg 27(1):42–44 Shin’oka T, Matsumura G, Hibino N, Naito Y, Watanabe M, Konuma T, Sakamoto T, Nagatsu M, Kurosawa H (2005) Midterm clinical result of tissue-engineered vascular autografts seeded with autologous bone marrow cells. J Thorac Cardiovasc Surg 129(6):1330–1338 Smith DJ, Chakravarthy D, Pulfer S, Simmons ML, Hrabie JA, Citro ML, Saavedra JE, Davies KM, Hutsell TC, Mooradian DL (1996) Nitric oxide-releasing polymers containing the [N (O) NO]group. J Med Chem 39(5):1148–1156 Smith MPW, Spark JI, Patel JV, Border DJ, Oswal D, Lewington AJ (2009) Massive aneurysmal dilatation of a depopulated ureteric hemodialysis xenograft. Hemodial Int 13(1):3–5

574

J. Wang et al.

Solan A, Mitchell S, Moses M, Niklason L (2003) Effect of pulse rate on collagen deposition in the tissue-engineered blood vessel. Tissue Eng 9(4):579–586 Song HG, Rumma RT, Ozaki CK, Edelman ER, Chen CS (2018) Vascular tissue engineering: progress, challenges, and clinical promise. Cell Stem Cell 22(3):340–354 Stegemann JP, Hong H, Nerem RM (2005) Mechanical, biochemical, and extracellular matrix effects on vascular smooth muscle cell phenotype. J Appl Physiol 98(6):2321–2327 Syedain ZH, Weinberg JS, Tranquillo RT (2008) Cyclic distension of fibrin-based tissue constructs: evidence of adaptation during growth of engineered connective tissue. Proc Natl Acad Sci 105(18):6537–6542 Syedain ZH, Meier LA, Bjork JW, Lee A, Tranquillo RT (2011) Implantable arterial grafts from human fibroblasts and fibrin using a multi-graft pulsed flow-stretch bioreactor with noninvasive strength monitoring. Biomaterials 32(3):714–722 System USRD (2010) Atlas of chronic kidney disease & end-stage renal disease in the United States. Saunders, Philadelphia van de Weijer MA, Kruse RR, Schamp K, Zeebregts CJ, Reijnen MM (2015) Morbidity of femoropopliteal bypass surgery. Semin Vasc Surg 28(2):112–121. Elsevier Wystrychowski W, McAllister TN, Zagalski K, Dusserre N, Cierpka L, L'heureux N (2014) First human use of an allogeneic tissue-engineered vascular graft for hemodialysis access. J Vasc Surg 60(5):1353–1357 Zhang P, Huang A, Ferruzzi J, Mecham RP, Starcher BC, Tellides G, Humphrey JD, Giordano FJ, Niklason LE, Sessa WC (2012) Inhibition of microRNA-29 enhances elastin levels in cells haploinsufficient for elastin and in bioengineered vessels–brief report. Arterioscler Thromb Vasc Biol 32(3):756–759 Zibari GB, Gadallah MF, Landreneau M, McMillan R, Bridges RM, Costley K, Work J, McDonald JC (1997) Preoperative vancomycin prophylaxis decreases incidence of postoperative hemodialysis vascular access infections. Am J Kidney Dis 30(3):343–348

Index

A Abdominal aortic aneurysm (AAA), 102 Accelerated testing, 52 Acellularized arteries, 150 Acellular scaffold, 505–512 Acetic acid, 214, 218 Acidic fibroblast growth factor (aFGF), 151 Acids, 214, 218 Acquired vascular disease, 6 Adhesion, 300 Adipose-derived stromal cells (ADSCs), 474 Adsorption, 163 Adventitia, 97, 171, 381, 397 Agarose, 327 Ageing and vascular tissue engineering, 112, 113 Agitation, 217 Air-impedance, 160 Albumin, 169, 170 Alkaline, 219 Alkaline phosphatase (ALP), 443 Allogeneic cells, 305 Allograft, 496 α-actin, 243 α1 and α2 chains of collagen VI, 243 Ammonium hydroxide, 214 Amorphous, 147 Aneurysms, 102, 246 formation, 190, 274 dilatation, 139 and dissection, 102 Angiogenesis, 275, 470–473 Angiopoietin, 471 Animal model, 248, 250–252, 504, 508, 510, 511 Animal model, TEVGs abdominal aorta of small rodents, 84 acquisition of experimental animals, 75 animal housing and management, 75

