Engineering Strategies for Regenerative Medicine [1 ed.] 012816221X, 9780128162217

Table of contents :
Front matter
Copyright
Contributors
Engineering strategies for regenerative medicine
References
Pluripotent stem cell biology and engineering
Stem cells
Pluripotent stem cells
Embryonic stem cells
Induced pluripotent stem cells
Applications of hPSCs
Stem cell systems
The cell: The functional unit of the system
From the microenvironment to tissues and organs: Inputs of the system
Biochemical signals
Biophysical signals
The multiorgan system
Stem cell engineering
Engineering at the single-cell level
Reverse engineering using omics technologies
Forward engineering using synthetic biology
Engineering at the tissue level
Engineering biophysical signals
Delivery of biochemical signals
Organ-on-a-chip
Engineering at the multiorgan level
Body-on-a-chip
Concluding remarks
References
Process development and manufacturing approaches for mesenchymal stem cell therapies
Introduction
Isolation
Monolayer expansion
Bioreactor-based hMSC expansion
Stirred-tank bioreactors
Rocking motion bioreactors
Rotating-bed bioreactor
Hollow fiber bioreactor
Fixed- or packed-bed bioreactor
Vertical-wheel bioreactor
Bioprocess monitoring and control: parameters for optimization
Temperature and pH
Oxygen supply and aeration strategy
Medium formulation and supplements
Harvesting
Manufacturing paradigms: Autologous and allogeneic
Extracellular vesicle production using hMSCs
Therapeutic product formulation
Approved hMSC-based products
Conclusion
References
Further reading
Bioinspired materials and tissue engineering approaches applied to the regeneration of musculoskeletal tissues
Musculoskeletal tissues
Physiology and function
Bone
Cartilage
Tendon
Response to injury and healing mechanisms
Cartilage regeneration
Scaffold/hydrogel-based approaches
Cell-based approaches using mesenchymal stem cells
Cell-free therapies
Strategies for bone regeneration
Bioinspired materials for bone tissue engineering
Injectable bone substitutes
Scaffold-based approaches
Tendon regeneration
Biomaterial processing technologies to meet tendon function and properties
Fiber-based technologies
3D bioprinting technologies
Current applications and clinical potential
Future perspectives and concluding remarks
Acknowledgments
References
Bioengineering strategies for gene delivery
Introduction
Current prospects of gene therapies
Vector design
Viral gene therapy
Retroviral and lentivectors
Adenovectors
Adeno-associated vectors
Nonviral gene therapy
Delivery strategies
Synthetic materials
Poly(ethylene glycol)
Poly(lactic- co -glycolic) acid
Naturally occurring materials
Alginate
Fibrin
Collagen and gelatin
Gene delivery via biomaterial strategies
Revascularization
Lentivectors
Adenovectors and adeno-associated vectors
Nonviral vectors
Neurodegenerative disease
Lentivectors
Nonviral
Current therapeutic outlook
References
Advanced microtechnologies for high-throughput screening
Introduction
Microfabrication techniques and design considerations
Photolithography versus soft lithography
Replica molding
Microcontact printing
Robotic printing
Microscale technologies for HTS
Cell-based microarray platforms
Micropillar/microwell system for HTS assays
Microfluidic HTS platforms for probing in vivo microenvironments
Microfluidic HTS platforms for probing cytotoxic effects
Future opportunities
Acknowledgments
References
Inductive factors for generation of pluripotent stem cell-derived cardiomyocytes
Introduction
Human pluripotent stem cell-derived cardiomyocytes in cardiac regenerative medicine
Heart development
Cardiomyocyte differentiation from pluripotent stem cells
Cell-extrinsic factors
Cells
Growth factors/cytokines
TGF- β superfamily
Wnt/ β -catenin ligands/inhibitors
Fibro blast growth factors
Vascular endothelial growth factor
Other growth factors, cytokines, and hormones
Small molecules
Wnt signaling activators
Wnt signaling inhibitors
BMP signaling inhibitors
TGF- β /activin/nodal inhibitors
Retinoic acid
Icariin and peroxisome proliferator-activated receptor alpha agonist
SB203580, a p38 MAPK inhibitor
NO donor
Cyclosporin A
Vitamins
Lipids
Other small molecules
Physical cues
Extracellular matrix
Substrate elasticity
Electrical stimulation
Substrate topology
Stress
Other physical cues
Cell-intrinsic factors
Genes
Octamer-binding transcription factor 3/4
Mesoderm posterior BHLH transcription factor 1
GATA4
Na + /H + exchanger isoform 1
Inducible nitric oxide synthase
Prodynorphin and fibronectin type 3 domain-containing 5 protein
Islet1
MicroRNAs
Short hairpin RNAs
Epigenetic modulators
Conclusion
Acknowledgments
References
Pluripotent cells for the assessment of chemically induced teratogenesis and developmental toxicology
Background
Pluripotent stem cells
Embryonic stem cell test
Embryoid body systems
In vitro endpoints
Morphological endpoints
Cellular differentiation endpoints
“Omics” endpoints
Signaling and developmental pathway endpoints
Physiological and metabolic endpoints
Conclusions and future directions
References
Conclusions and closing remarks
References
Index
A
B
C
D
E
F
G
H
I
L
M
N
O
P
Q
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S
T
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Citation preview

ENGINEERING STRATEGIES FOR REGENERATIVE MEDICINE

ENGINEERING STRATEGIES FOR REGENERATIVE MEDICINE Edited by

Tiago G. Fernandes Maria Margarida Diogo Joaquim M. S. Cabral

Academic Press is an imprint of Elsevier 125 London Wall, London EC2Y 5AS, United Kingdom 525 B Street, Suite 1650, San Diego, CA 92101, United States 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom Copyright © 2020 Tiago G. Fernandes, Maria Margarida Diogo & Joaquim M. S. Cabral. Published by Elsevier Inc. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, ­electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek ­permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical ­treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in ­evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of ­products liability, negligence or otherwise, or from any use or operation of any methods, products, ­instructions, or ideas contained in the material herein. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN 978-0-12-816221-7 For information on all Academic Press publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Andre G. Wolff Acquisition Editor: Glyn Jones Editorial Project Manager: Anna Dubnow Production Project Manager: Poulouse Joseph Cover Designer: Miles Hitchen Typeset by SPi Global, India

Contributors Pedro S. Babo  3B’s Research Group, I3Bs—Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, AvePark, Parque de Ciência e Tecnologia, Zona Industrial da Gandra, Guimarães; ICVS/3B’s—PT Government Associate Laboratory, Braga/Guimarães, Portugal Alexey Bersenev  Cell Therapy Laboratory at Yale-New Haven Hospital, Yale University, New Haven, CT, United States Adara Bochanis  Department of Pharmaceutical Sciences, University of Connecticut, Storrs, CT, United States Joaquim M.S. Cabral  Department of Bioengineering and iBB-Institute for Bioengineering and Biosciences; The Discoveries Centre for Regenerative and Precision Medicine, Lisbon Campus, Instituto Superior Técnico (IST), Universidade de Lisboa, Lisbon, Portugal João P. Cotovio  Department of Bioengineering and iBB-Institute for Bioengineering and Biosciences; The Discoveries Centre for Regenerative and Precision Medicine, Lisbon Campus, Instituto Superior Técnico (IST), Universidade de Lisboa, Lisbon, Portugal Pedro Silva Couto  Advanced Centre for Biochemical Engineering, Department of Biochemical Engineering, University College London, London, United Kingdom Maria Margarida Diogo  Department of Bioengineering and iBB-Institute for Bioengineering and Biosciences; The Discoveries Centre for Regenerative and Precision Medicine, Lisbon Campus, Instituto Superior Técnico (IST), Universidade de Lisboa, Lisbon, Portugal Jonathan S. Dordick  Department of Chemical and Biological Engineering and Center for Biotechnology & Interdisciplinary Studies, Rensselaer Polytechnic Institute, Troy, NY, United States Tiago G. Fernandes  Department of Bioengineering and iBB-Institute for Bioengineering and Biosciences; The Discoveries Centre for Regenerative and Precision Medicine, Lisbon Campus, Instituto Superior Técnico (IST), Universidade de Lisboa, Lisbon, Portugal Manuela E. Gomes  3B’s Research Group, I3Bs—Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, AvePark, Parque de Ciência e Tecnologia, Zona Industrial da Gandra; ICVS/3B’s—PT Government Associate Laboratory, Braga/Guimarães; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimarães, Portugal

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Contributors

Manuel Gomez-Florit  3B’s Research Group, I3Bs—Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, AvePark, Parque de Ciência e Tecnologia, Zona Industrial da Gandra, Guimarães; ICVS/3B’s—PT Government Associate Laboratory, Braga/Guimarães, Portugal Ana I. Gonçalves  3B’s Research Group, I3Bs—Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, AvePark, Parque de Ciência e Tecnologia, Zona Industrial da Gandra, Guimarães; ICVS/3B’s—PT Government Associate Laboratory, Braga/Guimarães, Portugal Gyuhyung Jin  University of Wisconsin-Madison, Chemical and Biological Engineering, Madison, WI, United States Soowan Lee  Department of Pharmaceutical Sciences, University of Connecticut, Storrs, CT, United States Sean P. Palecek  University of Wisconsin-Madison, Chemical and Biological Engineering, Madison, WI, United States Qasim A. Rafiq  Advanced Centre for Biochemical Engineering, Department of Biochemical Engineering, University College London, London, United Kingdom Theodore P. Rasmussen  Department of Pharmaceutical Sciences; Institute for Systems Genomics, Storrs/Farmington; University of Connecticut Stem Cell Institute, University of Connecticut, Storrs, CT, United States Rui L. Reis  3B’s Research Group, I3Bs—Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, AvePark, Parque de Ciência e Tecnologia, Zona Industrial da Gandra; ICVS/3B’s—PT Government Associate Laboratory, Braga/Guimarães; The Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimarães, Portugal Márcia T. Rodrigues  3B’s Research Group, I3Bs—Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, AvePark, Parque de Ciência e Tecnologia, Zona Industrial da Gandra; ICVS/3B’s—PT Government Associate Laboratory, Braga/Guimarães; The ­Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimarães, Portugal André L. Rodrigues  Department of Bioengineering and iBB-Institute for Bioengineering and Biosciences; The Discoveries Centre for Regenerative and Precision Medicine, Lisbon Campus, Instituto Superior Técnico (IST), Universidade de Lisboa, Lisbon, Portugal Shahin Shams  Department of Biomedical Engineering, University of California, Davis, Davis, CA, United States Eduardo A. Silva  Department of Biomedical Engineering, University of California, Davis, Davis, CA, United States

Engineering strategies for regenerative medicine In Greek mythology Prometheus was a Titan who was credited with the forging of the first humans out of clay, and later on with providing mankind the gift of fire after tricking and stealing it from the gods. For that action, once believed to allow the progress and advancement of civilization, he was eternally punished by Zeus, the king of the Olympian Gods. According to this myth, Prometheus was chained to a rock and each day an eagle, the symbol of Zeus, was sent to feed on his liver, which would then grow back overnight to be eaten again the following day. This myth is often used to illustrate the remarkable capacity of certain organs and tissues to regenerate themselves. In humans, this capacity if frequently incomplete and, after damage or necrosis, comes fibrosis or the formation of scar tissue [1]. However, some organisms like salamanders and newts are able to regenerate entire limbs, tails, and a variety of other body structures [2]. Regenerative medicine, an interdisciplinary field that applies engineering and life science principles to promote regeneration, can potentially restore diseased and injured tissues and whole organs in humans [3]. This convergence between engineering and medicine thus aims to maintain, recover, and improve the function of damaged tissues and organs, as well as the production of new engineered tissues that may be used to improve the quality of life in chronic patients. Therefore, regenerative medicine integrates several domains, particularly the biology and engineering of stem cells, tissue and organ engineering, genetic and cellular therapies, biomaterial science and technology, micro- and nanotechnologies, and bioprocess engineering [4]. Regenerative medicine also offers unique technologies and treatments for unmet clinical needs. Some of the potential applications include transplantation of stem cells, progenitors, tissues, or organs [5], stimulation of endogenous repairing mechanisms [6], the use of cells as gene expression vehicles for the controlled delivery of cytokines and other biological molecules [7], and cellular engineering/synthetic biology [8]. This represents a market predicted to reach $67.5 billion by 2020, according to the report by Allied Market Research [9]. For example, immunotherapy is also quickly moving to the bedside, and advanced manufacturing has the potential to recapitulate the native microenvironment and produce cellular surrogates with high degrees of functionality [10].

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In this context, this book starts by providing an in-depth overview of pluripotent stem cell biology and the engineering tools that can harness their full potential for regenerative medicine applications. Cotovio and coworkers have explored in Chapter 1 some of these questions and summarized key biological features of human pluripotent stem cells (hPSCs), some of the foreseen applications in cellular therapy, disease modeling and pharmaceutical testing, the bioengineering approaches used to explore stem cell systems at the cellular, multicellular and multiorgan levels, and how this progress could potentially revolutionize clinical practice and drug development. Following this, Couto, Bersenev, and Rafiq have discussed process development and manufacturing approaches for mesenchymal stem/stromal cell therapies in Chapter 2. In fact, one of the cell types under evaluation in numerous clinical trials are human mesenchymal stem/stromal cells (hMSCs) that possess both immunomodulatory properties and in  vitro differentiation capability. These properties have been explored for conditions such as graft vs host disease, acute myocardial infarction and Crohn disease, among others. One serious limitation, however, is the inability to generate clinically relevant number of cells to meet the dose criteria, while retaining the product’s critical quality attributes. Therefore, several challenges that manufacturers face during each stage of the hMSC handling, from isolation to expansion, downstream processing, and finally product formulation, are explored in this chapter. In Chapter 3, Rodrigues and coworkers have discussed how biomaterials and tissue engineering approaches can be used to help the regeneration of musculoskeletal tissues. Since these tissues are highly prone to injuries, conditions afflicting them have a great impact on the quality of life of patients worldwide. Particularly, with the increase in life expectancy and the maintenance of an active lifestyle by the aging population, the cumulative musculoskeletal conditions potentiate disability throughout life. Therefore, the economic impact of the United States alone was estimated to be $796.3 billion in 2009–11 [11]. The authors have discussed how the exploitation of materials and scaffolds to modulate cell responses potentially contribute to the functional regeneration of bone, cartilage, and tendon after injury. In the following chapter (Chapter 4), Shams and Silva have further highlighted how materials can be used in regenerative medicine. Particularly, they looked at d ­ elivery strategies using biomaterials for gene therapy applications focusing on vascular regeneration and neurodegenerative diseases. Stem cell-based microscale platforms, particularly the ones using ­patient-specific hPSCs, may prompt the development of phenotypic drug discovery assays that would allow the identification of personalized drug responses. In Chapter 5, Rodrigues and coworkers have outlined several microscale high-throughput screening platforms that have been employed in the regenerative medicine field. They have also provided examples of



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specific studies employing stem cells and derived progeny for screening drug compounds and microenvironmental cues influencing cell behavior. The impact of microscale technologies on early identification of potential new drug candidates may ultimately decrease both attrition rates and capital investment in the drug development process, and further help the progress of regenerative medicine applications. In the following chapter (Chapter  6), Jin and Palecek have provided an overview on how hPSC-derived cardiomyocytes have emerged as a promising cell source to regenerate damaged heart tissue. Ischemic heart disease is one of the leading causes of death worldwide and even though some therapeutic options are available following heart attack, such as pharmacological therapies, bypass surgery, valve surgery, stent implantation, and pacemaker implantation, most of these treatments are designed to prevent recurring heart failure, attenuate symptoms, and assist cardiac function. Since they do not provide restoration of physiological heart function, numerous efforts in the field of cardiac regenerative medicine have been made to replace necrotic cardiomyocytes, thereby restoring the function to the damaged heart. Particularly, understanding the effects and underlying mechanism of inductive factors on cardiomyocyte differentiation from hPSCs may provide important clues to address some of the existing challenges and finally provide regenerative therapies for devastating and lethal heart conditions. Lastly, one additional application explored in Chapter 7 by Lee, Bochanis, and Rasmussen is the use of hPSCs as models for early embryogenesis, and to precisely evaluate the risk of fetal exposure to teratogens. The development of in vitro assays for the detection of teratogens and developmental toxins is highlighted in this chapter, as well as the current state of PSC-based platforms for the detection of compounds that pose a risk to human developmental programs, and for the ascertainment of teratogenic mechanisms at cellular and molecular levels. These examples summarize some of the most compelling and innovative approaches being explored in regenerative medicine. While the list is not overly comprehensive, it for sure provides a collective impression that medicinal practice can now focus on regenerating diseased tissues and recover lost functionality. It was our aim to provide relevant and updated content on these topics, and we hope that the reader will be more informed about this exciting field. Tiago G. Fernandes Maria Margarida Diogo Joaquim M.S. Cabral Department of Bioengineering, Instituto Superior Técnico, The University of Lisbon, Lisbon, Portugal

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References [1] Atala A, Irvine DJ, Moses M, Shaunak S. Wound healing versus regeneration: role of the tissue environment in regenerative medicine. MRS Bull 2010;35(8):597–606. https://doi. org/10.1557/mrs2010.528. [2] Brockes JP, Kumar A. Plasticity and reprogramming of differentiated cells in amphibian regeneration. Nat Rev Mol Cell Biol 2002;3(8):566–74. https://doi.org/10.1038/nrm881. [3] Mao AS, Mooney DJ. Regenerative medicine: current therapies and future directions. Proc Natl Acad Sci 2015;112(47):14452–9. https://doi.org/10.1073/pnas.1508520112. [4] Tewary M, Shakiba N, Zandstra PW. Stem cell bioengineering: building from stem cell biology. Nat Rev Genet 2018;19(10):595–614. https://doi.org/10.1038/s41576-018-0040-z. [5] Bajaj P, Schweller RM, Khademhosseini A, West JL, Bashir R. 3D biofabrication strategies for tissue engineering and regenerative medicine. Annu Rev Biomed Eng 2014;16(1):247–76. https://doi.org/10.1146/annurev-bioeng-071813-105155. [6] Kami  D, Gojo  S. Tuning cell fate. Organogenesis 2014;10(2):231–40. https://doi. org/10.4161/org.28816. [7] Nowakowski  A, Drela  K, Rozycka  J, Janowski  M, Lukomska  B. Engineered mesenchymal stem cells as an anti-cancer trojan horse. Stem Cells Dev 2016;25(20):1513–31. https://doi.org/10.1089/scd.2016.0120. [8] Pawlowski M, Ortmann D, Bertero A, Tavares JM, Pedersen RA, Vallier L, Kotter MRN. Inducible and deterministic forward programming of human pluripotent stem cells into neurons, skeletal myocytes, and oligodendrocytes. Stem Cell Rep 2017;8(4):803–12. https://doi.org/10.1016/j.stemcr.2017.02.016. [9] Allickson  J. Emerging translation of regenerative therapies. Clin Pharmacol Ther 2017;101(1):28–30. https://doi.org/10.1002/cpt.549. [10] Lipsitz YY, Timmins NE, Zandstra PW. Quality cell therapy manufacturing by design. Nat Biotechnol 2016;34(4):393–400. https://doi.org/10.1038/nbt.3525. [11] Yelin  E, Weinstein  S, King  T. The burden of musculoskeletal diseases in the United States. Semin Arthritis Rheum 2016;46(3):259–60. https://doi.org/10.1016/ j.semarthrit.2016.07.013.

C H A P T E R

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Pluripotent stem cell biology and engineering João P. Cotovioa,b, Tiago G. Fernandesa,b, Maria Margarida Diogoa,b, Joaquim M.S. Cabrala,b a

Department of Bioengineering and iBB-Institute for Bioengineering and Biosciences, Instituto Superior Técnico (IST), Universidade de Lisboa, Lisbon, Portugal, bThe Discoveries Centre for Regenerative and Precision Medicine, Lisbon Campus, Instituto Superior Técnico (IST), Universidade de Lisboa, Lisbon, Portugal

1  Stem cells In 1868 Ernst Haeckel [1], a notable biologist from Germany, came up with the term Stammzelle to describe the unicellular ancestor from which all multicellular organisms evolved, a concept very different from the one existing today. Later, he used the same term to describe the fertilized egg, or the zygote, capable of giving rise to all cell types of an organism [2]. This was the cradle for the English term “stem cell”  (Fig.  1.1) [3]. After previous studies on the continuity of the germ plasm and on the origin of the hematopoietic system, Till and McCulloch proposed in the 1960s what are still today the two gold standard features of stem cells: (1) undifferentiated cells that are capable of self-renewal and (2) the production of specialized progeny through differentiation [4]. To accomplish such attributes, it is now known that stem cells undergo asymmetric cell division, by which the cell divides to generate one stem cell and one differentiating cell. Therefore, it may be added that stem cells are of major importance in the maintenance of homeostasis through a balance between self-renewal and differentiation [5–7]. From conception to death, as cells develop derived from embryonic tissue, they become progressively restricted in their developmental potency, reaching the point when each cell can only differentiate into a s­ ingle ­specific Copyright © 2020 Tiago G. Fernandes, Maria Margarida Diogo & Joaquim M. S. Cabral. Published by Elsevier Inc. All rights reserved. https://doi.org/10.1016/B978-0-12-816221-7.00001-X

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FIG.  1.1  Stem cell research timeline. Key events and technological breakthroughs in stem cell research. EC, embryonal carcinoma; hESCs, human embryonic stem cells; hiPSCs, ­human-induced pluripotent stem cells; iPSCs, induced pluripotent stem cells; mESCs, mouse embryonic stem cells.

cell type. In the beginning, the earliest cells in ontogeny are totipotent, giving rise, in mammals, to all embryonic and extraembryonic tissues, that is, only a totipotent cell can originate an entire organism [8, 9]. Through embryogenesis, when the pluripotent state is reached, a pluripotent stem cell (PSC) can originate all the cells from all the tissues of the body, although the contributions to the extraembryonic membranes or placenta are limited [8]. On the other hand, a multipotent stem cell is restricted to the generation of the mature cell type of its tissue of origin and finally, a unipotent stem cell displays limited developmental potential, giving rise to only a single-cell type [8]. In an adult organism, stem cells can be found in most tissues throughout the body, even within relatively dormant tissues. These stem cells experience low or no division in normal homeostasis, remaining quiescent for extended periods of time. However, these cells can respond efficiently to stimuli upon initiation of homeostasis or injury [10]. Altogether, in both plant and animal kingdoms, the multicellularity of highly regulated tissues is dependent of the generation of new cells for growth and repair. Therefore, biological systems are driven by a balance between cell death and cell proliferation, preserving form and function in tissues. From this point of view, stem cells are the units of the following attributes: development, regeneration, and evolution [9, 11].

