12 j. A literature review. Small-diameter vascular tissue engineering. R. van Lith. November 2004 BMTE Part I of MSc-thesis

Transcription

1 12 j Small-diameter vascular tissue engineering A literature review R. van Lith November 2004 BMTE04.57 Part I of MSc-thesis Supervisors: Prof. Dr. L.H.E.H.Snoeckx Dr. Ir. M.C.M. Rutten Ir. M. Stekelenburg Eindhoven University of Technology Faculty of Biomedical Engineering 1

2 Table of contents Introduction... 4 Chapter 1: Nature of vascular disease Atherosclerosis Endothelium dysfunction Inflammatory smooth muscle cell infiltration and proliferation Role of shear stress Cell layers Tunica adventitia Tunica media Tunica intima Biochemical properties Mechanical properties Collagen Elastin Coronary artery specific properties Chapter 3: Tissue engineered blood vessels Requirements of tissue-engineered vessels Mechanical Properties Vasoactivity Non-thrombogeneicity Approaches A-cellular constructs Non-biodegradable synthetic constructs Biodegradable synthetic constructs Biological constructs Concluding remarks Cellular constructs Non-biodegradable synthetic constructs Biodegradable synthetic constructs Biological constructs Concluding remarks Other techniques Self-assembly method Body cavities as bioreactors Synthetic protein based polymers Endovascular grafts Concluding remarks Chapter 4: Cell seeded synthetic grafts Criteria Favourable surface chemistry Sufficient mechanical compatibility Porous structure Commercial potential Scaffold materials Non-biodegradable

3 Dacron (PET) Gore-Tex (PTFE) Polyurethane (PU) Biodegradable PCL PLLA P4HB Dexon (PGA) Hyaff Copolymers Surface modifications Coatings Preclotting Chemical bondings Surface treatment Concluding remarks Production methods Fiber bonding Solvent casting and particulate leaching Membrane lamination Melt molding Three-dimensional printing Electrospinning Endothelial cell seeding Choice of cell source Phenotypic appearance Proliferative capacity Antigenic variety Function Seeding of the scaffold Gravitational seeding Hydrostatic seeding Biological glue Electrostatic seeding Characterization of cells Phenotypical markers Functional markers Preconditioning and dynamical loading Mechanical protocols Results so far Concluding remarks Chapter 5: Future directions List of references Uncited references Appendix A: List of abbreviations Appendix B: Leaders in the field

4 Introduction Deterioration of the cardiovascular system is still the most common cause of death in the western world, and though recognized as a severe problem, it remains a growing socio-economical problem. Atherosclerosis, the most common disease associated with these cardiovascular problems, is a process that leads to narrowing of the arteries by an expanding plaque within the intima of a vessel [61]. When it affects the coronary arteries, the myocardium is weakened, eventually leading to a myocardial infarct. The typical way to treat coronary disease is a coronary artery bypass graft surgery (CABG), which usually involves the replacement of the diseased artery by a substitute vein or artery (figure 1), which is harvested from the patient s leg or, in some cases, arm. The human saphenous vein (HSV) and internal mammary artery (IMA) are the most popular conduits [19]. Clearly though, removal of a functional vessel from its location in the body is suboptimal. Furthermore, in a fairly large number of patients, other problems are present like disease of the replacement vessel itself, usage in previous surgeries or even the need for multiple bypasses. Figure 1: Typical location for a coronary artery bypass graft. A recently emerged area of research is the field of tissue engineering (TE), defined as the application of knowledge from both engineering and life sciences to ultimately create biological substitutes to restore, maintain or even improve tissue function. Although this definition has been narrowed by some researchers, in this article I would like to keep the original definition in mind. The above mentioned issues concerning CABG have led to a so-called holy grail amongst researchers in the field of tissue engineering, being the development of a readily available small diameter tissue-engineered blood vessel (TEBV) that mimics the native vessel in such a way that long-term patency can be achieved [61], [20]. This article review aims at addressing recent developments in the field of small-diameter blood vessel tissue engineering, questioning several proposed approaches that have been subject of intensive research. A close look will be taken with regard to used techniques, results and relevant conclusions. Furthermore, current opinions on the requirements of a vascular graft for small-diameter vessels will be discussed. Finally, a view on future developments and relevant areas of research with respect to small caliber vascular grafts will be given. 4

5 Chapter 1: Nature of vascular disease This chapter aims at clarifying some basic issues associated with common vascular pathologies to which TEBV can provide a therapeutic answer. Also, it should give insight in what features a TEBV should have to prevent problems like thrombosis, aneurysms and stenosis as much as possible after implantation. The most abundant pathology associated with the need for bypass surgery is atherosclerosis, the primary cause of death in the western world. This disease is part of the group of endotheliopathies (i.e. ailment of the endothelium) and will be discussed in the next paragraph. 1.1 Atherosclerosis Atherosclerosis is an inflammatory process leading to a thickening of the intima and is accelerated in coronary artery bypass grafting (CABG). Besides being responsible for the need for CABG in most cases, many implanted grafts suffer from occlusion within 10 years after implantation, associated with atherosclerosis. This atherosclerosis is again associated with neointimal thickening. At present it is widely accepted that atherosclerosis can be considered as a response to vascular injury, initiated by circulating factors and modulated by local anatomy and hemodynamics [63], [84], [49] One of the substances responsible for vascular injury is oxidized low density lipid (oxldl). Infiltration into the vascular wall and subsequent oxidation is influenced by many factors, including blood pressure, circulating concentration of LDL, endothelial integrity (infiltration) and oxygen free radicals from endothelial cells (ECs), smooth muscle cells (SMCs) or inflammatory cells (oxidation) [49], [84], [46]. Regions where loss (of part) of the endothelial barrier occurs are very prone to high rates of LDL infiltration [46]. Most LDL infiltration occurs due to transcytosis though [75]. Usually this LDL is filtered through the vessel wall, but a hydrodynamic resistance to fluid flow is one factor that might delay this filtration. The hydrodynamic resistance, a property of the extracellular matrix in an artery (ECM), is increased during vasoconstriction [10]. Hence flow-mediated constriction mediated by nitric oxide [65] and prostacyclin (PGI2) might explain the low LDL accumulation rates in regions where the vessel wall is subjected to high shear. Vessel wall cells are capable to induce oxidation of LDL when there is a lack of antioxidants and/or when transition metal ions are present. When a sufficient number of LDL molecules is oxidized, the death of ECs and SMCs can be provoked [59,44]. Furthermore, oxldl can stimulate the expression of adhesion molecules on ECs and production of monocyte chemotactic peptide-1 (MCP-1) from SMCs. 5

6 1.2 Endothelium dysfunction Usually, endothelium favors vasodilatation when all factors and mediators are correctly balanced. A dysfunctional endothelium however has a preference for vasoconstriction as well as mitogenesis. Whenever endothelial cells are damaged the result is an inability to synthesize the vasodilatory factors nitric oxide [65] and prostacyclin (PGI2), resulting in a lack of anti-constrictive action. Dysfunctional endothelium will ultimately start releasing endothelins, angiotensin II and endothelial-derived constriction factor (EDCF), three vasoconstrictive mediators. Together these events will lead to an intense vasoconstrictive cascade. Furthermore, the potential of aggregating platelets to produce vasoconstrictors at the site of injury promotes an even larger vasoconstriction. The above stated is generally true for most vascular diseases. In atherosclerosis, a decreased production of NO has been demonstrated as well as an increase in the endothelin level. In the initial stage of atherosclerosis, endothelial dysfunction is one of the first events that can be observed. This dysfunction is illustrated by a loss of endothelium-dependent vasodilatation and an increased expression of adhesion molecules. An impaired ability of relaxation has been ascribed to decreased NO levels [117]. In bypass grafting, endothelial regrowth occurs rapidly, but impairment of endothelium-dependent vasodilatation in the regrown endothelium suggests a defect in receptor coupling to NO production [99]. The effects of NO are at least twofold: firstly NO inhibits neointima formation and SMC proliferation and secondly it inhibits leukocyte adhesion to ECs [90]. The second feature of endothelial dysfunction associated with atherosclerosis, i.e. the increased expression of adhesion molecules, is illustrated by induction of vascular cell adhesion molecule-1 (VCAM-1) and intracellular adhesion moleule-1 (ICAM-1), responsible for binding of leukocytes. This endothelial activation is not only associated with VCAM-1 and ICAM-1, but also with the production of other products like von Willebrand factor (vwf) and plasminogen activator inhibitor-1 (PAI-1) and release of them into the circulation. 1.3 Inflammatory smooth muscle cell infiltration and proliferation Inflammatory insudation, found in models of vein-grafting [87], remains unclear. Probably oxldl promotes infiltration after adhesion of monocytes. The discovery of platelet-derived growth factor (PDGF) has played a key role in the response to injury hypothesis. PDGF is produced by aggregating platelets and acts as a potent mitogen for SMCs. PDGF is a chemo-attractant for SMCs and thus stimulates SMC hyperplasia by directing SMCs from the media to the intima. Also it promotes matrix deposition. In vein grafting models, PDGF is considered the most important agent to direct cells into the intima [22] Ultimately, thrombotic occlusion is likely to occur in atherosclerotic coronary arteries 6