aorta, 84 AV-shunt, 85 coronary artery, 86 dogs, 80 ethical uses and responsible animal use, 74 graft patency and graft integrity, 77–78 iliac and femoral arteries, 85 implant locations for conduit evaluation, 83–86 large animal models, 80–83 mouse models, 78 non-human primates, 82–83 observation period, 77 pigs, 82 postoperative care, 76 preoperative treatment, 75–76 pulmonary artery, 84 rabbit model, 79–80 rat model, 79 selection criteria and considerations for graft testing, 73–74 sheep and goat, 81–82 small animal models, 78–80 surgical interventions, 76 vascular graft testing, 71–73 Animals, 286 Anisotropic strain, 368 Anisotropy, 37 Anti-calcification, 254 Anticoagulants, 76, 560 Anticoagulation therapy, 76, 438, 456 Anti-Gal antibody, 243 Anti-Gal IgG molecules, 243 Antigen response, 243 Anti-inflammatory properties, 254 Antimicrobial(s), 17–18 effect, 17 properties, 254 Antithrombin III (AT), 176

© Springer Nature Switzerland AG 2020 B. H. Walpoth et al. (eds.), Tissue-Engineered Vascular Grafts, Reference Series in Biomedical Engineering, https://doi.org/10.1007/978-3-030-05336-9

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576 Anti-thrombogenicity, 16 Anti-thrombogenic performance, 16 Anti-thrombotic, 254 coating, 198 Aorta, 227 Aortic stenosis, 438 Aprotinin, 216 Artegraft ®, 246 Arterial graft, 366 Arterial homografts, 65 Arterial revascularization, 8 Arteries, 97 Arterioles, 99 Arteriosclerosis, 190, 191 Arteriovenous graft (AVG), 175, 248, 370 Arteriovenous grafting, 496 Arteriovenous (AV) shunts, 248, 307 Artificial blood vessels, 412 Artificial intelligence, 450 Atheroma, 432 Atherosclerosis (ATH), 6, 100 arterial and ectopic calcification, 432 CABG, 435 co-culture, 440 definition, 432 endochondral bone formation, 432 flow cytometry, 441 imaging techniques, 433 inflammation, 433 inflammatory biomarkers, 436 in vitro models, 440 lipid accumulation, 433 macrophages, 434, 435 MGP, 439, 440 plaques, 432, 439 SMCs, 434, 435 Atherosclerotic mouse models, 72 Atomic force microscopy (AFM), 165 Autografts, 9, 151 Autologous, 272 bone marrow cells, 543 cells, 304, 305 tissue, 65, 287 vessels, 490, 491, 495 B Baboons, 152, 175 Bare metal stents, 7 Bases, 214 Basic fibroblast growth factor (bFGF), 471 Benzonase, 244 β-actin, 222

Index Bioartificial artery (BAA), 367, 368 Biocompatibility, 198 Biodegradable conduits, 306 Biodegradable grafts, 21 Biodegradable polymeric carriers, 472 Biodegradable polymers, 498, 499, 506 Biodegradable scaffolds, 189 Biodegradable tubular scaffolds, 248, 543 Biodegradation, 139 Bioengineered acellular vessels anticoagulant properties, 559 biomechanical control, 556 circumferential and axial stretching, 556 design criteria, 553 mechanical property, 558 scaffold, 554 storage method, 558 Biofabrication, 470 Biohybrid concept, 343 Biohybrid vascular grafts, 352 tissue-engineering, 344–355 vascular design considerations, 343, 344 Bioinks, 323, 324, 327, 329, 335 Bioinspired principles, 357 Biological and biomechanical integrity, 255 Biological and biomechanical performance, of decellularized vascular grafts, 227–246 Biological treatments, 221–223 Biological vascular grafts, 551 Biomacromolecules, 163 Biomaterial, 277 Biomimetic, 353 Biomolecules, incorporation of, 422 Biopolymer gel-based scaffolds, 371 Biopolymer scaffolds, 367 Bioprosthetic grafts autografts, 9 homografts, 9 xenografts, 10 Bioprosthetic heart valve, 438 Bioreactor, 273, 301, 307, 393, 553 testing, 51 Bioresorbable polymers, 507 co-polymers and polymer combinations, 153, 154, 156 PCL, 148, 149 PDO, 149–151 PGA, 147 PGS, 151–153 PLA, 147, 148 Bioresorption, 139 Biorubber, 151 Biotubes, 279

