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VASCULAR GRAFTS
Pınar ZORLUTUNA
INTRODUCTION
The pathology that effects the small and medium sized blood vessels is one of the major causes of death in Western society [1]. Atherosclerosis is the major disease of the blood vessels and affects the large and medium sized blood vessels that contain an intima [2]. The artherosclerotic lesion consists of a raised focal plaque within the intima consisting of a lipid core, surrounded by an extracellular matrix and smooth muscel cells and covered by a fibrous cap. As it increases in size through intimal hyperplasia it restricts blood flow and eventually block vessel [3]. Many of the advances in vascular surgery have been made possible by the development of vascular grafts which are used to reconstruct diseased or injured arteries and veins. As yet the ideal vascular graft has not been found. Autogenous grafts are the best condiuts currently in use but are not always available or appropriate.
STRUCTURE AND FUNCTION OF BLOOD VESSELS
Blood vessels function to carry blood from and to the heart and to and to from the tissues and organs.They form a branched system of arteries and veins that vary in size , mechanical properties , biochemical and cellular contetnt , and ultrastructural organization , depending on their location and spesific function ( for example ,volume of the blood to be carried , flow rate , vasoactivity ). The largest arteries such as the aorta function to transport blood originating from the heart. The small diameter muscular arteries deliver the blood from large arteries to the tissue and the organs. These then branch into small arterioles and capillaries , which function to distribute blood whitin the tissues and organs. Blood is returned to heart through venules , which combine to form veins. The large and medium sized arteries have distinct structural features, primarily the intima, media and adventitia , although these are less obvious in small arteriols and do not exist in capillaries. The intima, forms the layer closest to the blood flow and consists of a lining of endothelial cells attached to a connective tissue bed of basment membrane and matrix molecules. The mechanical properties critical to blood vessel function include the tensile stiffness, elasticity, compressibility and viscoelesticity [3].
TYPES OF VASCULAR GRAFTS AND THEIR CHARACTERISTICS
Vascular grafts can be classified into for broad groups: autogenous grafts, prosthetic grafts, biological prosthetic grafts and tissue engineered vascular grafts. Autogenous grafts are harvested from same individual into which the graft is placed , and may be either arterial or venous. Prosthetic grafts are wholly manufactured and are made of a yarn, Dacron or an extruded polymer (ePTFE and polyurethane). Biological grafts can also be obtained from a human source (allografts) or from non-human species (heterografts) .These grafts have to under go special praperation before they can be utilised in humans and have therefor been classified as prosthetic grafts. The scaffolds used for tissue engineered blood vessels can be biodegradable, partially or completaly non-biodegradable.
1.Autografts
Arterial autografts has many appealing features which make them ideal as arterial substitutes. They retain their vaibility due to an intact intrinsic blood supply, demostrate proportional arterial growth when used in children, do not degenerate with time, heal in an infected field, and exhibit normal flexibility at joints. The more widespread use of arterial autografts is limited by lack of avaibility and short lenght. The commonest autogenous condiut in clinical use is vein. It is freely available, easily harvested, available in adequate lenght and adapts well to placement in arterial circulation. The long saphenous vein is the first choice because of adequate calibre, being relatively thick walled and the longest vain in the body.
2.Synthetic prosthetic grafts
In cardiac and peripheral bypass surgery, autologous veins and sometimes autologous arteries are used as replacements. However, many patients do not have appropriate blood vessels for use as replacements, either because of diseased blood vessels or because the blood vessels were used in a previous surgery. In these cases, synthetic vascular grafts have been used successfylly in treating the pathology of large diameter arteries ( > 6 mm inner diameter) [3]. On exposure to blood flow, the luminal surface of a vascular prosthesis is immediately coated by a layer of protein, principally fibrinogen. The difference in protein absorbtion is releted to the characteristics of the graft surface. Surfaces that are irregular and have many electrochemical active sites tend to absorb proteins more readily than those with smooth, inert surfaces. Within minutes, platelets adhere to the flow surface, usually to a degree directly proportional to the concentration of the adherent fibrinogen. This results in the activation of coagulation system with deposition of fibrin and red blood cells. On the other hand the higher the porosity of the graft the greater the perigraft response, which means faster perigraft healing.
