- Email: [email protected]
Tissue engineering of biological cardiovascular system surrogates
Current Review Tissue Engineering of Biological Cardiovascular System Surrogates Catherine E. Sarraf, PhD, MRCPath, Andrew B. Harris, BSc, Andrew D. McCulloch, BEng and Mark Eastwood, PhD, MIMechE Centre for Tissue Engineering Research, Department of Biomedical Sciences, University of Westminster, London, UK
Cardiovascular diseases are common in ageing communities globally. This fact is most striking in the industrialised world where the aged population makes up a large proportion of society. Elderly patients are frequently treated surgically with grafts to replace damaged tissues and vessels. The number of human-donated components is insufficient and synthetic surrogates are sought. These might be wholly mechanical, wholly biological, or tissue engineered complexes of cells and their products growing in a scaffold. At present, many such composites exist with potential for use as substitutes for specific blood vessels. The challenges of producing tissue engineered heart valves are now being widely explored. Neotissues must provide an effective, durable, non-thrombogenic and non-immunogenic substitute that will fulfil the purpose of the natural tissue. The aims and scope of this paper are to review current and novel concepts in the field of tissue engineering of biological cardiovascular system surrogates. Mechanical stresses and strains on cardiovascular cells in vitro have been recognised and can be measured by a culture force monitor. Physiological stresses can be generated by a tensioning culture force monitor and applied to engineered tissue, aligning the cells and mimicking arterial wall architecture. The hydrostatic forces a vessel experiences and mechanical parameters of blood vessels can be studied in the tubular culture system of a multi-cue bioreactor. (Heart, Lung and Circulation 2002; 11: 142–150) Key words: biological matrix, bioreactor, tissue engineering.
T
issue engineering is a growing field of science in medicine that integrates relevant engineering principles and biology to produce organs or structures for intracorporeal implantation. In the large ageing populations of the industrialised world, cardiovascular disease is a predominant killer. In the USA, in addition to patients with end-stage heart failure who receive heart transplants, deaths from arterial disease and valvular disease claim more than 40 000 and 19 000 people per annum, respectively.1 Demands on cardiovascular replacement surgery greatly outstrip availability of Correspondence: Catherine Sarraf, Centre for Tissue Engineering Research, Department of Biomedical Sciences, University of Westminster, 115, New Cavendish Street, London W1W 6UW, UK. Email: [email protected]
donor organs and components, and here we review the goals of synthesising vascular and heart valve prostheses from biomedical materials and cultured cells. The major disorders of the vascular system are aneurysm of the ascending aorta and coronary artery occlusion, while the main causes of heart valve malfunction are rheumatic valvular disease and calcific aortic valve disease. Complete aortic root replacement can be done in cases of narrow aortic annulus,2 coronary bypass grafting using arterial conduits is common3 and valve replacement is widely used for end-stage valvular heart disease.4 The potential clinical application of engineered tissues determines the choice of a support matrix. Advances in cardiovascular tissue engineering and scaffold technology might ultimately lead to developing a whole tissue engineered heart.5
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Non-Biological Prostheses The type of implant chosen must provide the most advantageous result for the patient. It should be effective, long lasting, amenable to surgery, have a low thromboembolic potential and be non-immunogenic. Although a composition as close as possible to the natural tissue would seem preferable, considerable success has been consistently achieved in the cardiovascular system using non-biological prostheses. In the case of valves, success has been reported using polyurethanes,6 but for arterial implants, small diameter vessels made of plastic polymers alone are subject to occlusion.7
Wholly Biological Surrogates In contrast to totally synthetic implants, cellular complexes have been used with no artificial or modified biological extracellular matrix. In a dog model, fresh pericardial tissue has been fashioned into small diameter arterial grafts.8 In a different study,9 a tubular styrene mandrel was first covered with an acellular collagenous membrane, then with human vascular smooth muscle cells that had been cultured with ascorbic acid. A sheet of human fibroblasts was wrapped around this ‘media’, followed by endothelial cells, to make a three-layered co-culture. The cells produced extracellular matrix proteins (collagen and elastin), desmin, von Willebrand factor and prostacyclin appropriately. The endothelium was anti-thrombotic and inhibited platelet adhesion. A burst strength of 2000 mmHg was reported, comparable to that of human vessels. The engineered tissues took 3 months to grow in culture. In a lamb model,10 it has been shown that tissue engineered heart valves synthesised using cells from femoral arteries developed into acceptable leaflets.
Bioengineered Tissue of Cells on a Scaffold A compound approach is to use a bioengineered tissue equivalent seeded with cells that would be found in the native environment: endothelium, smooth muscle cells, myofibroblasts and fibroblasts. A considerable contribution expected from these cells is that they should selfassemble an extracellular matrix of collagen and elastin within the framework provided. The first attempt to construct such a cells-plus-scaffold blood vessel was made more than 15 years ago.11 Bovine smooth muscle cells in a collagen culture medium were cast around an annular mould. When this culture was mature, a Dacron® (non-absorbable plastic) mesh sleeve (Du Pont, Wilmington, DE, USA) was rolled over for mechanical support, followed by seeding with adventitial fibroblasts.