Index Blalock-Taussig shunt (BT shunt), 5 Blood flow assessment, 77 Blood outgrowth endothelial cells (BOECs), 176 Blood vessels, 37, 340 diffusion limits, 322 ECM molecules, 322 transport nutrients and oxygen, 322 stenosis/occlusion, 208 vascular cells, 322 Blow spinning principle, 355 tissue-engineered vascular grafts, 355 Bone marrow cells, 445 Bone marrow-derived mononuclear cells, 543 Bone marrow mononuclear cells, 504 Bone mesenchymal stem cells (BMSCs), 475 Bone morphogenic proteins (BMPs), 437 Bottom-up approach, 356, 357 Bovine carotid artery, 248 Bovine mesenteric vein, 248 Bovine type I collagen, 367 Bovine ureter, 248 Burst pressure, 47, 296, 307

C Ca-alginate, 356 Calcific aortic valve disease (CAVD) animal models, 442 bioprosthetic aortic valve calcification, 441 calcification, 437 cardiac cycle, 436 ECM properties, 436 fibrosis, 437 heart valve replacement, 438 inflammation, 437 in vitro models, 442, 443 in vivo models, 441 stenosis symptoms, 438 TAVR, 439 Calcification, 145, 254 resistance, 457 Calciphylaxis, 435 Calcium hydroxide, 214 Calcium phosphate, 432 Calcium phosphate cement (CPC), 475 Caliber, 21–23 Campbell, G.R., 15 Camptothecin, 223 Canine jugular vein, 250 Cannular tissues, 333

577 Caprine carotid arteries, 254 Carbodiimide, 254 Cardiac myocytes, 99 Cardiovascular calcification animal models, 431 calcium deposition, 430 in vitro and in vivo models, 431 monolayer cell culture, 431 tissue-engineered models, 444–457 valvular diseases, 432–439 vascular diseases, 432–443 Cardiovascular disease (CVD), 64, 138, 208, 340, 490, 491 Carotid interposition graft model, 85 Cell-cell/cell-matrix interactions, 431 Cell-derived matrix, 307 Cell-derived products, 309 Cell digestion, 161 Cell distribution, 301 Cell-free grafts, 413 Cell-free implants, 347 Cell phenotype regulation, 541 Cell scraping, 161 Cell seeding, 297, 536, 537, 543 dynamic techniques, 301–303 efficiency, 297 quasi-static techniques, 298–300 static, 297–298 Cell sources, 566 Cellular ingrowth, 139, 454 Cellularization in vascular grafts, 420 Cellularized autologous tissue, 272 Cellular residues, 243 CHAPS, 215, 220, 226 Chelating agents, 215, 220–221 Chemical decellularization, 224 Chemical decellularization techniques, 517 Chemical treatments, 213–221 Chitosan (CS), 350, 539 Chorionic plate, 226 Chromosomal and plasmid DNA, 219 Chronic graft rejection, 243 Chronic inflammation, 243, 254 Circumferential alignment, 368, 369 Circumferential orientation, 153, 171 Circumferential tensile modulus, 368 Clinically available decellularized grafts, 247 Clinical translation, 345 Clinical trials, 499, 535, 543, 544 Clopidogrel, 175 Coaxial electrospinning, 167, 168 Co-electrospinning, 154, 167 Collagen, 284, 348, 390, 393, 395, 397

578 Commercially available decellularized vascular grafts, 246–247 Completely tissue engineered blood vessel (TEBV), 14–15 Compliance, 44, 344 mismatch, 19, 140, 152 vascular grafts, 18–21 Composite structures, 341 Congenital disease, 5–6 Congenital heart defect, 534 Congenital heart disease (CHD), 492 Congenital heart surgery, 535, 543 Congenital malformations, 492 Contractile phenotype, 159 Controlled fiber deposition, 354 Copolymers, 153, 154, 156, 539, 540, 543 Core/shell morphology, 168 CoreoGraft ®, 248 Coronary arteries, 99 Coronary artery bypass grafts/grafting (CABG), 366, 435 Coronary artery disease, 7–8 Covalent linking, 163 of biomolecules, 162–164 C-reactive protein (CRP), 436 Creep behavior, 49 Creep/recovery tests, 50 Crimping, 456 CRISPR, 457 Crosslinker, 172 Crosslinking, 249, 447 Cryopreserved grafts, 504 Crystalline, 147 Crystallinity, 447 Cyclic compression, 453 Cyclic stretch, 449, 453 Cyclic stretching, 368, 556 Cyclic test, 50, 51