Textile synthetic grafts
Dacron grafts: Dacron yarn is a multifilament polyester yarn consisting of small continuous filaments which make the graft soft, elastic and easy to handle. The yarn can be fashioned into a prosthetic graft by weaving or knitting. In woven grafts, fabrict theareds interlaced in a simple over and under pattern both in lenghtwise circumferential directions. Woven fabric grafts have little or no stretch in any direction. The grafts are tightly constructed and are of low porosity, relatively strong and relatively stiff. The small interstices reduce bleeding and the relative strenght decreases the likelihood of elongation and dilatation. However, the graft has less desirable handling characteristics. Because of their low porosity, woven grafts has reduced perigraft healing. In a knit structure, the yarn is orientated in a predominantly longitudinal (warp knitted) or circumferential (weft knitted) direction. Since they are more stable warp knitted grafts most currently manufactured. In the manifacturing process the spacing of the yarn and hence the pore dimensions can be varied. The result is that knitted grafts generally have a higher porosity, with theoretically enhanced healing, improved compliance, superior handling characteristics and a lesser tendency to fray, compared to woven grafts. On the downside is need to preclot the graft and the reduced graft strength which can lead to long term dilatation [4]. Graft impregnation with substances such as collagen, albumin or gelatin, are being increasingly used in the manufacture of Dacron grafts. The addition of these biological substances render the graft leak proof without the loss of the favourable handling characteristics of the underlying porous material. Upon implantation the sealant is completely digested, thereby allowing normal healing of the underlying porous material. Some avidance suggests that early thrombogenicity of the graft surface may be lessened by such graft coatings with an expected improvement in graft patency [5]. A more significant advantage may be their ability to bond with antibiotics. At present all coated grafts are sinificantly more expensive than standart textile grafts.
Non-textile synthetic grafts
Expanded polytetrafluoroethylene (ePTFE)grafts: ePTFE is a fluorocarbon polymer form into sheets by a paste extrution process, producing a porous material that has solid nodes inter connected by fine fibriles. The intranodal distance can be varied to chance the graft porosity. The grafts in clinical use are impervious to blood, resistant to dilatation, and are chemically inert, highly electronegative and highly hydrophobic. The grafts are available in different wall ticknesses. The ones with thinner walls have easier handling characteristics, better conformability and improved compliance. ePTFE is better than Dacron for venous reconstraction if autogenous vein is not available.
Polyurethane grafts: There has been considerable effort made to develop polyurethane graft suitable for graft implantation. Theoretical adventages include a very smooth non-trombogenic inner surface, a thin walled graft with some compliance and improved handling characteristics. Degradetion of the polyurethane polymer leading to weakening and aneurysm has been a problem. Despite the theoritical advantages, patency rates have not been high [6].
Prosthetic graft developments
Mechanical matching of vascular grafts and host vessels has been suggested to be important in determining graft patency rates. Biomaterials that successfully integrate into surrounding tissue should match not only the tissue's mechanical properties, but also its topography. The cellular response to a biomaterial may be enhanced in synthetic polymer formulations by mimicking the surface roughness created by the associated nano-structured extra-cellular matrix components of natural tissue. As a first step towards this endeavor, an in vitro study was made to use these design parameters to develop a synthetic, nano-structured, polymeric biomaterial that promotes cell adhesion and growth for vascular applications. In a novel manner, poly(lactic-co-glycolic acid) (PLGA) (50/50 wt% mix) was synthesized to possess a range (from micron to nanometer) of surface features. Reduction of surface features was accomplished by treating conventional PLGA with various concentrations of NaOH for select periods of time. Results from cell experiments indicated that, compared to conventional PLGA, NaOH treated PLGA enhanced vascular smooth muscle cell adhesion and proliferation. However, PLGA prepared by soaking in NaOH decreased endothelial cell adhesion and proliferation compared to conventional PLGA. After further investigation, this finding was determined to be a result of chemical (and not topographical) changes during polymer synthesis. Surface chemistry effects were removed while retaining nano-structured topography by using polymer/elastomer casting methods. Results demonstrated that endothelial and smooth muscle cell densities increased on nano-structured cast PLGA. For these reasons, this in vitro study provided the first evidence that nano-structured surface features can significantly improve vascular cell densities; such design criteria can be used in the synthesis of the next-generation of more successful tissue-engineered vascular grafts [7].
