TY - JOUR
AU - Homer-Vanniasinkam, S
AB - Abstract Background Autologous conduits are not available in up to 40 per cent of patients with arteriopathy who require coronary or lower limb revascularization, and access sites for renal dialysis may eventually become exhausted. Synthetic prostheses achieve a poor patency rate in small-calibre anastomoses. This review examines how vascular tissue engineering may be used to address these issues. Methods A Medline search was performed, using the keywords ‘vascular tissue engineering’, ‘small diameter vascular conduit’, ‘vascular cell biology’, ‘biomechanics’, ‘cell seeding’ and ‘graft endothelialization’. Key references were hand-searched for relevant papers. Results and conclusion In vitro and in vivo approaches are currently being used for guided cell repopulation of both biological and synthetic scaffolds. The major clinical problem has been extended culture time (approximately 6 weeks), which precludes their use in the acute setting. However, recent advances have led not only to improved patency rates for prostheses, but also to a potential reduction in culture time. In addition, increased mobilization of endothelial progenitor cells in the presence of ischaemic tissue may increase the autologous cell yield for scaffold reseeding with further reduction in culture time. Introduction Vascular tissue engineering has been described as ‘an interdisciplinary field that applies the principles and methods of engineering and the life sciences towards the development of biological substitutes that restore, maintain and improve tissue function’1. It aims to develop biocompatible scaffolds that mimic the mechanical properties of autogenous conduits, while providing a framework for guided cell repopulation to enable integration as a functional cardiovascular conduit. Its recent advances developed from an appreciation of the need for viable alternatives to autologous cardiovascular conduits. Despite recent advances in secondary prevention and interventional cardiology2–4, coronary artery bypass remains an important therapeutic option, and the incidence of peripheral arterial disease and end-stage renal failure requiring vascular access is increasing5,6. Autologous vein is not available in 10–40 per cent of patients as a result of trauma, disease such as varicose veins, or previous surgery7. The problem is further compounded by the chronicity of these disease states, resulting in a cohort of patients who will inevitably require reoperation. Synthetic alternatives to vein have poor intermediate and long-term patency rates in vessels with a calibre of less than 6 mm8. This review describes the various approaches in vascular tissue engineering, the different scaffolds available, the mechanisms of graft failure and the ongoing matters of graft design, with particular attention to small-calibre grafts. It focuses on the integration of vascular cell biology into graft design, including graft endothelialization and endothelial progenitor cell research. Methods An OVID Medline search was performed under the headings ‘vascular tissue engineering’, ‘small calibre/diameter vascular substitute’, ‘vascular cell biology’, ‘biomechanics’, ‘graft endothelialization’ and ‘cell seeding’ from the years 1966 to 2005. Bibliographies of landmark papers were hand-searched for further relevant articles. Tissue engineering approaches General approach Vascular tissue engineering aims to obviate the twin obstacles of thrombogenicity and immunogenicity by using autologous cells to repopulate biocompatible scaffolds. These cells include endothelial cells, vascular smooth muscle cells (SMCs) and fibroblasts. Remodelling of the graft wall in response to haemodynamic stresses either in vivo or in vitro would then occur at the outset, strengthening the extracellular matrix (ECM) and graft wall. Hence, the development of a viable tissue-engineered blood vessel depends on a suitable scaffold for cellular attachment and reliable methods of repopulating the scaffold. Tissue engineering scaffolds Vascular tissue engineering scaffolds are tubular conduits or materials capable of being fashioned into blood vessel configurations that are neither cytotoxic nor thrombogenic. Lack of immunogenicity is facilitated by the acellularity of the scaffold. The scaffold functions to bring the various cells of the vasculature in close proximity to one another to enable intercellular signalling, which in turn facilitates cell adhesion, migration and differentiation9. In the in vivo approach, discussed below, the scaffold should also provide the initial requisite mechanical strength to withstand in vivo haemodynamic forces until vascular SMCs and fibroblasts reinforce the ECM of the vessel wall. Scaffolds can be either synthetic polymers or biological matrices. Most synthetic polymers involve materials that are already in widespread use in surgery, such as in sutures. These include polyglycolic acid (PGA), poly-l-lactic acid, poly-d,l-lactic-co-glycolic acid and polyurethanes9–11. The first three of these materials are biodegradable, with the aim of in vivo vascular cell remodelling to enable the final functional conduit to perform the structural role as well. The degradation must be understood and allow remodelling to confer adequate strength to the conduit if aneurysmal degeneration and rupture are to be