1.1  Pluripotent stem cells Pluripotency can be defined as a transient property of cells within the early embryo, where PSCs have the capacity to form tissues of all three germ layers of the developing embryo and, later, the organs of the adult organism—ectoderm, mesoderm, and endoderm—and still the germ lineage. As previously mentioned, PSCs typically provide little or no contribution to the trophoblast layers of the placenta [8, 12]. The first PSCs to be isolated and investigated in culture were derived from mouse teratocarcinomas—a tumor of germ cell origin that maintain a



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wide variety of diversely differentiated tissues—known as embryonal carcinoma (EC) cells [8, 13]. Nevertheless, PSCs can be isolated from several sources through development [14], as murine [15, 16] and human blastocyst [17] or even from the postimplantation epiblast [18, 19] or germ line [20, 21]. Also, pluripotency can be recapitulated in vitro by reprogramming somatic cells to become induced pluripotent stem cells (iPSCs) [22–24]. There are specific molecular mechanisms that characterize PSCs anchored by a selected set of core transcription factors essential to establish pluripotency. As part of the core pluripotency transcription factors encoding genes are octamer-binding transcription factor 4 (OCT4), SRY-box  2 (SOX2), and NANOG. In certain circumstances, the loss of SOX2 or NANOG or their substitution can be tolerated [8, 12]. Despite that, PSCs can be classified into different states of pluripotency based on the molecular signatures, with the terms “naïve” and “primed” being introduced to describe early and late phases of ontogeny, respectively [25]. Pluripotency can be suggested by such molecular signatures, but only functional assays can reveal the developmental potential of a cell. Functional assays to assess pluripotency include: differentiation into three germ layers in vitro, teratoma formation in vivo, chimaera formation, germline transmission through blastocyst injection, tetraploid complementation, and single-cell chimaera formation [8]. For human pluripotent stem cells (hPSCs), teratoma formation remains the gold standard of functional assays. 1.1.1  Embryonic stem cells Mammalian embryogenesis starts with a single totipotent cell, the zygote. After the first cell division, the two-cell embryo is composed by two equal blastomeres. In the earlier stages, including two- and four-cell embryos, cells are still considered totipotent. Later, in what is called blastocyst, it is possible to distinguish the extraembryonic trophectoderm (TE) on the outside and the inner cell mass (ICM) [14, 26]. It is in the ICM that pluripotent cells first arise. The ICM cells, cultured in conditions that allow indefinite self-renewal and maintenance of the pluripotent state, are known as embryonic stem cells (ESCs) and were first derived from mouse (mESCs) in 1981 by Martin Evans [15] and Gail Martin [16]. These cultures proved to have all the properties previously established for EC cell cultures, as well as a completely normal karyotype [27]. Only by the year 1998, Thomson derived the first ESC lines from human blastocysts [17], the so-called human embryonic stem cells (hESCs). Throughout normal development, the amount of ESCs is limited and their existence is constrained in the time course of development, being present for only a short period of time. In contrast, tissue culture allows the generation and maintenance of millions of ESCs indefinitely, preserving their pluripotent state [27].

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1.1.2  Induced pluripotent stem cells The molecular mechanisms “by which the genes of the genotype bring about phenotypic effects”—the epigenetics concept—was captured by Conrad Waddington [28] in the iconic image of the epigenetic landscape that influences cellular fate during development, analogously to the movement of a marble (Fig. 1.2). Since then, the possibility that cells can change their identity has fascinated scientists [29]. This notion was first suggested by Sir John Gurdon in 1958, establishing that in vivo plasticity of the differentiated state can be induced artificially by directly manipulating cells and their environment [30]. It was demonstrated that the marble can be rolled back to the top of the hill, that is, committed or differentiated cells can be reprogrammed back to a wider developmental potential (dedifferentiation). As the possibility to reprogram cells, not by transplanting their nuclei, but by introducing pluripotency factors into cells became a reality, cells with a gene expression profile and developmental potential similar to ESCs were generated in 2006. This accomplishment was reached using mouse somatic cells together with a cocktail of four transcription factors [23]. The resulting reprogrammed cells were termed iPSCs and were generated after retrovirally introducing four transcription factors encoding genes, OCT4, SOX2, Kruppel-like factor 4 (KLF4) and the MYC proto-oncogene, bHLH transcription factor (MYC)—the “Yamanaka factors.” After successfully generating mouse iPSCs, in 2007 iPSCs were generated from human fibroblasts, using the same four factors and alternatively NANOG and Lin-28 homolog A (LIN28) instead of KLF4 and MYC [22, 24]. It was the establishment of human induced pluripotent stem cells (hiPSCs). Since then, cellular reprogramming became a robust method to convert differentiated cells to a PSC state [31, 32].

FIG.  1.2  Cell fate plasticity. Contemporary version of Waddington landscape depicting an analogy between a marble rolling downhill as development leads undifferentiated cells to a mature state. Cellular reprogramming has shown that it is possible to make the marble roll back to the top of the hill as mature cells can be reprogrammed back to a wider developmental potential.



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Afterwards, besides the initially used retroviral or lentiviral vectors, nonintegrating methods have been developed and include reprogramming using episomal DNAs, adenovirus, Sendai virus, PiggyBac transposons, minicircles, recombinant proteins, synthetically modified mRNAs, microRNAs (miRNAs), and more recently, small molecules [31]. These new techniques, in addition to lower variability between cell lines, can lead to safer reprogramming of iPSCs and to more suitable cells for clinical applications by avoiding insertional mutagenesis and transgene reactivation.

1.2  Applications of hPSCs Since the isolation of ESCs from human embryos, the use of PSCs as a potential tool for research and medicine has been growing (Fig.  1.3). Besides that, after finding that somatic cells can revert all the way back to an ESC state through transcription factor activation, manipulation of signaling pathways aiming for cell differentiation has been studied contributing to hiPSCs applications in biomedicine. Accordingly, several protocols have been described for in vitro direct differentiation of neurons, hematopoietic cells, hepatocytes, smooth muscle cells, and cardiomyocytes, among other cell types across the three germ layers [33]. An obvious application of hPSCs in medicine is in cell therapy. Regenerative medicine strategies based on the use of stem cells to promote regeneration or to replace damaged tissues after cellular transplantation has been shown to successfully induce functional recoveries [31]. In fact, several clinical trials were already established using hPSC-based therapies [34]. In the particular case of hiPSCs, an important advantage of ­using

FIG. 1.3  Clinical applications of human induced pluripotent stem cells for precision medicine. After isolation of patient somatic cells, these cells can be cultured and reprogrammed into patient-specific induced pluripotent stem cells (iPSCs). This is a promising cell source for cell therapy as it is possible to silence mutations carried by patient-specific cells using novel gene-editing tools (like CRISPR-Cas9) for the generation of corrected iPSCs that can be used in regenerative medicine approaches. Also, the differentiation of patient-specific iPSCs makes disease modeling and drug screening a possibility.

6

1.  Pluripotent stem cell biology and engineering

these cells is the capability to generate autologous differentiated cells, that is, patient-specific cells, theoretically suppressing the risk of immune rejection. For instance, the first clinical study using hiPSC-derived products was performed in 2014 by Masayo Takahashi and Yasuo Kurimoto, in which these two Japanese physicians successfully transplanted autologous retinal pigment epithelium sheets derived from hiPSCs into a woman with macular degeneration [35]. Besides all this progress in ­hPSC-based therapies, the acquisition of chromosomal aberrations, due to the reprograming process and subsequent culture, represents one of the disadvantages of these cells [36]. Moreover, due to hPSC tumorigenicity, it is critical to ensure that the transplanted product does not contain undifferentiated cells with the potential to generate teratomas [31]. Another important biomedical application of hPSCs is in disease modeling [37]. It is expected that in vitro hPSC-based disease models help to identify the pathological mechanisms underlying human diseases. Both hESCs and hiPSCs have been used for modeling human genetic diseases, establishing isogenic cell lines with novel gene-editing tools (e.g., CRISPRCas9), to induce disease-causing mutations or to silence mutations carried by patient-specific cells [38, 39]. Modeling of human diseases is motivated by the necessity of developing novel therapeutic agents allowing the diseases to be treated, alleviated, or cured. Therefore, drug screening and toxicological assays is also considered as a potential application of hPSCs [37]. Animal models have been used in drug screening but differences from the actual human setting lead to an inaccurate forecasting of their effects. Moreover, animal models are not suitable for high-throughput screening of small-molecule ­libraries [38, 40]. Until now, many drug screens have been conducted using hiPSC-based models and potential drug candidates have been identified. Also, it is not only important to assess efficacy but also toxicity, predicting the likelihood of candidate drugs to cause serious side effects [31]. A specific patient has a specific genetic background and this fact implies different responses to medication for each individual. Accordingly, hiPSC-based drug screening is the key for a personalized therapy, an emerging approach known as precision medicine [40]. Just as new technologies are being developed, the greater will be the potential applicability of hPSCs in the emerging fields of regenerative medicine, disease modeling, and drug screening.

2  Stem cell systems For a widespread use of stem cells in biomedical applications, it is essential to understand the complexity and the dynamics of the stem cell biological system as well as the information that flows between each layer



2  Stem cell systems

7

FIG. 1.4  Layers of complexity within biological systems. Different layers of complexity of the biological systems from the cell level to the whole organism and respective engineering approaches. These approaches can help to understand each biological layer, and to a great extent, they can help to create tissues that resemble the in vivo anatomy and physiology of the human body for biomedical applications.

of it. In this section, a reductionist view of the different layers of complexity of the biological systems is presented (Fig. 1.4), exploring this question from the cell level to the whole organism.

2.1  The cell: The functional unit of the system The cell by itself can be considered the functional unit of a biological system. Furthermore, within each cell and as a first layer of regulation, there are collections of molecular species that interact and constitute what are called gene regulatory networks (GRNs), a term first coined in sea urchin developmental studies [41]. The GRNs, based on the microenvironmental signals (input), are able to control gene expression levels and, consequently, gene product abundance (output). Thus, the cellular GRN is responsible for the development potential of the cell by regulating its transcriptional activity [42, 43]. For example, the pluripotent state of a PSC is determined by a set of pluripotency transcription factors, namely OCT4, SOX2, and NANOG. These transcription factors are the core of the pluripotency GRN and, therefore, they cooperate and regulate different elements in the genome, including their own promoters [12, 44]. Besides that, additional factors that constitute regulatory inputs to the pluripotency GRN can include proteins, both activators and repressors, which will regulate gene expression at the transcription initiation stage. Moreover, another form of transcriptional regulation is based on the epigenetic information, namely chemical modifications or changes in the 3D organization of the chromatin [45]. These regulatory mechanisms are responsible for the unique epigenetic plasticity of PSCs shaping the transcriptional events during lineage commitment. Finally, RNA-based regulatory mechanisms also take part in the pluripotency GRN. Examples

8

1.  Pluripotent stem cell biology and engineering

of that are the miRNAs, long noncoding RNAs (lncRNAs), and alternative splicing [12]. Overall, one might say that these protein-based and epigenetic regulatory strategies, as well as regulatory RNAs, integrate the pluripotency GRN of the cell taking control of its transcriptional activity and, consequently, of the gene product abundance.

2.2  From the microenvironment to tissues and organs: Inputs of the system From a multicellular point of view, stem cells constitute a unique cell population capable to form, maintain, and restore an organized structure or even a given tissue and, in a broader perspective, an organ. This process of creation, which gives rise to an organized form, is called morphogenesis. In his opus “On Growth and Form” [46], Sir D’Arcy Thompson laid the foundations for the scientific explanation of this process. However, some pertinent questions still remained, particularly: how are stem cells regulated trough differentiation and morphogenesis until a mature status of the tissues and organs is reached? In truth, the regulation of the system is highly determined by the so-called niche—the local tissue microenvironment that hosts and influences the behavior or characteristics of stem cells [10, 47]. The concept of a stem cell niche has its foundation in the late 1970s with the regulatory microenvironment of adult mammalian hematopoietic stem cells (HSCs) [48]. The whole subject, as far as one can tell, remained elusive until an in-depth characterization of the stem cell niches of Drosophila melanogaster [49]. However, regardless of whether cells retain their stemness or are already committed into a specific lineage, aside from the intrinsic mechanisms already discussed, their fate and subsequent morphogenic processes are further regulated by a combination of extrinsic inputs. Despite their intrinsic differences, the architecture of different tissues shares several defining characteristics. Most notably, it is known that the signals provided by the microenvironment are received by the cell in the form of both biochemical and biophysical inputs (Fig. 1.5). These inputs are then transmitted within the cells via a cascade of molecular signal transduction events activating many other signaling pathways, producing controlled functional responses according to the specific physiological needs [43]. 2.2.1  Biochemical signals Generically, the molecular communication between cells in the form of biochemical signals is largely carried out by proteins, globally referred as signaling proteins. Examples of this include juxtacrine (­contact-dependent) and paracrine (secreted) signals received from adjacent or neighboring cells. A specific type of paracrine signals is autocrine signals, that occur when the secreted signal is received by the secretory cell [43].



2  Stem cell systems

9

FIG. 1.5  Regulatory microenvironment of biological systems and its biophysical and biochemical inputs. Biochemical signals include juxtacrine (contact-dependent) and paracrine (secreted) signals received from adjacent or neighboring cells, while biophysical signals include cell-cell, cell-extracellular matrix (ECM) interactions, and physiological factors (oxygen and pH). The GRN convert these signals (inputs) in controlled functional responses (output) according to the physiological needs.

As a prime example of a paracrine signal, one can consider a morphogen—a diffusible molecule that governs cell fate specification by its concentration through a gradient [50]. Paracrine signals that often act as morphogens can be grouped in four major families: the fibroblast growth factor (FGF) family, the Hedgehog family, the Wnt family, and the transforming growth factor (TGF)-β superfamily (that includes the TGF-β family, the activin family, the bone morphogenic proteins (BMPs), the Nodal proteins, and the Vg1 family). By contrast, juxtacrine signals do not make use of diffusible molecules. Instead, proteins interact with receptors of adjacent cells [50]. From all the juxtacrine signals, the Notch signaling proteins are the most well studied. 2.2.2  Biophysical signals On the other hand, biophysical signals include mechanical stimuli from cell-cell and cell-extracellular matrix (ECM) interactions as well as physiological factors (e.g., oxygen and pH) [43]. A large spectrum of cell adhesion molecules (CAMs) can mediate cell adhesion in cell-cell interactions, but among them, cadherins are the major ones [51]. In addition, cell-ECM interactions are influenced by chemical composition, structural organization, and mechanical properties of the ECM [52]. Regarding chemical composition, the ECM is composed of about 300 proteins including collagen, proteoglycans (PGs), and glycoproteins [53]. In what concerns to physical properties, force and geometry, can be sensed by mechanotransduction systems that can then transduce the physical signals into biochemical responses [54–56]. These ­mechanotransduction

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systems rely mainly on transmembrane receptors, such as integrins, which connect extracellular and intracellular structures, distributing the biophysical signals to the cytoskeleton, translating the mechanical forces of the ECM into morphogenetic processes, defining tissue architecture, and driving specific cell differentiation programs [52].

2.3  The multiorgan system It is a common fallacy to envision organs as isolated identities and discrete parts with distinct highly specialized functions. In truth, an organism is not a system made of parts, but a continuous and organic structure where each organ depends on each other as a single living entity—it is a multiorgan system. Cell communication across different tissues and organs do not end with development nor development ends with embryogenesis. This biological process of communication continues throughout the natural life span of living things. In fact, for many organs, its functionality depends on the continuous interorgan cross talk where communication between distal organs through paracrine signals is vital for homeostasis [57]. This is a topic that has been discussed at some, yet quite insufficient, length. Accordingly, the basic components of interorgan communication are now beginning to be identified, molecular species that are surely integrated in what is, after all, a broader interorgan communication network (ICN).

3  Stem cell engineering For the past decades, since the isolation and culture of hPSCs in vitro, the intention has been to engineer every cell, niche, tissue, and organ-like structures, recreating the complexity and architecture of the human body. This technology has grown immensely since then, bringing an exciting new era for the fields of disease modeling, drug discovery, and regenerative medicine, but nevertheless many more questions need to be further addressed. It is in the intersection of stem cell biology and engineering that new advances are being accomplished, applying bioengineering strategies to study each of the layers of the biological complexity of the human body. In this section, the main bioengineering approaches used at the cellular, multicellular, and multiorgan levels are explored.

3.1  Engineering at the single-cell level 3.1.1  Reverse engineering using omics technologies In the past, only a few molecular techniques were able to provide information at the single-cell level, like patch-clamping electrophysiology [58], fluorescence in situ hybridization [59], or flow cytometry [60], ­potentially



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analyzing 1–3 parameters from a given cell. Nevertheless, major breakthroughs in molecular biology like next-generation sequencing and novel omics technologies have provided transformative ways to comprehensively analyze the single cell at the molecular level in the recent years [61]. Consequently, these technologies have been used in favor of stem cell engineering, generating large datasets not only by capturing transcriptional information, but also by disclosing protein interactions, enabling computational techniques to construct and simulate cellular GRNs, and to increase the efficiency resolution of differentiation processes of so many stem cell derivatives. In this context, the use of this information to build improved in vitro cellular systems is known as reverse engineering [43]. Bellow, the most recent technological advances and main omics technologies that feed this kind of approach are summarized (Table 1.1). TABLE 1.1  Engineering at the cell level Approach Reverse engineering

Examples

References using stem cells

Transcriptomics

Plate-based scRNA-seq   STRT-seq [62]   Smart-seq [63]   Smart-seq2 [64] Commercial microfluidic approaches   Fluidigm C1 Pooled approaches   CEL-seq [65]   MARS-seq [66]   SCRB-seq [67]   CEL-seq2 [68]   10X Genomics [69]   SPLiT-seq [70] Massively parallel approaches   Drop-seq [71]   inDrop [72]

[73–77]

Epigenomics

Chromatin accessibility   DNase-seq [78]   ATAC-seq [79]   THS-seq [80] DNA methylation   RRBS [81]   WGBS [82]

[83–85]

Proteomics

Semiquantitative methods   Flow cytometry   Mass cytometry (CyTOF)   Single-cell western blot [86] Quantitative methods   Microengraving chips [87]   SCBC [88]

[89–92]

Continued

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1.  Pluripotent stem cell biology and engineering

TABLE 1.1  Engineering at the cell level—cont’d Approach

Forward engineering

Examples

References using stem cells

Multiomics

Transcriptomics + epigenomics   MT-seq [93]   NMT-seq [94]   CAR [95] Transcriptomics + proteomics   CITE-seq [96]   REAP-seq [97] Epigenomics + proteomics   Pi-ATAC [98] Chromatin accessibility + DNA methylation   COOL-seq [99]   NOMe-seq [100]

[93, 99]

Synthetic biology

Engineering of endogenous GRNs Addition of SGNs

[101–104]

ATAC-seq, assay for transposase accessible chromatin with sequencing; CAR, chromatin accessibility and mRNA; CEL-seq, cell expression by linear amplification and sequencing; CITE-seq, cellular indexing of transcriptomes and epitopes by sequencing; COOL-seq, chromatin overall omic-scale landscape sequencing; CyTOF, cytometry by time of flight; MARS-seq, massively parallel RNA single-cell sequencing; MT-seq, methylome and transcriptome sequencing; NMT-seq, nucleosome, methylation and transcription sequencing; NOMe-seq, nucleosome occupancy and methylome sequencing; Pi-ATAC, protein-indexed assay of transposase accessible chromatin with sequencing; REAP-seq, RNA expression and protein sequencing assay; RRBS, reduced representation bisulfite sequencing; SCBC, single cell barcode chip; SCRB-seq, single-cell RNA barcoding and sequencing; SPLiT-seq, split-pool ligation-based transcriptome sequencing; STRT-seq, single-cell tagged reverse transcription sequencing; THS-seq, transposome hypersensitive site sequencing; WGBS, whole-genome bisulfite sequencing.