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The constructs required 2 weeks to become fully contracted and then samples were lined with endothelial cells. Electron microscopy showed that the smooth muscle cells secreted collagen in the manner of fibroblasts. This system, however, had low burst pressure. Elastin had not been incorporated in the matrix, and models made with only one Dacron® mesh ruptured at less than 10 mmHg. Although strength to withstand up to 180 mmHg (still considerably lower than that required for a human blood vessel implant) could be achieved, it was only with the use of numerous Dacron® meshes, the whole becoming too stiff to be effective. As the L’Heureux9 group found in their wholly biological neovessel more than 10 years later, the proliferated endothelial layer successfully produced von Willebrand factor and inhibited platelet aggregation. In subsequent years, advances were made in the principle of using native cells with a variety of fabrics. Hydrolytic degradable constructs,12–14 absorbable patches,15 biologically modified native tissues,7 collagen sponge matrices,16 and third generation biomedical materials were tested. Engineering development of biomaterials for scaffolds, particularly techniques appropriate for mass production, are essential in the worldwide market. After improvement of scaffolds and co-cultures in the laboratory, to aid patients, these innovations must have practical reliability. Constituents must always be readily available, fully functional and reproducible. Pragmatically, products must have a sufficient shelf-life and be able to withstand shipping and storage.17 In the cardiovascular system, mechanical properties of the matrices themselves are crucial. They must be capable of withstanding complex mechanical and dynamic environments, and honeycomb-like polycaprolactone structures have been compared for compressive stiffness and offset yield strength.18 In the case of engineered heart valves, precise shape of the graft is important. Stereolithic models derived from X-ray computed tomography with appropriate software have been used to generate biocompatible grafts.19
Hydrolytic Degradable Scaffolds The Dacron® mesh used by Weinberg and Bell11 was a physical support for the neovessel only, with no pretension toward contributing either chemically or physiologically to tissue synthesis. But in 1998, sheets of polyglycolic acid (PGA) mesh were used as a substrate for seeding human cells.12 Such a biodegradable framework gradually deteriorates, provokes little or no immune response in the host and minimises any potential chronic inflammatory reaction. The PGA–cell complex allows native cells to grow, remodel the graft and contribute to repair.
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Polyglycolic acid has been used as a substrate in an ovine model of autologous heart valve leaflet replacement.13 Fibroblasts, then endothelial cells, were cultured onto the polymer. The synthetic complexes were grafted into lambs, with acellular PGA implanted into control hosts. In the controls, grafts degraded but all cell-seeded categories integrated well. Also in the ovine model, a heart valve equivalent was manufactured with nonwoven PGA plus a thin layer of poly-4-hydroxybutrate.14 Here, the constructs were cultured for 2 weeks in vitro in a pulse duplicator under gradually increasing flow and pressures, before being resected back into lambs. Induced biomimetic flow improved valve function.
Absorbable Patches Absorbable patches of poly beta-hydroxybutrate have been used in calves to close the pericardium after cardiopulmonary bypass. At investigatory re-operation, patch material contained pronounced macrophage infiltration, but no regenerative mesothelium.15 Adhesions in the healing pericardium might be caused by secretions of mesothelial cells.
Biologically Modified Native Tissues To engineer a narrow diameter artery in the rabbit, a porcine cell-free collagen layer was used,7 derived from the submucosa of the small intestine. To reduce thrombogenicity, luminal surfaces of grafts were treated with heparin-bezalkonium chloride cross-linked with 1-ethyl3-(3-dimethylaminopropyl) carbodiimide hydrochloride. Even so, without anticoagulant therapy, vessels were unstable. The technique was modified by deposition of bovine type I collagen on the lumen, which improved patency. The cell-free grafts were intact up to 90 days post surgery, and although they consisted of porcine and bovine collagen components, immunogenicity was minimal. There was considerable rabbit cell infiltration into the graft. Implants were reactive to experimental vascular agonists, had a non-thrombogenic surface, and had a burst resistance of up to 1000 mmHg. For heart valves, a technique of implanting a complex composed of combined fresh non-vascular tissue and cells was pioneered by Manothaya et al.20 Dura mater was used as the tissue source for frame-supported valves, and short-term results were reported as being satisfactory.
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collagen and 4–9% type III collagen). They are biocompatible with host tissue and are biodegradable. To prepare the sponges, solutions of extracted collagen are freeze-dried, resulting in a porous matrix with a mean pore size of 58 ± 24 µm. In our laboratory, similar calf skin-derived sponges have been seeded with porcine aortic wall myofibroblasts.21 Cells were infiltrative and proliferative, and porosity of the sponge provided excellent conduits for cell migration. Surface myofibroblasts were densely packed, lying in stratified layers, four to six cells deep, their long axes distinctively at right angles to those of the infiltrative population. Mouse fibroblasts have been found to proliferate to a constant cell density independent of seeding level,22 but here, developmental pattern was related to seeding number. Surface myofibroblasts generated sparse amounts of matrical proteins, but central sponge zones were replete with dense aggregates of banded collagen and elastic fibres, emanating from the well-spaced cells (Figs 1,2). This difference could be because nutrient and gas transfer may be restricted at distance from the liquid culture solutions, but the change in phenotype is possibly myofibroblast differentiation. This is completely in accord with normal myofibroblast behaviour in a native connective tissue, where cells are separated by a multiplicity of extracellular components, and myofibroblasts readily change direction of metabolic activity.23,24 Transforming growth factor β, for example, stimulates fibroblast culture behaviour, which is related to applied mechanical loading.25
Collagen Sponge Matrices Collagen sponges cross-linked or non-cross-linked with diphenylphosphorylazide seeded with bovine foetal chrondrocytes have been used to reconstruct cartilaginous tissues.16 Sponges are manufactured from collagen extracted from calf skin (90–95% dry weight type I
Figure 1. High-power transmission electron micrograph of the plasma membrane of a porcine myofibroblast grown in a collagen sponge. Collagen fibres (arrows) are clearly seen emerging from the cell into the extracellular matrix. Length of cell culture time: 4 weeks. Bar, 0.6 µm.
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Third Generation Biomedical Materials First generation biomedical materials for implantation have been defined as being ‘inert’, and those of the second generation as having controlled chemical selfbreakdown and resorption.26 Extrapolation of properties of these materials lies in development of tissue substitutes that are auxiliary to, and able to integrate with, the host body. Stimulatory proteins can be bound to polymer surfaces; adapted polymers have been produced that influence interactions with endothelial cells.27 Cellular growth factors that stimulate proliferation and development of cardiovascular cells can be chemically combined with structural matrices. Angiogenic patches of human dermal fibroblasts have been successfully used to cover infarcted cardiac tissue in immunodeficient mice.28 Exciting prospects lie in the possibility to manufacture novel, non-immunogenic bioengineered scaffolds incorporating molecules that stimulate appropriate genes in neighbouring host body cells. Subcellular localisation of gene products can be demonstrated by techniques such as electron microscopy, immunocytochemistry and in situ hybridisation.29 Supplementation of myofibroblast cultures with ascorbic acid for heart valve tissue engineering has already been used to increase collagen production.30 Native proteins can be produced to augment cell production and differentiation, repair the defect, heal the wound, and control inflammatory complications. Growth factor-stimulated tissue growth in arteries and valves, however, is an issue with which one must be concerned. In wound healing, there is limited and revers-
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ible hypertrophy. Excessive stimulation of cell proliferation could all too easily result in uncontrolled neoplasia. A list of some materials used in bioengineered scaffolds can be found in Table 1.