Fig 1. PLGA casting technique. (a) PLGA is first treated with NaOH for 1 h to increase nanometer surface roughness. (b) Silastic is poured onto the PLGA surface to create a negative mold. (c) The silastic mold is removed from the PLGA, inverted, placed into another petri dish, and additional silastic is poured around the mold to create a well for casting. (d) PLGA is then added to the mold. In this manner, PLGA with nanometer surface roughness is created without NaOH surface hydrolysis [7].
Fig 2. Representative AFM pictures of treated and cast PLGA. Images demonstrated that compared to (a) conventional untreated PLGA, (c) nano-structured NaOH treated PLGA had a higher degree of nanometer surface roughness. Moreover, NaOH untreated/treated PLGA (a and c) and cast PLGA (b and d) exhibited similar respective surface features, thus, providing evidence of the successful transfer of surface topography from treated PLGA to cast PLGA without a change in surface chemistry. Scale bar is 0.2 
m in the x and y directions and 30 nm in the z direction [7].
Fig 3. Representative SEM pictures of treated and cast PLGA. Images demonstrated that compared to (a) conventional untreated PLGA, (c) nano-structured treated PLGA had a higher degree of nanometer surface roughness. Again, untreated/treated PLGA (a and c) and cast PLGA (b and d) exhibited similar surface features, respectively. Scale bar (lower right) is 10 m [7].
In their progression towards clinical acceptance, any new synthetic vascular grafts under development must undisputedly prove that the chemistry and structure used in the construction of the prostheses is safe and that their biocompatibility and performance as arterial substitutes are satisfactory without degradation or weakening of the device.
Numerous synthetic biomaterials have been developed as vascular substitutes, but all materials, currently available, yield lower patency rates than do autogenous vascular grafts. Future investigations have to solve two problems concerning permanent or temporary artificial vascular substitution: 1. Graft healing reactions in the interface of biomaterial and surrounding human tissue. 2. Activation of blood clotting system and body defense reactions by contact of the artificial surface with blood. Concerning the first, the graft healing consists of a series of complex events including the incorporation by the perigraft fibrous tissue response. The degree of host tissue infiltration into the biomaterial depends on pore size, surface texture, anatomical location and the material's biocompatibility. Concerning the second, all biomaterials investigated so far are more or less thrombogenic and initiate the bodys defense reaction. Only the endothelial cell is non thrombogenic and achieves this property by active metabolic processes. To develop an optimal vascular graft there are two possibilities: first raw materials without thrombogenic effect have to be found, second the artificial surface has to be modified in a nonthrombogenic one consisting of new chemical groups or host's own endothelial cells [8].
Antithrombogenic coatings: Carbon coating of grafts imparts a negative charge on the intraluminal surface should decrease thrombogenicity. Animal studies using carbon-coated ePTFE grafts (Impra-Carboflo) have shown reduced platelet deposition [9]. Expanded polytetrafluoroethylene (ePTFE) remains the most commonly utilized synthetic graft material for infrainguinal arterial reconstruction. However, patency rates of ePTFE bypass grafts are inferior to those observed with autogenous vein grafts. Modification of the luminal surface of ePTFE grafts such as coating with carbon or heparin, may prevent early graft failures and improve overall patency rates. Fluoropolymers have also been used to coat Dacron grafts. This fluoropassivated, gelatin-sealed polyester grafts has been shown to cause less thrombogenisity and tissue reaction in an animal model [10]. Heparin-bonded, small calibre, collagen-sealed polyester grafts have also been developed. This reduces platelet agregation in short term but there is a small risk that platelet agregation may increase in sensitised individuals.
Endothelial cell seeding: The major differences between patency rates reported for natural venous conduits as compared to prosthetic grafts may be caused by the absence of an endothelial monolayer on the graft surface. Especially, the small diameter (
6 mm ID) synthetic vascular grafts, which are used as lower-limb vessel replacements in patients without suitable autologous saphenous veins, have a failure rate of 53% after 4 year. Graft failure is due to thrombosis and intimal hyperplasia, an increase in smooth muscle cells in the lumen of the vessel which leads to progressive closing and ultimate occlusion of the vessel.