From among a bewildering palette of omics technologies, transcriptomic analysis is being widely used at the moment. Gene expression analysis at the single-cell level has its origins back in the 1990s [104a], but single-cell RNA sequencing (scRNA-seq) revolutionized the way one can profile gene expression of PSCs and its derivatives at a given time [105]. The underlying technique of scRNA-seq has been subjected to a remarkable amount of improvements, and new approaches are now being reported very frequently with low-cost methods and increased number of sequenced cells arising at the scRNA-seq landscape (for review: Refs. [106, 107]). Standard approaches for scRNA-seq include: plate-based approaches using micropipettes or fluorescence-activated cell sorting (FACS); the commercial microfluidic approach Fluidigm C1; pooled approaches applying a barcode to the cells; and massively parallel approaches that isolate single cells into droplets thanks to the advances in droplet microfluidics enabling the profiling of thousands of cells in a single experiment [107]. In a recent publication, the need for ­single-cell isolation required by previous methods was even ­eliminated



3  Stem cell engineering

13

by not ­partitioning cells into individual compartments but relying on cells themselves as compartments [70]. Moreover, due to the complex data output of these procedures, numerous algorithms have been developed to analyze the amount of multidimensional data generated [108, 109]. These technologies and algorithms have led to ambitious largescale sequencing projects such as “The Human Cell Atlas,” aiming the sequencing of all cell types in the human body [110]. But scRNA-seq goes far beyond that and has proved to be a successful analytical tool in the stem cell field, as will be described below. These techniques have provided valuable information to dissect the molecular complexity of the pluripotency GRN, analyzing, for example, the different pluripotent states and resolving pluripotency heterogeneity of PSCs that differ in their developmental potency both in  vitro [75] and in  vivo [74]. Similarly, scRNA-seq was recently used to even study trajectories and major molecular events during reprogramming of iPSCs, helping in the development of faster reprogramming systems [77]. Besides this characterization of distinct cell states and pluripotency heterogeneity, an attractive application of scRNA-seq is lineage tracing. During cell differentiation and lineage commitment from pluripotency to a given cell type, scRNA-seq can capture different transcriptional states of the cells in different developmental stages. This analysis provides information of the cellular decision-making process leading to the prediction and reconstruction of the differentiation trajectories of the cells and at larger extent of the mapping of their developmental progression [111]. Prime examples of that are illustrated by the study of pairwise choices, or bifurcating lineage choices, to predict the developmental roadmap of mesodermal lineages from pluripotency to bone and heart [76], and to study cardiomyocyte maturation [73]. Besides transcriptomics, epigenomic analysis aiming at studding genome regulation can now be performed with an unprecedented resolution [112]. Techniques to analyze single-cell chromatin, such as the assay for transposase accessible chromatin with sequencing (ATAC-seq) [79, 113], and single-cell DNA methylation, like whole-genome bisulfite sequencing (WGBS) [82], have been published in the last years, and as a result, such approaches are starting to be applied toward the understanding of GRNs within stem cells. An example of that is the study of the regulatory role of chromatin dynamics during embryogenesis at single-cell resolution using ATAC-seq [83], revealing tissue-specific regulatory elements that shape gene expression and bringing insights into developmental programs. This type of studies surely contributes to the understanding of cell differentiation and lineage commitment and potentially, to the in vitro generation of different cell types. In fact, single-cell epigenomics have already been used to characterize transcriptional regulation of more specific cell types, or more precisely to characterize chromatin accessibility in developing

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1.  Pluripotent stem cell biology and engineering

brains [84]. Besides chromatin accessibility, recent studies have already focused on DNA methylation of stem cells, providing information about another layer of epigenetic regulation. For example, WGBS was already used to access DNA methylation dynamics during intermediate states of iPSCs reprogramming, where demethylation of pluripotency promoters and enhancers takes place, while somatic-specific regulatory elements are remethylated [85]. High-throughput single-cell transcriptomics and epigenomics are revolutionizing the way it is possible to study the molecular complexity of cells. However, these approaches do not provide phenotypic information. Therefore, proteomic analysis is also essential to understand GRNs as they are mediated by the proteins those genes encode [114]. Accounting the most recent advances in this area, the state of the art of single-cell proteomics can be divided in qualitative methods, identifying proteins produced by certain cells; semiquantitative methods, measuring protein abundance in relative units (e.g., flow cytometry, mass cytometry, and single-cell western blot); and quantitative methods, translating analytical signals into protein concentration (e.g., single-cell barcode chip (SCBC), microengraving chips) [61, 115]. Flow cytometry is a well-established single-cell protein assay, but some of the other techniques emerging in the field add increased multiplexing capability (i.e., higher number of proteins assayed). Mass cytometry, or cytometry by time of flight (CyTOF), also rely on the use of antibodies to label proteins of interest but with transition metal-containing mass tags rather than fluorophore tags, increasing the multiplexing capacity up to 30 proteins [116, 117]. Actually, mass cytometry was already applied in stem cell studies being used, for example, to evaluate the process of iPSC reprogramming in vitro [89, 92], to study myogenic lineage progression during muscle regeneration in vivo [91], or more recently to study human erythropoiesis [90]. As changes in the epigenome, transcriptome, and proteome occur on different time scales, methods for measuring regulatory dynamics, gene expression, and protein information from the same cell have been developed giving rise to single-cell multiomics [114, 118]. Examples of that are the parallel transcriptome profiling and chromatin accessibility and/or DNA methylation analysis [93, 94]. Interestingly, one of these studies using methylome and transcriptome sequencing (MT-seq) was able to reveal unknown associations between heterogeneously methylated regulatory elements and the transcription of pluripotency genes [93]. Paired protein and transcriptomic readouts can also be achieved [96, 97] and in addition, there are techniques that combine single-cell chromatin and proteomic profiling [98]. These are novel techniques that are still unexplored in the stem cell field but the combination of such omics are starting to arise with pioneering work being developed [119, 120]. Overall, these single-cell omics are emerging at a lightning pace providing a transformative view of the biology of the cell and the inner GRNs,



3  Stem cell engineering

15

and consequently, bringing insights into how to dissect stem cell regulatory complexity for bioengineers to take control over the stem cell system in a reverse engineering manner. 3.1.2  Forward engineering using synthetic biology While reverse engineering makes use of in  vitro studies to understand genetic circuits and further improve current cellular systems based on this information, the engineering of such genetic circuits, with in silico predictions that can be latter tested in vitro, is called forward engineering [121]. With this in mind, synthetic biology is a growing field that as greatly contributed for this approach, aiming to engineer artificial genetic networks to program cell functions with high spatiotemporal precision by making use of genetic engineering and molecular biology [122]. This process can include not only the engineering of the endogenous GRNs of a cell but also the addition of new synthetic genetic networks (SGNs), that can coexist or replace the natural GRNs [43]. These SGNs are composed by multiple interconnected gene switches allowing bioengineers to precisely control gene expression in a time- and context-dependent manner, that is, for a given controllable input, bioengineers can precisely predict and generate a desired output [122]. Therefore, synthetic biology opens the door for a wide range of applications in the stem cell field being a step further toward advancements in personalized medicine. The engineering of SGNs makes possible to drive stem cells into a desired cell fate improving the stem cell differentiation process for further medical application. An example of that is the differentiation of hiPSCs into β-like cells [104]. By using synthetic gene switches with differential sensitivity to a molecule known as vanillic acid, it was possible to design an SGN (or as they called it, a lineage control network) capable to control the expression of three transcription factors in hiPSC-derived pancreatic progenitor cells. This mechanism led to the generation of ­insulin-secreting β-like cells with a higher efficiency than the achieved by other conventional techniques. Other publications have reported different strategies for the overexpression of key transcription factors, such as doxycycline inducible system used already to differentiate hPSCs into neurons, skeletal myocytes, and oligodendrocytes [103]. Interestingly, this kind of approach can also hold great potential for regenerative medicine itself. Since cell transplantation requires pure populations of differentiated cells without the tumorigenicity potential of reminiscent undifferentiated hPSCs, technologies have been developed to control the engrafted cells by the elimination of the undesired ones [101, 102]. In addition, future SGNs may allow such synthetic networks to regulate cell proliferation, secretion activity, and apoptosis during regenerative medicine approaches [123].

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3.2  Engineering at the tissue level The growing knowledge about hiPSC differentiation triggered the development of new three-dimensional (3D) culture technologies that bring together multiple organ-specific cell types and to which the name organoids was given [124, 125]. Nowadays, an organoid can be considered a stem cell-derived 3D structure that through a self-organization process can recapitulate biological parameters like the spatial arrangement, cellcell, and cell-ECM interactions, providing a better model of the in  vivo anatomy and physiology of a given organ [126, 127]. Organoid technology has its origins in the 1970s with two-dimensional (2D) cocultures of primary human keratinocytes and 3T3 fibroblasts originating epithelial colonies resembling human epidermis [128]. Since then, the emergence of 3D culture systems has stimulated organoid research to the stage of what is today considered an organoid. Pioneer work was developed with 3D organoids of mammary gland by Mina Bissell [129] but, the real step forward in this field was given by the work of the groups of Yoshiki Sasai and Hans Clevers on optic cup [130] and intestinal organoids [131], respectively. The knowledge to create such structures enabled scientists to generate organoids from a multiplicity of different cell sources and continues to nurture numerous new efforts, driving forward the complexity of this technology (for review of main organoids types already developed: Ref. [126]). Organoid self-organization occurs through self-assembly, s­ elf-patterning, and self-morphogenesis [132, 133], but for the successful induction of these processes in vitro, culture conditions are determinant. Accordingly, biophysical characteristics must be taken into consideration. One can use solid ECMs to entrap 3D cell aggregates, which has been done with intestinal [131], cerebral [134], and gastric organoids [135], or follow a different strategy by simply using a scaffold-free approach, like in optic cup [130], cerebellar [136], and liver bud organoid generation [137]. Regarding biochemical signals, it is important to understand how self-governing the formation of a specific organoid is, since it can rely exclusively on endogenous signals, or may depend on the addition of exogenous cues at different time points and lengths. In fact, some organoids are exclusively driven by endogenous signals, for example, mouse optic cup organoids [130], while others are initially inducted by exogenous signals followed by self-organization solely relying on endogenous cues, for example, human kidney organoids [138], or alternatively be dependent of the continuous supplementation of exogenous signals, for example, human gastric organoid [135]. Finally, another critical parameter is the starting cell population, since organoids can be derived from a single cell, for example, intestinal organoids [131], from a homogeneous cell aggregate, for example, optic cup organoids [130], or from a coculture of different cell types, for example, liver bud organoids [137].



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3  Stem cell engineering

TABLE 1.2  Engineering at the tissue level Approach Biophysical signals

Biochemical signals

Examples

References using stem cells

Chemically defined scaffolds

Hyaluronic acid Polyethylene glycol

[139–142]

Topographically structured scaffolds

Micropatterning   Soft lithography   Robotic printing Nanopatterning   Electro-beam lithography   Electrospinning

[143–146]

Spatiotemporal dynamic scaffolds

Temporal control   Chemical, light, magnetic, or thermal stimuli Spatial control   Mechanical gradients

[147]

Spatial positioned cells

Bioprinting

[148]

Vascularized tissues

Integration of endothelial cells

[137, 144, 149]

Spatiotemporal delivered ligands

Light-triggered activation Microbeads Nanoparticles Microfluidics

[150]

Controlled juxtacrine signaling

Droplet-based microfluidics Micropatterning

[151–153]

Bioengineering can provide new tools and technologies for organoid generation. Particularly, it can help addressing the unmet need of creating tissues and organoids that mimic the anatomical and physiological features of the different organs in the human body so they can be used in disease modeling and drug discovery [37]. Below are summarized the most recent technological innovations in organoid engineering (Table 1.2). 3.2.1  Engineering biophysical signals Matrigel is an obvious choice when supporting ECMs are used for organoid generation, but this matrix presents some critical limitations. Besides not being well defined in what concerns to chemical composition and facing lot-to-lot variation, Matrigel does not cooperate with morphogenesis, spatially limiting the development of complex structures [126, 127]. Thus, to improve organoid architecture beyond the limits of s­ elf-organization, the bioengineering of chemically defined scaffolds and hydrogels with tuneable stiffness are of great interest. Some of the materials already used

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1.  Pluripotent stem cell biology and engineering

include hyaluronic acid (HA), tested in cerebral ­organoids [140], or the synthetic polymer polyethylene glycol (PEG), tested for neural tube morphogenesis in organoids [142]. In fact, it is well known that the surrounding ECM can regulate stem cell fate [154]. Studies on the differentiation of hESCs into mesoderm derivatives were performed varying the stiffness of the scaffold and leading to distinct outcomes in terms of differentiation [141]. In addition, some synthetic hydrogels are already being tested to transplant human intestinal organoids into mouse models [139]. Topographical features can also be added to these materials to mimic those found in the in vivo environment. To this end, micropatterning and microfabrication technologies play an important role controlling tissue geometry and environmental factors [43]. Examples of these technologies are soft lithography, using microcontact printing, replica molding and photolithography techniques, and robotic printing (for a detailed review of microtechnologies, see Chapter 7). A classic example is the use of microwells, usually made of poly(dimethylsiloxane) (PDMS) by soft lithography, to control organoid shape and size [155]. For a more personalized design, microstructured collagen gels are a good example of the use of these technologies, where a soft lithography approach using the replica molding technique is used to resemble the crypt architecture of the small intestine [145, 146]. On the other hand, ECM design at the nanoscale can be achieved using electro-beam lithography, to create computer-guided surfaces with nanopatterns, or electrospinning, responsible for the formation of nanofibrous substrates [127]. Besides topographical features, it is known that in  vivo biophysical signals are presented in a spatiotemporal fashion, shaping the developmental process of tissues. For that reason, dynamic hydrogels whose biophysical properties can be modulated, in both space and time, have been developed [126]. For scaffolds to match different developmental stages overtime mechanically, several techniques have been applied based on chemical, light, magnetic, or thermal stimuli [156]. An example of that was published by Lutolf and coworkers using a chemically induced change by hydrolysis from a static PEG to a mechanically dynamic PEG, allowing the alleviation of accumulated compressive forces [147]. Accordingly, this method initially favored cell expansion and then supported organogenesis in intestinal organoids by temporal modulation of the matrix. Along with temporal modulation, some spatial control has also been tested over hydrogel mechanical properties using, for example, mechanical gradients [156]. Spatial positioning of the different cell types within the generated tissue is another parameter on which bioengineers are focused, particularly trying to improve architecture and mimicking in vivo anatomical arrangements. Strategies like bioprinting are able to control not only cell-ECM but also cell-cell interactions as well, using cells together with a b ­ iomaterial,



3  Stem cell engineering

19

the so-called bioink, to precisely positioning them in a layer-by-layer ­manner [127, 157]. Indeed, this technology has already been applied to many stem cell derivatives [158] and a good example of that is the bioprinting of hiPSC-derived hepatic tissue [148]. In this work, the assembly of hiPSC-derived hepatocytes with other supportive cells, featured in the human liver, was performed using a microscale hexagonal architecture. This strategy led to improved morphological organization, upregulation of liver-specific genes, and enhanced functionality. 3.2.2  Delivery of biochemical signals Most cellular functions, from the embryo to adulthood, depend on the continuous supply of nutrients and soluble factors to the cells. Analogously, bioengineers have used different strategies to properly deliver such factors to stem cell-derived tissues and organoids [127]. One possible solution is the integration of endothelial cells (ECs) or their progenitors in the engineered constructs or tissues [126]. As an illustrative example, recent progresses have been made in liver organoids with the first report using a progenitor population of adult mouse liver [149] and, in a subsequent study, a human ­liver-like organoid structure called liver bud being generated by mixing three cell populations: hepatic endodermal cells derived from hiPSCs, human ECs and human mesenchymal cells [137]. This approach could recapitulate the early steps of hepatogenesis ending in a vascularized human liver bud organoid that showed improved functionality in producing key liver enzymes. More recently, the same authors were able to generate vascularized liver bud organoids by only using hiPSCs as a cell source for the tree cell types [144]. It is important to highlight that this kind of approach also improves the biophysical properties of the tissue by facilitating oxygen delivery to the cells. In fact, vascularization facilitates the delivery of nutrients and soluble factors through the tissue, but desired soluble factors can be delivered as well using special platforms designed for this intent. Traditionally, these soluble signaling molecules are presented in a uniform fashion to the cells cultured in  vitro but, once again, spatiotemporal control of biochemical signals is crucial. For example, delivery of such cues can be achieved by means of light-triggered activation of caged molecules that are masked by a photodegradable moiety in a hydrogel [150, 156]. Also, microbeads and nanoparticles, in all possible forms, play an important role in the spatiotemporal control of biochemical signals. These platforms can be loaded with required signaling molecules and conjugated on the surface of the cells [127]. The interesting part is that they can release its content in a bioresponsive manner, that is, the content is only released when proper molecular cues are present. Microfluidic devices can also be used to deliver soluble ligands mimicking signaling gradients present in vivo. This technology not only has

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the ability to manipulate the flow rate and flow profile, but also recent advances have shown its possible high-throughput nature [155]. In the past, several microfluidic devices were developed [159], to culture cells from blood vessels [160], muscles [161], bones [162], airways [163], liver [164, 165], brain [166], gut [167], and kidney [168]. All the above strategies are mainly focused on the delivery of diffusible molecules, but some biochemical signals like juxtacrine signals are contact dependent. In order to study these signals in an insulated manner, ­droplet-based microfluidic devices are an alternative approach creating ordered cellular structures by controlled encapsulation and pairing of single cells in selected ECMs [155, 169]. These techniques were tested in the past ­combining two different blood progenitor cell lines [152] and more recently combining mesenchymal stem/stromal cells (MSCs) and human umbilical vein endothelial cells (HUVECs), as proof-of-concept studies on the utility of such platform to interrogate intercellular communications [153]. In addition, micropatterning technologies can also be applied to study juxtracrine signals. Proof of this are cell-cell interaction arrays using DNA-programmed adhesion, used to study the dynamics of single adult neural stem cell fate decisions, based on the simultaneous presentation of juxtacrine signals [151]. 3.2.3 Organ-on-a-chip One outstanding question that remains is how can bioengineers design, using biological principles, meaningful in  vitro models in a scale that matches those found in human tissues and embracing technological innovation. Accordingly, combining microfluidic devices with tissue engineering, the so-called organ-on-chip concept has emerged. This term was coined in 2010 by Donald Ingber who described the design of a microfluidic chip capable of reconstructing in vitro the functions of the human lung [170]. These devices, also called microphysiological systems (MPSs), allow the culture of living cells in continuously perfused chambers at the microscale and they can model the physiological functions of a given tissue or organ [171, 172]. For that purpose, they typically involve spatial confinement of multicellular cocultures with fluid inlets and outlets in a controlled environment with incorporated sensors for physiological readouts [159]. At the present, one of the main focus of research in organ-on-a-chip technology is the replication of tissue barriers between different tissues that allow particular organ functions. They can model epithelial barriers like the ones in the lungs, gut and kidney, as well as lymphatic and vascular barriers like the blood-brain barrier, arteries, and microvascular networks. Another focus of organ-on-a-chip technology is the reproduction of the tissue-level organization of parenchymal cells that help in the homeostatic function of organs. This topic is of major concern when r­ eproducing



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organs like the liver or cardiac and skeletal muscle (for detailed list of published organs-on-a-chip: Ref. [159]). However, in organ-on-a-chip devices, cellular composition and tissue structure is normally oversimplified since the idea is not to create a whole organ but to establish a minimal functional unity. For that reason, to accomplish high fidelity in the in vitro modeling of human organs, the synergistic engineering of organoids within organ-on-a-chip devices can be of great significance [173].

3.3  Engineering at the multiorgan level 3.3.1 Body-on-a-chip As mentioned before, the human body can be perceived as a ­multiorgan system whose normal functioning depends on the constant cross talk across multiple organs. To mimic this complexity, organ-on-a-chip technology can model the interaction of a large spectrum of organs, bringing all together in what can be called a body-on-a-chip, or multi-MPS [43, 159]. Therefore, it is expected from these devices to model significant parts of the complexity and dynamics of the human body, using a design that works at the microscale. To study multiple organ interactions in these devices, body-on-a-chip cannot rely solely on individual designs, but in a more complex interconnected platform capable of mimicking vascular perfusion and recapitulating homeostasis [172]. This goal can be achieved mainly by a unidirectional single-loop perfusion or by recirculation of common media capable of simultaneous maintenance of all organ-like structures. Indeed, the study of systemic interaction is older than the concept of organ- and body-on-a-chip, with lung/liver cross talk being studied in 2004 [174]. However, at the moment, the use of body-on-a-chip was already reported with 14 chambers, representing up to 13 different organs [175], with complex devices and greater functional longevity being reported [176]. In the future, improvements to this technology should also have into account interorgan cell migration, and not only biochemical signals, for a more realistic modeling of the human body physiology. This could be important to incorporate the trafficking of cells in the immune system, or to model the migration of cancer cells during metastasis [159]. In regard to cell source, the vast majority of these studies were performed using cell lines and the use of differentiated PSCs in body-on-a-chip is still not a common practice. Bearing in mind the concept of body-on-a-chip, it is easy to understand that these devices are also an interesting platform for drug development and toxicology, providing insights into adsorption, distribution, metabolism, elimination, and toxicity (ADMET), mathematical pharmacokinetics (PK), pharmacodynamics (PD), and drug efficacy [171]. Moreover, the

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­ esire for better in vitro biological models for human disease strongly mod tivates the development of body-on-a-chip devices. In fact, it was already reported the analysis of an intertissue cross talk between gut and liver during inflammation processes [177]. Body-on-a-chip is indeed a milestone in the evolution of tissue engineering [159].