Mechanical Forces and Bioreactors In vivo, the architecture of heart and blood vessel connective tissue is not haphazard. The mesenchymal cells are sensitive to forces applied to them from their environment and integrate both mechanical and biochemical signals for growth, maintenance and repair. In turn, development and function of the endothelium requires intimate intertissue communication. First, pulsatile fluid flow gives rise to a radial (and longitudinal) strain; second, steady state flow (with no pulsatile component) causes shear stress on luminal wall surfaces, which results in endothelial cells aligning and bearing in the direction pertinent to the shear; and third, axial loading causes pure longitudinal strain to develop. Bioreactors have been used in the field of cardiovascular tissue engineering since the prototype blood vessel was constructed.11 Their role initially was as a culture vessel, to disperse cell density evenly by rotation and to provide a sterile, nutrient rich environment in which the preformed bioengineered culture could develop. Recently, however, mechanical forces derived by the cells have been measured,31–34 and externally driven forces have been applied to cultures, resulting in streamlined tissues with increased strength and patency.
Table 1. Examples of some materials used in bioengineered scaffolds
Figure 2. Transmission electron micrograph of a porcine myofibroblast grown in a collagen sponge, associated with elastic fibres (arrows), synthesised during the period of cell culture. Length of culture time: 4 weeks. Bar, 1.1 µm.
Synthetic polymers used as scaffolds in tissue engineering15,16 Poly beta-hydroxybutrate Polyglycolic acid Polylactic acid Polyglycolic acid-poly L-lactic acid co-polymer Polyethylene Polyethylene oxide Carrier materials of natural origin used as scaffolds in tissue engineering16 Collagen sponges Collagen gel Alginate Fibrin Hyaluronic acid Gelatin Collagen–glycosaminoglycan matrices
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Culture Force Monitor and Tensioning Culture Force Monitor Extensile cells create forces and also respond to stimuli applied to them from their surroundings. Such forces have been accurately measured on an apparatus called a culture force monitor (CFM) and have been provoked and measured by a tensioning-culture force monitor (t-CFM).31 Culture force monitor. A CFM was developed in 199431 in our laboratory. It can be used to study 3-D mechanical forces in different cell types in a variety of scaffolds. In addition, the CFM can be used to record cell responses to therapeutic drugs and/or growth factors. The CFM is an improvement on previous similar devices because its force transducer is coupled directly to a computer via an analogue-to-digital converter (Fig. 3). Movement of fibroblasts suspended in a collagen matrix has been measured to 10–6 m. The average force produced by a single fibroblast was 10–10 Newtons. In a period of contraction, the majority of force has been shown to be generated in the first 8 h, and morphological studies illustrate that changes in cell shape (extension of cytoplasmic processes, development of attachments, cytoskeletal rearrangement) initiate the increments.35,36 In recessive dystrophic epidermolysis bullosa (RDEB), in untethered fibroblast-populated collagen lattices, forces of normal fibroblasts and RDEB fibroblasts were quantified and compared. RDEB fibroblasts were hypercontractile, generating 2.5-fold the force of normal fibroblasts.37 The effect of burn blister fluid on fibroblast contraction showed that burn fluid stimulated fibroblasts to contract, predominantly in a period of 24–48 h after application,38 although chronic wound fluid inhibited cell proliferation.39 An adaptation of the CFM to measure forces generated by tissue cohesion after injury, an adhesion
Figure 3. Structure of the culture force monitor (CFM), indicating relationships of the force transducer and the culture well.
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CFM, has demonstrated levels of fibroblast-mediated adhesion in vitro between contrasting tissue interfaces.40 Formation of post surgical adhesions are also important in human primary and re-operative pericardium.41 Tensioning culture force monitor. The t-CFM is a derivative of the CFM and allows the application of precise mechanical loads of physiological dimension to be applied to fibroblast populated lattices.31 For the t-CFM (Fig. 4), a micro-stepping motor was added to the monitoring devices of the CFM, and a theoretical positioning accuracy of 10–8 m has been achieved. Fibroblast populated collagen matrices can now be loaded, experimental forces of predetermined magnitude engaged, and cellular responses in terms of both matrix tension and biochemical constitution measured. The t-CFM applies minute, unidirectional loads to cell–collagen complexes with specific loading patterns of striking and repeatable cell alignment reliably predicted and produced.31 Fibroblasts rank themselves along lines of maximum principal isostrain32 and orientate in consistent responsive patterns (Figs 5,6). Investigating homeostasis in dermal fibroblasts, the t-CFM has measured and quantified mechanical responses to precisely defined tensional loads.33 Fibroblasts actively migrated across restrained gels more energetically than across untethered gels.42 When cyclical loads were applied to produce unidirectional strain, cells which had been spherical at seeding were induced into characteristic bipolar-shaped fibroblasts aligned parallel to the loading axis of the neotissue. Dependence of cell alignment on strain contours implied that it was not only magnitude and direction of loading that dictated alignment, but also physical shape of the material subject to the load.
Figure 4. Structure of the tensioning culture force monitor (t-CFM), showing addition of a micro-stepping motor system to the basic CFM design.
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Figure 5. Scanning electron micrograph of freeze-fractured gel with non-aligned fibres. Reproduced with permission from Eastwood et al. 1998.32 Bar, 10 µm.