7 and vein grafts. This usually takes place by thrombus formation on a plaques surface denuded of endothelium or in a plaque that has ruptured [116]. 1.4 Role of shear stress Although atherosclerosis is associated with systemic risk factors like high blood pressure and high circulating LDL levels, the disease has a preference to develop at specific locations in the vascular tree. Such arteries like the carotid artery bifurcation and infrarenal, femoral and coronary arteries have certain regions of low shear stress in common. Studies conducted in the last 3 decades have confirmed the low-shear hypothesis of atherosclerosis first proposed by Caro and coworkers [10]. Direct measurements of shear stress values in the susceptible regions have revealed values in the order of 2-10 dynes/cm2, while in the rest of the arterial tree shear stress generally exceeds 15 dynes/cm2. It is now thought that physiological shear stress values keep the phenotype of ECs in such a way that the production of vasodilators is increased and that of adhesion molecules is decreased when compared to subphysiological shear stress, for example. Shear stress of high enough values thus keeps the endothelium in a atheroprotective state, being anti-proliferative, antioxidant and anti-thrombotic [50,74].In this chapter, it has undoubtedly become clear why a functional endothelial monolayer is essential when addressing the concept of tissue-engineered blood vessels. If the pathophysiological processes accompanying the vascular response to injury can be grasped, our total understanding of blood vessel responses to a variety of stimuli can be elevated to a level that makes us capable of producing true tissue equivalents. 7

8 Chapter 2: Blood vessel physiology To be able to place all used techniques and proposed methods in the right perspective, a basic understanding of the physiological characteristics of blood vessels is required. Although complex differences exist in the cardiovascular system, regionally as well as within organ systems, a common organization can be observed. In light of this review however, the physiological characteristics of an artery will be the starting point. 2.1 Cell layers Generally spoken, an artery consists of three distinct regions, all of them containing specialized cells and extracellular matrix (ECM). Anatomically, from the outside of the vessel to the inside, these regions are called the tunica adventitia, tunica media and tunica intima. Furthermore, the tunica intima is lined with a specialized, single layer of endothelial cells (see figure 2) Tunica adventitia The tunica adventitia or shortly adventitia, the outermost layer of a blood vessel, consists of mainly fibro-elastic connective tissue and ECM, supplying most of the mechanical strength of the vessel and the structural integrity. Main components of the adventitia are fibroblasts and elastin (hence fibro-elastic). Furthermore, the vasa vasorum are observed in the adventitia, a network of small thin-walled blood vessels needed to provide the larger arteries with the necessary nutrients and oxygen, since simple diffusion of the latter from the lumen of the artery to the outside is limited. Especially the media is dependent on the vasa vasorum. The ECM surrounding vascular cells is a complex structure consisting mainly of collagen I and III, elastin fibers, proteoglycans, hyaluronan and glycoproteins (e.g. laminin, fibronectin, thrombospondin and tenascin). The external elastic lamina serves as a boundary layer between the adventitia the smooth muscle cells of the medial layer. Figure 2: An overview of vascular anatomy.in the left picture,the structure of an artery (left) and avein (right) are drawn. On the right hand side is a cross-sectional view of an artery as seen with two different staining techniques. 8

9 2.1.2 Tunica media The tunica media or media contains mainly smooth muscle cells and elastin fibers. Especially the smooth muscle cells are highly abundant and make up the bulk of the vessel wall thickness. These cell layers are usually highly organized in larger arteries due to the function they have to fulfill in the movement of large volumes of intravascular blood (i.e. compliance of a vessel). Smooth muscle cells are orientated circumferentially, making it possible to strongly constrict. The elastin however should not be underestimated for its role in compliance, since this material, as we shall see later in this review, plays an important role in the visco-elastic behavior of a vessel. For structural support, the media rests on an internal elastic lamina, separating it from the innermost layer, the tunica intima Tunica intima The tunica intima or intima is usually the thinnest structural layer present in a vessel and is made of a single layer of endothelial cells (ECs) mounted on a basement membrane. These endothelial (squamous) cells are oriented in a longitudinal way, aligned with the direction of the main flow. Underneath this layer are a subendothelial fibro-elastic connective tissue layer and an organized layer of internal elastic lamina that provides flexibility and stability for the endothelial cells. Another type of cells found in close proximity to endothelial cells is the class of pericytes. These cells are multipotent stem cells that possess the ability to differentiate into several different cell types. They therefore provide the endothelial cell layer with a balanced cellular micro-environment. Other cells can also be present in the intimal layer, such as lymphocytes and macrophages. The surface of the EC layer express glycoproteins, together called the glycocalyx, who prohibit blood cells and plasma proteins to migrate into the vessel wall under normal conditions, hence it is a charged barrier [105]. The endothelial cell layer is nowadays considered as possibly the most important agonist in the vessel wall. Instead of being a passive player, acting as merely a sievelike physical barrier against unwanted substances to penetrate into the vessel wall from the bloodstream, it is now accepted that it has a critical role in regulating thrombolysis as well as in coagulation, inflammatory and immunological processes. The endothelial monolayer present in all vessels of the vasculature will be discussed in greater detail in paragraph 2.2 [77,93]. 2.2 Biochemical properties The endothelium is involved in various physiological processes, such as hemostasis maintenance, vasomotor tone regulation, inflammatory and immunological regulation and the modulation of angiogenesis (see figure 3). 9

10 Figure 3: Anti-thrombogenic properties of endothelial cells(figure adopted from Mitchell, 2003). Regarding hemostasis, the endothelial layer already in a quiescent state releases many agents that act against the formation of thrombi. Some of them include heparan sulfate, prostacyclin (PGI2), nitric oxide [65] and ADPases. These all inhibit platelet aggregation. Other factors that influence the coagulation process are tissuetype plasminogen activator (t-pa) and plasminogen activator inhibitor-1 (PAI-1), which have a positive effect on the digestion of platelet-fibrin clots. Besides counteracting the thrombosis-cascade, the ECs also participate in regulating thrombus-formation locally. The von Willebrand factor (vwf), a glycoprotein responsible for carrying factor VIII in the plasma, mediates the adhesion of platelets and their incorporation into evolving thrombi. The second key process in which the endothelium is involved is the modulation of vasoactivity [56]. Usually, endothelial layers favor dilatation of the vessel wall. Many mediators are released that act as potent vasodilators, such as endothelium-derived relaxation factor (EDRF), NO, PGI2 and endothelium-derived hyperpolarizing factor (EDHF). Except NO, these substances are synthesized de novo when an appropriate trigger is induced (see figure 4). Figure 4: Basic vasoactive mechanism in a vessel wall (Figure adopted from Mitchell, 2003). 10

11 Aside from vasodilatation, the endothelium also takes care of different situations when rapid vasoconstriction is needed. Vasoconstrictive agents are mainly the endothelins. These peptides are secreted by the endothelial cells primarily on their abluminal side. When released, they cause an intense and prolonged constriction of the vessel wall. A second constriction mediator is endothelium-derived constriction factor (EDCF), which plays a minor role, together with angiotensin II, another stimulating agent. It should also be mentioned that besides producing and releasing factors themselves, endothelial cells promotes vasoactivity indirectly by binding substances that are secreted elsewhere in the body, such as serotonin, acetylcholine, histamine, insulin and of course norepinephrine. When bound, the action is translated to the smooth muscle cells by the endothelium[56]. The third key process in which the endothelium plays a crucial role is the regulation of inflammatory and immunological processes. A wide variety of cellular adhesion molecules (CAMs) has been discovered, expressed on the glycocalyx surface of several cells like leukocytes, platelets and vascular endothelial cells [40]. Whenever cellular damage and/or infection occur, the CAMs get involved in the binding of leukocytes and inflammatory mediators. Although an elaborate discussion of these CAMs is beyond the scope of this review, it should be noted that they are vital for a correct functioning of a vessel. Table 1 summarizes the known CAMs at present and their function in the vessel [3,85]. Table 1: Overview of known vascular CAMs. The fourth and last key process in which endothelial cells in the vessel wall are involved is angiogenesis or new vessel formation from pre-existing vessels. Though the exact mechanisms still remain unclear, it is assumed that angiogenesis is controlled by a balance of angiogenic and anti-angiogenic cytokines [15,65]. Vascular endothelial growth factor (VEGF) and fibroblast growth factor (FGF) promote angiogenesis [74]. Both promote mitosis of cells, and are produced by smooth muscle cells and pericytes. Angiostatin (made from plasminogen) and endostatin inhibit angiogenesis. Upon stimulation of angiogenic cytokines, both the endothelium and SMCs proliferate and migrate in the direction of the angiogenic stimulus to form a new intima and media. 2.3 Mechanical properties 11

12 The structural and mechanical properties of a blood vessel determine its strength and reaction to hemodynamic forces. In the light of this review and in particular the part of it coping with the tissue engineering [107] issues, these characteristics of vessels will be covered first. Blood vessels and in particular arteries bear three distinct mechanical characteristics, Figure 5: Typical stressstrain curve of native tissue and intended graf (Adopted from Shum- Ttm, 1999). being elasticity, determined by the elastin content, tensile stiffness, determined by the collagen content and compressibility which is taken care of by the glucosaminoglycans (GAGs). The degree of contribution of elastin and collagen is not only determined by the amounts of these proteins, but also by the orientation of the fibers and the degree off crosslinking between individual fibers, meant to stabilize the fibers. Collagen and elastin are part of the extracellular matrix (ECM) and may be secreted by vascular smooth muscle cells in the medial layer of the artery wall and/or fibroblasts in the adventitial layer. Together with the GAGs they make that an artery has a visco-elastic nature, responding to stress as indicated by the following curve (see figure 5) [101]. For clarification, the anticipated development of the curve for a vessel conduit after implantation is depicted as well. Arterial burst strength and compliance are both closely connected with the viscoelastic behavior. The burst strength determines whether a vessel is capable of withstanding the physiological pressures that exist in the cardiovascular system, while the compliance of a vessel determines the way pulsatile wave energy is dissipated and thus how the blood enters the downstream vasculature. Further in this review we will focus on mechanical values that are relevant for tissue engineering of vessels. First let us return to the two basic mediators of vascular mechanical strength in a vessel Collagen Mechanical strength of a blood vessel is mainly derived from collagen and elastin, both produced by the SMCs in the media. The collagen fibers are the main providers of the tensile stiffness of a vessel and therefore maintain the structural integrity, 12