The source of cells for this procedure is critical. Endothelial cells may be harvested from large vessels and grown in tissue culture [11].A better approach, is the use of mirovessel endothelium derived from fatty tissue, has provided a convenient and successful cell source allowing isolation of sufficient number of endothelial cells [12]. These cells than can be incubated with the luminal surface of a prosthetic graft to produce a stable endothelial monolayer. Initial animal experimentation with endothelial cell seeded grafts produced very good results with increased graft patency and reduced thrombogenisity [13]. Unfortunately, these encouraging results were not seen in the early clinical trials which reported no difference in patency between the seeded and unseeded protheses [14]. The initial failures of clinical seeding trials were largly method related, and recently several studies have reported encouraging results with improved endothelial coverage and patency rates [15]. Using ePTFE as the sythetic substrate and vessel derived endothelial cells, has shown a 68% patency in a clinical study of peripheral implants [16]. At present, endothelial seeding is too technically demanding to warrant widespread use. However, future advances in cell culture and recombinant DNA technology may allow endothelial cells to be used as the vehicle for specifically targeted gene therapy aimed at reducing graft thrombogenicity and myointimal hyperplasia [17].
The hydrophobicity, low surface energy, and weak electrical charge of expanded polytetrafluoroethylene (ePTFE) and polyurethane provides conditions which are not optimal for endothelial cell attachment. In a hybrid vascular graft, stable ECs adhesion on a polymer surface is required under a flow shear stress generated by luminal blood flow [18,19]. Improvement in cell attachment is an essential condition for more efficient seeding of prosthetic surfaces.
Recent studies demonstrated that the arginine-glycine-aspartic acid (RGD) sequence of many extracellular matrix proteins interacts with the integrin family of cell matrix receptors in endothelial cells, which is known to be the mediating major receptor of cell attachment to matrix proteins. After seeding, grafts were exposed to shear stress for one hour, at flow rates of 100 ml/min, in an artificial flow circuit rinsed with tissue culture medium. EC attachment after seeding and retention after perfusion was assessed by image analysis using scanning electron microscopy. Both EC attachment and retention were significantly decreased on uncoated PTFE graft surfaces whereas they were increased by the coating of PTFE with fibronectin. Graft coating with an adhesion promoting RGD-peptide lead to the highest increase in EC attachment and retention after shear stress compared with fibronectin coated and uncoated PTFE grafts [20].
For making surface more suitable, vacuum ultraviolet (VUV) modification of ePTFE on endothelial cell adhesion and proliferation was investigated. Pieces of ePTFE graft material were exposed to 10, 20 or 40 W VUV radiation for 10, 20 or 40 min using a UV excimer lamp. Half of the pieces were precoated with fibronectin (20 g/ml). VUV modification had no effect on cell adhesion for all power levels studied. In addition, it appears that cell adhesion is independent of the presence of fibronectin. Cell proliferation, on the other hand, is augmented by modification, especially in the presence of fibronectin. These results suggest that VUV modification may provide a better surface for endothelial cell colonization of synthetic vascular grafts [19]. In development of small artery (less than 3 mm) prostheses, the reserchers used segmented polyurethane (SPU), which has good biocompatibility, as a graft substrate, and modified its surface by plasma or carbon treatment. The endothelial cells used are not capable of adhering to non-treated SPU surface, but the surface treatment brought about a drastic improvement in cell adhesion. The cells were almost completely retained on the carbon-deposited SPU surface after imposition of a shear stress of 2 Pa. The increase in the shear stress to 6 Pa resulted in detachment of only 20% of the cells. Furthermore, approximately 90 % of the BAECs on the plasma-treated SPU were retained after imposition of a shear stress of 9 Pa. Trypsin resistivity of the endothelial cells have also been invesatigated. Resistivity of the endothelial cells cultured on the plasma-treated SPU surface was also higher than on the carbon-deposited SPU surface. The results from this study suggest that surface-treated SPU can be used as a substrate of hybrid vascular grafts with a small caliber less than 3 mm [18].
1. Controlled Release of growth factors: Further more, growth of endothelial cells seeded on the luminal surface of synthetic vascular grafts, coated with a matrix suitable for cell seeding (e.g. collagen), can be accelerated by local, sustained release of basic fibroblast growth factor [21]. With further studies on the same subject, It is concluded that heparinized, crosslinked collagen, pre-loaded with basic fibroblast growth factor is a good synthetic vascular graft coating for in vivo endothelial cell seeding [22,23].