Biochemical changes took place in fibroblasts subjected to mechanical loading, and chemical properties were modulated by changes in applied stress. Effects of mechanical stimulation on production of matrix metalloproteases (MMP) were investigated. MMP-9 synthesis was more sensitive to tension than MMP-2 for both induction and activation of enzymes, and also, stromelysin activity decreased with tension.43 Quantitative reverse transcription (RT)–polymerase chain reaction (PCR) revealed twofold, fourfold and threefold increases in MMP-1, MMP-2 and MMP-3, respectively, in nonaligned strain zones relative to aligned strain zones during contraction in wound healing.44
Pulsatile Bioreactors Prototype pulsatile bioreactors have been developed by the groups of Hoerstrup et al.,14 Zeltinger et al.45 and Eastwood et al.21 In general, they consist of the bioreaction chamber, a culture media exchange mechanism, a means of generating pulsatile flow (micro-stepping motors), force transducers and computerised control of all reactions and forces applied. Multiple loading regimes are available to condition the cell–scaffold complex. For vascular specimens, a tubular culture vessel is used. This type of tissue engineered construct has either to have been cultured on an annular mould, or a planar matrix has been cell seeded, rolled, sutured and finally the edges bonded together with an effective fibrin sealant. Pulsatile flow causes cyclic fluid shear stress on the inner wall of the neovessel in addition to generating
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Figure 6. Scanning electron micrograph of freeze-fractured gel with aligned fibres (large arrow) emanating from a fibroblast (small arrow). Reproduced with permission from Eastwood et al. 1998.32 Bar, 10 µm.
radial and longitudinal stresses within the wall. These stresses and associated strains have the effect of orientating all cells within and appropriately aligning any fresh collagen and/or elastin that they might synthesise.
Multi-Cue Bioreactor A development of the original pulsatile bioreactor is the multi-cue bioreactor (M-CB)21 from which a range of types of stimulus (cue) can be delivered via a hermetically sealed electric motor linked to a rack and pinion straining system and a peristaltic pump head. The strain unit has a mathematical resolution of 2 × 10–6 m and total displacement of 70 mm. Pressures in the range of 450 mmHg can be generated. Strain induction and pressure gradients can also be measured across a structure’s wall.33,34,43,46 In addition to inducing conditioning strain, at any time during a procedure, in situ non-destructive tensile testing can be performed at key intervals. The M-CB computer controlled stepper motor units are programmed through X-Ware® software (Parker Hannifin, Poole, Dorset, UK) whose parameters are able to direct fluid velocity and acceleration, as well as axis of rotation. Further, mechanical cue drivers can be used in conjunction with a transducer array to perform feedback conditioning, where programs are adjusted in response to force/pressure readings, as occurs in vivo. Versatility and adjustability of the M-CB provide the flexible range of variables needed to condition a simple cell-seeded matrix from a mechanically unstructured mass into a viable, implantable and biologically sufficient neotissue.
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Cell Origin and Graft Success
Conclusion
To manufacture an effective, durable, tissue engineered composite graft for human implantation, the biological source of the seeded cells is of the highest importance. To avoid immune system complications, cells are ideally autologous (from the patient him/herself), but for practical purposes (such as time required to grow a sufficient culture), the nearest immunological match, allogenic cells, might have to be resorted to. The use of autologous cells has many advantages (not least that ethical considerations are not incurred), and tissue engineering had its first successes more than 20 years ago,47 when patients’ own skin cells were first used to cover wounds. This approach might be technically taxing for cardiovascular grafting at present. Allogenic cardiovascular cells have been used successfully in animal studies of vessel and valve implants13,14 but, once more, ethical constraints prevent similar work in humans. A source of dermal fibroblasts is available from human foreskin.48 Such cells have been shown to be acceptable immunologically, but dermal fibroblasts have proven to be inferior to cardiovascular fibroblasts in potential heart prostheses.10 Similarly, human umbilical cord endothelium is a source of suitable human cells,49 but neonatal myofibroblasts and embryonic fibroblast stem cells for allogenic surgical purposes, although useful animal experimental or veterinary tools, cannot be considered for the purpose of making human grafts. Populating a bioengineered tissue scaffold with adult human stem cells (autologous or allogenic) might be possible. The potential for stimulating maturation of adult stem cell progeny along predetermined lines has been widely demonstrated.50 In the mouse, myofibroblasts in colonic lamina propria have been derived from bone marrow stem cells, transplanted between adults,51 so the possibility of producing vascular myofibroblasts is raised. Harvesting and transplanting human bone marrow are common techniques, and selectively directing differentiation outcome in an engrafted, bioengineered scaffold could solve problems of effectiveness and durability of synthetic connective tissue surrogates. Bioengineered scaffolds with or without seeded cells, but porous enough to permit rapid infiltration of local native cells into the graft would be ideal. Huynh et al.7 found their porcine–bovine collagen complex immunologically acceptable in rabbit hosts, and the implant rapidly became colonised by rabbit cells. Fibroblasts, myofibroblasts, and endothelial cells have the genetic ability to express proteins controlling cell migration and cell proliferation: these activities are among their normal repertoire in wound healing.
Consummate tissue engineered cardiovascular replacements must possess all of the properties of the natural tissue. Technically, there are numerous and consistent advances in the fields of matrix commercial availability, linking scientific discovery with practical medical application. Scaffolds, whether synthetic or biological, must satisfy mechanical demands in addition to being non-immunogenic and non-thrombogenic. Mechanical stresses and strains on blood vessel and heart valve cells have been recognised and measured, and can be reproduced to condition tissue constructs in vitro. Seeded and cultured cells must be appropriate and effective for the purpose. In the case of fibroblasts or myofibroblasts, their innate ability to produce requisite extracellular matrix proteins themselves is an invaluable aid for the tissue substitute. The possibility of incorporating cellular growth factors into a potential implant is attractive, although this should be pursued with caution. Vascular endothelial growth factor, the fibroblast growth factors, plasma proteins and cytokines all stimulate vascular cell growth, and synthesis of collagen and elastin. When a chosen scaffold slowly disintegrates in a controlled fashion after implantation, while being replaced by totally sufficient, natural structural cells and fibres, the outcome will be a replaced tissue rather than a replacement tissue.
Acknowledgements We would like to thank the Engineering and Physical Science Research Council, the Phoenix Trust, and Pearl Assurance for financial support for this work.
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44. Mudera VC, Pleass R, Eastwood M et al. Molecular responses of human dermal fibroblasts to dual cues: contact guidance and mechanical load. Cell Motil. Cytoskel. 2000; 45: 1–9. 45. Zeltinger J, Landeen LK, Alexander HG, Kidd ID, Sibanda B. Development and characterization of tissue-engineered aortic valves. Tissue Eng. 2001; 7: 9–22. 46. Prajapati R, Eastwood M, Brown R. Duration and orientation of mechanical loads determine fibroblast cyto-mechanical activation: monitored by protease release. Wound Report Regen. 2000; 8: 239–47.