13 largely determining the burst pressure. The synthesis of collagen is positively affected by growth factors and cyclic strain. The fact that collagen is such a prevalent protein in the vessel wall stimulated many research groups to use collagen substrates as scaffolds for tissue engineered blood vessels (TEBV). However, collagen fibers only provide mechanical strength when they are crosslinked in a highly organized way (see figure 6). In coronary arteries, the collagen content varies between 40 % and 27%, depending on age (less when the person is older) and coronary artery (left or right) [70]. Figure 6: Organization of collagen and elastin in a vessel Elastin Elastin, the other ECM protein secreted by the SMCs, is responsible for the elastic properties of a blood vessel. It acts as a recoil protein that stretches and pulls the vessel back to its original diameter, and thus contributes to a large extent to the compliance of vessels. Vascular dilatation, which would be caused by the pressures of blood flow, is prevented in vivo after implantation. Also, adequate transmission of pulsatile wave energy is hindered when the compliance of the TEBV is too low, which is often the case when synthetic grafts are used. The in vitro production of elastin fibers still remains a challenge, although some synthesis of elastin after implantation has been observed [64]. Consequently, the elasticity index remains below requirement level. The elastin level in human coronary arteries varies between 1 to 12 %, again being lower at higher age. The ratio of collagen and elastin in the right coronary artery decreases from 3.8 to 1.7 and in the left coronary artery from 5.4 to 2.6, similar reductions of about 50 % between the age of 20 and 60 [70] Coronary artery specific properties In the previous paragraphs, an attempt has been made to clarify the basic properties of blood vessels in general. However, this review is specifically aimed at the research done in the field of small-diameter vessel replacement. Therefore it is useful to briefly describe the specific features of human coronary arteries and their hemodynamics which are relevant to the TEBV construction process. 13

14 First of all the physical characteristics, important for the construction of the graft: a typical coronary artery has a diameter that varies from mm and a wall thickness that varies from mm, according to van Andel and coworkers [115]. Two important mechanical properties, essential for TEBVs to match as good as possible, are the burst pressure, being about 2000 mmhg [79] and the vascular compliance, varying between mmhgand mm 2 /mmhg at a mean arterial pressure of 100 mmhg, with a mean compliance of 0.02 mm 2 /mmhg [121]. The hemodynamic characteristics of the coronary circulation are highly relevant to the TE approach. Blood pressure varies between 80 and 120 mmhg with a mean of 100 mmhg by generalization, whereas the mean coronary blood flow rate is about ml/min [121]. The pulsatile flow has the effect on the coronary arteries of imposing a cyclic radial strain of 4-10 % per 100 mmhg (Greisler & Zilla, 1999), giving a circumferential stress of MPa and a shear stress on the endothelium of dynes/cm 2 with a mean of 6.8 dynes/cm 2 [17]. The latter is interesting, since the shear stress values in the rest of the arterial circulation is much higher, being generally spoken higher than 15 dynes/cm 2. The shear stress in the coronary circulation is closer to that in veins, which could explain the success of saphenous vein grafts, even though the higher arterial blood pressure is detrimental [48]. 14

15 Chapter 3: Tissue engineered blood vessels With the ongoing increase in percentage of the western population suffering from cardiovascular disease, also the incidence number of coronary artery bypass grafting is getting larger and larger. Therefore the development of an artificial construct that mimics the native coronary artery after implantation seems increasingly interesting, also from a commercial point of view. Although considerable success has been achieved with such constructs for replacement of larger vessels, when applied to small-diameter arteries with a radius of less than 3 mm, outcomes have been disappointing. Hence a tissue engineered blood vessel (TEBV) that can replace an autologous vein or artery without the at present still accompanying higher costs (apart from still far from satisfactory results) is desirable. Though kind of macabre the increasing number of diseased is likely to further reduce the costs of TEBVs. Hence the future is bright for groups working on this domain! However, as noted the results are still far from optimal and many aspects still have to be improved before a coronary artery equivalent is produced that can be approved by the FDA. To accomplish these goals several issues have to be addressed. In the next chapters some of these issues and requirements will be discussed and elucidated. 3.1 Requirements of tissue-engineered vessels Four basic TEBV requirements have to be met in order to mimic the functional characteristics of a living blood vessel. TEBVs should be non-thrombogenic, nonimmunogenic, exhibit vasoactivity and possess mechanical properties matching those of the native vessel [108,61,79]. To achieve this, Nerem et al. describes three essential technologies to be developed [61]. 1. Cell technology This core technology concerns the source of cells to be used and the manipulation of cell function. Obviously, the best choice would be autologous cells, although usually allogenic cells are used in current research. For smooth muscle cells this may not be a problem, as we will see in paragraph 4.3, but for ECs there should be a solution to enhance immune acceptance. Control and manipulation of cell function, which is achieved by either altering the cell s environment or changing the cell s genetic features, are both an option. The first can be easily understood because SMCs and ECs are all very much under influence of their environment, both biochemically and mechanically. The second feature includes the change of a cell s genetic ability for the secretion of anti-thrombogenic substances, expression of elastin and immunological acceptance. 15

16 2. Construct technology In order to gain success in the field of TE, a neovessel should optimally resemble the native vessel to be replaced. With regards to this matter it is not only important to be able to engineer constructs that provide the right 3-dimensional structure and mechanical properties, e.g. burst pressure, elasticity etc. Also, the matter of reproducibility should be considered if in the future a methodology has to be commercialized. 3. Integration into living system For this, one should consider the choice of animal model, control of biological responses that occur when implanted into a living system and immunogenic acceptance. Is it possible to control the reaction mechanisms and in-vivo evolution of the construct? The accomplishment of these technologies should result in a functional TEBV that fulfills the requirements for a blood vessel substitute (see table 2 [60]), being nonthrombogeneicity, vasoactivity and proper mechanical properties. Table 2: Requirements for a blood vessel substitute (Adopted from Nerem, 2003) Mechanical Properties The mechanical properties problem is the first to be addressed, since a TEBV, being a polymeric graft, a reconstructed neovessel or a combination of both, has to bear a load immediately after implantation. The hemodynamic conditions inside the human body require a blood vessel to have a certain strength (in terms of burst pressure) and compliance [25]. In order too prevent rupture of the neovessel the burst pressure should be at least 2000 mmhg. Although physiological pressures generally do not exceed 250 mmhg, a large safety margin is desirable. For a correct energy dissipation and prevention of thrombosis, the radial elasticity or compliance of the implanted graft should match that of the native artery as close as possible [1,47]. Compliance of a vessel is defined as the ratio of change in diameter over change in blood pressure, expressed as the percentage diameter change per mmhg (%/mmhg x 10-2 ). Writing this definition in mathematical form, this gives: 16

17 ( Ds Dd ) C = (1) [( P P ) D ] s d s In this equation of compliance C, D S and D D depict the end-systolic and end-diastolic diameter, respectively, whereas P S and P D describe the corresponding pressures. A typical curve of the compliance of a human coronary artery in relation to arterial pressure is depicted in figure 7. In this graph, the compliance is depicted as change in cross-sectional surface per mmhg, another way to define this parameter. The declining slope found in coronary arteries is lacking in grafts by definition [126]. Figure 7: Variation of compliance with pressure in coronary arteries (Reproduced from Zilla, 1999). A normal radial elasticity of a human coronary artery results in a dilatational strain of about 6% per 100 mmhg [126]. When this compliance property can not be achieved when engineering a TEBV, the patency rate reduces greatly with increasing compliance mismatch. Early trials with Dacron and PTFE grafts showed a compliance mismatch much larger than with autografts like saphenous vein (see figure 8). Patency (%) R 2 = Compliance (%/mm Hg) Figure 8: Relation between compliance of native coronary artery or coronary grafts and patency rate. Increasing values of compliance belong respectively to PTFE grafts, Dacron grafts, bovine heterografts, umbilical veins, saphenous veins and native arteries (graph data taken over from Zilla & Greisler, 1999). 17