D Dacron ® grafts, 10–11, 20, 142, 346 See also Polyethylene terephthalate 3D bio-fabrication techniques, 478–480 3D bioprinting, 446, 451, 452, 479 2D culture, 444 3D culture, 444 Decellularization, 370, 456, 477, 557 methods, 558 Decellularized biologic tubular structures, 192 Decellularized bioreactor engineered vessel, 190

Index Decellularized blood vessels, 245, 551 Decellularized constructs, 307 Decellularized extracellular matrices (dECMs), 323 Decellularized native tissues, 367 Decellularized natural matrix, 66, 189, 190 Decellularized scaffolds, 254 Decellularized small blood vessels, 226 Decellularized tissue engineered matrix-based grafts, 514–516 Decellularized tissues, 500, 535 Decellularized vascular grafts, 214 acid and alkaline treatments, 218–219 chelating agents, 220–221 combined treatments, 225–227 commercially available decellularized vascular grafts, 246–247 detergent treatments, 219–220 enzymes, 221–222 histological and biochemical properties, 244–245 hypotonic/hypertonic treatments, 213–218 immunogenicity, 227–243 mechanical properties, 245–246 pressure treatments, 224 protease inhibitors, 223 supercritical-CO2 treatments, 224–225 toxins, 222–223 in vivo performance of, 247–254 Decellularized vein allografts, 249 Decellularized vessels, 298, 306 Decellularized xenogeneic umbilical artery, 249 Degradable textiles, 347 Degradation, 371 period, 542 Dermal fibroblasts, 370 Detergent treatments, 219–220 Diabetes and vascular tissue engineering, 114–115 Digital light project (DLP), 480 Digital Light Processing, 195 Diisocyanate, 144 Disinfection, 228–242 Disuccinimidyl suberate (DSS), 158 4D light sheet imaging, 452 DNA, 222 removal, 226 DNase, 216, 221, 222, 243 Double-Raschel knitting machine, 349 3D printing, 190, 450 See also See also Three-dimensional (3D) bioprinting Drug development, 431

Index 3D scaffolds, 470 Dynamic mechanical analysis (DMA), 53 Dynamic seeding, 301–303 E EC-binding peptide, 415 ECM, see Extracellular matrix (ECM) Economic paradigm, 458 Ectopic calcification, 432 Electrostatic, 25 Elastic arteries, 97 Elasticity, 342 Elastic modulus, 43, 218 Elastic properties vascular grafts, 18 Elastin, 173, 280, 285 animal-derived, 395 animal vessels/natural matrices decellularization, 390–392 biochemical cues, 389 biological signaling, 381 bioreactor, 390 cardiac mechanical energy, 383 coating, 392–393 culturing stiffness, 388 decellularized conduits, 392 deposition induction by biological graft, 390 ECM, 381 fiber synthesis, 384 inducing deposition in situ, 389–390 inducing deposition in vitro, 387–389 molding, 393–394 natural vascular system, 381–383 network, 566 production, 388 pulsatile flow, 388 seeding cells, 391 sourcing, 384–386 tropoelastin, 384 vascular ECM, 382 vascular graft, 380–381 ElastinGrafts, 347, 349 Elastin-like peptides (ELP), 397–398 Elastin-like recombinamers (ELR), 347 Elastogenesis, 369, 385 Elastography, 54 Electric field, 158 Electrospinning, 158–160, 193, 194, 447, 536 co-blend, 395–397 cross-linking, 396 electrospun fibers, 394, 396 multiple layers, 397 quality affecting parameters, 394 tropoelastin, 395