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47. Eldad A, Burt A, Clarke JA. Cultured epithelium as a skin substitute. Burns 1987; 13: 173–80. 48. Harding KG, Morris HL, Patel GK. Healing chronic wounds. BMJ 2002; 324: 160–63. 49. Barker JN, Wagner JE. Umbilical cord blood transplantation: current state of the art. Curr. Opin. Oncol. 2002; 14: 160–64. 50. Alison MR, Poulsom R, Jeffery R et al. Hepatocytes from nonhepatic adult stem cells. Nature 2000; 406: 257–8. 51. Brittan M, Hunt T, Jeffery R et al. Bone marrow derivation of pericryptal myofibroblasts in the mouse and human small intestine and colon. Gut 2002; 50: 752–7.
Cardiovascular diseases are common in ageing communities globally. This fact is most striking in the industrialised world where the aged population makes up a large proportion of society. Elderly patients are frequently treated surgically with grafts to replace damaged tissues and vessels. The number of human-donated components is insufficient and synthetic surrogates are sought. These might be wholly mechanical, wholly biological, or tissue engineered complexes of cells and their products growing in a scaffold. At present, many such composites exist with potential for use as substitutes for specific blood vessels. The challenges of producing tissue engineered heart valves are now being widely explored. Neotissues must provide an effective, durable, non-thrombogenic and non-immunogenic substitute that will fulfil the purpose of the natural tissue. The aims and scope of this paper are to review current and novel concepts in the field of tissue engineering of biological cardiovascular system surrogates. Mechanical stresses and strains on cardiovascular cells in vitro have been recognised and can be measured by a culture force monitor. Physiological stresses can be generated by a tensioning culture force monitor and applied to engineered tissue, aligning the cells and mimicking arterial wall architecture. The hydrostatic forces a vessel experiences and mechanical parameters of blood vessels can be studied in the tubular culture system of a multi-cue bioreactor. (Heart, Lung and Circulation 2002; 11: 142–150) Key words: biological matrix, bioreactor, tissue engineering.
T
issue engineering is a growing field of science in medicine that integrates relevant engineering principles and biology to produce organs or structures for intracorporeal implantation. In the large ageing populations of the industrialised world, cardiovascular disease is a predominant killer. In the USA, in addition to patients with end-stage heart failure who receive heart transplants, deaths from arterial disease and valvular disease claim more than 40 000 and 19 000 people per annum, respectively.1 Demands on cardiovascular replacement surgery greatly outstrip availability of Correspondence: Catherine Sarraf, Centre for Tissue Engineering Research, Department of Biomedical Sciences, University of Westminster, 115, New Cavendish Street, London W1W 6UW, UK. Email: [email protected]
donor organs and components, and here we review the goals of synthesising vascular and heart valve prostheses from biomedical materials and cultured cells. The major disorders of the vascular system are aneurysm of the ascending aorta and coronary artery occlusion, while the main causes of heart valve malfunction are rheumatic valvular disease and calcific aortic valve disease. Complete aortic root replacement can be done in cases of narrow aortic annulus,2 coronary bypass grafting using arterial conduits is common3 and valve replacement is widely used for end-stage valvular heart disease.4 The potential clinical application of engineered tissues determines the choice of a support matrix. Advances in cardiovascular tissue engineering and scaffold technology might ultimately lead to developing a whole tissue engineered heart.5
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Non-Biological Prostheses The type of implant chosen must provide the most advantageous result for the patient. It should be effective, long lasting, amenable to surgery, have a low thromboembolic potential and be non-immunogenic. Although a composition as close as possible to the natural tissue would seem preferable, considerable success has been consistently achieved in the cardiovascular system using non-biological prostheses. In the case of valves, success has been reported using polyurethanes,6 but for arterial implants, small diameter vessels made of plastic polymers alone are subject to occlusion.7
Wholly Biological Surrogates In contrast to totally synthetic implants, cellular complexes have been used with no artificial or modified biological extracellular matrix. In a dog model, fresh pericardial tissue has been fashioned into small diameter arterial grafts.8 In a different study,9 a tubular styrene mandrel was first covered with an acellular collagenous membrane, then with human vascular smooth muscle cells that had been cultured with ascorbic acid. A sheet of human fibroblasts was wrapped around this ‘media’, followed by endothelial cells, to make a three-layered co-culture. The cells produced extracellular matrix proteins (collagen and elastin), desmin, von Willebrand factor and prostacyclin appropriately. The endothelium was anti-thrombotic and inhibited platelet adhesion. A burst strength of 2000 mmHg was reported, comparable to that of human vessels. The engineered tissues took 3 months to grow in culture. In a lamb model,10 it has been shown that tissue engineered heart valves synthesised using cells from femoral arteries developed into acceptable leaflets.
Bioengineered Tissue of Cells on a Scaffold A compound approach is to use a bioengineered tissue equivalent seeded with cells that would be found in the native environment: endothelium, smooth muscle cells, myofibroblasts and fibroblasts. A considerable contribution expected from these cells is that they should selfassemble an extracellular matrix of collagen and elastin within the framework provided. The first attempt to construct such a cells-plus-scaffold blood vessel was made more than 15 years ago.11 Bovine smooth muscle cells in a collagen culture medium were cast around an annular mould. When this culture was mature, a Dacron® (non-absorbable plastic) mesh sleeve (Du Pont, Wilmington, DE, USA) was rolled over for mechanical support, followed by seeding with adventitial fibroblasts.