4  Concluding remarks The presented technologies have made major contributions to stem cell research, sparking the imagination of countless scientists with astonishing discoveries and inspiring the future generation of curious minds. Indeed, it is interesting to contemplate the human body, made of many organs and many tissues, “and to reflect that these elaborately constructed forms, so different from each other, and dependent on each other in so complex a manner, have all been produced by laws acting around us”—Charles Darwin, On the Origin of Species. The purpose of stem cell research is to dissect these laws, recreating cells, niches, tissues, and organ-like structures in such a way that they are able to replicate structural and functional characteristics of their in  vivo counterparts. In fact, these tissue and organ surrogates would present similar physiological functions, making them promising sources for medical applications in regenerative medicine, disease modeling, and drug screening. This progress would critically allow the development of precision medicine approaches, potentially revolutionizing clinical practice and drug development [31, 123].

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C H A P T E R

2

Process development and manufacturing approaches for mesenchymal stem cell therapies Pedro Silva Coutoa, Alexey Bersenevb, Qasim A. Rafiqa a

Advanced Centre for Biochemical Engineering, Department of Biochemical Engineering, University College London, London, United Kingdom, b Cell Therapy Laboratory at Yale-New Haven Hospital, Yale University, New Haven, CT, United States

1  Introduction There is a new therapeutic paradigm in medicine which utilizes whole cells, either somatic or genetically modified, as the active ingredient of the therapy (Fig.  2.1). This is the cornerstone of the cell and gene therapy (CGT) industry, a nascent field which now includes over 20 approved products and several hundreds of clinical trials using multiple different cells types to target a myriad of diseases [1]. Cell types being considered as CGT candidates include human mesenchymal stem/stromal cells (hMSCs), gene-modified T cells, hematopoietic stem cells (HSCs), neural stem cells (NSCs), progenitors derived from embryonic stem cells (ESCs), or induced pluripotent stem cells (iPSCs) [1–3]. In addition to whole cells, other therapeutic interventions, such as extracellular vesicles (EVs) (e.g., exosomes [4]) are also being considered. The range of products that fall under the CGT spectrum are broad and will require distinct bioprocessing, manufacturing, and delivery strategies. While somatic cells are usually isolated from patients and then expanded before being reinjected into the patient (e.g., hMSC or HSC), genetically modified or reprogrammed cells are engineered in vitro to express a functional receptor (e.g., CAR-T cells) or are differentiated toward a defined lineage (iPSCs). Exosomes are Copyright © 2020 Tiago G. Fernandes, Maria Margarida Diogo & Joaquim M. S. Cabral. Published by Elsevier Inc. All rights reserved. https://doi.org/10.1016/B978-0-12-816221-7.00002-1

34

2.  Process development and manufacturing approaches for MSC therapies

Cell and gene therapy products

Somatic cells

Genetically modified or reprogrammed cells

Exosomes

FIG. 2.1  Schematic representation of different CGT products.

cell-free products secreted by many cell types and contain small membrane vesicles that contain are proteins, mRNA or miRNA which are key mediators of the intercellular communication [5–7]. Process development and manufacturing strategies are required to improve efficiency and scalability of each of the key unit operations such as cell isolation [8, 9], transduction/transfection [10], expansion [11, 12], purification [13, 14], and formulation [15]. Due to their immunomodulatory properties and in vitro differentiation ability, hMSCs are a key candidate for CGT applications with approved products on the market using hMSCs and are the focus of many clinical trials investigating their use for a range of clinical indications [1,16–26]. As many hMSC products are in late-stage clinical trials, there is a pressing need to develop large scale, controlled, and robust manufacturing methods for hMSC production. This is especially critical for allogeneic products where the approach favors economies of scale which can be achieved by using bioreactor technologies to improve scalability as well as reducing the dependency on human operators and increasing process monitoring and control capability [27–29]. Three main strategies have been employed for the production of hMSCs for either research and/or clinical purposes: (1) monolayer culture systems, (2) spheroid culture, and (3) adherent bioreactor culture. Given that hMSCs are anchorage dependent [30], the lack of a matrix for the cells to attach will result in the cells undergoing a form of programmed cell death called anoikis [31]. Monolayer culture systems are routinely used in research and even during the manufacturing process used in clinical trials [32–34], but they have several disadvantages. Monolayer culture platforms are traditionally open systems with limited monitoring and control capability and are often reliant on incubators to maintain pH and CO2 conditions. Moreover, they are restricted by the available surface area which presents scalability challenges and space constraints. Such systems also



2 Isolation

35

require human operators to perform medium exchanges or cell passaging which inherently increases the risk of variation and contamination. Due to the open nature of monolayer culture systems, they need to be housed in a cleanroom environment which ultimately increases overall production costs [35–38]. Spheroid hMSC cultures involve the production of hMSCs as aggregates which can reach 1 mm in diameter and have been reported to demonstrate improved immunomodulatory properties [39,40]. However, spheroid culture presents critical challenges for manufacture, not least the inability to control their size during the expansion process. This is especially critical because an increase in spheroid diameter leads to mass transfer problems which may trigger cell necrosis and inhomogeneity in the final cell product. Moreover, disaggregation or generation of a single-cell population of cells ready for administration can be challenging, with little research demonstrating effective and reproducible methods for single-cell generation from large-scale hMSC aggregate culture. In addition to the expansion methods outlined above, another manufacturing approach investigated for the large-scale production of hMSCs is the use of bioreactors. Multiple types of bioreactor have been used to expand hMSCs [12,38,41–50], with each system possessing advantages and disadvantages (described later). Given the anchorage-dependent nature of hMSCs, a key challenge for any bioreactor is to provide a surface for the cells to attach, in addition to controlling and modulating other key parameters (Fig. 2.2). These surfaces can be in the form of an external substrate which is added to the bioreactor (such as microcarriers using in stirredtank bioreactors (STRs) [43,51–54] or rocking motion bioreactors [45, 55]) or a physical structure incorporated as part of the bioreactor system by default (hollow-fiber bioreactors [50, 56, 57], fixed- or rotating-bed bioreactors [58–60]). The purpose of this chapter is to detail the process development and manufacturing approaches for hMSC production and outline the challenges, translational bottlenecks, and strategies to address these moving forward.

2  Isolation Different tissue sources can be used to isolate hMSCs. Traditionally, bone marrow (BM) has been the predominant source hMSCs, but today it is possible to isolate hMSC from a myriad of other adult and perinatal tissues. Several groups have successfully isolated hMSCs from adult sources such as adipose tissue (AT) [61], synovial fluid [53], peripheral blood (PB), and also perinatal ones such as umbilical cord tissue (UCT) [62, 63], and umbilical cord blood (UCB) [61] among others. While this

36

Key parameters to control/modulate during microcarrier-based expansion in bioreactors

Seeding density

Oxygen supply and aeration strategy

Microcarrier

pH, DO

Feeding regime Temperature

Agitation speed

Medium formulation and supplements used Bioreactor type

FIG. 2.2  Schematic representation of the parameters to be studied to increase cell yield of a microcarrier-based hMSC manufacturing processing.

2.  Process development and manufacturing approaches for MSC therapies

hMSC source



37

2 Isolation

­ otentially presents multiple sources of hMSCs, a major question arises as p to whether the cells, from a biological perspective, are the same or have different characteristics. Several authors have reported that different tissues yield cells with distinct growth kinetics, metabolic profile, differentiation ability, immunophenotype, and gene expression [64, 65]. From a bioprocessing perspective, this presents an additional complexity in that different isolation protocols need to be established for each source used. Although for nontissue sources (such as BM, PB, or UCB) adherence methods are usually employed, it is also possible to start the bioprocess by cell sorting. For tissue-based sources (e.g., AT or UCT), the isolation procedure starts with the separation of the cellular fraction from the tissue which can be performed by using enzymatic or nonenzymatic methods such as using explants or semiautomated methods (Fig.  2.3) [66, 67]. It was recently reported that the use of collagenase is directly associated with lower cell yields when compared to nonenzymatic methods [66]. The same study also reported that the immunophenotype was negatively affected by the use of enzymes suggesting that use of collagenase might affect the in vivo function of the isolated cell populations [66]. To avoid using enzymes and to move toward closed and automated systems, some companies have developed enzyme-free, closed, automated, tissue isolation systems such as the AC:Px system (Auxocell Laboratories Inc., United States) [66] and StromaCell system (MicroAire, United States) [68], e.g.,. Others have developed automated, closed tissue isolation systems which require the use of enzymes, such as the Celution system (Cytori Therapeutics, United States) [69] and the Sepax system (Biosafe-GE HealthCare, Switzerland) [70]. For bioprocessing and quality control purposes, it is crucial to assess viability and identity so as to ensure that following isolation, the desired cell population has been successfully isolated without compromising cell quality.

Enzymatic methods Explants

Isolation Nonenzymatic methods

Semiautomated

QC/QC: Isolation success rate, viability, identity, potency.

FIG. 2.3  Options available during the isolation step and critical quality attributes (CQA) to be studied.

38

2.  Process development and manufacturing approaches for MSC therapies

3  Monolayer expansion The quantity of cells obtained after isolation can vary (Table 2.1), but in most instances, are insufficient to meet the dose criteria for clinical applications; expansion of hMSCs is therefore required to meet the dose requirements. Traditional culture methods involve the use of monolayer expansion systems to either produce enough cells for small-scale clinical trials for autologous applications [1, 18, 20–26], or to generate a sufficient number of cells to inoculate a larger-scale bioreactor, where initial cell densities studies range from 4000 to 8000 cell/cm2 [43, 44, 51, 75]. Several studies have reported a decrease in growth kinetics, differentiation ability and potency of hMSC after extensive monolayer culture [76–79], as such, the number of population doublings should be kept to a minimum to avoid detrimental impact on the cell product quality. Population doublings is an important measurement that must be tracked as it establishes the number of cellular divisions (Fig.  2.4). However, in research and clinical practice, passage number is routinely used as an indicator for cell “age” in vitro. Multiple groups who have undertaken bioreactor studies have expanded the hMSCs in monolayer systems prior to bioreactor inoculation and have reported using cells at P4 [12, 44, 80] up to P8 [81] or P10 [82, 83]. In some cases, such as the operation of hollow fiber bioreactors, there is no need to preexpand the cells in tissue culture flasks as this aspect of the bioprocess can be performed directly in the bioreactor [41, 57, 84]. The advantage of using this approach is to reduce the use of multiple devices and human handling. The traditional monolayer culture systems used for 2D expansion such as T-flasks or cell stacks/cell factories are heavily dependent on human handling and require other pieces of equipment to maintain temperature and CO2 concentration. For quality control purposes it is a key to evaluate growth rate, metabolic production and consumption TABLE 2.1  Summary of the quantity of cells isolated per gram of UCT or AT processed in multiple studies Tissue source UCT

Cells per gram of tissue processed (cell/g) 6

(2.9 ± 1.4) × 10

[8] 5

(7.20 ± 0.53) × 10

[66]

5

[71]

(4.89 ± 3.2) × 10

5

[72]

(4.29 ± 0.46) × 10

5

[73]

4

[74]

(1.75 ± 0.94) × 10 AT

References

6.6 × 10



39

4  Bioreactor-based hMSC expansion

High doubling level

Potency and genetic stability

Expansion in monolayer Low doubling levels

Number of doublings

QC/QC: Growth rate, metabolic production and consumption rates, identity, potency and genetic stability.

FIG. 2.4  Options for monolayer expansion and important CQA to assess at that stage.

rates, identity as well as genetic stability and potency that to assess if the cells are not senescent and chromosomal stability was not affected.

4  Bioreactor-based hMSC expansion When selecting a bioreactor with which to expand hMSCs, several criteria should be taken into consideration. Ideally, the bioreactor should be a closed system, automatable, controlled, and single-use system [27]. While the majority of bioreactor studies focus on hMSC expansion in large-scale STR platforms using microcarriers, it is important to recognize that are other options besides STRs including rocking motion bioreactors, hollow fiber bioreactors, packed-bed bioreactors, and vertical-wheel bioreactors. A comparison of the systems’ characteristics is provided in Table 2.2.

4.1  Stirred-tank bioreactors Although the majority of clinical trials uses monolayer expansion systems [32–34, 85], several research groups have published ­microcarrier-based expansion protocols in STRs [12, 43–46, 51, 53, 82, 86–88]. While in the early days of development, hMSC STR-microcarrier processes were p ­ erformed TABLE 2.2  Summary of the final cell densities obtained and surface area available in rotating bed and hollow-fiber bioreactors Type of bioreactor

Area available (cm2)

Final cell density (cell/cm2) 5

Rotating-bed bioreactor

6000

5.8 × 10

2000

1.2 × 10

Hollow-fiber bioreactor

21,000

5.6 × 105

References [58]

5

[57] 5

[84]

2

[50]

4

[41]

4.67 × 10 9.53 × 10 2.16 × 10

40

2.  Process development and manufacturing approaches for MSC therapies

at small scale (100 mL) [45, 51, 53, 87, 89], the most recent studies reported have employed working volumes of 35 [46] and 50 L [86]. An overview of the final cell density, bioreactor working volume and hMSC source for the various hMSC STR-microcarrier studies is provided in Fig. 2.5. Regarding the tissue source, the two studies that achieved highest cell densities used synovial membrane [87] and AT [53] achieving 8.8 ± 0.2 × 105 and 12.5 ± 0.5 × 105 cells/mL, respectively. It is notable that the higher cell yields are obtained at the lower working volumes, likely to be the case as multiple runs can be conducted at these scales, but over time, it will be necessary to translate these final cell yields obtained in these smallscale studies to the multiliter-scale large-scale STRs. This is expected to improve as research groups generate more data and understanding of the bioprocess. It is a key to highlight that working volume and the target of the therapeutic product are closely related; while for allogeneic products larger scales will be used to produced batches for several different patients, in an autologous context, every single patient will have a personalized batch which will therefore necessitate different manufacturing and business model paradigms. It should also be mentioned that the majority of the studies were conducted with adult hMSC sources whereas only a few used perinatal hMSC sources [12, 51, 82, 87, 90–93]. As perinatal hMSC have higher proliferative capability than their adult counterparts in monolayer, it is expected that bioreactor-based expansion will achieve higher cell yields with perinatal cells compared to adult ones [64, 76, 94].

FIG.  2.5  Graphical representation of the maximum cell yields reported in literature together with working volume in the studies.



4  Bioreactor-based hMSC expansion

41

Several different commercially available STR systems are currently available and many have been used for hMSC-microcarrier culture including the Biostat B (Sartorius) [44, 46, 49, 82, 95], Biostat Q Ambr15 (Sartorius) [96], Ambr250 (Sartorius), BioFlo/CelliGen (Eppendorf) [12, 54], DASGIP (Eppendorf) [54], UniVessel (Sartorius) [46, 88] and Mobius (Merck) [86], among others. Each of these STRs differs with respect to the vessel geometry, working volume, impeller design, and some offer different levels of control. Little research has been undertaken comparing the growth of hMSCs on microcarriers in different STR systems, however given the potential for different fluid dynamics in different STR systems, there may be observable differences with respect to cell growth and potentially cell functionality. When choosing a bioreactor to expand hMSC several criteria should be taken into consideration. Ideally, the bioreactor should be a closed system, to have at least inoculation and harvesting automated, being disposable avoiding sterilization issues before the process and low running costs [27]. Recently, in 2017 Rafiq et al. used the Ambr15 (Sartorius) [96] to module key parameters to increase cell yield. In this study it was reported a decrease of the lag phase, by improving the mixing of the bioreactor as well as improving cell attachment to the microcarriers. The study also reported a decrease in cell viability coefficient of variance by using the Ambr15 and changing from serum-based medium to serum free. This suggests that SF medium combined with automated and controlled bioreactors are a key to minimize process variability increasing batch-to-batch consistency.

4.2  Rocking motion bioreactors Rocking motion bioreactors (such as the WAVE Bioreactor or the BIOSTAT RM) can be either used to expand suspension or adherent cells. In 2016, Jossen et al. expanded AT-hMSC with a BIOSTAT RM at three different working volumes (0.5, 1.0, and 1.5 L) using polystyrene-based microcarriers. The authors were trying to predict the effects of shear stress using rocking angle and rocking rate at three different working volumes as studied parameters. The number of cumulative doublings using BIOSTAT RM was much lower than the control system (spinner flask with a working volume of 100 mL) [45]. Rocking motion bioreactors can also be used to expand hMSC in the spheroids form. This was the approach of Tsai et al., who reported that the smaller rocking angle the bigger hMSC spheroids size is. In addition to this correlation between rocking and hMSC spheroid size, the authors also reported that spheroid size can be regulated by seeding density and cultivation time [55]. With respect to functionality, the authors reported that spheroid-expanded hMSC had higher clonogenic ability and higher stemness-associated gene expression than compared to the monolayer controls [55]. Using the same type of bioreactor Timmins et  al. also developed a closed system to isolate and expand placenta-derived hMSC.

42

2.  Process development and manufacturing approaches for MSC therapies

The authors obtained 12-fold expansion after 7 days with a working volume of 0.5 L. Aiming to scale-up this system, the authors estimate to be possible reaching 7000 doses (using as 5×106 cells/kg as target dose) [93].

4.3  Rotating-bed bioreactor In 2013, Reichardt et al. successfully expanded hMSC isolated from the UC vein in a rotating-bed bioreactor. After 9 days the authors harvested 5.8×105 cell/cm2 corresponding to a fold expansion of 38.1 ± 6.1. To perform this work, a rotating bed of polycarbonate with a surface area of 6000 cm2 was used. Growth kinetics and metabolic rates of hMSC expanded in this system were comparable to the ones obtained with a T-25 used as control [58]. One year later (2014) developed a single use rotating-bed bioreactor with a total surface area of 2000 cm2. The authors expanded UCT-MSC and reported a total yield of 24.6 × 106 (~1.2 × 105 cell/cm2) with a total of 8.2 ± 0.8-fold expansion. The authors also reported that the expanded cells kept their immunophenotype as well differentiation ability [97].

4.4  Hollow fiber bioreactor In 2013, Rojewski et  al. expanded BM-hMSC using the automated ­ ollow-fiber Quantum bioreactor (Terumo BCT, United States). The study h reported a final cell density of 5.6 × 103 cell/cm2 after 5.9 days in culture, yielding a total of 118 × 106 cells. It must be noted that high cell densities were obtained because the hollow-fiber surface area is 2.1 m2. It was also possible to expand BM-hMSC right after cryopreservation using this bioreactor. By using the bioreactor immediately after thawing, monolayer culture systems are avoided which constitutes an advantage of the Quantum bioreactor [57]. In the same year, Nold et al., obtained up to 98 × 106 cells (corresponding to a 20-fold expansion) after 13 days of culture also using the Quantum bioreactor [84]. The expansion efficiency of the same bioreactor was also evaluated by Jones et al. They used several BM-hMSC donors and expanded the cells for four passages. To assess the clinical feasibility of using the Quantum, the genetic stability of the expanded cells was performed. Neither chromosomal aberrations nor DNA damage was found ensuring the genetic stability of the hMSC expanded with this bioreactor [50]. A multidonor study using the Quantum bioreactor was carried out by Lambrechts et al. in 2016, who reported cells yields of 316–444 million of cells after 8 days of culture [41]. It should be mentioned that as the hMSC growth attached to the hollow fiber, hMSC expanded using a hollow-fiber bioreactor might experience nutrient and oxygen gradients throughout the length of the fiber [98, 99]. Table 2.2 shows between the final cell yield achieved with adherence-based bioreactors reported in the literature.



4  Bioreactor-based hMSC expansion

43

4.5  Fixed- or packed-bed bioreactor Using a fixed-bed bioreactor Mizukami et al. achieved after 7 days of culture a total of 4.2 ± 0.8 × 108 cells representing a fold increase of 7.0 ± 1.4. In this study, the cells were preexpanded in monolayer before being transferred and expanded in the fixed-bed bioreactor. The authors also reported that the hMSC expansion process using the fixed-bed bioreactor did not change the cellular immunophenotype [59]. Using a packed-bed bioreactor and perfusion conditions, Osiecki et al. in 2015 successfully expanded placenta-derived hMSC. In this case authors did not preexpand the hMSCs in tissue culture flask, rather expanded them in the bioreactor immediately after isolation. They also suggested that expanding cells in the bioreactor postisolation may have a negative impact on keeping hMSC differentiation ability due to the large amounts of red blood cells and debris generated during the isolation step [60].