avoided. Synthetic scaffolds have the advantages of industrial-scale replication, and biochemical and mechanical qualities determined at the point of manufacture. These include degradation products, degradation time, compliance, tensile strength, conduit size and configuration10,12. An illustration of this point is the modification of polyurethane scaffolds, with the removal of an ether linkage to make it more resistant to in vivo degradation, as a result of concerns about premature degeneration. Its honeycomb middle layer also allowed a degree of compliance matching with that of artery12. The disadvantage of synthetic scaffolds is the lack of the cell attachment factors present in naturally derived scaffolds9. There is also concern that scaffold degradation remnants may cause intimal hyperplasia. Higgins et al.13 studied the effects of PGA remnants on vascular SMC cell behaviour and found that some breakdown products produced a change in SMCs from contractile to synthetic phenotype. This dedifferentiation effect was noted in intimal hyperplasia with synthetic SMCs laying down increased ECM components, resulting in intermediate to late graft occlusion87. Biological scaffolds can be divided into three broad categories: decellularized allogeneic and xenogeneic blood vessels14–18, decellularized non-vascular conduits19–24 and prefabricated ECMs25,26. The last includes collagen gels or matrices generated in vitro by seeded fibroblasts in a bioreactor. These scaffolds are then further seeded with SMCs and endothelial cells with dynamic conditioning in a bioreactor to produce a functional blood vessel. Native cells are removed by chemical or mechanical means. Chemical methods aim to produce native cell lysis, detergent removal of cell membrane lipids and removal of nuclear debris, which can stimulate calcification and graft degeneration. They involve a combination of enzymatic and detergent treatments, including trypsin–EDTA, hypotonic Tris buffer, sodium dodecyl sulphate and nucleases14,15,20,27. Mechanical methods involve eversion and surface abrasion of harvested tissue to produce an acellular layer composed primarily of type 1 collagen19,24. These are used primarily in the production of xenogenic small intestinal submucosa scaffolds. Decellularized biological scaffolds have the advantage of possessing an ECM rich in the cell signalling components essential for cell adhesion, migration, proliferation and differentiation. Preserved ECM components include collagen, elastin and glycosaminoglycans. Intact collagen and elastin fibres ensure retention of tensile strength and elastic recoil properties, and there is evidence to suggest that they inhibit SMC proliferation, an early step in intimal hyperplasia28,29. Glycosaminoglycans, which include chondroitin, dermatan, heparan and heparin sulphate, play a crucial role in endothelial cell adhesion and proliferation, inhibition of SMC proliferation and migration following injury or inflammation, in addition to their antithrombotic properties30,31. Small intestinal submucosa scaffolds also retain angiogenic growth factors, such as basic fibroblast growth factor and vascular endothelial growth factor32. Evidence for the effects of biological matrices on cell behaviour is seen in the work of Gillis et al.33, who found that the production of important endothelial mediators involved in coagulation homeostasis (prostacyclin, tissue plasminogen activator and plasminogen activator inhibitor 1) was significantly greater from human saphenous vein endothelial cells grown on biological matrices, such as collagen type 1 gel, fibrin glue and de-endothelialized porcine aorta, than in those grown on polytetrafluoroethylene (PTFE). Although there was no direct comparison with synthetic scaffolds, it is notable that Niklason et al.34, using a PGA scaffold, found endothelialization to fail in pockets of the graft surface, with persisting PGA remnants. This may have been due to cytotoxicity of PGA breakdown products or lack of cell signalling support in areas of persisting PGA. Biological vascular substitutes also display a greater resistance to infection compared with synthetic grafts. Jernigan et al.35 compared PTFE and small intestinal submucosa iliac patch angioplasties in a porcine model contaminated with faecal soiling. They found patch infection in 73 per cent and pseudoaneurysm formation in 25 per cent in PTFE patches, compared with 8 and 0 per cent respectively in submucosal patches. Lantz et al.19, on reviewing the work done on submucosa scaffolds as vascular conduits, suggested that earlier remodelling and development of vasa vasora in biological scaffolds enable a more rapid and effective recruitment of circulating immune system components, thereby preventing graft infection from taking hold. The early formation of a confluent endothelium also prevents formation of thrombi, which might promote bacterial adherence and the establishment of infection36. Concerns over the use of xenogenic scaffolds include transmission of animal retroviruses and cross-species infection. Porcine endogenous retrovirus (PERV) is widespread and not known to cause disease in pigs, but its effects in humans are uncertain. This concern should not be taken lightly as human immunodeficiency virus 1, which was not thought to produce any pathological state in monkeys, went on to cause serious