579 Electrospun fibers, 351 Electrostatic cell seeding, 162 Electrostatic deposition, 176 Electrostatic force, 300 Embedded extrusion bioprinting, 329–331 Embryoid bodies (EBs), 477 Embryonic stem cells (ESCs), 474, 501 Embryonic vascular system, 471 Emerging technologies, 356 Encapsulation, 298 Endocardium, 99 Endonucleases, 222, 244 Endothelial cell (EC), 99, 139, 154, 157, 161, 168, 299, 322, 367, 444, 470, 497, 499, 501, 503, 516, 535, 541, 542 adhesion, 16 seeding, 161, 162, 197 transplantation, 13 Endothelial denudation, 100 Endothelial dysfunction, 100 Endothelialization, 197, 249–252, 279, 353, 370 of vascular grafts, 414 Endothelial monolayer, 284 Endothelial progenitor cells (EPCs), 414, 474 Endothelial seeding, 14 Endothelial-to-mesenchymal transition (EndoMT), 418 End-stage renal disease (ESRD), 490, 494 Environmental conditions, 42 Enzymatic activity, 220 Enzymatic decellularization, 224 Enzymatic digestion, 161 Enzyme prodrug technique (EPT), 415 Enzymes, 215, 221–222, 246 EPC-specific peptide, 414 Ephrin-B2, 471 ePTFE, see Expanded polytetrafluoroethylene (ePTFE) Equine carotid arteries, 237, 238 1-Ethyl-3-(3-dimethyl aminopropyl)carbodiimide (EDC), 254 Ethylenediaminetetraacetic acid (EDTA), 215, 220, 221 Ethylene glycol-bis(β-aminoethyl ether)-N,N, N0 ,N0 -tetraacetic acid) (EGTA), 215, 220 Expanded polytetrafluoroethylene (ePTFE), 11–12, 24–25, 138, 142, 143, 189, 380, 494, 496, 497 prostheses, 412 External elastic lamina, 97 External stent, 22 External supports, 22

580 Extracardiac total cavopulmonary connection, 543 Extracellular matrix (ECM), 225, 322, 367, 381, 437, 497, 498, 512, 515, 535 decellularization process, 391 elastin signaling, 383 microstructure, 228 in regenerative medicine, 210–212 Extracellular vesicles (EVs), 433, 435 Extracorporeal membrane oxygenators (ECMO), 16 Extrusion bioprinting, 324 F Fabrication methods electrospinning, 158–160 knitting and weaving, 156, 157 sponges, 160, 161 False aneurysm, 102 Fatigue testing, 52 Fiber diameter, 510, 536, 542 Fiber formation, 351 Fibrillogenesis, 367 Fibrin, 174, 346, 389, 447 degradation, 373 gels, 367, 369, 371–373 stabilizing factor, 174 Fibrin-based bioartificial artery, 372 Fibrin-based constructs, 369 Fibrinogen, 174, 372 Fibrinolysis, 373 Fibroblast growth factors (FGF), 198 Fibroblasts, 277, 369 Fibronectin, 163 Fibrous cap, 433, 441 Fibrovascular tissue, 146 First Phase I/II clinical study, 561 Fontan procedure, 493, 496, 503 Foreign body response (FBR), 254, 272, 287 to biomaterials, 274, 275 Freeform, 329 Freeze-thawing, 217 Functional arterial replacements, 551 Functionalization, 249 G α-Gal, 211, 212, 222, 227, 243 Galactose-α-1,3-galactose, 211 Gelatin methacryloyl (GelMA), 327, 331, 333 Gel compaction, 368 Genipin, 254

Index Glutaraldehyde, 254 crosslinking, 254 Glycidal methacrylate-hyaluronic acid (GM-HA), 333 Glycolic acid, 538 Glycosaminoglycans (GAGs), 210, 218–221, 223, 226, 244, 436, 437 Gore-Tex ®, 11–12, 142 Graft dilatation, 21 Graft stenosis, 544 Granular hydrogel, 330 Graphene oxide-copper (GO-Cu), 475 Gravitation cell seeding, 162 Growth factors, 249 Growth/signaling factors, 470–473 H Heart failure, 431 Helical groove patterns, 344 Hemocompatible materials, 353 Hemodialysis, 280, 490, 494, 495, 497, 504 access shunt, 14 vascular access, 561, 562, 564 Hemostasis, 15 Heparin, 168, 176 Hepatocyte growth factor (HGF), 471 Heterologous grafts, 9 High-hydrostatic pressure (HHP), 217, 224 High-sensitivity assay for CRP (hsCRP), 436 High-throughput screening (HTS), 450, 453 Hollow fiber bioprinting, 330–333 Homografts, 9 Homopolymers, 540 HUMACYL ®, 248 Human acellular vessels (HAV), 190, 199, 552 regulatory perspective, 567 Human amniotic membrane, 241, 242 Human aortic endothelial cells (HAECs), 330, 444 Human aortic smooth muscle cells (HASMC), 177 Human cadaver femoral vein, 248 Human clinical trial, 248 Human dermal fibroblasts (hDFs), 175 Human great saphenous vein, 229 Human iPSCs (hiPSCs), 474 Human placenta, 226 Human umbilical arteries, 236, 252 Human umbilical vein endothelial cells (HUVECs), 176, 444, 474 Humoral response, 254 Hydrogel, 300