4.6  Vertical-wheel bioreactor Using a Vertical-Wheel bioreactor with a working volume of 2.2 L, Sousa et al. expanded BM-hMSC during 14 days, achieving a maximum cell yield of 3 × 105 cell/mL. The final cell yield reported was not higher than control group (STR-Biostat Qplus STR 0.25 L). Although expanded BM-hMSC showed similar growth rates and expansion factors in both bioreactors, a lower percentage of apoptotic cells were reported using the Vertical-Wheel bioreactor. The authors suggested that due to its specific design, the VerticalWheel bioreactor offers a more efficient mixing environment while offering a low shear stress environment. However, given the limited number of studies using this system, it is difficult at this stage to make this conclusion. No differences were reported in the clonogenic ability between the cells expanded with the two different bioreactors (Table 2.3) [100]. Bioprocesses involving the use of different expansion systems will likely have different measures of process efficiency which may be linked with the product’s CQAs. For example, a microcarrier-based STR expansion process will have an efficiency measure associated with the percentage of cell attachment, particularly in the early stages of culture, so as to measure and understand the extent of the lag phase [95, 96]. Cell growth kinetics, metabolic production, and consumption rates, differentiation ability immunophenotype and potency will need to be included in the CQA panel. While growth kinetics and metabolic rates will provide information about the cell growth and potential senescence of the cell ­population, differentiation and immunophenotype are a key to monitor cell identity. As potency is usually used as predictor of the clinical effectiveness of CGT [101,102], it should be carefully assessed not only after expansion but also after thaw just prior administration.

44

2.  Process development and manufacturing approaches for MSC therapies

TABLE 2.3  Summary of the differences between bioreactors used to expand hMSC Type of bioreactor

Characteristics

Microcarriers used?

References

Stirred tank bioreactor

• • • •

Available as single-use platform Easily scalable system Inline monitoring system Controlled gassing strategy

Yes

[44, 46, 49, 82, 87, 95, 96]

Rocking motion bioreactors

• Available as single-use platform • Scalable up to 500 L • Sampling system not as straightforward as other bioreactors • Microcarrier aggregate formation may occur

Yes

[45, 55, 93]

Rotating-bed bioreactor

• Limited scale-up • Cell sampling is not possible during culture

No

[58, 97]

Hollow-fiber bioreactor

• Available as single-use platform • Promotes cell-cell and cell-matrix interaction. • Spatial concentration gradients formed along the hollow-fibers • Cell sampling is not possible during culture

No

[41, 50, 57, 84]

Fixed- or packed-bed bioreactor

• Promotes cell-cell and cell-matrix interaction • High surface area • Cell sampling is not possible during culture

No

[59, 60]

Vertical wheel bioreactor

• Available as single-use platform • Scalable up to 500 L • Vertically orientated wheel instead of impeller • Low shear stress environment

Yes

[100]

Modified from Rodrigues CaV, Fernandes TG, Diogo MM, da Silva CL, Cabral JMS. Stem cell cultivation in bioreactors. Biotechnol Adv 2011;29(6):815–29. https://doi.org/10.1016/j.biotechadv.2011.06.009.

4.7  Bioprocess monitoring and control: parameters for optimization 4.7.1 Temperature and pH Temperature and pH are two key parameters that need to be controlled to optimize cell growth [103, 104] (Fig. 2.6). Commercially available bioreactors have different mechanisms to control temperature. While the Biostat B Plus (Sartorius) [44] and CelliGen (Brunswick) [12] use a water jacket,



4  Bioreactor-based hMSC expansion

45

heating blankets are used in the UniVessel (Sartorius) [46] and Mobius 3 L, 50 L (Applikon) [86] to keep temperature at 37°C. The rationale behind using this temperature is to try mimicking what happens in vivo. It should be mentioned that simpler systems such as spinner flasks rely on incubators to maintain the temperature. There are several examples from the biochemical engineering industry where a shift in temperature increased the production of protein [105] or fragments of antibodies [106] using CHO cell lines. It is expected that bioprocess optimization studies will be published soon about the manufacturing of exosomes [4] and although 37°C optimize cell growth, it is unclear if that is the case for exosomes too. From a bioprocessing perspective, to control pH several different options can be used: either using CO2 [44, 49, 81, 87, 95] or air [82] and or sodium bicarbonate [12, 54, 86] as a buffer. During glucose degradation into lactate and glutamine in ammonia, there is a drop in pH associated to the accumulation of these metabolic products. [107] It was reported by Schop et  al. that the accumulation of metabolic products inhibits cell growth [108]. While pH is not easy to monitor nor control in static cultures, most of the commercially available bioreactors have control strategies to prevent pH from reaching levels that inhibit cell growth or even cause functionality impairment. As pH is a key parameter to control during a biochemical engineering process, several authors have conducted research efforts to study its impact in hMSC manufacturing. Yuan et al. studied the impact of pH during the attachment phase and concluded that 8.0 was the pH level that optimized cell-microcarrier adherence [109]. Other authors have defined pH working ranges while running microcarrier-based STR expansion of hMSC. Rafiq et al. defined [44] 7.2 and 7.4 while Lawson et al. [88] chosen to control pH to be between 7.2 and 7.8. As referred to previously, Schop et al. investigated the impact of high concentrations of ammonia and lactate in hMSC growth reported that a concentration of 2.4 mM of ammonia and 35.4 mM of lactate without reporting any detrimental effect on differentiation ability [108]. When there is accumulation of lactate and ammonia, not only does the pH increase but also the osmolarity is expected to be higher. It is worth mentioning that the study from Schop et al. was conducted using with static monolayer systems which are poorly controlled when compared bioreactor-based expansion systems. 4.7.2  Oxygen supply and aeration strategy Although hMSCs have been routinely expanded with 100% dissolved oxygen (dO2) or under atmospheric conditions, this does not necessarily mimic what happens in  vivo because in certain niches, cells experience hypoxic conditions [110–112]. In 2010, dos Santos et  al. reported that growth rate and clonogenic ability of BM-hMSC was improved in hypoxia conditions, when compared to normoxia [113]. These results are aligned

46

2.  Process development and manufacturing approaches for MSC therapies

with the work reported by Feng et al. in 2014, who reported that AT-hMSC expanded in hypoxia conditions had the same differentiation ability yet lower levels of apoptosis than normoxia-expanded AT-hMSC [114]. While Krinner et  al. reported that BM-hMSC expanded in hypoxia conditions have an improved chondrogenic ability [115] compared to normoxia, Fang et al. published that AT-hMSC tend to adopt a smooth muscle-like phenotype when expanded in hypoxia [116]. This suggested that hMSC isolated from different sources might be differently impacted by hypoxia. Thus, from a cell biology perspective, it is a key to understand what is the impact of different oxygen concentrations during hMSC expansion as this parameter may change the characteristics of the final product. From a bioprocessing perspective the concentration, the challenge is to understand what is the most efficient gassing strategy (Fig.  2.6). The administration of gasses to the majority of bioreactor can be performed in two different ways: (1) sparging (gas administered directed in the medium) or (2) headspace, when the gas is provided in the headspace area at the top of bioreactor. Different bioreactors can be operated under one or more ­gassing options. While STRs can usually operate with either ­option,

Type of bioreactor

Using microcarriers

Without microcarriers

Stirred tank Rocking motion Vertical wheel Rotating bed Hollow fiber

Temperature and pH

Fixed or packed bed Normoxia Oxygen level

Hypoxia

Bioreactor expansion Oxygen supply and aeration strategy

Medium formulation and supplements

Harvesting

Headspace Gassing strategy

Sparging

FBS supplemented

Platelet lysate

Serum free

Xenofree Chemically defined

QC/QC: Growth rate, metabolic consumption and production rates, identity potency and genetic stability.

FIG. 2.6  Simplified diagram of the options available during the bioreactor-based hMSC expansion.



4  Bioreactor-based hMSC expansion

47

the most up to date version of the rocking motion platforms do not have the option of sparging. As bubble formation has detrimental effect on mammalian cell viability [117–119], sparging does not seem to be an essential gassing option to be included by bioreactor manufacturers that work in the CGT space. In the specific case of rocking motion bioreactors, dissolved oxygen can be controlled by modulating the rocking speed and angle or injecting gas in the headspace [120]. In 2013, Reichardt et al. also reported to inject gas directly into the headspace while using a r­ otating-bed bioreactor to expand hUCT-MSC [58]. In the Quantum bioreactor the gas control is ensured by using a hollow-fiber oxygenator [41, 56, 57]. 4.7.3  Medium formulation and supplements Several medium formulations can be used to support hMSC expansion (Fig.  2.6). Several authors have reported the use of DMEM while other groups choose α-MEM instead. As fetal bovine serum (FBS) contains a plethora of proteins and growth factors, it is currently used to support hMSC growth. Although FBS has been routinely used during the manufacturing of hMSC [85], there are several disadvantages associated to its use. Batch-to-batch variability and potential contamination with prions are some of the problems of using FBS. Moreover, as FBS is obtained from bovine fetus blood serious ethical issues can be raised while discussing its collection. For these reasons and to meet FDA and EMA [121] guidelines, the manufacturing industry is moving toward the use of using serum-free (SF) and xeno-free (XF) options expecting more consistency between batches. The alternatives to the use of FBS are: (1) human platelet lysate (hPL), (2) xenofree, and (3) chemically defined. hPL can be obtained by isolating plasma from blood that is submitted to freeze and thaw cycles after leukocyte removal [122, 123]. Although the use of hPL instead of FBS eliminates the ethical problems, the dependency of undefined components in the process and batch-to-batch variability still persist. With the goal of minimizing process variability, SF/XF and chemically defined medium have been the focus of significant research activity and will likely be the basis for any hMSC cell therapy product moving forward [124]. A major impediment to their use at present is the significant cost associated with SF/XF and chemically defined media. It is worth noting that different supplements or medium formulation will have an impact at the cell level. Multiple studies have reported that hPL-supplemented and/or SF/XF medium lead to higher proliferation levels of hMSC when compared to FBS-based medium [123, 125–127]. Although reporting higher proliferation level for hPL-supplemented medium, Fernandez-Rebollo et al. did not find any difference in the DNAmethylation profiles and the differences at transcriptomic level were also

48

2.  Process development and manufacturing approaches for MSC therapies

moderate. The authors suggest that hPL and FBS supplemented medium do not select different types of hMSC and that the morphological differences are interchangeable [127].

4.8 Harvesting As hMSCs are attached to surfaces during the expansion process (either microcarriers, hollow fiber or other types of structures) once the stationary stage of the culture is reached, the cells need to be harvested. Although there have been numerous studies investigating the expansion of hMSCs in various bioreactors, the literature pertaining to the harvesting procedures is somewhat limited. Various enzymatic reagents are used to harvest cells including Trypsin-EDTA and TrypLE. TrypLE is free of ­animal-derived components which are highly desirable to minimize potential regulatory challenges and to ensure process consistency. Room temperature stability of TrypLE is also an important advantage because trypsin’s stability decrease overtime at room temperature [11]. While dissociation reagents need to be used to detach cells from microcarriers, prolonged use of enzymes may damage cell surface markers in a time and concentration manner [128, 129]. In terms of the harvesting itself, different bioreactors will have different harvesting methods. In a hollow-fiber bioreactor the harvesting procedure is performed via a single-step process usually using a dissociation reagent such as Trypsin-EDTA and TrypLE. On the other hand, if microcarriers are being used, a two-step procedure needs to be performed: (1) to detach the hMSC from the microcarriers to generate a cell-microcarrier suspension and (2) to use a filter to obtain a microcarrier-free cell suspension. The first step is again performed using a dissociation reagent whereas the separation between hMSC and microcarriers is mediated by a filter with pore size smaller than microcarrier size (>125 μm) [11, 52]. The cell recovery from a cell-microcarrier suspension is function of filter pore size with filters. Bigger filter pore size yields higher cell recoveries and with viabilities when compared to filters with smaller pore size [130]. As the detachment of cells from microcarriers includes centrifugations and filtrations [89, 131], viability constitutes an important CQA. Nienow et  al. in 2014 reported that after harvesting, viability remained above 90%. Although viability remained high after the harvesting procedure, it would be interesting to study the impact of this step of the bioprocess in the expression of early apoptosis markers. Although centrifugation and filtration are routinely used during harvesting, this choice has limited scalability [132]. In 2015, Cunha et al. used tangential flow filtration (TFF) and a diafiltration module and obtained more than 80% of cell recovery with 95% of viability [131]. Although a lot of research has been conducted to optimize the upstream part of the MSC manufacturing process, several studies still need



49

4  Bioreactor-based hMSC expansion

to be conducted in the downstream section. Scalability challenges around scaling up downstream processing steps, as well as the recovery optimization of cell-based and cell-free products are some of the key hurdles to overcome (Fig. 2.7).

4.9  Manufacturing paradigms: Autologous and allogeneic Autologous and allogenic approaches are very different from a manufacturing, business, and supply chain model [133]. From a manufacturing standpoint, autologous and allogenic share several commonalities with respect to the initial part of manufacturing process, for example, the isolation of the tissue/material, transport of the tissue/material and isolation of the desired cell population (Fig. 2.8). Upon arrival at the manufacturing site, cells for an autologous or allogeneic manufacturing process would follow very different routes as the target batch sizes for each paradigm are very different [134]. In an allogeneic, universal donor bioprocess, after initial isolation, hMSCs are likely to be expanded in systems that can facilitate large-scale production, for example, hollow-fiber bioreactors or STRs. For personalized, autologous hMSC therapies, the aim is to generate sufficient biological material such that an individual patient can receive one or more infusions (depending on the clinical indication). In this instance, smaller-scale systems, such as Cell Factories, may be appropriate to generate the material required. With respect to the bioprocess itself, in an allogeneic context, a master and working cell bank would be created to generate sufficient supplies of cellular material, while in an autologous setting, any cellular material generated is likely to be delivered to the patient as part of the therapy, and any surplus material (if available) would be banked in case the patient needs further treatment.

Bench scale

Filtration

Centrifugation

Downstream processing

Scalable

Tangential flow filtration

Diafiltration

QC/QC: Viability, apoptosis markers and identity.

FIG. 2.7  Simplified diagram of the options available during downstream processing after the bioreactor-based hMSC expansion and CQA to be studied.

50

Allogeneic

Autologous

Collection of tissue/material from the patient

Isolation of the target cell population Master cell bank

Cell expansion

Expansion fo cell banking

Cell expansion

Harvesting

Downstream processing

Administration to the patient

Transfer to administration site

Short-term storage

Working cell bank

Harvesting

Downstream processing

Administration to the patients

Transfer to administration site

Long-term storage

Quality control

Quality control of the batch

Quality control

FIG. 2.8  Schematic representation of allogeneic (left side, in green (light gray in print version)) and autologous (right side, in red (dark gray in print version)) manufacturing processes of hMSC.

2.  Process development and manufacturing approaches for MSC therapies

Transport to the manufacturing facility



5  Extracellular vesicle production using hMSCs

51

Regarding product formulation, usually for autologous cell therapies, short-term biopreservation methods are employed, such as the use of reagents like HypoThermosol (BioLife Solutions, United States), and in some instances, the product is delivered fresh [135, 136], while for allogeneic products the product are usually cryopreserved [137, 138] (described in more detail in Section 6). As large quantities of hMSC biological material are generated after expansion in an allogeneic context, the quality control panel can be much more detailed and complex than for autologous products where a smaller number of cells is available (Table 2.4). Another key difference between allogeneic and autologous manufacture is related to the infrastructure required and the manufacturing facilities needed; while autologous therapies could potentially be produced at or near the patient bedside, allogeneic products require larger facilities and large-scale equipment [139–141]. From a cell biology perspective, a­ llogeneic and autologous hMSC products are also different. Generally, for autologous products, the cells undergo fewer expansion cycles and are less prone to become senescent [76, 142] or accumulating chromosomal errors [143, 144], which is not necessarily the case for allogeneic cell therapies. Moreover, the risk of immunogenicity is significantly reduced with an autologous therapies when compared to an allogeneic one, where there is a growing body of literature suggesting that allogeneic hMSCs can, to a certain extent, trigger an immune response in the recipient [145–147].

5  Extracellular vesicle production using hMSCs Although the majority of the studies focus on the production of hMSC for use as a cell-based product, a new type of therapeutic modality is emerging from stem cells and other cell types: cell-free products [148]. Whether the focus is to manufacture a EV product or one of the other subclasses (apoptotic vesicles, microparticles, or exosomes) [149] there are aspects of TABLE 2.4  Manufacturing and cell biology differences between autologous and allogeneic products Autologous

Allogenic

Patient specific

Off the shelf

Higher running costs per batch produced

Lower running costs per batch produced

Can be manufactured at the bedside

Needs large manufacturing facilities

No long-term storage needed

Cryopreservation needed

Lower number of doublings

Higher number of doublings

No incompatibility risks

Risk of immune response

52

2.  Process development and manufacturing approaches for MSC therapies

the bioprocess that need to be adapted toward the production of these cellfree products. It should be noted that although cell-free products are now the focus of several research and industry groups globally, EVs have been extensively studied as biomarkers for medical diagnostics [150, 151]. During the division process that cells undergo during expansion, EVs are produced and secreted to the liquid phase. While the challenge in cellbased products is to concentrate the cells in a small volume: the goal of EV-based bioprocesses is to remove cellular components and larger proteins to ensure that only EVs are isolated [152]. The key challenges for EV production are associated with scalability and purity as the majority of the downstream options for EVs are small scale and nonspecific [153–156]. There are several methods to isolate EVs with the most common being ultracentrifugation, density-gradient centrifugation, nonspecific precipitations, and benchtop commercially available options (Exo-Spin, Vivaspin, and ExoQuick among others). TFF has been used to overcome the scalability problems of the aforementioned techniques. To increase purity of the isolated populations, immunoaffinity methods are promising techniques to obtain subpopulations from highly heterogeneous noncellular fractions [153–156]. Ultracentrifugation steps are usually used sequentially when trying to isolate EVs population. While the objective of the first step is to remove any intact cells or debris in culture, the second step increases the centrifugation speed (around 100,000g) to promote EVs sedimentation [156, 157]. This is a highly nonspecific technique ideal to remove cellular fractions that are still present in the culture medium. Performing a ­density-gradient centrifugation step, the molecules will sediment through the gradient and will be separated by density, shape, and size [158, 159]. Through this technique, the final product is composed of a gradient of different populations, but has limited scalability. Nonspecific precipitations usually use polymers like polyethylene glycol (PEG) [14] and does not require any specific equipment which offers an advantage in terms of scalability. Because of its nonspecific nature, precipitations are usually succeeded by filtrations or low-speed centrifugations. As mentioned above, TFF holds a significant potential with respect to improved scalability of downstream processing for cell-free products as it can be combined with other size-based separation methods. An initial clarification step followed by TFF and ultrafiltration unit should remove the cell debris and major proteins [160, 161]. The biggest advantage of the size-based methods is their scalability. Immunoaffinity methods make use of the surface proteins expressed by EVs that will bind to specific antibodies that capture the EV subsets desired. Although this method is highly specific, it does have limitations with respect to scalability and the need to remove label postprocessing. Studies have attempted to compare the efficiency of different isolation of EVs. In 2014, Van Deun et  al. reported differences at a gene level of



6  Therapeutic product formulation

53

the exosomes isolated with four different methods (ultracentrifugation, ­density-gradient method, and a commercially available kit). The authors raised the issue of comparability between laboratories using different techniques, which is particularly critical from a diagnostics perspective [154]. Similar findings had been observed by Tauro et al. when using exosomes derived from a human colon cancer cell line [162]. The same authors reported that although centrifugation, density-based separation, and immunoaffinity successfully isolated EV expressing the same exosome markers, the latter method was able to enrich exosome markers in the fraction separated. Zlotogorski-Hurvitz et  al. reported that although commercially available ExoQuick-TC successfully isolated exosomes, the population obtained was less pure than the control using ultracentrifugation [163]. Using amniotic fluid-derived exosomes, Kosanović et  al. published that when compared to ultracentrifugation, ion exchange chromatography yielded a population of exosomes with higher purity and a higher separation efficiency based on CD markers expressed by the exosomes [164]. While different options will have their own advantages and drawbacks, the optimal solution will very likely be the combination of several processes rather than a single step. As different methods use distinct biological principles, it is important to assess all these techniques with respect to their separation efficiency, purity, identity, and functionality. For cellbased products cell number, viability, metabolism, immunophenotype, and potency are usually assessed, whereas for cell-free therapies, the analytics are likely focus on particle size distribution, protein, and gene expression together in addition to relevant potency assays. The biggest challenge in this case is the development of a complete bioprocess for the production of the two products simultaneously: cells and exosomes. With this particular goal in mind, the key focus for the downstream processing strategy is to separate the two products. Regarding product storage and formulation, there is also the need to explore different approaches. In the case of cell therapies, the long-term storage of the cell product is likely to be achieved via cryopreservation in an appropriate freezing solution, whereas exosomes could be formulated in a similar manner to other protein-based products and therefore be potentially amenable to freeze drying and other such formulations without the need of cryoprotectants (Fig. 2.9) [165, 166].

6  Therapeutic product formulation Although the goal of using microcarriers and bioreactor for hMSC is to bring cost down, the majority of clinical trials ongoing using this cell type are still using monolayer-based manufacturing systems [25, 167–175]. Regardless of the expansion method, products can be offered in one of

54

Bioreactor expansion

QA/QC

Downstream processing

QA/QC

QA/QC

Cryopreservation QA/QC

QA/QC

hMSC manufacturing gantt chart

Unit operations

Isolation of MSC Monolayer expansion Bioreactor expansion Downstream processing Cryopreservation 0

2

4

6

8

10

12

14

16

18

20

22

24 Days

FIG. 2.9  Simplified diagram and process flow chart of an hMSC manufacturing process.