18 It is clear that the patency rates of grafts will increase together with compliance in a linear fashion. This graph indicates the importance of a sufficient compliance of the graft to assure a good clinical outcome. These essential mechanical properties can be achieved through a visco-elastic nature of the newly constructed artery, which is mainly determined by the contents of collagen and elastin in the TEBV, as mentioned earlier. Finally an additional mechanical requirement should be mentioned. A good suturability seems to be obvious, but nonetheless very important. The anastomosis between the neovessel and the intact artery is often the location where intimal hyperplasia occurs, also because of a compliance mismatch [45]. This can lead to occlusion of the graft and therefore a shorter patency period. Furthermore, the suture should hold under circumferential as well as longitudinal tensions, and should retain axial and radial compliance and pulsatility Vasoactivity The ability of a neovessel to adapt to changing hemodynamic forces and chemical stimuli is directly correlated with the mechanical properties as mentioned above. Grafts that are unable to dilate or contract in an appropriate way will consequently lack the correct compliance upon changing hemodynamic situations. Therefore the medial layer is indispensable in a TEBV. The endothelial cell layer is usually the layer that transfers the stimuli to the medial layer since it provides a (not absolute) barrier between the blood with vasoactive factors and the smooth muscle cells. Therefore this layer seems indispensable as well. However, in vitro experiments have proven that smooth muscle cells can behave in a vasoactive way similar to their behavior in in vivo conditions, even in the absence of an endothelium Non-thrombogeneicity Possibly the most important requirement for successful substitution of a blood vessel is a confluent, adherent endothelial layer to provide a non-thrombogenic lining. Whenever blood comes into contact with another surface than the endothelium, there is a risk of thrombosis [7]. Furthermore, loosely attached ECs could detach right after implantation due to blood flow related shear stress. The EC layer also actively inhibits thrombosis. This is achieved by thrombomodulin receptors, heparan sulfate, proteoglycans and the secretion of NO, prostacyclin, protein S and t-pa, all of which inhibit the clotting process (see fig. 3 in Chapter 2). Aside these features, the endothelium also plays a role in blood pressure regulation, angiogenesis and adhesion and transmigration of inflammatory cells. Hence, it can be understood that the lack of a functional endothelium or an aberrant one can lead to problems such as atherogenesis, bleeding disorders or graft rejection. In conclusion, a TEBV for small-diameter arteries should behave vasoactively through a non-thrombogenic luminal lining and exhibit correct mechanical properties, matching the compliance of the native vessel as close as possible and possessing a 18

19 strength that allows a burst pressure of at least 2000 mmhg, thereby creating a sufficiently large safety margin. 3.2 Approaches The current constructs used for the development of TE neovessels can be classified in two main categories: a-cellular and cellular constructs. Another division can be made into synthetic or natural constructs. These classifications can be looked upon in a chronological perspective as well, since basically investigators started with just a- cellular constructs to replace blood vessels, whereas the need for a more adaptive approach with incorporation of living cells was recognized only in a later stage. In the next paragraph we will focus on the main disadvantages and advantages of the methods, clinical results, if any, and future directions. Also, other recently developed methods with great potential are briefly discussed A-cellular constructs Non-biodegradable synthetic constructs Synthetic polymer grafting has been very successful for the replacement of largediameter blood vessels. Such materials as Dacron (PET) and eptfe have extensively and successfully been used and were the standard biomaterials for synthetic grafts over the past decades. However, for the replacement of small-diameter blood vessels, they suffer from rapid occlusion, mainly because of the thrombogenic bloodcontacting surface. Basically, one could say that the healing process of polymeric synthetic grafts is quite unfavorable, leading to problems like incomplete endothelialization and intimal hyperplasia (IH), eventually resulting in the occlusion of the graft. Intimal hyperplasia is characterized by proliferation and migration of SMCs from the medial wall to the intima, with subsequent synthesis of ECM proteins and other matrix material [1,95]. The cause of IH is probably multifactorial, and factors involved are a compliance mismatch between native artery, the site of anastomosis and the graft, which leads to a disturbed flow pattern, turbulence and vessel wall injury. When implanted as a replacement for small blood vessels, blood and tissue reactions occur immediately, which leads to a cascade of events ultimately occluding the graft. Dynamic protein adsorption/desorption, followed by platelet adhesion, inflammatory cell infiltration, and EC and SMC migration are the first steps. These processes result in the deposition of a compact fibrin layer, the so-called pseudo-intima, and of foreign body giant cells, densely packed between the outer layer of the graft wall and surrounding connective tissue. These initial effects keep evolving until final failure of most grafts. An alternative for eptfe and Dacron grafts are the polyurethanes (PU), polymers that gained popularity because they have a better biocompatibility than eptfe and Dacron [112]. 19

20 Aside from this, PU, and especially segmented polyurethane (SPU) have far better elastic properties than the traditional polymers. Some types of PUs have reached the phase of clinical trial [112]. However, eventually the technique was abandoned because of varying problems like occlusion or rapid biodegradation. Most recently, a new PU named MyoLink has been produced, consisting of poly(carbonate) polyurethane. This material has shown superior compliance characteristics as well as adaptable mechanical properties. A final version of the MyoLink graft was compared with native vein and artery grafts, as well as eptfe and Dacron grafts (see figure 9). Figure 9: Comparison of several grafts, indicating a excellent compliance of the MyoLink graft (graph reproduced from Tiwari, 2002). This PU grafts show similar compliance properties as a native artery in the physiological range of mmhg [112]. Furthermore, this new PU better resists biodegradation than other PUs. Other experiments with PU prostheses as small as 1.5 mm ID were performed by Zhang et al [125]. PU grafts were implanted in rats and remained patent for minimal 8 weeks, indicating no initial occlusion problems. Nonetheless, still no conclusion can be drawn as to whether PU grafts are truly applicable for small blood vessel replacement. The significance of immunological reactions to the polyurethane, still being a foreign material to the human body, has yet to be proven irrelevant. Another major concern is the carcinogenic potential of the degradation products, although recent results indicate only a non-significant biodegradation [94]. Biodegradable synthetic constructs In the case of non-degradable synthetic grafts tissue-material interactions continue, and this eventually could lead to failure of the graft. Bioresorbable or biodegradable polymers disappear over time, hence excluding foreign-body reactions. Therefore, it should be theoretically possible to in situ engineer a neoartery if enough loadbearing capacity is present to resist initial dilatation, and if cells are attracted in such way that desirable physiologic characteristics are already present before degradation occurs. 20

21 Two intensively studied polymeric materials are polyglycolic acid (PGA) and polylactic acid (PLA). PLA is mainly used in its L-form (PLLA) because it has a high mechanical strength. PGA is highly crystalline and hydrophilic. PLLA is more hydrophobic and thus less prone to hydrolysis. In several clinical studies both polymers have shown to have a good potential to eventually develop into a viable alternative for diseased vessels. The main problem is the generation of a sufficiently strong neo-tissue before total resorption of the polymer occurs, leading to aneurismal dilatation. The combination of PGA or PLLA with other polymers was not successful in solving this problem. Although a relatively compliant material could be introduced into the body to induce circumferential orientation of the SMCs, tissue ingrowth compromises the compliance in many cases. Several methods have been developed to enhance the patency rates. Examples are the linking of heparin to graft surfaces so that the thrombogenic activity is reduced. The major problem is the duration of heparin activity due to premature release of the compound or the presence of a physical barrier, created by adherent blood components. Other modifications are the coating of the luminal surface with carbon so that electro negativity is improved and thus thrombus formation reduced. Another coating material is fibrin glue. It is believed to improve endothelialization and other physical and chemical variations. However, little clinical improvements have been observed. Biologica l constructs One of the main problems of a-cellular synthetic constructs is the danger of an immunological response to the bulk material or the degraded debris. Furthermore, it is impossible to mechanically remodel these constructs. The concept of natural (biological) materials was therefore proposed as early as 1985 when Weinberg and Bell proposed the concept of a collagen-based construct [120]. The first tested samples had an unfavorable mechanical strength, so that several, yet unsuccessful attempts have been made to reinforce these constructs, among others by improving the alignment [41,28] or through glycation of the collagen fibers [24]. Another a-cellular approach was described by Badylak and coworkers [4]. Badylak used the small intestinal submucosa (SIS), which is a cell-free layer of collagen derived from the small intestine. When rolled it has a typical burst pressure of about 3500 mmhg, making it very strong. A SIS graft is about half as compliant as a canine carotid artery, but much more compliant as a typical vein graft [82]. It is therefore a potential candidate for application. Badylak s group observed recruitment of a neointima and endothelial monolayer within one month after implantation [83]. Other researchers like Huynh [31] and Kim [38] observed patency rates of 13 and 8 weeks in rabbits and rats, respectively. Recently, investigators have been tempted to use constructs containing both collagen and elastin fibers. A lacking elastin is often the reason for insufficient compliance. 21

22 One could therefore argue that decellularized arteries may offer more suitable grafts than collagen grafts. Tamura et al. [106] reported a method for the decellularization of an artery, resulting in a construct with aligned elastin and collagen fibers and a burst pressure comparable to that of native vessels. Upon implantation very good results were obtained. The conduits had normal layers of SMC and a lining of endothelial cells covering the entire luminal surface. Even neovascularization was observed in the adventitia. A coating of heparin was required to prevent direct obstruction of the graft though. Instead of using arteries for decellularization, the group of Schaner [91] tried human saphenous veins. Advantages of veins include a thinner wall which is easy to decellularize and thus also a thinner ECM for better migration of recipient cells and nutrients. Besides, veins are easy to harvest from tissue donors. After decellularization, Schaner found excellent burst pressure levels and suture retention strength. Upon implantation in dogs an equally good patency time was observed. Although this group refrained from using human veins and did not test the constructs in a low-flow environment where thrombosis is more likely to occur, the model seemed to be feasible and interesting for further research. Concluding remarks The use of a-cellular constructs thus far resulted in quite disappointing results in terms of long-term patency, strength and immunogenic response. However, when their mechanical and chemical characteristics are properly adapted they can be used in practice. A-cellular constructs can be easily sterilized and time-consuming, expensive and difficult tissue engineering is avoided. Upon implantation, a major disadvantage is the incomprehensive recruitment of cells after implantation. This is poorly understood so that it is also difficult to predict the development of a functional neovessel. Another drawback is the lack of an endothelial lining when implanted, leading to exaggerated embolization and immunogenic responses Cellular constructs Non-biodegradable synthetic constructs One of the first attempts to improve function and patency of polymeric grafts was to seed endothelial cells onto the synthetic surface so that it became less thrombogenic. This technique has been applied to Dacron and eptfe grafts [16,92], although up till now only autologous endothelial cells have been used because of immunological reactions. Since the adhesion of EC to the polymer material was poor, several attempts have been made to alter the surface properties of the polymer, such as the use of a recombinant fibronectin-like adhesion factor [54] or of ammonia plasma treatment, both on eptfe grafts [102]. No satisfying improvements were observed. In 2001, Seifalian and co-workers showed that a combined binding of heparin and the RGD-sequence (amino acids sequence of arginine-glycin-aspartic acid) to the 22