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The constructs required 2 weeks to become fully contracted and then samples were lined with endothelial cells. Electron microscopy showed that the smooth muscle cells secreted collagen in the manner of fibroblasts. This system, however, had low burst pressure. Elastin had not been incorporated in the matrix, and models made with only one Dacron® mesh ruptured at less than 10 mmHg. Although strength to withstand up to 180 mmHg (still considerably lower than that required for a human blood vessel implant) could be achieved, it was only with the use of numerous Dacron® meshes, the whole becoming too stiff to be effective. As the L’Heureux9 group found in their wholly biological neovessel more than 10 years later, the proliferated endothelial layer successfully produced von Willebrand factor and inhibited platelet aggregation. In subsequent years, advances were made in the principle of using native cells with a variety of fabrics. Hydrolytic degradable constructs,12–14 absorbable patches,15 biologically modified native tissues,7 collagen sponge matrices,16 and third generation biomedical materials were tested. Engineering development of biomaterials for scaffolds, particularly techniques appropriate for mass production, are essential in the worldwide market. After improvement of scaffolds and co-cultures in the laboratory, to aid patients, these innovations must have practical reliability. Constituents must always be readily available, fully functional and reproducible. Pragmatically, products must have a sufficient shelf-life and be able to withstand shipping and storage.17 In the cardiovascular system, mechanical properties of the matrices themselves are crucial. They must be capable of withstanding complex mechanical and dynamic environments, and honeycomb-like polycaprolactone structures have been compared for compressive stiffness and offset yield strength.18 In the case of engineered heart valves, precise shape of the graft is important. Stereolithic models derived from X-ray computed tomography with appropriate software have been used to generate biocompatible grafts.19
Hydrolytic Degradable Scaffolds The Dacron® mesh used by Weinberg and Bell11 was a physical support for the neovessel only, with no pretension toward contributing either chemically or physiologically to tissue synthesis. But in 1998, sheets of polyglycolic acid (PGA) mesh were used as a substrate for seeding human cells.12 Such a biodegradable framework gradually deteriorates, provokes little or no immune response in the host and minimises any potential chronic inflammatory reaction. The PGA–cell complex allows native cells to grow, remodel the graft and contribute to repair.
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Polyglycolic acid has been used as a substrate in an ovine model of autologous heart valve leaflet replacement.13 Fibroblasts, then endothelial cells, were cultured onto the polymer. The synthetic complexes were grafted into lambs, with acellular PGA implanted into control hosts. In the controls, grafts degraded but all cell-seeded categories integrated well. Also in the ovine model, a heart valve equivalent was manufactured with nonwoven PGA plus a thin layer of poly-4-hydroxybutrate.14 Here, the constructs were cultured for 2 weeks in vitro in a pulse duplicator under gradually increasing flow and pressures, before being resected back into lambs. Induced biomimetic flow improved valve function.
Absorbable Patches Absorbable patches of poly beta-hydroxybutrate have been used in calves to close the pericardium after cardiopulmonary bypass. At investigatory re-operation, patch material contained pronounced macrophage infiltration, but no regenerative mesothelium.15 Adhesions in the healing pericardium might be caused by secretions of mesothelial cells.
Biologically Modified Native Tissues To engineer a narrow diameter artery in the rabbit, a porcine cell-free collagen layer was used,7 derived from the submucosa of the small intestine. To reduce thrombogenicity, luminal surfaces of grafts were treated with heparin-bezalkonium chloride cross-linked with 1-ethyl3-(3-dimethylaminopropyl) carbodiimide hydrochloride. Even so, without anticoagulant therapy, vessels were unstable. The technique was modified by deposition of bovine type I collagen on the lumen, which improved patency. The cell-free grafts were intact up to 90 days post surgery, and although they consisted of porcine and bovine collagen components, immunogenicity was minimal. There was considerable rabbit cell infiltration into the graft. Implants were reactive to experimental vascular agonists, had a non-thrombogenic surface, and had a burst resistance of up to 1000 mmHg. For heart valves, a technique of implanting a complex composed of combined fresh non-vascular tissue and cells was pioneered by Manothaya et al.20 Dura mater was used as the tissue source for frame-supported valves, and short-term results were reported as being satisfactory.
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collagen and 4–9% type III collagen). They are biocompatible with host tissue and are biodegradable. To prepare the sponges, solutions of extracted collagen are freeze-dried, resulting in a porous matrix with a mean pore size of 58 ± 24 µm. In our laboratory, similar calf skin-derived sponges have been seeded with porcine aortic wall myofibroblasts.21 Cells were infiltrative and proliferative, and porosity of the sponge provided excellent conduits for cell migration. Surface myofibroblasts were densely packed, lying in stratified layers, four to six cells deep, their long axes distinctively at right angles to those of the infiltrative population. Mouse fibroblasts have been found to proliferate to a constant cell density independent of seeding level,22 but here, developmental pattern was related to seeding number. Surface myofibroblasts generated sparse amounts of matrical proteins, but central sponge zones were replete with dense aggregates of banded collagen and elastic fibres, emanating from the well-spaced cells (Figs 1,2). This difference could be because nutrient and gas transfer may be restricted at distance from the liquid culture solutions, but the change in phenotype is possibly myofibroblast differentiation. This is completely in accord with normal myofibroblast behaviour in a native connective tissue, where cells are separated by a multiplicity of extracellular components, and myofibroblasts readily change direction of metabolic activity.23,24 Transforming growth factor β, for example, stimulates fibroblast culture behaviour, which is related to applied mechanical loading.25
Collagen Sponge Matrices Collagen sponges cross-linked or non-cross-linked with diphenylphosphorylazide seeded with bovine foetal chrondrocytes have been used to reconstruct cartilaginous tissues.16 Sponges are manufactured from collagen extracted from calf skin (90–95% dry weight type I
Figure 1. High-power transmission electron micrograph of the plasma membrane of a porcine myofibroblast grown in a collagen sponge. Collagen fibres (arrows) are clearly seen emerging from the cell into the extracellular matrix. Length of cell culture time: 4 weeks. Bar, 0.6 µm.