2.  Process development and manufacturing approaches for MSC therapies

Monolayer expansion

Isolation



55

6  Therapeutic product formulation

two different forms: cryopreserved [170, 176, 177] or fresh [172, 174]. Cryopreserved products have a logistic advantage over fresh ones in that they can be stored for longer periods and the transportation process does not seem to negatively affect cell properties as compared to keeping them at 4°C for a long period of time [178]. It is important to mention that cryopreservation protocols should be optimized not only by choosing the best cryoprotectant for the application but also through the optimization of the freezing protocol (Fig. 2.10) [179, 180]. In 2016, Gramlich et al. published that the immunosuppressive potential and the therapeutic effect were not affected by the cryopreservation process. It was also reported in the same publication that although viability decreased after a cryopreservation and thawing cycle compared to a fresh preparation of hMSC, it never decreased below 95% [181]. These conclusions were confirmed by Yuan et al. in 2016 who reported that postthaw viability was kept above 85% using a cryoprotectant solution composed by 10% dimethyl sulfoxide (DMSO) and 90% of human albumin [182]. Although the same publication reported that the homing capacity of hMSC test did not change with the cryopreservation process this was not verified with MSC transduced with lentiviral vector expressing tumor necrosis factors. In 2017, Kaplan et  al. published that cryopreservation does not affect BM-hMSC growth kinetics [183]. Although several authors have reported that cryopreservation does not affect cell properties, there are other studies that point toward the opposite direction. Guido Moll in 2014 has published that cryopreserved MSCs have reduced immunomodulatory properties reporting also a decrease in viability after thawing. François et al. in 2011 reported that T-cell inhibition proliferation of hMSC also decreases after a cryopreservation and thawing cycle [184]. In 2012, Ginis et  al. reported that the cryopreservation-thawing cycle did not compromise hMSC’s growth rate, immunophenotype nor osteogenic differentiation ability [185]. It should be highlighted that these studies used different cryopreservation conditions: different cryoprotectant formulation and different freezing protocol which ultimately lead to different results [179, 186]. Although the literature for exosome storage is still very scarce, they can be stored for a short period of time at 4°C while between −20°C and

Freezing protocol

Cryopreservation success

Cryopreservation solution

FIG. 2.10  Simplified diagram of the parameters to be considered to establish an efficient hMSC cryopreservation method.

56

2.  Process development and manufacturing approaches for MSC therapies

−80°C for long-term storage. Lee et  al. reported in 2016 that short-term storage at 4°C lead to major losses in CD63 expression [165]. At this point it is not clear what is the ideal type of product formulation for exosomes nor if the current short-/long-term storage affects their properties. While there are still no gold standard procedures to produce and store exosomes, the number of clinical trials using hMSC-derived extravesicles as therapeutic product is still very low (NCT02138331, NCT03384433, NCT03608631, NCT03437759, UMIN000011290, ChiCTR-INR-17010677, and IRCT2017101736840N1). It is important to mention that apart from their therapeutic potential, exomes have also been studied as biomarkers in the diagnostic fields.

6.1  Approved hMSC-based products As of Q4 2018, 10 hMSC-based products (Table  2.2) have been approved. Of these, four have been approved in South Korea by the KFDA, one in Canada and New Zealand (Health Canada & Medsafe approved), one in India (approved by the DCGI), four in Japan (with Japan MHLW conductional approval), and another mostly recently in Europe (EMA) (2018) [187] (Table  2.5). The products listed in Table  2.5 target different types of indications ranging from cardiovascular disorders, neurological or autoimmune diseases. In this group of approved products, six are using BM-MSCs, four use AT-MSCs, and one is using UCB-derived MSC. It should also be mentioned that five of these products were developed for autologous application while six of them are allogeneic cell therapies. These trends highlight that the approval of cell therapies is faster in Asia than in Europe and North America. It is expected that this list keeps growing as some of these companies have other products in their pipeline and once they have several clinical trials reaching phases 3 and 4. There are several challenges trying to develop a sustainable, commercially viable cell therapy industry. From a manufacturing perspective, it should be highlighted that some clinical trials still expand cells in monolayers and use FBS supplemented medium in their cultures. To move to bioreactor-based expansion as opposed to monolayer, it is expected to decrease the cost per batch, essential for autologous purposes. Another challenge that needs to be addressed is the identification of optimal hMSC donors. As hMSC seem to exhibit several paracrine effects in vitro (angiogenesis, immunomodulation, antiapoptotic, and antiscarring among others [188]) they might a response (and mechanism of action) for different conditions. This highlights the importance of conducting donor screenings for specific disorders as hMSC that perform well for a specific condition will not necessarily perform well in another (Fig. 2.11). While the first study using hBM-MSC expanded with microcarriers in spinner flask happened in 2009 [189], the world’s first hMSC product



TABLE 2.5  hMSC-based approved products around the world (as of March 2018) Company

Target indication

Cell source

Type of therapy

When approved

Where marketed

HearticellGram-AMI

FCB PharmiCell

Acute myocardial infarction

BM

Autologous

2011

S. Korea

Cartistem

Medipost

Degenerative arthritis

UCB

Allogeneic

2012

S. Korea

Cupistem

Anterogen

Anal fistula in Crohn’s disease

AT

Autologous

2012

S. Korea

Prochymal

Osiris Therapeutics

Acute graft vs host disease

BM

Allogeneic

2012

Canada, New Zealand

Nuronata-R

Corestem

Amyotrophic lateral sclerosis

BM

Allogeneic

2014

S. Korea

TemCell

JCR Pharm.

Acute graft vs host disease

BM

Allogeneic

2015

Japan

aJointStem

Biostar Stem Cell Research

Degenerative arthritis

AT

Autologous

2015

Japan

Stempeucel

Stempeutics

Critical limb ischemia

BM

Allogeneic

2016

India

Alofisel (Cx601)

TiGenix-Takeda

Anal fistula in Crohn’s disease

AT

Allogeneic

2018

EU

AstroStem

Biostar Stem Cell Research

Alzheimer’s disease

AT

Autologous

2018

Japan

Stemirac

NIPRO Corp

Spinal cord injury

BM

Autologous

2018

Japan

6  Therapeutic product formulation

Product name

57

58

2009

• First hBM-MSC microcarrier-based expansion in spinner.

Cartistem HeartiCellGram Cupistem Prochymal

2011

• hAT-MSC microcarrier-based expansion in spinner reported. • Study using XF medium in microcarrier-based expansion

Nuronata-R

2014

TemCell

Stempeucel

2016

• hUC-MSC microcarrier-based • Study using hPL expansion in spinner reported. sup. medium in microcarrier-based expansion.

2017

• Study using hSerum in microcarrier-based expansion. • 50 L scale study published

Manufacturing processes development

FIG. 2.11  Timeline to represent evolution of the hMSC-based therapies market together with the development of the manufacturing industry.

2.  Process development and manufacturing approaches for MSC therapies

hMSC production approved



References

59

only reached the market in 2011. Six more hMSC-based products got approved between 2012 and 2016: Cartistem, Cupistem and Prochymal-2012, Nuronata-R-2014, TemCell-2015, and Stempeucel-2016 while the first study using hPL [51] and running a 50-L scale [86] expansion process happened in 2016 and 2017. This is evidence of the time-lag between R&D activity and commercial production. As yields improve and a greater understanding of the bioprocess is generated, this time-lag is expected to shorten.

7  Conclusion Human mesenchymal stem/stromal cells (hMSCs) are a promising candidate for cell therapy and regenerative medicine applications due to their immunomodulatory properties and in vitro differentiation capability. Moreover, there is significant interest in isolating the EVs that the cells produce and administering these as a cell-free product to the potential therapeutic capability of these vesicles. However, in order for these cellbased and cell-free therapies to become a clinical and commercial reality, key bioprocessing and biomanufacturing challenges need to be addressed. Traditional planar-based culture systems are not suitable for large-scale manufacture and as such there is extensive research and development activity focusing on establishing scalable and controlled bioprocesses for hMSC expansion. In this chapter, we have outlined the key bioprocessing aspects that must be considered for scalable bioreactor culture and the key parameters that need to be monitored, controlled, and optimized to achieve a consistent hMSC-based therapy, or indeed, a cell-free EV product from a hMSC source.

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[159] Yu  LL, Zhu  J, Liu  JX, et  al. A comparison of traditional and novel methods for the separation of exosomes from human samples. Biomed Res Int 2018. https://doi. org/10.1155/2018/3634563. [160] Haraszti RA, Miller R, Stoppato M, et al. Exosomes produced from three-dimsensional cultures of mesenchymal stem cells by tangential flow filtration show higher yield and improved activity. Mol Ther 2018. https://doi.org/10.1093/gbe/evy043/4895092. [161] Heinemann  ML, Ilmer  M, Silva  LP, et  al. Benchtop isolation and characterization of functional exosomes by sequential filtration. J Chromatogr A 2014. https://doi. org/10.1016/j.chroma.2014.10.026. [162] Tauro BJ, Greening DW, Mathias RA, et al. Comparison of ultracentrifugation, density gradient separation, and immunoaffinity capture methods for isolating human colon cancer cell line LIM1863-derived exosomes. Methods 2012. https://doi.org/10.1016/j. ymeth.2012.01.002. [163] Zlotogorski-Hurvitz  A, Dayan  D, Chaushu  G, et  al. Human saliva-derived exosomes: comparing methods of isolation. J Histochem Cytochem 2015. https://doi. org/10.1369/0022155414564219. [164] Kosanović  M, Milutinović  B, Goč  S, Mitić  N, Janković  M. Ion-exchange chromatography purification of extracellular vesicles. Biotechniques 2017. https://doi. org/10.2144/000114575. [165] Lee  M, Ban  JJ, Im  W, Kim  M. Influence of storage condition on exosome recovery. Biotechnol Bioprocess Eng 2016;21(2):299–304. https://doi.org/10.1007/ s12257-015-0781-x. [166] Kumeda  N, Ogawa  Y, Akimoto  Y, Kawakami  H, Tsujimoto  M, Yanoshita  R. Characterization of membrane integrity and morphological stability of human salivary exosomes. Biol Pharm Bull 2017;40(8):1183–91. https://doi.org/10.1248/bpb.b16-00891. [167] Satti HS, Waheed A, Ahmed P, et al. Autologous mesenchymal stromal cell transplantation for spinal cord injury: a phase I pilot study. Cytotherapy 2016;18(4):518–22. https://doi.org/10.1016/j.jcyt.2016.01.004. [168] Sánchez-Guijo FM, Lõpez-Villar O, Lõpez-Anglada L, et al. Allogeneic mesenchymal stem cell therapy for refractory cytopenias after hematopoietic stem cell transplantation. Transfusion 2012;52(5):1086–91. https://doi.org/10.1111/j.1537-2995.2011.03400.x. [169] Šponer  P, Filip  S, Kučera  T, et  al. Utilizing autologous multipotent mesenchymal stromal cells and β -tricalcium phosphate scaffold in human bone defects: a prospective, controlled feasibility trial. Biomed Res Int 2016;2016: https://doi. org/10.1155/2016/2076061. [170] De La Portilla  F, Alba  F, García-Olmo  D, Herrerías  JM, González  FX, Galindo  A. Expanded allogeneic adipose-derived stem cells (eASCs) for the treatment of complex perianal fistula in Crohn’s disease: results from a multicenter phase I/IIa clinical trial. Int J Colorectal Dis 2013;28(3):313–23. https://doi.org/10.1007/s00384-012-1581-9. [171] Kim  SJ, Moon  GJ, Chang  WH, Kim  Y-H, Bang  OY. Intravenous transplantation of mesenchymal stem cells preconditioned with early phase stroke serum: current evidence and study protocol for a randomized trial. Trials 2013;14(1):317. https://doi. org/10.1186/1745-6215-14-317. [172] Lee J-W, Lee S-H, Youn Y-J, et al. A randomized, open-label, multicenter trial for the safety and efficacy of adult mesenchymal stem cells after acute myocardial infarction. J Korean Med Sci 2014;29(1):23. https://doi.org/10.3346/jkms.2014.29.1.23. [173] Cheng H, Liu X, Hua R, et al. Clinical observation of umbilical cord mesenchymal stem cell transplantation in treatment for sequelae of thoracolumbar spinal cord injury. J Transl Med 2014;12(1): https://doi.org/10.1186/s12967-014-0253-7. [174] Syková E, Rychmach P, Drahorádová I, et al. Transplantation of mesenchymal stromal cells in patients with amyotrophic lateral sclerosis: results of phase I/IIa clinical trial. Cell Transplant 2016. https://doi.org/10.3727/096368916X693716.

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[175] Pers Y-M, Rackwitz L, Ferreira R, et al. Adipose mesenchymal stromal cell-based therapy for severe osteoarthritis of the knee: a phase I dose-escalation trial. Stem Cells Transl Med 2016;5(7):847–56. https://doi.org/10.5966/sctm.2015-0245. [176] Butler J, Epstein SE, Greene SJ, et al. Intravenous allogeneic mesenchymal stem cells for nonischemic cardiomyopathynovelty and significance. Circ Res 2017;120(2):332– 40. https://doi.org/10.1161/CIRCRESAHA.116.309717. [177] Dai  G, Liu  X, Zhang  Z, et  al. Comparative analysis of curative effect of CT-guided stem cell transplantation and open surgical transplantation for sequelae of spinal cord injury. J Transl Med 2013;11(1): https://doi.org/10.1186/1479-5876-11-315. [178] Nikolaev NI, Liu Y, Hussein H, Williams DJ. The sensitivity of human mesenchymal stem cells to vibration and cold storage conditions representative of cold transportation. J R Soc Interface 2012;9(75):2503–15. https://doi.org/10.1098/rsif.2012.0271. [179] Shivakumar  SB, Bharti  D, Jang  SJ, et  al. Cryopreservation of human wharton’s ­jelly-derived mesenchymal stem cells following controlled rate freezing protocol using different cryoprotectants; a comparative study. Int J Stem Cells 2015;8(2): https://doi. org/10.15283/ijsc.2015.8.2.155. [180] De Lara Janz F, De Aguiar Debes A, De Cássia Cavaglieri R, et al. Evaluation of distinct freezing methods and cryoprotectants for human amniotic fluid stem cells cryopreservation. J Biomed Biotechnol 2012. https://doi.org/10.1155/2012/649353. [181] Gramlich  OW, Burand  AJ, Brown  AJ, Deutsch  RJ, Kuehn  MH, Ankrum  JA. Cryopreserved mesenchymal stromal cells maintain potency in a retinal ischemia/ reperfusion injury model: toward an off-the-shelf therapy. Sci Rep 2016;6: https://doi. org/10.1038/srep26463. [182] Yuan Z, Lourenco SDS, Sage EK, Kolluri KK, Lowdell MW, Janes SM. Cryopreservation of human mesenchymal stromal cells expressing TRAIL for human anti-cancer therapy. Cytotherapy 2016;18(7):860–9. https://doi.org/10.1016/j.jcyt.2016.04.005. [183] Kaplan A, Sackett K, Sumstad D, Kadidlo D, McKenna DH. Impact of starting material (fresh versus cryopreserved marrow) on mesenchymal stem cell culture. Transfusion 2017;57(9):2216–9. https://doi.org/10.1111/trf.14192. [184] François  M, Copland  IB, Yuan  S, Romieu-Mourez  R, Waller  EK, Galipeau  J. Cryopreserved mesenchymal stromal cells display impaired immunosuppressive properties as a result of heat-shock response and impaired interferon-γ licensing. Cytotherapy 2012;14(2):147–52. https://doi.org/10.3109/14653249.2011.623691. [185] Ginis  I, Grinblat  B, Shirvan  MH. Evaluation of bone marrow-derived mesenchymal stem cells after cryopreservation and hypothermic storage in clinically safe medium. Tissue Eng Part C Methods 2012;18(6):453–63. https://doi.org/10.1089/ten. tec.2011.0395. [186] Thirumala S, Zvonic S, Floyd E, Gimble JM, Devireddy RV. Effect of various freezing parameters on the immediate post-thaw membrane integrity of adipose tissue derived adult stem cells. Biotechnol Prog 2005;21(5):1511–24. https://doi.org/10.1021/ bp050007q. [187] Celltrials.org. https://celltrials.org/. [188] Meirelles LdS, Fontes AM, Covas DT, Caplan AI. Mechanisms involved in the therapeutic properties of mesenchymal stem cells. Cytokine Growth Factor Rev 2009;20 (5–6):419–27. https://doi.org/10.1016/j.cytogfr.2009.10.002. [189] Schop D, Van Dijkhuizen-Radersma R, Borgart E, et al. Expansion of human mesenchymal stromal cells on microcarriers: growth and metabolism. J Tissue Eng Regen Med 2010;4(2):131–40. https://doi.org/10.1002/term.224.



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Further reading [190] Ghorbani A, Jalali SA, Varedi M. Isolation of adipose tissue mesenchymal stem cells without tissue destruction: a non-enzymatic method. Tissue Cell 2014;46(1):54–8. https://doi.org/10.1016/j.tice.2013.11.002. [191] De Bruyn C, Najar M, Raicevic G, et al. A rapid, simple, and reproducible method for the isolation of mesenchymal stromal cells from Wharton’s jelly without enzymatic treatment. Stem Cells Dev 2011;20(3):547–57. https://doi.org/10.1089/scd.2010.0260. [192] Han Y-F, Tao R, Sun T-J, Chai J-K, Xu G, Liu J. Optimization of human umbilical cord mesenchymal stem cell isolation and culture methods. Cytotechnology 2013;(51). https://doi.org/10.1007/s10616-012-9528-0. [193] Vangsness CT, Sternberg H, Harris L. Umbilical cord tissue offers the greatest number of harvestable mesenchymal stem cells for research and clinical application: a literature review of different harvest sites. Arthroscopy 2015;31(9):1836–43. https://doi. org/10.1016/j.arthro.2015.03.014. [194] Priya  N, Sarcar  S, Sen  MA, Sundarraj  S. Explant culture: a simple, reproducible, efficient and economic technique for isolation of mesenchymal stromal cells from human adipose tissue and lipoaspirate. J Tissue Eng Regen Med 2012;1–9. https://doi. org/10.1002/term.1569. [195] Miyagi-Shiohira C, Kurima K, Kobayashi N, et al. Cryopreservation of adipose-­derived mesenchymal stem cells. Cell Med 2015;8(1–2):3–7. https://doi.org/10.3727/2155179 15X689100. [196] Heath N, Grant L, De Oliveira TM, et al. Rapid isolation and enrichment of extracellular vesicle preparations using anion exchange chromatography. Sci Rep 2018. https:// doi.org/10.1038/s41598-018-24163-y. [197] Corso G, Mäger I, Lee Y, et al. Reproducible and scalable purification of extracellular vesicles using combined bind-elute and size exclusion chromatography. Sci Rep 2017. https://doi.org/10.1038/s41598-017-10646-x.

C H A P T E R

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Bioinspired materials and tissue engineering approaches applied to the regeneration of musculoskeletal tissues ⁎

⁎

Márcia T. Rodriguesa,b,c, , Ana I. Gonçalvesa,b, , Pedro S. Baboa,b, Manuel Gomez-Florita,b, Rui L. Reisa,b,c, Manuela E. Gomesa,b,c a

3B's Research Group, I3Bs—Research Institute on Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the European Institute of Excellence on Tissue Engineering and Regenerative Medicine, AvePark, Parque de Ciência e Tecnologia, Zona Industrial da Gandra, Guimarães, Portugal, bICVS/3B’s—PT Government Associate Laboratory, Braga/Guimarães, Portugal, cThe Discoveries Centre for Regenerative and Precision Medicine, Headquarters at University of Minho, Guimarães, Portugal

1  Musculoskeletal tissues Musculoskeletal tissues are a complex system which provides form, support, stability, and movement to the body. It is also characterized by natural mechano-responsiveness and an efficient complementarity among muscles, tendons, ligaments, cartilage, and bones necessary for function and coordinated actions both for daily life and the most demanding sports activities. The musculoskeletal system relies on the proper articulation of soft and hard tissues tailored to accomplish specific functions (Fig. 3.1). *Authors equally contributed to this work. Copyright © 2020 Tiago G. Fernandes, Maria Margarida Diogo & Joaquim M. S. Cabral. Published by Elsevier Inc. All rights reserved. https://doi.org/10.1016/B978-0-12-816221-7.00003-3

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FIG.  3.1  Musculoskeletal system. Representation of the structural organization of the musculoskeletal tissues and summary of their specific features.

Skeletal muscles function to produce force and motion and are responsible not only for the articulated movements that allow the locomotion but also for the movement of internal organs and protection to external shocks. Tendons in the extremities of muscles transmit forces to bones while ligaments keep the stability of the joints and movements between bones. Bones provide structural support, locomotion, and protection of the internal organs as well as in the maintenance of acid-base balance, and calcium, magnesium, and phosphate homeostasis [1]. In joints, the end of the long bones is covered by cartilage, which protects the bones from friction and absorbs the forces over the skeleton. All these tissues work together in continuous adjustment to balance the body during locomotion and movements. With the increase in life expectancy and the maintenance of an active lifestyle by the aging population, the cumulative musculoskeletal



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c­ onditions potentiate disability throughout life. Although the pathological conditions result from a combination of genetic, physiological, environmental, and behavioral factors, aging and the degeneration of tissues and organs have a major impact in increased pain, reduced mobility, and lack of patient’s autonomy. Thus, it is imperative to understand the structural characteristics and physiological behavior of the tissues to develop improved tissue engineering and regenerative strategies in order to obtain better therapies.