disease in humans. Although decellularization reduces the risk of retroviral contamination, recent studies have shown that up to 2 per cent of donor DNA remains detectable on some scaffolds after decellularization37. PERV can infect human endothelial cells, vascular fibroblasts and mesangial cells in vitro38, and Paradis et al.39 found no evidence of PERV cross-infection in human patients who had received porcine tissue implants up to 12 years previously. This is supported by work in animal models37,40. Recellularization of scaffolds In vitro methods In this approach autologous cells are expanded in culture and then seeded on to the scaffold in vitro. The first completely tissue-engineered blood vessel was developed by Weinberg and Bell25 by culturing bovine endothelial cells, vascular SMCs and fibroblasts on a collagen gel scaffold that had been prefabricated by culturing fibroblasts on a mandril. The progressive addition of cellular constituents under dynamic culture conditions produced a small-calibre vascular conduit. However, even with reinforcement with Dacron mesh, it did not have the requisite burst strength, comparable to that of autologous conduits. The importance of this development, however, was its demonstration that dissociated cells could be recruited and cultured into a functioning tissue (endothelial cells established on the conduit stained positively for von Willebrand factor and prostacyclin). Further refinement by other groups involved the addition of ascorbic acid to the culture medium, improving burst pressure up to 3000 mmHg; human long saphenous vein has a burst pressure of approximately 2000 mmHg. Contractility in response to vasoactive agents was also demonstrated24,26. In the in vitro approach, the cultivation of the appropriate vascular SMC phenotype is important. Vascular SMCs, when harvested and grown in static culture, dedifferentiate and proliferate rapidly, acquiring ultrastructural features (large Golgi apparatus and widespread rough endoplasmic reticulum) that enable them to perform a synthetic role, namely the generation of ECM proteins. Contractile SMCs, in contrast, have a cytoplasm rich in myofilaments. This reversible change from the contractile (present in the healthy tunica media of adult arteries) to the synthetic phenotype is seen in atherosclerotic and intimal hyperplasia lesions41. Clearly, for an engineered blood vessel, a predominance of the contractile SMC is desirable. One of the important determinants of the reversion to a contractile phenotype is the exposure to shear stress and pulsatile flow41. This intuitive requirement for a pulsatile flow environment is borne out by several studies indicating that mechanical preconditioning is essential for endothelial cell adhesion and retention42–44, correct alignment of SMCs, maintenance of the contractile SMC phenotype45 and reinforcement of the vessel ECM46. This is an important design requirement for any bioreactor system in which tissue-engineered blood vessels will be developed. The in vitro approach has inevitably raised the issue of a suitable bioreactor system within which scaffolds can be cultivated with seeded vascular cells for extended periods under sterile conditions. Such a system was first developed successfully by Niklason et al.34. Vascular tissue engineering requires an enclosed, sterile system that enables perfusion of the scaffold with cell culture medium under pulsatile flow conditions. Successful systems are usually modular, with a range of port sites to enable change of medium and components, cell seeding, and mounting of scaffolds to occur with minimal disturbance to other parts of the system and minimal risk of microbial contamination47,48. There is also the logistical issue of extended culture time. The minimum time from harvest of autologous cells to culture of a mature and durable conduit is approximately 8 weeks34. Clearly, this approach limits the conduit to elective procedures and requires local expertise and bioreactor facilities. In vivo methods This approach has been widely investigated with decellularized biological scaffolds, such as small intestinal submucosa17, 19–23. It involves the direct implantation of the tissue engineering scaffold into the patient as a graft, with the onus on remodelling of the conduit in vivo. An important prerequisite is for the scaffold to meet most, if not all, of the features of an ideal vascular substitute before implantation. The scaffold must be biocompatible, with minimal thromboreactivity, which is the greatest immediate concern. It should also have the requisite biomechanical strength and properties to enable it to function adequately in vivo as an arterial substitute until remodelling takes place49. An intriguing in vivo example of a completely tissue-engineered blood vessel was reported by Campbell et al.50, in which four segments of Silastic tubing were inserted into the peritoneal cavity of a rabbit to generate an autologous conduit from the host inflammatory response. After 2 weeks, the tubing segments were removed and everted. An average of 50 per cent of the segments yielded viable trilayered conduits composed of a lumen confluent with mesothelium, a ‘media’ composed of ECM including collagen and elastin with contractile SMCs, and an ‘adventitia’ with obvious vasa vasora. The conduits were then implanted into the