23 surface of MyoLink improves cell attachment to over 75% [94]. Hence these investigators succeeded in obtaining a vascular conduit with an optimal compliance and sufficient EC-attachment and -retention properties. Although this CABG substitute seems to be appropriate, it remains to be proven if in situ remodeling of the graft is such that it continues to bear the correct properties, even with the polymeric framework still inside. It is possible that a correct regulation of vasotone is inhibited by the synthetic material, eventually leading to intimal hyperplasia. Biodegradable synthetic constructs A new approach was established using biodegradable polymeric scaffolds seeded with cells. The general idea is that bioresorbable or biodegradable polymers are seeded with (autologous or allogenic) cells in vitro. Further in vitro culture would stimulate the development of a neovessel. When inserted in the body, the polymer continues to degrade while the new tissue further develops into a functional vessel with an optimal cell orientation, neovascularization and vessel wall organization. For example, a cell-seeded polymer could be left to develop in vitro for one month, after which the rest of the development will take place in vivo, leading to the formation of an endothelium and a media containing collagen and elastin. Although at present this technique seems very promising, a number of issues have to be verified. For example, it has to be determined whether or not residual polymer fragments have a negative effect on the microenvironment under in vivo circumstances. Another issue is the way in which a cell-seeded scaffold has to be preconditioned before placement. It is clear that an organized tissue with correct mechanical and vasoactive properties can only be achieved when the appropriate mechanical conditions are imposed. Still it remains unclear what these conditions should ideally be. Several regimes have been investigated with various degrees of success, as far as mechanical strength and vasoactivity are concerned. Sometimes the preconditioning protocol seemed to be appropriate, but the in vivo hemodynamics resulted in an overshoot of ECM production. Although it is clear that ECM deposition is vital for the establishment of graft strength, the excessive matrix formation indicates undesirable tissue remodeling. The balance is thus very fragile, not well understood, and hence insufficiently controllable. This approach however has not been without success. Recently promising results were obtained by several groups. Niklason et al. [64,104] reported a burst pressure of more than 2000 mmhg in constructs made of PGA meshes seeded with SMCs. Under these conditions, the constructs were allowed to mature for 8 weeks after which ECs were seeded onto the lumen. The authors applied a dynamic culturing protocol by using pulsatile radial stresses (165 pulses per minute, maximal radial strain 5%). In addition, the culture medium was supplemented with 20% fetal bovine serum (FBS), ascorbic acid, proline, alanine, glycine and copper sulfate to ensure proper nutrition and biochemical signaling. Pulsed constructs had very high burst pressures, most likely because of high collagen content. Also, they exhibited some vaso-activity in that they contracted when stimulated with serotonin. Compared to 23

24 native vessels however, the response was less prominent. There was also no elastin present. In vivo studies with this type of constructs have already started; the first results seem to be promising. Even though it only concerns short-term results in animals and many aspects still have to be elucidated, it still seems very promising. Attempts are underway to induce elastin production by means of transfecting the SMCs with the tropo-elastin gene. At MIT in Boston, the group of Shum-Tim [101] used a PGA scaffold supplemented with an outer layer made of polyhydroxyoctanoate (PHO). They cultured mixed cell populations of FBs, ECs and SMCs and seeded such populations onto the luminal surface of the grafts. Seven days later the grafts were implanted in lambs. Evaluation of the constructs after a 3-5 months period revealed a structure of the implanted graft similar to that of native arteries. An endothelial lining was present, as well as that the cells in the TEBV generated collagen, elastic fibers and the von Willebrand factor, indicating a differentiated EC-type. Furthermore in TEBV s implanted for almost 6 months the stress-strain profile started to approach that of native tissue. Although optimization of cell attachment and degradation time need to be further established, the ability of a cell-seeded graft to remodel into a physiological vessel has been proven. Hoerstrup et al. [29] designed a tubular scaffold made of PGA coated with the novel material poly-4-hydroxy-butyrate (P4HB). The investigators seeded tubular constructs of 5 mm diameter with myofibroblasts. After 4 days of incubation, ECs were seeded onto the luminal surface. During the first week, an increasing flow and pressure was applied to minimize the wash-out effect of flow after seeding of the graft. An increasing burst pressure, collagen content and DNA content was observed when the constructs were pulsed compared to static controls. The feasibility of using this type of construct was demonstrated although exact culture conditions remain to be determined. The group of Opitz [67] made a P4HB scaffold using a mold the size of an ovine aorta. The polymer solution was poured in the mold, after which the dichloromethane was allowed to dissolve. Before cell loading, the scaffolds were precoated with collagen I, Matrigel, and gelatin and then placed in a roller mixer. This precoating provided a better cell attachment. vsmc loading was done by suspending 10 7 cells in 80 ml of medium and allowing them to attach overnight at 37 degrees in the roller mixer. A coating of ECs completed the in vitro culture. Pulsatile cultivation of the construct resulted in a complete colonization with vsmcs, with expression of SMC specific markers like α-sma, calponin and caldesmon, although the first two of these markers were less prominent as compared to native ovine arteries. Also collagen and elastin contents amounted to a small percentage of that in native arteries. In our opinion this undesirable effect could be prevented by varying the culture conditions. One of the most important aspects of the work of Opitz is the demonstration of redifferentiation of vsmcs, indicating a reversible phenotypic modulation of SMCs even when cultured in vitro. Since beneficial features of SMCs 24

25 are otherwise lost because of the in vitro culturing, this could be of great value. In summary, the latest progress in the field of biodegradable polymers as scaffolds for TEBVs has been very encouraging. Some questions remain to be answered. For example, what culturing conditions are best to promote the development of a compliance-matching graft? Furthermore, how could the issue of lack of elastin content be tackled? The same can be said about the long maturation time needed for optimizing a construct in vitro. Once these problems solved the capability of remodeling and vaso-activity makes this approach very appealing. Biologica l constructs Biological grafts on which cells are seeded could offer some advantages over the earlier described constructs. First off all the problem of remodeling associated with synthetic non-biodegradable grafts could be circumvented. Secondly the imminent danger of residual polymer fragments would be circumvented. As mentioned earlier, Weinberg & Bell developed a collagen-based constructs. More recently they produced an adventitia of fibroblasts and collagen, a media from SMCs and collagen and an intima from ECs. Nevertheless, due to inherent weakness of collagen gel additional Dacron support sleeves were necessary to withstand physiological pressures. This addition however limited both remodeling and vaso-activity. Therefore, investigators focused their efforts on the improvement of collagen-based constructs without using synthetic support materials. Some investigators aimed at improving strength with magnetic prealignment and mandrel contraction; however no sufficient strength could be achieved. Recently Girton et al. demonstrated that glycation of the collagen matrix through culturing in high glucose medium significantly improved the construct strength [24]. According to their measurements, a construct of 1 mm thickness and 3 mm diameter has a burst pressure of 225 mmhg. This strength is still far from the one observed in native arteries and insufficient for TEBVs to serve as arterial replacements. Berglund and his coworkers [6] proposed the use of a so-called construct sleevehybrid (CSH). To this end, an a-cellular collagen support sleeve was fabricated by dehydrating a tubular collagen suspension. A collagen suspension with human fibroblasts was then cast upon the support sleeve. As such the support sleeve provides the initial strength required for a vessel conduit, while during cell maturation the sleeve itself is remodeled to be incorporated in the construct as a whole. Berglund observed a positive effect of the a-cellular support sleeve on the strength of the CSH. The same was true for crosslinking of the collagen fibers through glutaraldehyde treatment. Both in supported and crosslinked constructs the burst pressures were as high as 700 mmhg. Although this is much higher than in earlier constructs, it still remains below the burst strength of native arteries, which lies in the order of 2000 mmhg. Aside from this the crosslinked constructs seemed to be more prone to abrupt failure than non-crosslinked constructs. 25