1.1  Physiology and function 1.1.1  Bone Bones are mineralized connective tissues that compose the skeleton of vertebrates, harboring the bone marrow which is the primary site of hematopoiesis. The extracellular (ECM) bone matrix is composed of an organic (40%) and an inorganic (60%) phase [1]. The organic matrix is mainly composed of collagenous proteins, predominantly type I collagen, noncollagenous proteins and growth factors [2]. Among the noncollagenous proteins, the osteocalcin, osteonectin, and osteopontin are the most abundant. Osteocalcin is associated with the regulation of the osteoclasts activities and is regarded as a marker of bone formation [3, 4], while osteonectin is related to the osteoblast growth and proliferation, and matrix mineralization [4]. The matrix cohesion and adhesion of bone tissues is maintained by osteopontin [5]. Growth factors and cytokines including insulin-like growth factors (IGFs), transforming growth factor β (TGF-β), bone morphogenetic proteins (BMPs), fibroblast growth factor (FGF), and platelet-derived growth factor (PDGF) are also present in the organic matrix in trace quantities, and modulate cell response to bone remodeling [6]. The organic phase forms a fibrillary matrix over which the mineral phase nucleates. The inorganic material of bone consists predominantly of crystals of a very insoluble calcium phosphate (CaP) salt, the hydroxyapatite [Ca10(PO4)6(OH)2] which is responsible for bone stiffness and load-­bearing properties. Significant amounts of bicarbonate, sodium, potassium, ­citrate, magnesium, carbonate, fluorite, zinc, barium, and strontium are also present in bone mineral phase [7]. Morphologically, the bone can be classified into the cortical or trabecular bone. Cortical bone is dense and solid and surrounds the marrow space, whereas trabecular bone is composed of a honeycomb-like network of trabecular plates and rods interspersed in the bone marrow compartment. Regarding the shape and size, bones are classified into four categories: (1) the long bones, including the clavicles, humeri and femurs; (2) the short bones, including the carpal and tarsal bones; (3) the flat bones, including the skull, mandible, and scapulae; and lastly, (4) the irregular bones which include the vertebrae, sacrum, and coccyx [7].

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The bone cells are the living elements responsible for bone homeostasis and response to external stimuli and comprise osteoblasts, osteocytes, bone lining cells, and osteoclasts. Osteoblasts, osteocytes, and bone lining cells derive from mesenchymal lineage while osteoclasts are originated from the hematopoietic lineage [1]. Osteoblasts are responsible for the production of the bone matrix constituents [3]. Conversely, the osteoclasts are responsible for bone matrix resorption [4]. Bone health depends on the coordination between the resorption activity promoted by the osteoclasts and the new-bone deposition promoted by the osteoblasts in a process termed remodeling. The osteocytes are terminally differentiated osteoblasts that reside within the bone matrix. These are the most abundant bone cells and, among other functions, are responsible for the mechanotransduction events that allow the bone to adapt to mechanical stress [8]. Bone lining cells are mainly metabolically inactive osteoblasts and form the periosteum and endosteum. Besides participating in the osteoclast differentiation [2], the bone lining cells contribute to the exchange of ions between bones and the surrounding tissues [9]. 1.1.2 Cartilage Cartilage is a highly specialized connective tissue composed of a single cell type, the chondrocyte, organized into groups of few cells within an ECM rich in collagen fibrils and proteoglycans. Three types of cartilage have been identified: fibrocartilage, articular, and elastic cartilage, differing in the amounts and organization of collagen and proteoglycan and in function. Fibrocartilage is found for instance at tendon/ligament junction with bone and is designed to stand compressive strength while the principal function of hyaline cartilage, often designed as articular cartilage available at the surface of diarthrodial joints, is to provide a smooth, lubricated surface for articulation and to facilitate the transmission of loads with a low frictional coefficient [10]. The ECM of elastic cartilage found in the trachea and ears is also rich in elastin fibers, which allows tolerating repetitive deformation. Compared to other connective tissues, cartilage has a very slow turnover and a limited capacity to undergo intrinsic healing. Hyaline cartilage is the most prevalent type of cartilage. Since cartilage is avascular and aneural, nutrition to chondrocytes is provided through diffusion, which is assisted by the fluid flow generated in the compression of the articular cartilage or in the flexion of the elastic cartilage. Articular cartilage comprises four zones. A superficial zone populated with a considerable high number of chondrocytes and where collagen fibers (primarily, type II and IX collagen) are packed tightly and aligned parallel to the articular surface. This zone is in contact with synovial fluid and is responsible for most of the tensile properties of cartilage, which enable it to resist the sheer, tensile, and compressive forces imposed by articulation [10]. Then, a middle zone constituted with collagen organized obliquely into thicker



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collagen fibrils with a low density of spherical chondrocytes. Functionally, the middle zone is the first line of resistance to compressive forces [10]. The deep zone is responsible for providing the greatest resistance to compressive forces, given the high proteoglycan content. The chondrocytes are typically arranged in columnar orientation, parallel to the collagen fibers and perpendicular to the joint line [10]. The tide mark distinguishes the deep zone from the calcified cartilage. The calcified layer plays an integral role in securing the cartilage to bone, by anchoring the collagen fibrils of the deep zone to subchondral bone. In this zone, the cell population is scarce and chondrocytes are hypertrophic [10]. The anisotropic nature and unique viscoelastic properties of cartilage in a structure of few millimeters thick highlights the challenges involved in the development of an artificial biomimetic and biofunctional cartilage substitute for regenerative medicine strategies. 1.1.3  Tendon Tendons and ligaments are similar dense fibrous connective tissues that connect muscle to the skeletal elements (bone) and bone to bone, respectively [11]. These tissues are characterized by the presence of few and dispersed fibroblasts/fibrocytes (ligament) or tenoblasts/tenocytes (tendon) within a collagen-rich ECM, resulting in a dense and hypocellular structure [11]. Tendon ECM is formed from the continual aggregation of the smallest structural unit, collagen, into an increasingly complex architecture [12]. Spontaneous aggregation of multiple collagen molecules results in the formation of collagen fibrils and, in turn, bundles of fibrils form larger primary fiber bundles called fascicles, groups of which associate to form tertiary fiber bundles [12, 13]. These are bound together by a thin layer of connective tissue named endotenon that contains blood vessels, lymphatics, and nerves. The multiple fiber bundles and endotenon are encompassed by the epitenon, which is covered by another layer of connective tissue called paratenon. Tendon tissues are crucial in all joint movements and the limitations of current surgical interventions motivate tissue engineering approaches to build patient-personalized biological substitutes for tendon repair. Despite tendon’s hypocellular nature, different cells co-exist constituting a heterogeneous population of tenocytes and tendon stem and progenitor cells [14, 15]. Although the collagen type I is the major component, tendon ECM also includes elastin fibers embedded in a hydrated proteoglycan matrix. The collagen fibers resist to tension forces applied to these tissues, while proteoglycans are responsible for the viscoelastic properties of tendons [16].

1.2  Response to injury and healing mechanisms The proportion of the population reporting musculoskeletal conditions increased from 28.0% in 1996–98 to 33.2% in 2009–11. The reality of these figures is only for the US population, suggesting that these figures can be

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significantly aggravated [17]. This implies that persons with musculoskeletal conditions accounted for an aggregate economic impact of $367.1 billion in 1996–98 and $796.3 billion in 2009–11, an increase of 117% in real terms [17]. The mechanical forces applied to these tissues can be both protective and detrimental. Although the mechanical stress is critical for the ­physio-anatomy of healthy tissues, misuse or overuse can inflict pathological alterations in biomolecular and cell-mediated mechanisms, thus contributing to impaired healing. The healing response of musculoskeletal tissues typically involves an acute inflammatory response with the production and release of several important molecules including cytokines that initiate a healing cascade. In bone, this response is accompanied by the recruitment of mesenchymal stem cells (MSCs) in order to generate a cartilaginous callus that undergoes mineralization and resorption. The callus provides biomechanical stability but requires further remodeling to fully restore the biomechanical properties of normal bone. The healing of bone fractures is a complex and efficient process that occurs without the development of a fibrous scar. Unlike bones, damaged articular cartilage and injured tendon/ligaments have limited intrinsic healing, likely because of the hypocellular nature and the lack of resident vascular and lymphatic vessels that restricts the access to cells and biochemical factors with regenerative action. Moreover, in adult tissues, mature chondrocytes and tenocytes that are responsible for ECM maintenance, have low mitotic and metabolic rates. In general, early-stage lesions in cartilage and tendon are asymptomatic which may relate to the absence of an intrinsic nerve network in these tissues. Moreover, these lesions are still manageable with antiinflammatory and analgesic drugs. When the damage is severe, patients become candidates for surgical interventions, in many cases resourcing to tissue grafts, with variable and often insufficient outcomes in the long term. The dissimilar healing responses of musculoskeletal tissues are thus related to their intrinsic properties and function, influencing tissue adaptation and response to mechanical forces (Fig. 3.2). Impaired tissue restoration or nonresolved healing favors the progression of lesions into chronic and degenerative conditions, and ­compromises nearby tissues that can culminate into a total joint destruction. Current strategies for the repair of defects and lesions in musculoskeletal tissues are generally unsatisfactory as the restored tissue does not meet the mechanical biofunctionalities of uninjured tissues.

2  Cartilage regeneration Cartilage lesions progress from asymptomatic lesions to injuries and diseases such as osteoarthritis (OA) that compromise the entire articular joint. Cartilage injuries are broadly classified as partial thickness or full



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FIG. 3.2  Schematic representation of the articulation of the musculoskeletal tissues from cell to the joint level in the adaptation and response to external mechanical forces.

thickness defects. Partial thickness (chondral) defects as clefts and fissures do not reach subchondral bone and fail to heal spontaneously, while full thickness (osteochondral) defects penetrate to the subchondral bone with a resource to the blood supply and progenitor cells and elicit of an intrinsic repair response. However, the fibrocartilaginous repair tissue formed fails to replace hyaline cartilage in organization and function. Current standard procedures for the treatment of cartilage defects include chondral resurfacing with abrasion, debridement, autologous chondrocyte implantation, and matrix-induced chondrocyte implantation, or osteochondral autologous transplantation. Cartilage surgical procedures can result in short- and mid-term clinical improvement of the joint but do not prevent degeneration of repaired tissue and are a major cause of morbidity of donor tissues, interfering with local biomechanics, and creating the need for replacement of donor cartilage in future years. Thus, the limited success of current treatments and the shortage of cartilage substitutes challenges for innovative approaches to improve cartilage treatments and favor tissue regeneration.

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2.1  Scaffold/hydrogel-based approaches During the last decades, hydrogels have been pursued as preferential candidates for cartilage tissue engineering and regenerative medicine (TERM) approaches. Hydrogels are hydrophilic networks formed by physical and/or chemical crosslinking of polymers providing versatile and highly desirable 3D environments for biological processes (Table 3.1). Hydrogels proposed for cartilage approaches have been designed with different shapes, complexity (e.g., single and multiple polymers and single and multiple crosslinked networks), and tunable properties (e.g., mechanical properties and responsiveness toward an external stimuli) to support and stimulate cells within a physical matrix [18, 19]. These hydrogels have been developed to act as cell carriers [20] and other therapeutic agents and to study chondrogenesis and associated mechanisms [18, 19]. In large defects, a supportive (hydrogel) matrix may assist the transport of therapeutic cells and enhance tissue regeneration [21] toward a complete integration with surrounding tissues. Furthermore, hydrogels offer the possibility for filling irregular defects using minimally invasive procedures for local and sustained delivery of therapeutics, including cells, drugs, and bioactive molecules such as growth factors and genetic material (revised by Liu et al. [22]) holding the promise for off-the-shelf products to use upon request in patient customized solutions [23]. The advent of personalized medicine brought the attention for bioprinting and plotting technologies to hydrogel fabrication. In the particular case of cartilage, the 3D layered deposition by printing technologies can recreate the tissue-specific anatomic-physiological design in shape and depth-dependent structure with zonal-like hierarchies assisting the fabrication of cartilage constructs with increased biomimicry. Cell-laden printing is thought to significantly enhance the interaction between cells and matrix and improve tissue regeneration. This technology has been explored using different cell types and materials; however, the translation of knowledge from traditional hydrogel fabrication into these inspirational technologies is hampered by the requirement for specific hydrogel properties. Besides a controlled gelation time and swelling or contraction, bioinks require stability and shear-thinning properties to allow their syringe uptake and application as well as self-healing features for a fast reassemble when shear forces are removed. Bioinks are often chosen for their printability but is also important to consider their influence in the biological response. In a study by Daly et al. the phenotype of bone marrow-derived mesenchymal stem cells (BM-MSCs) encapsulated in bioinks produced with different materials [agarose, alginate, gelatin (GelMA), and BioINK] resulted in the development of cartilaginous-like tissues with hyaline-like and fibrocartilage-like structures [24].



TABLE 3.1  Overview on hydrogels for cartilage tissue engineering and regeneration approaches Material

Manufacturing techniques Fibrous Electrospinning

Synthetic PEG pNIPAAm OPF

Organized structure Solid freeform fabrication (SFF): Bioprinting Plotting

Amorphous Cryogelation Freeze-drying

Advantages

Limitations

Physical methods (responsive to) Temperature pH Ionic concentration Electric fields Magnetic forces Electrostatic interaction Hydrophobic interactions Hydrogen bonding Stereo-complexation

High water content Similarities to natural ECM Porous framework Hydrogel morphologies Facilitates migration, adhesion, proliferation of cells Inclusion of other materials in the matrix as precursors (hybrid gels), particles (nanogels) Precise deposition of cells (SFF) Control of pore size and geometry (SFF) Minimally invasive procedures in injectable systems Match irregular defects

Unsuitable degradation kinetics Poor control of gelation kinetics Weak mechanical properties and instability (especially physical methods) Toxicity of several cross-linking agents (chemical methods) Requirement for shear-thinning and self-healing properties (bioinks)

Chemical methods Enzymatic crosslinking Schiff base cross-linking Michael additions Click chemistry Photo-crosslinking

2  Cartilage regeneration

Natural derived Chitosan Collagen/gelatin Hyaluronic acid Chondroitin sulphate Alginate Fibrin Keratin

Fabrication methods

Abbreviations: OPF, oligo(poly(ethylene) glycol) fumarate; PEG, polyethylene glycol; pNIPAAm, poly(N-isopropylacrylamide).

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Besides the 3D matrix support and stimulation, printing technologies are evolving into more complex and multifactorial systems to meet the natural requirements of tissues and guide the process of regeneration. Zhu et al. investigated the chondrogenic potential of human BM-MSC laden in a stereolithography-based 3D bioprinting matrix fabricated with gelatin and polyethylene glycol diacrylate (PEGDA) [25] incorporating core-shell nanospheres as TGF-β1 carriers. The developed system allowed maintaining cell viability and TGF-β1 bioactivity postprinting. Moreover, TGF-β1 improved the chondrogenic differentiation of MSCs with increased expression levels of Collagen II, SOX-9, and Aggrecan. In a proof of concept work, distinct cell types associated to cartilage tissue, namely articular cartilage-resident chondroprogenitor cells, BMMSCs, and chondrocytes, were cultured in gelatin-based hydrogels using multicompartment printed hydrogels to recapitulate the zonal stratification of native cartilage and assess for preferential zonal-like and for cartilage regenerative potential [26]. Despite variations found in the amount and quality of cartilage ECM synthetized by the different bioinks, this study proposes a bioprinted model for the exploitation of cell-cell and cell-matrix interactions within the layered distribution for cartilage regeneration.

2.2  Cell-based approaches using mesenchymal stem cells MSCs potential for cell-based therapies aiming at cartilage regeneration has been anticipated by numerous studies with different models [18, 21]. Clinical trials aiming at articular cartilage repair, especially focusing on the treatment of OA are still at early stages with preliminary aims to evaluate safety, feasibility, and efficacy (revised by Lee et al. [27]). Overall, these trials indicate pain relief but the renewal or improvement of cartilage tissue shows high degree of variability in human subjects. Recently, intraarticular administration of autologous human BM-MSC entered clinical trials as a safe and feasible procedure with improved outcomes for knee OA [28–30]. Autologous peripheral blood stem cells (AAPBSCs) were also investigated for the treatment of early OA. Three groups of 20 patients were 3 weekly treated with APPBSCs activated with platelet-rich plasma, with or without granulocyte colony-stimulating factor + IA-HA (intraarticular hyaluronic acid carrier) and IA-HA alone (control). Clinical scores showed statistically significant improvements at 6 and 12 months for the AAPBSC groups vs controls. Moreover, the differential effects of stimulated AAPBSCs were noted with an earlier onset of symptom alleviation throughout. At 12 months follow-up, AAPBSC groups also avoided the need for total knee arthroplasty [31]. The delivery of MSC to the injury site with a carrier system is an ­important procedure in clinical therapies to avoid cell leakage to other



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areas of the joint keeping the therapeutic potential of implanted biological agents where it is needed the most. Recent studies showed the success of magnetic targeting to deliver MSC system for the treatment of focal articular cartilage defects [32, 33], including a preliminary study with human induced pluripotent stem cells magnetically delivered to osteochondral defects in a rat femur model. The latter demonstrated that magnetic forces generated by a neodymium magnet improved the repair of the defects resulting in the formation of hyaline-like cartilage in comparison to nonmagnetic actuated models [20]. Interesting though is the fact that only when external magnetic forces were applied, the tumor formation (teratoma) was prevented [20]. In the clinical scenario, a significant improvement was found in the outcome scores of the knee defects in 5 human patients 48 weeks after magnetic delivery of autologous BM-MSC [33], suggesting the safety and efficacy of magnetic targeting as a minimally invasive treatment for cartilage repair. Magnetic actuation strategies also allow for more sophisticated ­real-time monitoring of the implanted systems while assisting a remote noninvasive control of therapeutic agents during follow up treatments. The ability to monitoring and control the evolution of tissue regeneration contributes to the possibility for medical intervention at critical stages of the treatment, thus, assisting better outcomes for the patients. 2.2.1  Cell-free therapies Increasing evidence suggests that the therapeutic efficacy of MSC relies on the paracrine signaling necessary to guarantee proper coordination among cells and to modulate the microenvironment surrounding the cells (revised by [34]). The trophic role of MSCs for cartilage repair was highlighted by Pleumeekers et al. [35] with a coculture of BM-MSC and OA chondrocytes. In this work, the extracellular vesicles (EVs) secreted by MSC showed an antiinflammatory effect on TNF-alpha-stimulated OA chondrocytes, inducing the stimulated OA chondrocytes to produce ECM. EVs and exosomes are part of the cell-to-cell communication and may provide novel opportunities as noninvasive tissue-oriented products of regenerative medicine. Toh et al. also reported that MSC exosome-treated rats displayed accelerated neotissue filling and enhanced matrix synthesis of type II collagen and sulfated glycosaminoglycan (s-GAGs) displaying features that resemble hyaline cartilage [36]. In another study performed in a rodent osteochondral model, human embryonic MSC exosome-treated animals revealed enhanced tissue repair at 2 weeks that persisted and extended to week 12 [37]. At 6 weeks, exosome-treated defects showed improved surface regularity and integration with the host cartilage, hyaline cartilage formation with chondrocytic cells, high expression of s-GAG and type II collagen, and low expression of type I collagen [37].

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Vonk et al. also investigated the role of EVs secreted by human BM-MSC in human OA cartilage repair [38]. EVs inhibited TNF-alpha-induced inflammatory effects associated to NF-κB signaling in chondrocytes derived from osteoarthritic patients but favored the production of proteoglycan and collagen II, thus evidencing important regenerative and immunoregulatory properties for the regeneration of OA cartilage [38]. The exploratory findings of these vesicles in cell biology and tissue homeostasis and disease models envision ready-to-use exosomal therapies translation to human patients. EVs and exosomes are acellular products and do not hold the ethical restrictions of a cell-based therapy. Despite the minimal risk of immunogenicity and toxicity as a result from their MSC origin, exosomes and EVs face some concerns on biosafety, efficacy, kinetics, and bio-distribution and clearance that require investigation in future randomized trials.

3  Strategies for bone regeneration The bone tissue is highly dynamic, responding to mechanical stimuli and displaying a remarkable regenerative potential. However, bone defects beyond a critical size, pathological conditions, such as osteogenesis imperfecta, osteoporosis, or autoimmune diseases, and glucocorticoids compromise bone regenerative ability. Moreover, the healing process is not similar for all types of bone, nor follows the same healing fashion. The healing of a long bone defect is generally faster than that of flat bone by approximately twofolds [39]. Unlike the flat bones, the long bone defects undergo endochondral regeneration, recapitulating their embryogenesis [39]. Bone-related diseases or bone loss have usually significant effects on a patient’s quality of life. The worldwide incidence of bone disorders and conditions has trended steeply upward and is expected to double by 2020, especially in populations where aging is coupled with increased obesity and poor physical activity. The ultimate goal of bone regeneration strategies is to provide reliable cost-effective substitutes to autologous bone grafts, the current gold standard treatment of large bone lesions. These substitute biomaterials aim to reproduce the osteogenic properties of bone grafts, while circumventing their associated drawbacks, such as limited availability and requirement for secondary surgery [40].