right carotid artery as an autologous graft, with 67 per cent remaining patent at 4 months. No perioperative heparin or antiplatelet agents were given. Although the study had only a short follow-up, did not report on complications such as aneurysm formation or rupture, or attempt any mechanical characterization, it illustrates another feature of tissue engineering, namely that it is the cell environment rather than cell source that determines eventual function. The mesothelium of the graft displayed immunohistochemical features of endothelium and the media responded to contractile agonists on explant. Thromboresistance and endothelialization The endothelium is not a smooth inert surface that facilitates laminar blood flow through the blood vessel, but a dynamic organ with an active role in coagulation homeostasis, the sensing and transduction of the haemodynamic forces of circulation, and the cellular metabolism of the vascular wall51. Indeed, it has been shown that endothelial cell loss from vascular injury plays a significant role in the local activation of the pathophysiological cascade leading to the development of intimal hyperplasia52,53. The need for a confluent endothelium on the graft luminal surface has recently been highlighted by the results of PTFE graft endothelialization studies54 that showed similar patency rates for endothelialized PTFE and autologous saphenous vein in infrainguinal bypasses55. These results suggest that surface endothelialization has a much greater role in intermediate and long-term patency than was previously thought54. It was widely considered that biomechanical and haemodynamic factors were largely responsible for intimal hyperplasia and graft occlusion, but, increasingly, there is evidence that a confluent endothelium is crucial in prevention of the initiation and progression of the process. One aspect of endothelial protection is the physical barrier it forms to prevent contact with subendothelial components of the arterial wall and activation of the coagulation cascade56. In addition, early events in the cascade, such as platelet degranulation following contact with type I collagen, have been shown to induce mitogenic factors such as transforming growth factor β41. Animal models in which endothelial injury was induced show that loss of an intact endothelium results in a change of SMC phenotype to a synthetic state57. Conversely, such a change was inhibited by the presence of endothelial cells57,58. Intriguingly, Nugent et al.58 showed that the modulating effect of endothelial cells can occur even without anatomical proximity with SMCs. Perivascular implants of endothelial cells cultured on Gelfoam® (Pharmacia & Upjohn, Peapack, New Jersey, USA) matrices reduce the formation of neointimal lesions of angioplasty-induced injury of porcine carotid arteries in vitro58. Immediate histological examination of vein grafts of patients who died within 10 days of coronary artery bypass shows that endothelial denudation is soon followed by a shift in phenotype from contractile to synthetic, with increasing deposition of such cells in the intimal region as early as 7 days after graft implantation59. It seems that, if tissue-engineered blood vessels are to be a viable and durable vascular conduit in vivo, a confluent endothelium is an important prerequisite. Previous non-endothelialized implants of decellularized biological conduits in animal models produced moderately acceptable patency rates when placed in sites of either high flow or low resistance, or both15,19,20,22,23. In one animal study21, small intestinal submucosa was implanted as a small-diameter vascular graft in high-flow anastomoses involving carotid and femoral arteries; all animals received warfarin and aspirin for up to 8 weeks after surgery. The overall patency rate was 75 per cent (48 weeks), the graft rupture rate 5 per cent (within 21 d), and aneurysmal dilatation was found in 11 per cent of grafts. The success of Meinhart and Zilla's graft endothelialization programme54 was achieved through improved techniques, including the concept of ‘cell sodding’ (use of high-density seeding and incubation) and a two-stage process in which harvested endothelial cells are first expanded in culture before seeding on to the graft. These advances must be viewed cautiously. Although the improved cell culture protocols reduced the cell harvest to implant time to a mean of 29 days (a reduction of 7 days compared with an earlier study60), this would still preclude the use of seeded grafts in the acute setting. This cultivation time was needed to coat a 700-mm long graft with an internal diameter of 6 mm. In addition, unless the technology becomes more widespread and cheaper, development of an on-site cell culture and seeding facility will be beyond the resources of most hospitals, even in affluent countries. As an alternative, work is under way to refine single-stage seeding protocols in which cells are harvested and seeded on to the graft within the duration of the bypass procedure. Clearly, such a process would require protocols that produce a better cell yield per unit of vein harvested. Currently, cell yields depend on the vein segments obtained, with the most commonly used sites being the external jugular and cephalic veins54,61. The saphenous vein is used only when it is deemed unsuitable as bypass