Index Hydrolysis, 147 Hydrophobicity, 15–16 Hydroproxyline, 244 Hydrostatic cell seeding, 162 Hydroxyapatite, 430, 435, 441, 443, 446 Hydroxyproline, 170 Hyperacute rejection, 243 Hypercholesterolemia, 442 Hyperplasia, 138, 139 Hypertension, 102 Hypertonic solution, 214 Hypotonic/hypertonic treatments, 213–218 Hypotonic solution, 214, 243 Hypoxia, 475 Hypoxia-inducible factor-1 (HIF-1), 475 Hypoxia-inducible factor-1α (HIF-1α), 475

I Ideal vascular graft, 138 Immune suppression, 254 Immunofluorescent staining, 172 Immunogenicity, 227–243 Implant remodeling and regeneration, 254 Incubation, 285 Induced pluripotent stem cells (iPSCs), 192, 335, 445, 457, 474, 502 Inferior vena cava interposition grafts, 543 Inflammation-mediated process, 542 Inflammatory potential, 246 Inflammatory process, 542 Inflammatory reaction, 199 Inflammatory response, 275, 420, 421 Infrainguinal arterial reconstruction, 248 Inkjet bioprinting, 324 In silico approaches, 449 In silico platforms, 117–119 In situ tissue engineering, 107, 413, 505 In-stent restenosis, 415 Insulin, 369 Integrin binding, 170 Intercellular adhesion molecule-1 (ICAM-1), 437 Interconnectivity, 26 Internal elastic lamina (IEL), 97 internodal distance (IND), 12, 24 Intima, 97, 170 Intimal hyperplasia, 4, 19, 100, 139, 190, 286, 366, 560 Intimal thickening, 100 In vitro platforms, 119–121 In vitro tissue engineering, 107

581 In vivo studies, of decellularized vascular grafts, 247–254 In vivo tissue engineered vascular grafts, see Tissue engineered vascular grafts (TEVG) In vivo tissue engineering, 107 Ionic detergents, 219, 226 Iron oxide, 299 Ischemic coronary artery disease, 7–8 Ischemic heart disease (IHD), 432, 491, 492 Ischemic stroke, 432 J Justification for the Use of Statins in Prevention: an Intervention Trial Evaluating Rosuvastatin (JUPITER), 436 K Keutel syndrome, 440 Kink diameter test, 47 Klotho, 440 Knitted textile vascular grafts, 10 Knitting, 156, 157 principle, 346 tissue-engineered vascular grafts, 346–350 L Lamellae, 342 Langer, S., 14 Large animal models, 122, 123 Large diameter grafts, 19–21 Laser-induced forward transfer, 324 Leukocytes, 432 L’Heureux, N., 14 Lipid–lipid interactions, 220 Lipid mediators, 421 Lipid–protein interactions, 220 Lipopolysaccharides, 443 Lipoprotein(a) (Lp(a)), 438 Living fibers, 356 L-lactic acid, 538 Longitudinal orientations, 171 Low-blood-flow vascular replacements, 247 M Machine learning, 450 Macrocalcifications, 434, 457 Macromolecules, 249