2. Genetically engineered endothelial cells: Seeding of vascular grafts with genetically engineered endothelial cells (EC) secreting anticoagulant or fibrinolytic agents offers a potential means of improving patency rates preventing thrombosis or limiting neointimal hyperplasia. Prior to the initiation of in vivo studies, researchers examined cell retention and tissue plasminogen activator (t-PA) secretion from small-diameter synthetic graft segments seeded with sheep venous EC genetically engineered to secrete human t-PA. Following retroviral-mediated gene transfer, EC were seeded at varying densities onto 4-mm-diameter synthetic graft segments of different composition, achieving confluent coverage of all materials. t-PA production from seeded grafts was evaluated under both static conditions and after flow exposure, for up to 3 days after seeding. t-PA secretion varied directly with increasing seeding density for all graft types, reaching a maximum of 20 ng/cm2/24 hr. For all graft types tested, approximately 50% of seeded cells were retained after exposure to flow in vitro. Retained EC remained viable as determined by t-PA secretion. The rate of t-PA secretion from collagen-impregnated Dacron grafts was higher than that obtained with other materials both under static conditions and after flow exposure. This higher rate was most likely due to the higher surface area presented by the Dacron grafts. These data demonstrate that small-diameter prosthetic graft materials can be coated with a layer of EC that (i) remains metabolically active and capable of secreting a fibrinolytic agent, and (ii) remains adherent to the graft surface after exposure to flow. These experiments provide a foundation for in vivo studies in which grafts are seeded with EC genetically engineered to increase local fibrinolysis [24]. Using this approach of seeding vascular grafts by recombinant ECs the promblem of artherial thrombosis can be solved too. Conventional antithrombotic treatments with antiplatelet, anticoagulant, or fibrinolytic drugs are not uniformly successful and are associated with hemorrhagic side effects. Thus, new approaches to the prevention and treatment of arterial thrombosis are desirable. The gene transfer approach is particularly attractive because of its unique ability to express an antithrombotic gene at selected sites of the vessel wall (where thrombosis is threatened) while avoiding systemic anticoagulation. Clinical conditions potentially amenable to antithrombotic gene therapy include coronary artery bypass grafting, percutaneous transluminal coronary angioplasty, peripheral artery angioplasty or thrombectomy, intravascular stenting, and vascular graft prostheses. Gene therapy may prove effective in preventing subacute thrombosis in these settings and, eventually, may play an adjuvant role to systemic thrombolysis in the treatment of acute arterial occlusion. The introduction of an antithrombotic gene into the arterial wall can be achieved either by direct in vivo gene transfer (e.g., by luminal administration of a viral vector) or by in vitro genetic manipulation of cells before their seeding onto vascular grafts, stents, or denuded arteries. The direct gene transfer approach has been used to deliver antithrombotic genes to animal arteries in vivo. Antithrombotic genes used to date include those encoding enzymes of the prostacyclin synthetic pathway, nitric oxide synthase, the thrombin inhibitor hirudin, and thrombomodulin. The in vitro gene transfer approach has been used to enhance the fibrinolytic activity of vascular grafts by overexpressing plasminogen activators. If the initial successes of gene therapy for thrombotic disease in animal models are confirmed by longer-term experiments, and if new vectors are developed which permit prolonged transgene expression without inflammation, human studies can be initiated [25].
It is also suggested that a cell-mediated extracellular matrix (ECM) modification of ePTFE would stimulate increased vascularization within the graft and thus promote lumenal endothelialization. ECM modified samples exhibited extensive ablumenal vascularization and tissue incorporation compared to nonmodified samples. Also, ECM modified grafts had a cellular lining, while the nonmodified grafts were void of a cellular lining except for a limited pannus ingrowth [26].
Fig 4 Light micrograph demonstrating cell growth on the ablumenal surface of 1-mm ePTFE. Hematoxylin and eosin-stained section of thin wall, 1-mm ePTFE following 8 days of HaCaT cell growth. Arrows indicate layer of cell growth on the surface of the polymer, and the low magnification demonstrates the continuous cell coverage of the surface [26].
Fig 5. Scanning electron micrographs of the lumenal surface of 1-mm vascular grafts. Nonmodified (A-B), HaCaT-modified (C-D), and II-4-modified (E-F) samples at 12× (A, C, E) and 1200× (B, D, F) magnification of mid-graft region. Bar = 1 mm and 10 m, respectively [26].
Healing of synthetic vascular grafts:
Filament wound synthetic prostheses have anisotropic material properties and are therefore able to match closely the elastic properties of the replaced host vessels. Highly porous prosthesis walls are required to allow ingrowth of capillar cells from the outer surface of the graft in order to increase endothelium coverage of the luminal surface. The coating of highly porous grafts with biodegradable polymers has been shown to result in a sealed structure at the time of implantation followed by controlled porosity during the healing process [27].