conduit. Other sources include omentum62,63 and subcutaneous fat64,65. All of these options require an additional painful surgical procedure for harvest. Omental cells can feasibly be obtained at laparoscopy, and subcutaneous fat through liposuction. One area of concern is the cell yield and quality from diabetics and smokers. On average there is a 5–27 per cent risk that seeded endothelial cells will not grow. The number of cells extracted from smokers is significantly lower than from non-smokers, and they require prolonged culture times60; in diabetics, cell retention on exposure to pulsatile flow is inferior to that in non-diabetic patients66. Another possible alternative to avoid prolonged culture times is to treat scaffolds with cell adhesion factors or peptides, to enhance endothelial cell adhesion and proliferation in vivo. One reason for the disappointing early results with low-density seeding was that cell attrition on exposure to pulsatile flow was high. Cell loss usually occurred within the first 45 min67, with retention rates as low as 4 per cent68. Treatments to increase endothelial cell retention employ components of the healthy arterial ECM, such as fibronectin, laminin, growth factors69 and adhesion peptides like the Arg-Gly-Asp sequence70. Some studies suggest that complex interactions exist within the ECM to regulate optimal cell adhesion and retention, vascular SMC migration and ECM protein production71,72. Biomechanical considerations: compliance match The concept of compliance mismatch as a cause of intimal hyperplasia and graft failure is controversial. The similar patency rates of endothelialized PTFE and autologous vein favour the view that graft thrombogenicity rather than compliance mismatch is the primary cause of graft failure. However, freshly harvested saphenous vein is fairly incompliant at arterial pressures12. Given that the initiation of the intimal hyperplasia cascade begins within 24 h of vascular injury (surgical anastomosis being one example), compliance mismatch may still have a significant role, if not in initiating the process then in maintaining events along the cascade, including vascular SMC modulation73. Compliance mismatch may cause excessive stretching of SMCs, resulting in their proliferation. Further evidence for this is the finding in humans that preoperative ultrasonographic measurement of human saphenous vein compliance is a predictor of femorodistal bypass stenosis74. The only limitation of this study was that it did not measure host artery compliance to assess the degree of mismatch between vein graft and artery. Other studies have shown conflicting results regarding the impact of compliance mismatch on distal anastomotic intimal hyperplasia. Okuhn et al.75, using an end-to-end anastomotic configuration, showed that compliance mismatch had no effect on its development. However, it is the end-to-side arrangement that is used most often in clinical situations and, in such a configuration, compliance mismatch has been shown to be associated with intimal hyperplasia76,77. A recent in vitro haemodynamic study by Ballyk et al.78 offers an explanation for these differing results in terms of altered suture-line stresses. The detrimental effect of compliance mismatch on graft patency is further supported by the work of Abbot et al.88, in which segments of native artery with equivalent surfaces were divided into a control (untreated) group and a treated group. Compliance of test segments was altered by glutaraldehyde treatment and they were reimplanted. Patency rates for less compliant grafts were significantly lower. One limitation to the general application of the Ballyk study78 was the use of interrupted sutures at the anastomoses; clinically, a continuous suture technique is generally used. Another argument in favour of grafts of similar compliance to host arteries is the notable differences in patency rates between vein grafts and internal mammary artery grafts when used for cardiac bypass (10-year patency rates of 50 and 80–90 per cent respectively79). Tai et al.12, in an assessment of the compliance of various vascular conduits, pointed out that, although human saphenous vein was anisotropic (a non-linear relationship in the degree of radial deformation was displayed in response to increased intraluminal pressure) like artery, it was incompliant at the high mean pressures of the arterial circulation. This was attributed to the fact that vein has comparatively less elastin and smooth muscle than artery, and so transfers its load at a lower strain (pressure) than artery. This means that a significant compliance mismatch occurs when vein grafts are first implanted and exposed to arterial haemodynamics (saphenous vein grafts, however, are more compliant than synthetics in that pressure range). It is also important to note that the intimal hyperplasia cascade is initiated as early as 24 h after exposure to the arterial circulation59. Although other biochemical reasons have been put forward for the differences in patencies between vein and arterial grafts in the heart80, there is still a cogent argument for compliance mismatch playing a significant role in graft failure. Availability The issues of off-the-shelf availability, and ease of harvest and propagation of autologous cells, are the greatest obstacles to the widespread acceptance of tissue