582 Macrophage colony-stimulating factor (M-CSF), 433 Macrophage polarization, 420, 421 Macrophages, 148, 161, 275, 433, 435, 445, 537, 542 Magnetic field, 299 Magnetic seeding, 299 Major histocompatibility complex class I (MHC I), 227, 243 Major histocompatibility complex class II (MHC II), 227 Manufacturing time, TEVGs, 567 Matrix GLA protein (MGP), 439 Matrix metalloproteinases (MMPs), 437, 446 Matrix-vesicle, 441 Mechanical decellularization, 518 Mechanical properties, 245–246, 249 Mechanical strain, 19 Mechanical strength, 368 Mechanical tests basic material parameters, 39 classification, 39 fundamental behavior, 40 physiological range, 40 regulations, 41 standards, 40 Mechano-transduction phenomenon, 199 Media, 97, 171 Melt electrospinning, 344 Melt electrowriting (MEW), 344 principle, 354 tissue-engineered vascular grafts, 354, 355 Membrane bulge test, 47 Mesenchymal stem cells (MSCs), 159, 304, 445, 474, 501 Methacrylated gelatin and methacrylated hyaluronic acid (GelMA/HAMA), 446 Mice models, 442 Microcalcifications, 434, 445, 451 Microchannel, 478 Micro-contact printing, 452 Micro-emboli, 21 Microenvironment, 431, 448, 454 Microfluidic technology, 356 Micropatterned PDMS, 167 MicroRNAs (miRNAs), 419 Mineralization, 434, 440, 441, 443, 447, 452, 457 Minor histocompatibility loci, 227 M2 macrophages, 420 Modes of loading, 38 Modulus, 369 Molecular pathways, 286

Index Monkey immune system, 243 Mouse models, 78 Multi-layer concentric printheads, 333 Multi-material bioprinting, 332, 333 Myocardial infarction (MI), 431, 432 Myocardium, 99 Myofibroblast, 197, 277, 437 N Nano/micro-fibers, 193 Nano/micromechanics testing, 55 Narrowing, lumen, 97 Native-like tissue, 343 Natural based polymer scaffolds, 191 Natural biomaterials, 477 Natural materials, 348 Natural polymers, 499, 505 Neoangiogenesis, 149 Neointima, 100, 104, 139 Neointimal hyperplasia, 537, 540, 541 Neotissue, 536, 537, 540–542, 544 Neovascularization, 470, 472, 473 N-glycolylneuraminic acid (Neu5Gc), 211, 212, 222, 227, 243 Niklason, E., 14 Niklason vessel bioreactor, 554 Nitric oxide (NO), 415 Nogo-B, 418 Non-degradable, 347 Nondestructive tests, 54 Non-human primates, 82–83 Non-immunogenicity, 565 Non-immunogenic scaffolds, 255 Non-ionic detergents, 220 Non-resorbable polymers ePTFE, 142, 143 PET, 142 PU, 143–146 Non-woven, 351 NO-releasing grafts, 415 No-touch aortic techniques, 8 No-touch versus endoscopic vein harvesting, 8 Nucleases, 218 O Obstruction, lumen, 97 Off-the-shelf, 495, 505, 510, 515, 516, 518, 540 available vascular grafts, 422 products, 554 Optical coherence tomography (OCT), 452 Osteogenesis, 434

Index Osteogenic media (OM), 443 Ovine aorta, 232 Ovine carotid artery, 228, 250 Oxidative stress, 449

P Para-anastomotic hypercompliance zone (PHZ), 140 Paracrine effect, 305 Patency, 250–252 PCL, see Poly(ε-capralactone) (PCL) PECAM-1, 154 Peptides, 249 Peracetic acid, 214, 218 Perfusion, 217, 302 Peripheral artery/arterial disease (PAD), 6–7, 366, 432, 564 Peripheral vascular trauma, 564 Perivascular cells, 470 Permeability, 23 Permeance, 23 Permittivity, 23 PGA, see Polyglycolic acid (PGA) PGS, see Poly(glycerol-sebacate) (PGS) Phagocytosis, 148 Pharmacogenetics, 111 Phenotypic switching, 435 Phenylmethylsulfonyl fluoride (PMSF), 216, 223 Phospholipase A2, 243 Photolithography, 165 Photo-oxidizing agents, 254 Physical decellularization, 518 Physical treatments, 217, 223–225 PIII, see Plasma immersion ion implantation (PIII) PLA, see Polylactic acid (PLA) Planar biaxial test, 46 Plaques, 100 Plasma immersion ion implantation (PIII), 392 Platelet deposition, 16 Platelet-derived growth factor (PDGF), 471 Pluronic microfibrous, 327 PNGase F, 216, 222 Poly(ε-capralactone) (PCL), 148, 149, 191, 414, 477, 499, 536–538, 540 Polyclycolic acid (PGA), 191 Poly (dimethylsiloxane) (PDMS), 478 Polydioxanone (PDO/PDS), 149–151, 191 Polyester (PET), 10, 23–24, 148 Polyethylene glycol (PEG), 163, 446, 477