The healing performance of polyester vascular prostheses have been investigated. Uncleaned polyester grafts had shown poor healing and gelatin-impregnated polyester grafts showed delayed but satisfactory healing, whereas commercially cleaned polyester grafts had demonstrated excellent healing [28].
The healing of xenogeneic small diameter grafts (3.5 to 5.0 mm diameter) made from porcine small intestine submucosa (SIS) has also been compared with expanded polytetrafluorethylene (ePTFE) implanted in the contralateral carotid in 8 dogs. Two dogs were sacrificed for graft evaluation at 7, 28, 90, and 180 days after surgery. Only one SIS graft was occluded at 28 days and the other 7 were patent. Six of 8 ePTFE grafts were occluded with thrombi. One was patent at 7 and one at 90 days. At 7 days postimplant, the luminal surface of the SIS graft was covered by a thick, compact fibrin meshwork. By 28 days endothelial cells were seen completely covering the fibrin meshwork. Smooth muscle cells were observed in the neointima. Most ePTFE grafts had fibrin on the luminal surface which formed fibrin thrombi with platelets and numerous red blood cells. Complete endothelial coverage of the ePTFE grafts was not observed by 180 days. There was not a pronounced neointima seen on the luminal surface of the graft. The vasa vasorum was present in the fibrous capsule surrounding the ePTFE graft, but it did not penetrate into the graft as seen in the SIS graft. At 90 days the SIS vascular graft had the histological appearance similar to a normal artery [29].
Following implantation different cell types interact with synthetic vascular prostheses resulting in a complex immuno-inflammatory response. Dendritic cells are responsible for activating the primary T-lymphocyte immune response in various pathological conditions by their role as antigen-presenters. Dendritic cells which accumulated within synthetic grafts were found to co-localise with T-lymphocytes. Based on these observations, it was speculated that dendritic cells may be involved in the immuno-inflammatory responses following the implantation of synthetic vascular prostheses through their interaction with T-lymphocytes [30].
Antibiotic bonding: Prosthetic vascular graft infection is a serious problem, particularly if an aortic graft is involved. Most graft infections are due to the implantation of bacteria at the time of operation. The aim of bonding antibiotics to a graft is to prevent bacteria adhering at the time of surgery and for a few days there after. The antibiotic has to have an appropriate activity spectrum and adhere to the graft for long enough to be effective.
3. Biological prosthetic grafts
Arterial allografts: Fresh or preserved arterial allografts were used as vascular substitutes in the 1940s and 1950s. Fresh allografts underwent rapid rejection and thrombosis. Allografts preserved by formalin, gamma irradiation or freeze-drying gived better results in the short term but subsequently underwent atheromatous degeneration and aneurysm formation [31]. Problems of unsatisfactory small-calibre synthetic grafts, revision surgery and graft infection has led to a renewed interest in allogrfts. Cryopreservation with lliquid nitrogen and 15% dimethyl sulphoxide, an oxygen radical scavenger, may decrease the host immunoglogical responce and increase patency but late degeneration remains a problem [32].
Umbilical vein allografts: The glutaraldehyde-tanned human umbilical vein graft (Biograft-Biovascular Inc.) was introduced in 1975. Despite excellent results from some centers [33], there is a littel evidence that the biografts have superior patency than small-calibre prosthetic grafts. Aneurysmal dilatation occures despite the use of a Dacron mesh wrap [34]. The manufacturing is improved in 1989 but dilatation is still a long term problem [35].
Heterografts: Animal heterografts suffer from the same problems of rejection and degeneration as allografts. Glutaraldehyde-tanned bovine carotid arteries have been used for angioaccess [36]. The risk of viral transmission is a further complication to their use.
4. Tissue engineered vascular grafts
Although synthetic vascular grafts such as Dacron and ePTFE have been successfully used in treating the pathology of large arteries ( > 6 mm inner diameter), these have generally not proved successful in replacing the smaller diameter vessels. There is therefore, a massive clinical need for an alternative supply of vessels to replace diseased arteries. Tissue engineering offers the potential of providing vessel that can be used to replace diseased and damaged native blood vessels [3].