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Third Generation Biomedical Materials First generation biomedical materials for implantation have been defined as being ‘inert’, and those of the second generation as having controlled chemical selfbreakdown and resorption.26 Extrapolation of properties of these materials lies in development of tissue substitutes that are auxiliary to, and able to integrate with, the host body. Stimulatory proteins can be bound to polymer surfaces; adapted polymers have been produced that influence interactions with endothelial cells.27 Cellular growth factors that stimulate proliferation and development of cardiovascular cells can be chemically combined with structural matrices. Angiogenic patches of human dermal fibroblasts have been successfully used to cover infarcted cardiac tissue in immunodeficient mice.28 Exciting prospects lie in the possibility to manufacture novel, non-immunogenic bioengineered scaffolds incorporating molecules that stimulate appropriate genes in neighbouring host body cells. Subcellular localisation of gene products can be demonstrated by techniques such as electron microscopy, immunocytochemistry and in situ hybridisation.29 Supplementation of myofibroblast cultures with ascorbic acid for heart valve tissue engineering has already been used to increase collagen production.30 Native proteins can be produced to augment cell production and differentiation, repair the defect, heal the wound, and control inflammatory complications. Growth factor-stimulated tissue growth in arteries and valves, however, is an issue with which one must be concerned. In wound healing, there is limited and revers-
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ible hypertrophy. Excessive stimulation of cell proliferation could all too easily result in uncontrolled neoplasia. A list of some materials used in bioengineered scaffolds can be found in Table 1.
Mechanical Forces and Bioreactors In vivo, the architecture of heart and blood vessel connective tissue is not haphazard. The mesenchymal cells are sensitive to forces applied to them from their environment and integrate both mechanical and biochemical signals for growth, maintenance and repair. In turn, development and function of the endothelium requires intimate intertissue communication. First, pulsatile fluid flow gives rise to a radial (and longitudinal) strain; second, steady state flow (with no pulsatile component) causes shear stress on luminal wall surfaces, which results in endothelial cells aligning and bearing in the direction pertinent to the shear; and third, axial loading causes pure longitudinal strain to develop. Bioreactors have been used in the field of cardiovascular tissue engineering since the prototype blood vessel was constructed.11 Their role initially was as a culture vessel, to disperse cell density evenly by rotation and to provide a sterile, nutrient rich environment in which the preformed bioengineered culture could develop. Recently, however, mechanical forces derived by the cells have been measured,31–34 and externally driven forces have been applied to cultures, resulting in streamlined tissues with increased strength and patency.
Table 1. Examples of some materials used in bioengineered scaffolds
Figure 2. Transmission electron micrograph of a porcine myofibroblast grown in a collagen sponge, associated with elastic fibres (arrows), synthesised during the period of cell culture. Length of culture time: 4 weeks. Bar, 1.1 µm.
Synthetic polymers used as scaffolds in tissue engineering15,16 Poly beta-hydroxybutrate Polyglycolic acid Polylactic acid Polyglycolic acid-poly L-lactic acid co-polymer Polyethylene Polyethylene oxide Carrier materials of natural origin used as scaffolds in tissue engineering16 Collagen sponges Collagen gel Alginate Fibrin Hyaluronic acid Gelatin Collagen–glycosaminoglycan matrices
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Culture Force Monitor and Tensioning Culture Force Monitor Extensile cells create forces and also respond to stimuli applied to them from their surroundings. Such forces have been accurately measured on an apparatus called a culture force monitor (CFM) and have been provoked and measured by a tensioning-culture force monitor (t-CFM).31 Culture force monitor. A CFM was developed in 199431 in our laboratory. It can be used to study 3-D mechanical forces in different cell types in a variety of scaffolds. In addition, the CFM can be used to record cell responses to therapeutic drugs and/or growth factors. The CFM is an improvement on previous similar devices because its force transducer is coupled directly to a computer via an analogue-to-digital converter (Fig. 3). Movement of fibroblasts suspended in a collagen matrix has been measured to 10–6 m. The average force produced by a single fibroblast was 10–10 Newtons. In a period of contraction, the majority of force has been shown to be generated in the first 8 h, and morphological studies illustrate that changes in cell shape (extension of cytoplasmic processes, development of attachments, cytoskeletal rearrangement) initiate the increments.35,36 In recessive dystrophic epidermolysis bullosa (RDEB), in untethered fibroblast-populated collagen lattices, forces of normal fibroblasts and RDEB fibroblasts were quantified and compared. RDEB fibroblasts were hypercontractile, generating 2.5-fold the force of normal fibroblasts.37 The effect of burn blister fluid on fibroblast contraction showed that burn fluid stimulated fibroblasts to contract, predominantly in a period of 24–48 h after application,38 although chronic wound fluid inhibited cell proliferation.39 An adaptation of the CFM to measure forces generated by tissue cohesion after injury, an adhesion
Figure 3. Structure of the culture force monitor (CFM), indicating relationships of the force transducer and the culture well.
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CFM, has demonstrated levels of fibroblast-mediated adhesion in vitro between contrasting tissue interfaces.40 Formation of post surgical adhesions are also important in human primary and re-operative pericardium.41 Tensioning culture force monitor. The t-CFM is a derivative of the CFM and allows the application of precise mechanical loads of physiological dimension to be applied to fibroblast populated lattices.31 For the t-CFM (Fig. 4), a micro-stepping motor was added to the monitoring devices of the CFM, and a theoretical positioning accuracy of 10–8 m has been achieved. Fibroblast populated collagen matrices can now be loaded, experimental forces of predetermined magnitude engaged, and cellular responses in terms of both matrix tension and biochemical constitution measured. The t-CFM applies minute, unidirectional loads to cell–collagen complexes with specific loading patterns of striking and repeatable cell alignment reliably predicted and produced.31 Fibroblasts rank themselves along lines of maximum principal isostrain32 and orientate in consistent responsive patterns (Figs 5,6). Investigating homeostasis in dermal fibroblasts, the t-CFM has measured and quantified mechanical responses to precisely defined tensional loads.33 Fibroblasts actively migrated across restrained gels more energetically than across untethered gels.42 When cyclical loads were applied to produce unidirectional strain, cells which had been spherical at seeding were induced into characteristic bipolar-shaped fibroblasts aligned parallel to the loading axis of the neotissue. Dependence of cell alignment on strain contours implied that it was not only magnitude and direction of loading that dictated alignment, but also physical shape of the material subject to the load.
Figure 4. Structure of the tensioning culture force monitor (t-CFM), showing addition of a micro-stepping motor system to the basic CFM design.
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Figure 5. Scanning electron micrograph of freeze-fractured gel with non-aligned fibres. Reproduced with permission from Eastwood et al. 1998.32 Bar, 10 µm.