583 Polyethylene terephthalate (PET), 10–11, 138, 142, 189, 340, 355, 380 Poly(glycerol-sebacate) (PGS), 151–153, 191, 538–539 Polyglycidyl ether, 254 Polyglycolic acid (PGA), 147, 345, 472, 477, 499, 537–538, 540, 543 Poly-4-hydroxybutyrate (P4HB), 191 Polylactic acid (PLA), 147, 148, 191, 537, 538, 540, 543 Poly (lactic-co-glycolic acid) (PLGA), 472, 474 Poly(lactide/caprolactone) (PLCL), 499 Poly (L-lactic acid) (PLLA), 477 Poly(l-lactide-co-ε-caprolactone) (PLCL), 536, 540, 543 Polymer combinations, 153, 154, 156 Polyol, 144 Polytetrafluoroethylene (PTFE), 10 grafts, 551 Polyurethane (PU), 12, 25, 143–146, 191, 499 Polyvinylidene fluoride (PVDF), 347 Porcine aorta, 230, 232–234, 240, 241, 253 sheet, 229, 232, 239 Porcine arteries, 152, 235 Porcine carotid arteries, 228, 229, 231, 234, 252, 253 Porcine radial arteries, 231, 253 Porcine thoracic aortas, 230, 237, 238 Porcine ureter, 235, 243, 252 Pore sizes, 536, 542 Porogen, 152 Porosity, 23–26, 499, 508, 510 Positive Allen’s test, 8 Preclinical studies, 561 Pre-clotting, 15, 169 Preconditioning, 42, 169, 448 Pre-seeding, 500, 504, 505, 507 Pressure treatments, 224 Pre-vascularized network, 478 Procalcifying media (PM), 443 ProCol ®, 246 Procyanidins, 254 Progenitor cells, 305, 306 Pro-inflammatory cytokines, 437 Prostheses, 65 Prosthetic grafts, 272 Protease inhibitors, 216, 223 Proteases, 222 Protein-protein interactions, 220 Protein structure, 220 Proteoglycans, 221 Proteomics, 450 Pseudoaneurysm, 246

584 Pseudointima, 104 Pseudo plaque, 441 PTFE grafts, 246 Pulmonary valve, 243 Pulsatile perfusion bioreactor, 172 Pulse rate, 557

Q Quasi-static seeding, 298–300 Quasi-static tests, 44

R Rabbit model, 79–80 Rapid prototyping (RP) technique, 479 Rat abdominal aorta, 239, 251 Rat abdominal artery, 251 Rat carotid artery, 251 Rat thoracic aorta, 239, 251 Recellularization, 370, 371 Recombinant tropoelastin, 210, 385, 386 Regenerative medicine, 210–212 Remodeling, 250–252, 278, 347 entrapped tissue cells, 372 Revascularization, 7 RGD, 415 groups, 164 RNA, 222 RNase, 216, 221, 222, 243 Rotational seeding, 301 Ruthenium-catalyzed photo-crosslinking, 373

S Sacrificial bioprinting, 327–329 Sacrificial fibers, 159 Saphenous vein (SV), 340 Scaffolds, 211, 213, 221, 222, 224–227, 245, 249, 254, 446 architecture, 510 -based methods, 189 decellularized, 254, 255 degradation, 497, 508, 510 design, 198 functionalization, 511 non-immunogenic, 255 xenogeneic tissue, 211, 212 Scanning electron microscopy (SEM), 158, 195 Sclerosis, 438 Seeding efficiency, 299, 301 Self-assembly approaches, 189

Index Sericin, 539 Shear stresses, 448, 453 Sheep, 175 Sheep pulmonary aorta, 250 Shinoka, T., 14 Sialic acid, 243 Silastic tube mandrill, 15 Silk, 349, 351 fibroin, 539 Siloxanes, 145 Single stage seeding, 13 Site of implantation, 248 Skeletal integrity, 459 Small animal models, 121 Small-diameter (