3.1  Bioinspired materials for bone tissue engineering In order to mimic bone structure and function, the proper control of micro-architecture, mechanical properties, and degradability of the scaffolds/substrates is of utmost relevance to achieve bone regeneration



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(Fig.  3.3A) [48]. For instance, the scaffolds for bone TE should provide an interconnected micro- and macro-porosity (100 μm, respectively) for nutrient diffusion and new bone ingrowth [49]. Moreover, the stiffness of the substrates can determine cell fate [50], therefore, the stiffness of the scaffolds should resemble that of natural bone. Finally, biomaterials for bone TE approaches should maintain shape stability during the healing/regeneration process, in order to promote the formation of the anticipated bone volume and not to lose the bridging effect between bone margins. Different biomaterial-based approaches have shown potential to support bone TE; the use of soft (injectable) polymer-based materials that resemble the bone ECM, and the use of stiff scaffolds that mimic the bone mechanical support or a combination of both approaches.

FIG.  3.3  Biomimetic bone tissue engineering strategies: (A) micro-computed tomography (micro-CT) and histological sections images of trabecular bone [41]. Injectable biomaterials: (B) micro-CT images of injectable CaP and hyaluronic acid microparticles loaded with platelet lysate nanocomposite cement [42]; (C) histological sections from the reconstructed bone using an injectable alginate/hyaluronic acid hydrogel loaded with BMP2 [43]. Scaffold-based approaches: (D) biofabrication of mandible bone using PCL and cellladen hydrogels [44]; (E) micro-CT and SEM images of an ultra-porous TiO2 scaffold [45]; and (F) scanning electron microscope images of collagen/HAp scaffolds derived from marine sources produced using freeze-drying [46, 47]. All figures reproduced with permission from the copyright holders.

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3.1.1  Injectable bone substitutes The injectable biomaterials, capable of setting or crosslinking in situ, have been explored as bone substitutes. Their ability to be injected into complex defects and reshaped to the original bone anatomy before setting is a major advantage over the precasted bone fillers. Therefore, these systems are of particular interest to reduce bone fractures or the filling of bone defects caused by trauma or neoplasia [51]. The injectable calcium phosphate (CaP) cements (CPC) or CaP/polymer composites have found wide applications in bone repair or bone implant fixation [51, 52]. The injectable CPCs have been preferred over the CaP/ polymer composites [e.g., polymethyl methacrylate (PMMA)] given their milder curing temperatures, which reduce the local site necrosis [52]. Injectable CPCs are often a two-component system composed by a solid and a liquid phase which, after being mixed, form a self-setting paste. The solid phase usually contains hydroxyapatite and other calcium salts that can originate apatite or brushite cement, depending on the original composition, the first being more soluble than the second (more information in [52]). The similarity with bone mineral matrix composition makes the injectable CPCs highly biocompatible and osteoinductive. Moreover, their compressive strength, ranging from 0.2 to 180 MPa [53], spans that of human trabecular (4–12 MPa) and cortical bone (130–180 MPa) [54]. For filling of bone fractures, or implant fixation [51] the injectable CPCs are good alternatives to bone grafts. Nevertheless, the low bioresorbability and bone ingrowth rates of the injectable CPCs or the polymeric bone cement impair their application for the regeneration of large bone defects. The incorporation of micro- and macro-porosity has been successfully attempted for accelerating the degradation of injectable cement [42] and improving bone tissue ingrowth and osteointegration of the cement. Moreover, the porogenic elements can be used as growth factors’ delivery vehicles [42, 55]. In a recent work, Babo and coworkers incorporated platelet lysate (PL) into injectable CPCs, both directly into the cement paste or laden in hyaluronic acid (HA) microparticles [42] (Fig.  3.3B). The incorporation of platelet-rich hemoderivatives, namely PL, into regenerative medicine approaches intends to benefit from their richness in cytokines and growth factors involved in the orchestration of wound healing [56]. The incorporation of PL directly into the cement paste or laden into HA microparticles was shown to modulate the release of specific growth factors and, consequently, the osteogenic potential of the injectable CPCs composites [42]. Conversely, the increase of the porosity makes the cement brittle and decreases the load-bearing capacity [42, 53]. The low stability of these cement composites compromise their therapeutic potential, particularly for flat bone regeneration [57–59].



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Produced by the crosslinking of natural or synthetic polymers, the mesh of the injectable hydrogels closely mimics the ECM of connective tissues. Moreover, the mild conditions used to crosslink the hydrogels and the aqueous environment, allow for the loading and sustained release of biomolecules and cell encapsulation [43, 60, 61]. Jung and coworkers proposed a hybrid hydrogel composed of the natural origin polymers alginate and HA, which gelled in situ with Ca2+, for the delivery of the pro-osteogenic BMP-2 and BM-MSCs [43]. The combined effect of BMP-2 delivery and BM-MSCs promoted the regeneration of a mandibular defect in guinea pigs (Fig. 3.3C) [43]. Other versatile crosslinking chemistries have been proposed for hydrogels production with potential for bone regeneration. In our group, methacrylated glucosaminoglycans have been studied for the production of PL-laden hydrogels [60, 62, 63]. These PL-laden hydrogels were shown to be stable and released growth factors in a sustained manner for long time frames [60, 62, 63]. Moreover, the incorporation of PL into HA hydrogels enhanced the osteogenic differentiation of human dental pulp stem cells [64]. Likewise, synthetic materials have been explored to produce injectable hydrogels for the delivery of BMP-2 [65]. Nanoparticles of the thermosensitive poly(phosphazene) were modified with PEG to enhance the affinity of BMP-2 [65]. The hydrogels produced by thermal actuation over nanoparticles solutions were able to release BMP-2 up to 3 weeks promoting ectopic and orthotopic bone formation in mice [65]. The incorporation of hydroxyapatite nanoparticles into the hydrogel matrices, emulating the bone ECM has also been explored [66]. Hydroxyapatite nanocomposite hydrogels of silk are more osteoinductive and promote larger bone regeneration in preclinical models than the injectable silk hydrogels alone [66]. 3.1.2  Scaffold-based approaches In scaffold-based approaches, the structural characteristics, such as roughness, scaffold porosity, pore structure, and interconnectivity, play a crucial role to provide optimal conditions for bone tissue formation in vitro and in vivo [67]. However, in load-bearing applications, the scaffold is also expected to provide sufficient mechanical support during the bone healing process and substitute the lacking mechanical function of the missing or damaged bone tissue. In order to retain its pore architectural structure under physiological loading and to support and transfer the appropriate mechanical stimulation to the bone forming cells within the scaffold and to the host bone, a porous bone scaffold is required to exhibit initial mechanical strength and stiffness that is comparable to the native bone tissue [68]. Therefore, the ideal scaffold should display good mechanical strength, biocompatibility, osteoconductivity, and optimal size and interconnected porous spaces for the bone cells homing.

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Natural and synthetic materials, or their combinations, have been investigated as bone scaffolds (for detailed recent reviews, readers are addressed to Ref. [69, 70]). On one hand, natural polymers, such as collagen, chitosan, silk fibroin, and HA, tend to exhibit excellent biomimicry, biocompatibility, and biodegradability. These materials are particularly suitable to fabricate hydrogels, cryogels, and freeze-dried scaffolds. Nevertheless, the scaffolds created using these methods tend to show limited mechanical properties, which may impair their use in load-bearing bone repair. On the other hand, synthetic polymers, such as polycaprolactone (PCL), poly-lactic acid (PLA), and PEG, offer suitable solutions for mechanical challenges faced by the natural polymers. These have been used to produce bone scaffolds through techniques such as porogen leaching, gas foaming, phase separation, fiber meshing, supercritical fluid processing, microsphere sintering, and 3D printing. Yet, synthetic polymers are characterized by poorer cell attachment properties and slower degradation rates. Blending of natural and synthetic polymers has offered suitable solutions for the shortcomings of both types of polymers. For example, our group developed blends of starch and PCL or PLA to produce fiber-mesh scaffolds with adequate porosity and mechanical properties to support cell adhesion, proliferation, and differentiation into the osteoblast lineage in vitro and bone formation in vivo [71–74]. In a biofabrication approach, Atala’s Group combined the mechanical integrity of PCL with the superior biological performance of cell-laden gelatin/fibrinogen/HA hydrogels to create perfusable constructs using a multihead 3D printer (Fig. 3.3D) [44]. The authors showed the feasibility to produce clinical-relevant size constructs matured into vascularized functional tissues assessed in mandible, calvarial bone, cartilage, and skeletal defects in rodents. Ceramic scaffolds are typically derived from biocompatible and osteoconductive inorganic materials, such as CaPs, bioglass, and titanium oxide (TiO2). The chemical composition of CaPs, such as hydroxyapatite (HAp) and tricalcium phosphate (TCP), is close to the inorganic phase of bone, which makes these materials very attractive for bone scaffolds. In this regard, nearly 60% of the commercially available synthetic bone graft substitutes involve ceramic materials [75]. Porous ceramic scaffolds have been classically produced by different methods including bone decellularization, sponge replication, and gas foaming. Furthermore, two recent studies have explored the possibility of 3D print ceramic-based materials (TCP and HAp) [76, 77]. These ceramic inks allowed rapid manufacturing of scaffolds with micro- and macroporosity needed for bone regeneration. Using a biocompatible and bioactive material, Haugen and colleagues developed bone biomimetic ultraporous TiO2 scaffolds with compressive strength above 2.5 MPa using the foam replication process



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(Fig.  3.3E). These scaffolds promoted osteogenic differentiation in  vitro and bone formation using different in  vivo models [45, 78–80]. Our group has been working with marine species such as coral skeletons, sea urchins, and sponges as biomorphic scaffolds and templates for bone TE, since they represent a promising, inexpensive, and biomimetic alternative to engineered scaffolds [81, 82]. In particular, sponges interconnected porous architecture together with their high content in biosilica have been shown to mimic ideal bone scaffolds and to stimulate bone formation and mineralization [83]. Although ceramic scaffolds might present a biomimetic architecture while exhibiting high stiffness and compressive strength, there is some concern regarding the brittle nature of these materials. Creating composite materials addresses the challenges experienced by single material and has yielded more optimal materials and functionalized scaffolds [69]. Generally, ceramic and bioglass minerals are added to the natural and synthetic polymers to create scaffolds with enhanced mechanical and biological performance. While the presence of CaP-based ceramics in the composites improves compressive strength, degradability rate, and osteogenic capacity of the scaffold, the polymers are credited for maintaining good elastic strength and providing a crosslinking mechanism. Composites such as TCP/polymer, PLA/CaP, HAp/starch, HAp/collagen, and PCL/HAp are frequently reported in the literature [69]. Using novel biomaterials sources, that is, by-products of the fishing industry such as fish skin and bones, collagen/CaP scaffolds produced using freeze-drying supported attachment and proliferation of ­osteoblast-like cells (Fig.  3.3F) [46]. A very interesting and recent composite biomaterial for bone regeneration is the hyperelastic “bone,” developed by Jakus et al. (90 wt% HAp and 10 wt% PCL) [47]. This material was rapidly 3D printed into personalized bone scaffolds with excellent elastic mechanical properties (~32%–67% strain to failure, ~4–11 MPa elastic modulus). Furthermore, it induced osteogenic differentiation of human BM-MSCs cultured in  vitro without exogenous supplementation of ­osteo-inducing factors in the medium and supported new bone growth in vivo. More recently, nanocomposites involving biopolymeric matrices and bioactive nanosized fillers have gained a considerable amount of attention due to their capacity to mimic the nano-sized features of the natural bone mineral [84, 85]. Minardi et  al. fabricated a biologically inspired nanocrystalline magnesium-doped HAp/collagen type I composite scaffold in order to mimic the composition and structure of the osteogenic niche [86]. These scaffolds increased the expression of osteogenic markers in  vitro, when compared with nonmineralized ­collagen-based scaffolds, and allowed the formation of trabecular and cortical bone in vivo.

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4  Tendon regeneration Tendons and ligaments hold a critical role in the musculoskeletal system, transmitting forces, and stabilizing the joints, being able to withstand the high tensile forces upon which locomotion is entirely dependent [16, 87, 88]. The anatomical location, architecture, and function make these tissues highly prone to injuries with limited endogenous resolution. The development of engineered functional tendons is greatly dependent on the mimicry of tendon mechanical behavior and structural components in order to recapitulate the tendon matrix toward the development of functional substitutes. Moreover, the replication of native tendon microenvironment, requiring a highly aligned architecture and oriented cell morphology [89] is a key aspect in biomaterials design. Therefore, the challenge to engineer advanced functional biomaterials holding mechanical and structural cues highly depends on the proper combination of material type, processing technique, and structure design. In terms of mechanical properties, tendons exhibit a unique crimp pattern and viscoelastic properties akin to a spring that enables tendon to effectively store and subsequently release mechanical energy [90]. The profile of a typical tendon stress-strain curve is composed of different regions: at the toe region of strain up to 2% the tendon retains a characteristic crimped structure; the linear region in which the strain remains lower than 4% and the tendon behaves in an elastic fashion being able to lengthen its crimped collagen fibers and withstand forces. The linear region is representative of the physiological range of the tendon and the slope of the curve defines the Young’s modulus of the tissue. Stretching over 4% results in microscopic tearing and tendinopathy can develop, whereas repeated micro-tears and strain beyond 8%–10% leads to macroscopic failure and tendon rupture [12, 91, 92]. Thus, in order to engineer a tendon mimetic scaffold, the characteristic nonlinear biomechanical behavior of tendons and the characteristic anisotropic hierarchical structure (as described in Section 1.1) must be combined.

4.1  Biomaterial processing technologies to meet tendon function and properties 4.1.1  Fiber-based technologies The complex fibrous hierarchical structure of tendon instigates engineers in the development of materials that ensure enough strength under uniaxial tension and, at the same time, viscoelastic properties in order to optimize stiffness under different loading environments [93]. A growing number of publications resources to aligned scaffolds that can exert influence on cell morphology and tenogenic differentiation in both in vitro and



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in vivo models [94–101]. A recent example of the development of aligned scaffolds for tendon TE was proposed by Zheng and coworkers [102] in which a macroporous 3D aligned collagen/silk scaffold was investigated in a rabbit massive rotator cuff tear model. Aligned collagen/silk scaffolds were fabricated using 12 yarns of silk fibers and also resourcing to a unidirectional freezing technology. These scaffolds presented a profound influence on the cellular morphology and arrangement of rabbit tendon stem/progenitor cells. Moreover, the in  vivo performance revealed abundant organized bundles of collagen fibers formed in the outer zone of the macroporous 3D aligned collagen/silk scaffold and evidence of a denser, matured and organized regenerative tissue with cell infiltration [102]. Indeed, aligned fiber-based scaffolds were suggested to guide cell response from repair to healing [95]. Lee et al. also evaluated the effect of fiber diameter of unaligned meshes and fiber alignment on human tendon fibroblast attachment, organization, growth, and phenotype, as models of connective tissue repair and healing. Unaligned fibers with nanometer diameters promoted cell proliferation and matrix deposition as well as the expression and activity of RhoA and Rac1, characteristic of the initial, proliferative phase of wound repair. Moreover, the mature repair model represented by unaligned micron-sized fibers supported cell organization and adhesion, while suppressing cell growth and ECM biosynthesis, indicative of the remodeling phase of tissue repair. The nanofiber model showed matrix alignment as a critical design factor for circumventing scar formation and promoting biological healing of soft tissue injuries [95]. Conventional fiber fabrication techniques toward the replication of tendon structure relied on spinning-based methods, such as wet- and melt-spinning [103–105]. In recent years, electrospinning systems have been increasingly used for the production of nano- to micro-fibrous anisotropically aligned biomaterials from a wide range of polymer matrices (Table 3.2). Electrospinning technique consists of a capillary system through which the spinning dope solution to be electrospun is forced, a high-voltage source, and a grounded collector [89]. The potential of electrospun nanofiber scaffolds to modulate cells behavior has been extensively reviewed [114–116]. Generally, the produced electrospun fibers are 2D matrices that need to be assembled into hierarchical scaffolds to mimic tendon architecture. One option that has also been considered in tendon TE is the use of textile technologies such as knitting, weaving, or braiding [117, 118] for assembling fibrous structures into 3D tendon mimetic scaffolds. A recent study by Rothrauff et al. provides a representative example of this strategy for tendon and ligament TE [108]. Multilayered scaffolds of aligned electrospun nanofibers were produced using two designs, stacking or braiding. For multilayered scaffold fabrication, the PCL and poly-l-lactic acid (PLLA) nanofibrous sheets were either manually stacked or rolled and braided.

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TABLE 3.2  Materials used in electrospinning technique for tendon applications Study type

Main outcomes

References

Chitosan (CHT) Polycaprolactone (PCL) Cellulose nanocrystals (CNC)

In vitro

The use of CNCs significantly reinforces the mechanical properties of aligned nanofiber scaffolds; The nanotopography and microstructure of hierarchical assemblies of nanofiber threads promoted tenogenic differentiation of hASCs.

[106, 107]

Polycaprolactone (PCL) Poly-l-lactic acid (PLLA)

In vitro

Aligned nanofibrous scaffolds-stacked versus braided: braided constructs upregulated tenogenic differentiation of hMSCs to a greater degree than stacked constructs.

[108]

Polycaprolactone (PCL)

In vitro

Tube-shaped scaffolds with bi-axially aligned fibers based on a “Chinese-fingertrap” design showed nonlinear mechanical response. hMSCs adhered, proliferated, and aligned along fibers. The scaffold geometry encouraged the differentiation of MSCs toward tendon.

[98]

Polycaprolactone (PCL) Poly-lactic acid (PLA)

In vitro

Nanofiber yarns were processed into plain-weaving fabrics, which could guide hASCs and tenocytes alignment and the synthesis of a tendon-like ECM containing oriented fibers of TNMD and COL1 by contact guidance.

[109]

Poly-l-lactic acid (PLLA) Polyethylene oxide (PEO)

In vitro In vivo

Trichostatin A-incorporated aligned fibers modulated HDAC activity, maintaining Scx expression and promoting tenogenesis of TSPCs. Moreover, the in situ implantation study in a rat model further confirmed that the aligned TSA scaffold promoted the structural and mechanical properties of the regenerated Achilles tendon.

[97]

Polycaprolactone (PCL)

In vivo

Fibers were manually twisted into yarns and the effects of gamma and ethanol sterilization were evaluated over a 6-week time period in murine tendon model. Neither techniques had an observable effect on the functionality of the scaffold when compared to the autograft control.

[110]

3.  Musculoskeletal tissue engineering approaches

Material(s)



Polycaprolactone (PCL) Polyamide 6 (PA6) Silica nanoparticles

In vitro

Electrospun bead-on-string fibrous nanocomposite scaffolds (incorporating silica particles) showed greater cell spreading, proliferation, activity, and ECM deposition compared to the pristine polymeric ones.

[111]

Polycaprolactone (PCL) Hydroxyapatite (HA)

In vitro

The hybrid electrospinning and electrospraying process provided an efficient method to achieve a gradient structure with a controllable layer thickness customizable for tendon-bone interface.

[112]

Silk fibroin (SF) Polycaprolactone (PCL)

In vitro In vivo

The combination of biochemical cues from SF and physical cues from fiber alignment had a positive effect on the direction of RDFBs migration, proliferation, and upregulation of gene expression of tendon-specific ECM proteins. Histological and immunohistochemical analysis revealed the production and deposition of collagen and tenascin C in Achilles tendon defect.

[113]

4  Tendon regeneration

Abbreviations: COL1, collagen type I; ECM, extracellular matrix; hASCs, human adipose stem cells; HDAC, histone deacetylases; hMSCs, human mesenchymal stem cells; RDFBs, rabbit dermal fibroblasts; Scx, scleraxis; TNMD, tenomodulin; TSA, trichostatin A; TSPCs, tendon stem/progenitor cells.

93

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Braiding technique increased suture-retention and tensile strength, but decreased cell infiltration and proliferation compared to stacked constructs. Despite this, both multilayered scaffolds supported tenogenic differentiation of seeded MSCs by expression of tenogenic markers [108]. Recently, our group developed aligned nanofibrous scaffolds [106, 107] aimed at tendon TE using electrospinning technique. Electrospun nanofiber scaffolds combining chitosan (CHT), a natural polymer, and PCL, a synthetic polymer, were reinforced with cellulose nanocrystals (CNCs) [106], and the electrospinning was conducted using a home built disk electrospinning unit. The topography of anisotropically aligned scaffolds, as opposed to randomly oriented scaffolds, promoted a remarkable uniaxial cell orientation and induced elongated tendon cells morphology. Moreover, the incorporation CNCs into electrospun natural/synthetic polymer (PCL/CHT) nanofiber bundles significantly improved mechanical properties in tendon/ligament relevant range (σ = 39.3 ± 1.9 MPa and E = 540.5 ± 83.7 MPa, P