26 Berglund s experiments clearly show the feasibility of tailoring a support sleeve made of collagen such that compliance and mechanical strength are optimal, whereas the biodegradation process of co-cultured layers of a-cellular and cellular collagen is comparable to that in synthetic polymer-based constructs, eventually leading to a completely cellularized vessel conduit. In summary, at present collagen-based biological constructs in combination with an a-cellular collagen support sleeve seem to have the best potential for use as vascular graft, although they still lack the desired strength. The consequence of the absence of any synthetic material is a reduced strength. Concluding remarks By reviewing the concept of cell-seeded scaffolds, whether synthetic or biological, it becomes clear that biodegradable materials are superior over others both from an immunological and remodeling point of view. Although potentially ideal many questions remain to be answered. Correct culturing conditions are essential for the in-vitro development into a vessel-like conduit. These conditions have a profound effect on cell orientation, strength, compliance and visco-elastic behavior of the construct, on the formation of ECM and the expression of other proteins. Furthermore, the degradation process should be characterized in more detail in order to predict the level of the eventually desired strength. Hence the choice of scaffold material, whether coated, modified or otherwise enhanced, is crucial. All factors have to be adequately tuned in order to obtain a TEBV suitable for commercial use. Even with all these drawbacks, the polymeric biodegradable scaffold seeded with living cells offers a great potential and has the benefit over biological constructs in terms of higher initial strength and better controllability Other techniques Self-assembly method In 1998 L Heureux et al. proposed an exciting new methodology, now known as the self-assembly method [42]. In short the procedure involves a number of consecutive steps, i.e. 1the construction of an inner membrane consisting of a dehydrated fibroblast layer, 2 the wrapping of another layer around the inner membrane, providing the media, and 3 a second layer of fibroblasts to make up the adventitial layer and also intended to provide the necessary strength. Finally, the three layers were wrapped around an inert cylindrical mandrel made of PTFE. Each layer was created by culturing the required cell source (FBs, SMCs etc.) in a flask for a certain amount of time. When a sheet was formed, it was rolled around the mandrel. New layers were added as described above. In some cases the intraluminal side was covered with endothelial cells. After the formation of a cohesive construct, L Heureux and colleagues tested the mechanical strength and/or integrity. Very 26

27 encouraging results were obtained as a supraphysiologocal burst pressure of over 2000 mmhg was reached. Furthermore, constructs were organized strikingly similar to native human arteries. Observed was a dense collagenous matrix, ECM proteins such as fibronectin and laminin, and even elastin fibers produced in the adventitial layer. Especially the presence of elastin could be of utmost importance in favoring this new methodology above traditional ones. Furthermore, desmin was produced by the SMCs, which is a feature normally lost in cultured SMCs. Some questions however still remain to be answered, such as about the right conditions for obtaining optimal compliance and visco-elastic properties. The major drawback of this technique is the extremely long in-vitro maturation period of at least three months, which could hamper the progress towards clinical applications. Body cavities as bioreactors Recently, Julie Campbell s group [11,14,8] published the first results of a new alternative. They investigated the possibility of body cavities such as the peritoneal cavity to behave as a native bioreactor for TEBV maturation. The underlying assumption is that foreign body material invokes an inflammatory reaction. Haematopoetic cells which float in the peritoneal cavity attach to the foreign body, differentiate into myofibroblasts and form a capsule through the production of ECM. In the peritoneal cavity of dogs polymer tubes of up to 25 cm long were implanted for a period of three weeks. The tubes, around which a thick homogeneous capsule had formed was tested for mechanical properties and characterized biochemically. Burst pressure exceeded 2500 mmhg, while α-sma and collagen as well as small amounts of elastin, were found. Campbell reimplanted these conduits as replacements for femoral arteries in the same dogs, which was associated with remarkably good patency rates. Harvested TEBVs not only exhibited α-sma expression as before, but also synthesized myosin and smoothelin. Vasa vasorum were localized in the adventitia. In a minor group thrombosis was found. Even so, the possibility of growing tubes suitable for vascular replacement inside the receivers own body with the use of minimal invasive surgery has been proven. More research is needed to understand all processes taking place in these cavities, as well as to evaluate this procedure in human cavities. Synthetic protein based polymers Recently, a new class of polymers based on synthetic proteins, has been described. For instance, recombinant DNA technology allows the production of a polymer based on a core sequence of elastin. Furthermore, it has been shown that the degradation rate of such materials can theoretically be controlled by an incorporated chemical clock, achieved through the addition of certain chemical components. Also the attachment characteristics are better controlled by hooking special RGD sequences to the polymer [124]. 27

28 Endovascular grafts Similar to the replacement of a diseased or occluded vessel segment, it is theoretically feasible to insert a stent-like graft inside a diseased vessel [124]. After introduction, gradual deterioration of the native vessel segment and optimization of the endovascular graft might occur. However, some aspects remain unclear. For instance, these grafts need to be particularly thin-walled for fitting into a delivery sheet or catheter. Also, porosity and tissue reactions might be very different from conventional graft placing Concluding remarks Although eptfe and Dacron are still the material of choice for large-diameter grafts, their application in small-diameter grafts is still problematic. Recent developments have been such that in the next decade a living vascular graft with controllable, predictable and desirable characteristics is likely to be constructed. This can be achieved by culturing blood vessel cells on scaffolds, whether biological or synthetic. The in vitro development of scaffolds and suitable grafts has proven to be most encouraging. The correct hemodynamic and biomechanical preconditioning in combination with possible incorporation of bioactive agents and/or genetic engineering can lead to an optimal prepared neovessel at the time it needs to be implanted. At present the development of correct preconditioning protocols is the main objective of investigation in most laboratories. Even though recent results indicate a possible role for biological scaffolds, at this moment synthetic scaffolds are more likely to result in a fast clinically applicable tool to replace autologous native veins in bypass surgery. 28

29 Chapter 4: Cell seeded synthetic grafts The technique of cell-seeded synthetic grafts seems the most promising and especially in recent years promising results have been achieved by several research groups. Synthetic materials are preferable due to a higher controllability, e.g. a highly predictable lot-to-lot uniformity. Synthetic polymers can be further engineered to give a wider range of properties. These two advantages over natural materials is the reason why in this review focus will be laid on this methodology. Several aspects that are of major influence in possible success in TEBV for bypass surgery will be discussed. The scaffold material at first seems to have the biggest influence, since this defines the attachibility of cells to the scaffold surface, degradation time in the body (if any) and development of the new to be formed vessel. Therefore it is important to define the graft criteria that deserve the most attention before looking at the various biomaterials that could be used. 4.1 Criteria A biodegradable graft to be used for the seeding of cells which are subsequently preconditioned should meet the following criteria: 1. Favourable surface chemistry 2. Sufficient mechanical compatibility 3. Porous structure 4. Commercial potential Favourable surface chemistry This seems to be the most obvious criterion and has been addressed extensively in the previous chapters. In short, the polymer surface should promote endothelial cell attachment and facilitate growth and proliferation of desirable cell types as SMCs and fibroblasts. The polymer and its degradation products should not be harmful in any way, i.e. they should not elicit any form of immunological reaction or inflammation, leading to such unwanted effects as intimal hyperplasia (an important cause of graft failure). The polymer is preferably inert and anti-coagulant, making the problem of incomplete endothelial monolayer formation less problematic. 29

30 4.1.2 Sufficient mechanical compatibility Although the range of acceptable mechanical properties of the scaffold material does not seem to impose large restraints on its design, the importance should not be underestimated. At first, when cells are seeded onto the scaffold surface and conditioning is started, the scaffold itself needs to be sufficiently strong to withstand the imposed conditioning regime. Furthermore, upon graft implantation into the body, mechanical support is vital. In that case the graft is directly forced to react correctly to a pulsatile pressure. This means usually the ability to withstand physiological pressure of mmhg and a pulsatile circumferential stretch. Under physiological conditions the radial elasticity of coronary arteries should counteract a radial strain of about 6%. Also, since the compliance of a vessel is essential for successful graft applications, the visco-elastic behavior should be matched as good as possible. Another aspect that is part of the mechanical compatibility criterion are the degradation kinetics of the biomaterial. Hypothetically, the combination of the polymer s mechanical properties with those of the newly formed cell/vessel material remains constant over time if the scaffold material degrades in an ideal way (see figure 10). Figure 10: Ideal combination of tissue formation and polymer degradation of cell-seeded polymeric constructs after implantation (figure modified from Berglund, 2003). Slightly modifying this criterion means that the degradation rate has to be such that tissue integration is optimized and that the strength of the graft remains above a minimal threshold. This modification is especially necessary when part of the degradation and neovessel formation has to take place in vivo, which is most likely the case in future applications. 30

31 4.1.3 Porous structure An underestimated feature of scaffold material is its porous structure. Since the 60s it has been realized that a porous structure exhibits a crucial effect on the neovessel formation. However, still no conclusive reports have been published on the used structure in terms of optimal porosity, pore size and spatial distribution. It is widely accepted that a porous inner surface is necessary for the proper anchoring of a neointima, while a porous outer surface attracts tissue to infiltrate. Recently it has been demonstrated that a minimum pore size of 20 µm is required for cells in general to migrate into a scaffold, whereas a pore size of over 60 µm is less favorable [88]. Hence tissue ingrowth has to be encouraged by the biomaterial through a sufficiently large pore size. A pore size that is too big bears a higher risk for inflammation (if exposed to the blood of course, what is generally not intended when using a TEBV) and is less prone to cellular adhesion. While cellular ingrowth is promoted by pore sizes between 20 and 60 µm, neovascularization, a desirable event after graft implantation, is favored by pore sizes around 5 µm [88]. A combination of these pore sizes does not seem to be feasible, although a gradually increasing or decreasing pore size from the inside to the outside of a polymer graft is desirable. It might be concluded that also in terms of porous structure an optimum exists which has to be determined for every novel scaffold material Commercial potential This fourth and final criterion may seem inappropriate to assess at this stage, since the primary goal of the whole TEBV approach is to create an improvement in coronary artery bypass grafting. However, even if a methodology which meets all the requirements can be developed, its application will depend solely on the industrial interest. This industrial interest is usually guided by the trivial question: how profitable can it be? Since acceptable interventions exist, a graft not only needs to lead to substantial improvement when compared to conventional grafts; it should also be relatively cheap. A successful outcome is therefore based on the cost-benefit analysis which in turn - depends on several factors like processability, off-the-shelf availability and sterilizability. 4.2 Scaffold materials The choice of scaffold material is essential when aiming at a cell-seeded TEBV construct, since the material s properties determine many critical aspects like mechanical properties, stability, cell attachment, migration and proliferation and inflammation. There are two options for polymeric scaffold materials, biodegradable and non-biodegradable. 31