engineering technology in clinical practice. The problem is significant in vascular surgery where much of the workload involves treatment of the acutely ischaemic lower limb. This limits the ‘bench to bedside’ applicability of the in vitro approach because of its long cultivation times. The limited conduit length generated in current bioreactors and the need for additional surgical interventions for sourcing autologous cells (vein harvest, liposuction, laparoscopy and bone marrow aspiration)81 are other issues that must be considered. The recent discovery of circulating bone marrow-derived endothelial progenitor cells (EPCs)82 may potentially alleviate some of these concerns. EPCs are CD34 + cells derived from adult bone marrow that are committed to a vascular lineage. They can be isolated and cultured from the mononuclear fraction of circulating leucocytes. Unlike pure stem cells, by definition they have little or no potential for self-renewal83, but they retain the ability to differentiate into endothelial cells and in some situations vascular SMCs18. EPCs are mobilized into the circulation in response to a variety of stimuli, notably vessel injury, hypoxia and tissue ischaemia. It seems that they may be a significant source of endothelial cells in the face of vascular injury, forming new blood vessels via vasculogenesis rather than angiogenesis. The exciting potential of EPCs was demonstrated recently by Tepper et al.84 in a rat model of soft tissue ischaemia. EPCs were mobilized into the circulation and subsequently formed new vessels at the site of the ischaemic insult. The development of an ischaemic gradient along a murine skin flap correlated with the density of EPC-driven endothelialization and new vessel formation. These findings have several important implications. First, the mobilization of EPCs in response to ischaemic tissue means that a larger circulating load is present in the peripheral blood of patients who require their presence most, that is patients with an ischaemic insult such as an ischaemic limb or a myocardial infarction. EPCs are easily harvested by venesection. Second, EPCs can be obtained from a ‘non-depleting, self-renewing resource’ (peripheral blood) for tissue engineering. In 2001 Kaushal et al.18 used EPCs to seed a decellularized porcine scaffold, with promising results. When implanted as a xenogenic interposition graft in a sheep common carotid artery model, the engineered vessels had a 100 per cent patency rate (seven of seven grafts) either at 15 or 130 days after implantation, whereas non-seeded conduits thrombosed. The animals received aspirin, but not warfarin, for up to 7 days after surgery. On explantation, the grafts showed confluent endothelial cell layers at 15 days with a media repopulated with SMCs at 130 days. The grafts also displayed promising vasoactive properties, including contraction in response to noradrenaline and serotonin, high nitric oxide production and nitric oxide-dependent relaxation. These were comparable to the properties of native sheep carotid artery and better than those of saphenous vein. Such a finding is encouraging as the last two properties are implicated in the better long-term patency rates achieved with native internal mammary artery than with vein grafts in cardiac bypass. The approximate time from harvest to implantation in the above work was 49 days. It should be appreciated, however, that the animals that were the source of cells were healthy, with no ischaemic injury to stimulate EPC mobilization from the bone marrow; the cell yield per unit of blood is likely to have been low. Recent findings by Tepper's group84 suggest that cultivation times may be shortened further. Pharmacological augmentation of endothelialization is a further consideration: erythropoietin and hydroxymethylglutaryl-coenzyme A reductase inhibitors (statins) can boost mobilization and homing of EPCs to areas of denuded vessels85,86. Statins are often taken by arteriopaths as a cholesterol-lowering agent. Even with all of these developments, the immediately available, off-the-shelf, tissue-engineered, small-calibre vascular graft has remained unachievable. It is tempting to extrapolate the above findings to imagine a chemically coated scaffold that enhances in vivo endothelialization with the aid of circulating EPCs augmented by pharmacological means. As an alternative, developments in cell culture techniques may result in greater cell yields and faster cell expansion times for ex vivo graft repopulation within the duration of a vascular surgical procedure. 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Google Scholar Crossref Search ADS PubMed WorldCat Copyright © 2006 British Journal of Surgery Society Ltd. Published by John Wiley & Sons, Ltd. This article is published and distributed under the terms of the Oxford University Press, Standard Journals Publication Model (https://academic.oup.com/journals/pages/open_access/funder_policies/chorus/standard_publication_model) Copyright © 2006 British Journal of Surgery Society Ltd. Published by John Wiley & Sons, Ltd.
TI - Tissue engineering of vascular conduits
JF - British Journal of Surgery
DO - 10.1002/bjs.5343
DA - 2006-05-16
UR - https://www.deepdyve.com/lp/oxford-university-press/tissue-engineering-of-vascular-conduits-eThK1vM8Ex
SP - 652
EP - 661
VL - 93
IS - 6
DP - DeepDyve
ER -