Biochemical changes took place in fibroblasts subjected to mechanical loading, and chemical properties were modulated by changes in applied stress. Effects of mechanical stimulation on production of matrix metalloproteases (MMP) were investigated. MMP-9 synthesis was more sensitive to tension than MMP-2 for both induction and activation of enzymes, and also, stromelysin activity decreased with tension.43 Quantitative reverse transcription (RT)–polymerase chain reaction (PCR) revealed twofold, fourfold and threefold increases in MMP-1, MMP-2 and MMP-3, respectively, in nonaligned strain zones relative to aligned strain zones during contraction in wound healing.44
Pulsatile Bioreactors Prototype pulsatile bioreactors have been developed by the groups of Hoerstrup et al.,14 Zeltinger et al.45 and Eastwood et al.21 In general, they consist of the bioreaction chamber, a culture media exchange mechanism, a means of generating pulsatile flow (micro-stepping motors), force transducers and computerised control of all reactions and forces applied. Multiple loading regimes are available to condition the cell–scaffold complex. For vascular specimens, a tubular culture vessel is used. This type of tissue engineered construct has either to have been cultured on an annular mould, or a planar matrix has been cell seeded, rolled, sutured and finally the edges bonded together with an effective fibrin sealant. Pulsatile flow causes cyclic fluid shear stress on the inner wall of the neovessel in addition to generating
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Figure 6. Scanning electron micrograph of freeze-fractured gel with aligned fibres (large arrow) emanating from a fibroblast (small arrow). Reproduced with permission from Eastwood et al. 1998.32 Bar, 10 µm.
radial and longitudinal stresses within the wall. These stresses and associated strains have the effect of orientating all cells within and appropriately aligning any fresh collagen and/or elastin that they might synthesise.
Multi-Cue Bioreactor A development of the original pulsatile bioreactor is the multi-cue bioreactor (M-CB)21 from which a range of types of stimulus (cue) can be delivered via a hermetically sealed electric motor linked to a rack and pinion straining system and a peristaltic pump head. The strain unit has a mathematical resolution of 2 × 10–6 m and total displacement of 70 mm. Pressures in the range of 450 mmHg can be generated. Strain induction and pressure gradients can also be measured across a structure’s wall.33,34,43,46 In addition to inducing conditioning strain, at any time during a procedure, in situ non-destructive tensile testing can be performed at key intervals. The M-CB computer controlled stepper motor units are programmed through X-Ware® software (Parker Hannifin, Poole, Dorset, UK) whose parameters are able to direct fluid velocity and acceleration, as well as axis of rotation. Further, mechanical cue drivers can be used in conjunction with a transducer array to perform feedback conditioning, where programs are adjusted in response to force/pressure readings, as occurs in vivo. Versatility and adjustability of the M-CB provide the flexible range of variables needed to condition a simple cell-seeded matrix from a mechanically unstructured mass into a viable, implantable and biologically sufficient neotissue.
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Cell Origin and Graft Success
Conclusion
To manufacture an effective, durable, tissue engineered composite graft for human implantation, the biological source of the seeded cells is of the highest importance. To avoid immune system complications, cells are ideally autologous (from the patient him/herself), but for practical purposes (such as time required to grow a sufficient culture), the nearest immunological match, allogenic cells, might have to be resorted to. The use of autologous cells has many advantages (not least that ethical considerations are not incurred), and tissue engineering had its first successes more than 20 years ago,47 when patients’ own skin cells were first used to cover wounds. This approach might be technically taxing for cardiovascular grafting at present. Allogenic cardiovascular cells have been used successfully in animal studies of vessel and valve implants13,14 but, once more, ethical constraints prevent similar work in humans. A source of dermal fibroblasts is available from human foreskin.48 Such cells have been shown to be acceptable immunologically, but dermal fibroblasts have proven to be inferior to cardiovascular fibroblasts in potential heart prostheses.10 Similarly, human umbilical cord endothelium is a source of suitable human cells,49 but neonatal myofibroblasts and embryonic fibroblast stem cells for allogenic surgical purposes, although useful animal experimental or veterinary tools, cannot be considered for the purpose of making human grafts. Populating a bioengineered tissue scaffold with adult human stem cells (autologous or allogenic) might be possible. The potential for stimulating maturation of adult stem cell progeny along predetermined lines has been widely demonstrated.50 In the mouse, myofibroblasts in colonic lamina propria have been derived from bone marrow stem cells, transplanted between adults,51 so the possibility of producing vascular myofibroblasts is raised. Harvesting and transplanting human bone marrow are common techniques, and selectively directing differentiation outcome in an engrafted, bioengineered scaffold could solve problems of effectiveness and durability of synthetic connective tissue surrogates. Bioengineered scaffolds with or without seeded cells, but porous enough to permit rapid infiltration of local native cells into the graft would be ideal. Huynh et al.7 found their porcine–bovine collagen complex immunologically acceptable in rabbit hosts, and the implant rapidly became colonised by rabbit cells. Fibroblasts, myofibroblasts, and endothelial cells have the genetic ability to express proteins controlling cell migration and cell proliferation: these activities are among their normal repertoire in wound healing.
Consummate tissue engineered cardiovascular replacements must possess all of the properties of the natural tissue. Technically, there are numerous and consistent advances in the fields of matrix commercial availability, linking scientific discovery with practical medical application. Scaffolds, whether synthetic or biological, must satisfy mechanical demands in addition to being non-immunogenic and non-thrombogenic. Mechanical stresses and strains on blood vessel and heart valve cells have been recognised and measured, and can be reproduced to condition tissue constructs in vitro. Seeded and cultured cells must be appropriate and effective for the purpose. In the case of fibroblasts or myofibroblasts, their innate ability to produce requisite extracellular matrix proteins themselves is an invaluable aid for the tissue substitute. The possibility of incorporating cellular growth factors into a potential implant is attractive, although this should be pursued with caution. Vascular endothelial growth factor, the fibroblast growth factors, plasma proteins and cytokines all stimulate vascular cell growth, and synthesis of collagen and elastin. When a chosen scaffold slowly disintegrates in a controlled fashion after implantation, while being replaced by totally sufficient, natural structural cells and fibres, the outcome will be a replaced tissue rather than a replacement tissue.
Acknowledgements We would like to thank the Engineering and Physical Science Research Council, the Phoenix Trust, and Pearl Assurance for financial support for this work.
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