32 4.2.1 Non-biodegradable Dacron (PET) Polyethyleneterephtalate (PET), better known as Dacron, has been one of the traditional polymers used for constructing vascular grafts. It is a strong material with a tensile strength of about 175 MPa and a tensile modulus of about 14 GPa. It is available in two forms, woven or knitted, the latter having a much higher porosity and greater distensibility. Impregnation of Dacron knitted grafts with albumin, collagen or gelatin to seal the pores for prevention against blood leakage is possible [51]. Due to its relative stability the material was considered to be a good candidate for vascular reconstruction. The highly crystalline and hydrophobic nature of Dacron both prevent hydrolysis of a graft, leading to a potential of residing inside the human body for decades. In large diameter grafting, Dacron has been used extensively and with considerable result. In small-diameter grafts however, Dacron grafts suffer from fast occlusion because of high reactivity with blood and vascular tissue, which in turn leads to inflammation and neo-intimal proliferation. Also the endothelial retention rate when seeded with ECs was poor, as observed by Turner [114]. Furthermore, knitted Dacron grafts are prone to dilatation when implanted in an arterial environment [66]. Nowadays, Dacron is generally considered to be inappropriate for small-diameter grafting, although surface modifications and possibly coatings could improve endothelial retention rates and decrease inflammation responses. Figure 11: Structural formula of poly ethylene terephtalate (PET). Gore-Tex (PTFE) Gore-Tex, or PTFE, is by far the most commonly used material in implants, but in an extended form (eptfe). This is mainly due to its excellent biostability and biocompatibility in vivo. Usually, biological deterioration is absent when eptfe is used for grafting. When used in tubular grafts the electronegative surface of eptfe minimizes the reaction with blood components. Again, just like in Dacron, the highly crystalline and hydrophobic nature yields a stable product by preventing hydrolysis. Tubular grafts made from eptfe are produced by an extrusion, drawing and sintering process [43] and consist of fibrils and nodules, controllable to different pore sizes. Typical material properties are a tensile strength of about 14 MPa and a tensile modulus of 0.5 GPa. The controllable pore size is a major benefit in grafting, since this pore size substantially affects the process of endothelialization. However, still this 32

33 endothelialization remains a problem. Even the highest levels of endothelialization are far from sufficient, with levels of only 14 % retention (close to that of Dacron grafts [114] and the thrombogenecity of the graft prevents high patency rates in smalldiameter grafts. Figure 12: Structural formula of poly tetrafluoroethylene (PTFE). Although used massively in large- and medium-diameter grafts, for small-diameter surgery, eptfe grafts seem inappropriate. Still many researchers believe to be able to modify the surface of eptfe in such a way that endothelialization becomes easier an thrombogenecity decreases [5], [111], [124]. Polyurethane (PU) Polyurethanes comprise a large group of polymers with very diverse characteristics. Besides the urethane group present in PUs, other groups are usually present. Usually PUs are copolymers made of three different regions to differentiate mechanical properties. These regions consist of a hard domain derived from a diisocyanate, a chain extender and a soft domain, usually polyol. The soft part is responsible for flexibility, whereas the hard region imparts strength. Good elastic and compliant properties and acceptable biocompatibility, together with relatively easy possibilities for modification makes PU an extremely attractive material for vascular grafts. PUs have a wide range of mechanical properties due to the presence of soft, tough elastomers and strong, rigid polymers. When PUs are considered as permanent implants, possible in vivo deterioration is present because of hydrolytical instability and oxidative degradation of the soft segment. Potential carcinogenic effect of its degradation products come into play here, implying that care has to be taken before using PUs in clinical trials. Relatively hard polymers will not suffer from this problem much though. Recently, oxidatively and hydrolytically stable polycarbonatebased PUs have been proposed for vascular grafts. In a non-woven PU graft it was demonstrated that endothelialization, early stabilization of neo-intimal proliferation and a neo-intima thickness was better than in eptfe grafts [36]. At this very moment this type of graft is undergoing clinical trials. Yet it remains undecided whether PUs of any kind are favorable when compared to Dacron and eptfe although their mechanical properties are generally superior to that of Dacron and eptfe. Optimal porosity can also be achieved and controlled in certain PUs [125]. The largest problem remains the potential instability and carcinogenicity of the PU debris. 33

34 4.2.2 Biodegradable Of course many biodegradable polymers have been tested for their potential in vascular scaffolding, but most of them have been abandoned in an early stage of the development process. Some however continue to be at interest of researchers. PCL Poly ε-caprolactone (PCL) is a semi-crystalline polymer, synthesized following ringopening polymerization of the monoester. Since it is tissue compatible, PCL has achieved a tremendous success as a biodegradable suture. Recently it also attracted the attention of investigators for use as vascular scaffold, since it exhibits a better elasticity than PGA or PLLA. The elastomeric properties required for certain tissues are not present though. Recently Serrano and colleagues [97] demonstrated good adhesion, growth, viability, morphology and mitochondrial activity of adhered endothelial cells, indicating potential for vascular tissue engineering. A disadvantage of PCL is the degradation time, which is in the order of two years. Figure 13: Structural formula of poly-ε-caprolactone (PCL). PLLA Polylactic acid (PLLA) is another potential scaffold material. The L-homopolymer is a semi crystalline, natural occurring hydrolytic product, and has a high tensile strength and low elongation characteristics. This makes PLLA more suitable for load-bearing applications such as in the cardiovascular system. PLLA is also slightly hydrophobic, so that it is more soluble than PGA in organic solvents and thus easier to process. Also the hydrolytic degradation rate is low due this hydrophobic nature. The typical degradation period of PLLA is in the order of two years. P4HB A new biomaterial is poly-4-hydroxybutyrate (P4HB) [103]. P4HB is not only a naturally occurring polymer, but also easily synthesized in vitro. It is a substance belonging to the group of polyhydroxy-alkanoates and is characterized by a rough, porous surface, ideal for cellular attachment. Its pore sizes can be varied from nm, a range well suited for different TE applications. The degradation rate in vivo is in the order of months, making it exceptionally attractive for vascular TE. As described only recently, the potentials of P4HB are largely unknown. When used as a coating P4HB yielded good results [29]. Figure 14: Structural formula of polyhydroxy butyrate (PHB). 34

35 Dexon (PGA) Dexon or poly-glycolic acid (PGA) has been used for decades as a suture material. It is highly crystalline in nature, and is insoluble in most organic solvents. Furthermore, the porosity of PGA scaffolds can be up as high as 90%, so that cellular ingrowth can easily be promoted. After implantation PGA scaffolds tend to loose their mechanical strength over a period of only 2-4 weeks and are completely absorbed after 4-6 months, due to their hydrophilic nature. On PGA grafts implanted in rabbits a confluent layer of ECs and myofibroblasts (MFs) was formed within 4 weeks. Within one year 10 % of the implanted grafts showed aneurysmal dilatation [26]. The implantation of PGA is associated with a considerable inflammatory response, a problem which should be considered when using PGA in small-diameter reconstructed vessels. Figure 15: Structural formula of polyflycolic acid Hyaff-11 A novel material is Hyaff-11, an esterified form of hyaluronan. Hyaff-11 is a highly hydrophobic polymer which can easily be shaped into a 3-dimensional scaffold by weaving or spinning. Hyaff-11 does not provoke inflammation upon degradation in contrast to other materials like PGA and Dacron. Furthermore, upon hydrolysis the polymer becomes more and more hydrophilic, eventually turning into a gel similar to native hyaluronan. In about 2 months Hyaff-11 is completely degraded, an excellent time-frame for tissue engineering applications. Although still in an early stage of investigation, Turner and co-workers [114] have demonstrated that Hyaff-11 supports attachment, proliferation and migration of ECs through the scaffold. The latter occurred when the fabric was not compressed, resulting in pore sizes of about µm. When these scaffolds are compressed cells cannot penetrate, so that it is more suitable for the formation of an endothelial monolayer. Also, the degradation time of compressed scaffolds is longer than that of the not compressed scaffolds because of the production of more ECM, which holds the fibers together. EC retention and proliferation rates were excellent with over 1.5 times the initial seeding concentration after 10 days, making it an attractive material to investigate more extensively in the future. Copolymers As some homopolymers do lack some required features for application in vascular grafting, the problem can be overcome through the combination of different types of homopolymers. 35

36 A copolymer of glycolide and PCL, named PGCL was first used by Lee and coworkers [43]. PGCL has a degradation time appropriate for vascular grafting, namely in the order of a few months. Lee showed that PGCL is very elastic in nature, making it interesting to achieve a desirable compliance in tubular constructs. Another copolymer, made out of a combination of PLLA and PCL (ratio 50:50) has been investigated [57,123,69,35] and even applied clinically for the reconstruction of a part of the pulmonary artery. At 8 weeks after implantation an excellent patency was observed [27]. A great variety of other copolymers have been investigated of which combinations of PLLA and PU and PGA and PLLA were most popular, but still no ultimate blend has been found Surface modifications In spite of the wide range of biopolymers that has been used or experimented with to improve biocompatibility, mechanical structure and processability, cellular adhesion rate is still poor. Cells seeded onto the lumen of polymer grafts, especially endothelial cells, generally adhere poorly to the surfaces of all above-mentioned polymers. Furthermore, upon implantation, the endothelial cells are exposed to a pulsatile flow, which causes endothelial cell losses of up to 70 %. Cellular retention rates need to be as high as possible, since every lost EC will increase the risk of thrombogenic effects. Methods used so far for improving this retention include the preconditioning of the seeded scaffold with an increasing (pulsatile) flow and electrostatic charging. However, the best results have been made with surface modifications of the graft to improve EC adhesion and retention. These surface modifications can consist of a variety of methods, like the coating of a graft before seeding with a substance that promotes EC-graft adhesion forces, preclotting the graft or treatment of its surface [122]. Coatings A truly vast variety of coating substances have been used by TE groups, with the number continuously increasing. However, a short summary of the most commonly used chemical coatings includes mainly well-known natural materials. Popular for example is the coating of polymer grafts with collagen and fibronectin, but also laminin, gelatin and other ECM components. In the case of collagen, the material can provide a fibrous matrix and cell-attachment promoting matrix proteins. Laminin, an ECM glycoprotein that regulates a variety of tissue processes, has been inconsistently found to enhance adhesion of ECs. Another large natural-occurring molecule, gelatin, also yielded various different outcomes. The application of not only a part of the ECM like collagen, but the complete ECM has been applied by Kidd and co-workers [89], who found a much higher patency in small-diameter eptfe grafts precoated with ECM. 36

37 Fibronectin was the most successful peptide up till now. This glycoprotein is synthesized by many cells, including ECs, and it is thought to facilitate attachment and expansion up to confluency of ECs. Furthermore, the effect of other coatings like collagen and gelatin will be enhanced when combined with fibronectin. Preclotting It is thought that cellular attachment can be improved by preclotting the graft. To preclot the vessel, one uses the patient s own plasma, blood, serum or a fibrin glue [39]. The effects of these substances have been compared with coatings like fibronectin and laminin, with varying results. The use of fibrin glue as a kind of embedding substance has yielded very promising results when used on eptfe and implanted in the infrainguinal position, giving evidence for usage in small-diameter grafts. Chemical bondings Chemical bonding of such surface molecules as peptides provides a more or less ofthe-shelf availability of vascular grafts. This method yields promising results. The RGD-sequence for example is a recognition site used by ECs to attach to the ECM. RGD used in chemical bonding has provided very good results, comparable to those of fibronectin. Also, RGD enhances DNA activity of ECs, and thus promotes endothelial function [30]. Heparin has also been investigated and shown to enhance cellular adhesion [39]. Surface treatment Besides modifying a graft by adding a substance, it is also possible to physically alter graft surfaces by treating the top layer only. Instead of altering the entire graft material, the surface can be modified by using irradiation or plasma treatment. Bacakova and coworkers found promising results after fluorine ion radiation of polystyrene [81]. Another possibility is to change surface properties by treating it with certain gases like argon or even oxygen [76] Concluding remarks Several materials such as P4HB or copolymers of PCL, PLLA, PGA and/or PUs seem to be appropriate, but at present no ideal substance can be identified. Besides the convential polymers, other, new materials have been discovered and found to be potentially suitable in arterial reconstruction [34]. Furthermore, surface modifications can enhance graft performance through alternation of their boundary structure and character. The search for an ideal scaffold material will probably continue for many years, although it is quite probable that several materials are suitable for usage, depending on the intended application. 37

38 4.3 Production methods Good scaffolds should allow and promote cells to attach, proliferate and in some cases migrate. Success not only depends on a proper selection of the scaffold material, but also on the processing technique to createa 3-dimensional construct. The processing technique should certainly not decrease the biocompatibility of the material. Ideally, the technique should allow to control the porous structure and to yield reproducible scaffolds. A variety of methodologies have been developed to fit the intended applications Fiber bonding To promote structural stability of a scaffold two ways of fiber bonding have been developed. These techniques have mainly been applied to bonding of PGA fibers. The first method involves casting a polymer solution (e.g. PLLA) over a PGA mesh and dissolving the solvent afterwards. Subsequently, the PGA fibers embedded in the PLLA matrix is heated for a given period of time, resulting in bonding of the PGA fibers. The PLLA matrix is required to ascertain a fiber-like structure of the PGA by confinement and to prevent the collapse of the PGA mesh [55]. Problems involved in this technique are choice of solvent, immiscibility of the polymers and lack of control of porosity. The second technique is a coating technique. By spraying a PGA mesh with a polymer solution (e.g. PLLA), a layer of PLLA is deposited on the PGA, thereby interconnecting PGA fibers [58]. This methodology has successfully been applied to create hollow tubes and is therefore interesting when vascular constructs are concerned Solvent casting and particulate leaching A new technique to overcome some of the drawbacks of the fiber bonding techniques was developed by Mikos in 1993 [55]. With this method, porosity, crystallinity and the surface:volume ratio is controllable. Briefly, sieved slat particles are dispersed in a polymer solution and cast into a glass container, after which the solvent is left to evaporate. The salt is then removed from the resulting polymer/salt composite by either leaching out the salt by placing it in water or by first heating the composite and subsequent annealing by cooling at a slow pace and then leaching out the salt. The porosity of these membranes can be controlled by choosing the appropriate size of the salt particles and the amount of salt used. The crystallinity can be controlled by using the second method to remove the salt from the composite. One major disadvantage of this methodology however, is the fact that only thin wafers of membranes can be produced. By polyethyleneglycol incorporation the material can be made more flexible [119], so that it can be rolled up to form a tubular construct which 38

39 still bears the same porous features. The mechanical characteristics of the material however, are inherently changed. Another concern is the seeding of cells onto this material. The material may be highly hydrophobic, leading to bad attachment of cells. This can be circumvented by pretreating the material by soaking in PVA [58] or by prewetting the material through soaking it ethanol for one hour Membrane lamination This relatively simple technique is used for the creation of three-dimensional scaffolds from thin membranes which are made by using the particulate leaching technique. Membranes are basically glued on top of each other by coating a small quantity of chloroform (or any solvent appropriate) on the boundary layer. The top part of each layer will be solved shortly and will solidify again, hereby connecting the individual membranes. By cutting the membranes, the layer-by-layer fabrication results in any 3-dimensional shape desired. However, this technique is only possible when the porous structure of the original membrane is preserved Melt molding This technique offers many advantages over the membrane lamination method. A mixture of the desired polymer with (sieved) gelatin or salt particles is poured into a Teflon mold in the desired shape (e.g. tubular). By heating the mold with polymer composite above glass temperature, a porous structure is formed. After removal of the salt or gelatin by placing the construct in water, a polymer structure the shape of the mold is left [109]. Likewise to the particulate leaching method, porosity of the structure can be controlled by varying the particle size and amount Three-dimensional printing Three-dimensional printing aims at creating a shape similar to the tissue that needs to be replaced. The three-dimensional printing (3DP) technique belongs to the group of the solid free-form techniques [72]. In short, in 3DP a thin layer of polymer powder is spread over a piston. Then, an inkjet printer (or other suitable printing device) prints a liquid binder ( glue ) like chloroform onto the powder layer. The piston is then lowered and a new powder layer is dropped onto the former one and the process repeats itself. Powder spraying and glue printing can be well controlled so that 3- dimensional constructs can be fabricated very precisely. Again, incorporation of salt in the powder again is possible to achieve porous constructs. 39

40 4.3.6 Electrospinning Electrospinning (ELSP) is a rather novel technique to produce fibers of diameter ranging from micrometers to as small as 100 nm under a high-voltage electrostatic field operated between the metal tip of a syringe and a metallic collector (see figure 16). A polymer solution is deposited randomly by a projected jet from the needle onto a metal collector. The metal capillary is charged to induce the electrostatic field [73]. The physical properties of the fibers can easily be controlled by changing the tip diameter, the distance between tip and collector, concentration of the polymer solution and the flow rate of the solution into the needle tip. Translation, rotation of the needle and/or collector can be used to control the shape of the constructs. In this way the construction of tubular constructs of any kind of polymer (when a suitable solvent is available) is possible. Recently, even the electrospinning of vascular polymers like collagen has been investigated [53]. Novel methods, such as alternating solutions or combining multiple needles with different polymer solutions at the same time [37], provide means to use mixed constructs and layered tubular constructs. Figure 16: schematic representation of the electrospinning process. The target substrate can be any form, e.g. a sheet, tubular mandrel or even more complex shapes like valvular. In conclusion, at present various techniques are available for the fabrication of polymer scaffolds, although no ideal method can be identified. Each technique has advantages and drawbacks, leading to different choices of method based on intentions. Otherwise, for the fabrication of porous tubular constructs many techniques can be abandoned. In fact, only 3-DP, melt molding and electrospinning seem to be appropriate. Electrospinning and 3DP seem to be the most promising techniques in incorporating superior controllability of the constructs characteristics. In terms of handling and costs of methods, electrospinning seems to offer more potential. The printing technique might turn out to yield better results in terms of reproducible constructs. 40