The success of a tissue-based graft depends on its ability to meet several requirements. First, a graft must possess a confluent, adherent and quiescent endothelium to resist thrombosis in vivo. The mechanical behavior of the graft must mimic the mechanical properties of a native vessel. Hence, a graft must have a highly organized collagen matrix to impart tissue strength. Finally, a graft must contain an elastin network to provide compliance and recoil [37].
Fig 6. Collagen and elastin are secreted by smooth muscle cells. Cross-linking stabilizes collagen and elastin, making them less susceptible to proteolysis. Well-organized layers of insoluble collagen and elastin result in a strong, compliant vessel [37].
Fig 7. Development of (a) collagen gel-based, (b) rolled sheet and (c) degradable scaffold vascular grafts [37].
Fig 8. Elastin immunostaining in a native (A) and an explanted tissue engineered artery (B). Arrows mark elastin staining [37].
Tissue engineering, using either polymer or biological based scaffolds, represents the newest approach to overcoming limitations of small diameter prosthetic vascular grafts. Their disadvantages include thromboembolism and thrombosis, anticoagulant related haemorrhage, compliance mismatch, neointimal hyperplasia, as well as aneurysm formation [38].
The challenge of tissue engineering blood vessels with the mechanical properties of native vessels, and with the anti-thrombotic properties required is immense. Recent advances, however, indicate that the goal of providing a tissue-engineered vascular graft that will remain patent in vivo for substantial periods of time, is achievable. For instance, collagen gels have been used to fabricate a tissue in vitro that is representative of a native vessel: an acellular collagen tubular structure, when implanted as a vascular graft, was able to function, and to become populated with host cells. A completely cellular approach culturing cells into tissue sheets and wrapping these around a mandel was able to form a layered tubular structure with impressive strength. Culture of cells onto a biodegradable scaffold within a dynamic bioreactor, generated a tissue-engineered vascular graft with substantial stiffness and, when lined with endothelial cells, was able to remain patent for up to 4 weeks in vivo. Use of a non-degradable polyurethane scaffold and culture with smooth muscle cells generated a construct with mechanical properties similar to native vessels has been tried. This composite tissue engineered vascular graft with an endothelial layer formed using fluid shear stress to align the endothelial cells, was able to remain patent with an neointima for up to 4 weeks. These results show that tissue engineering of vascular grafts has true potential for application in the clinical situation [3].
One of the first attempts to tissue engineering vascular grafts is to develop a biocompatible and mechanically stable vascular graft combining human cells and a xenogenic acellular matrix. Design/Materials decellularised matrix tubes were obtained by enzymatic cell extraction of native porcine aortas. Endothelial cells and myofibroblasts were isolated from human saphenous veins and grown in cell cultures. The inner surface of the tubes was seeded with endothelial cells or myofibroblasts and exposed to pulsatile flow. After cell extraction, the absence of cellular components, as well as the maintenance of matrix integrity, was demonstrated. Furthermore, the porcine matrix was successfully seeded with human endothelial cells, which grew to a monolayer under flow conditions. Stable biomechanical properties were achieved at physiological perfusion pressures in vitro [39].
Ovine pulmonary arteries have been tissue-engineered from autologous cells and biodegradable polyglycolic acid (PGA)-polyglactin copolymers. Use of this cell-polymer construct in the systemic circulation resulted in aneurysm formation. So tissue-engineered vascular graft in the systemic circulation which is based on a new copolymer of PGA and polyhydroxyalkanoate (PHA) was tried. Tissue engineered conduits were evaluated for collagen content, deoxyribonucleic acid (DNA) content, structural and ultrastructural examination, mechanical strength, and matrix metalloproteinase (MMP) activity. The percent collagen and DNA contents approached the native aorta over time. Also the mechanical strain-stress curve of the tissue engineered aorta approached that of the native vessel. Autologous aortic grafts with biological characteristics resembling the native aorta can be created using TE approach. This may allow the development of "live" vascular grafts [40].
Fig 9. Echographic appearance of a tissue-engineered aorta [40].
Fig10. (A) Two control explants thrombosed acutely (1 and 2 days) and marked by a fresh clot within the lumen (top). Two other control aorta occluded at 55 and 101 days. Note the lack of tissue formation within the polymer (bottom). (B) Tissue formation was visible 10 days after implantation (top). All tissue-engineered conduits were patent at 3 and 5 months postoperatively with one developing mild stenosis in the proximal anastomosis (bottom ) [40].
Fig11. (A) Structural appearance of tissue-engineered and native aorta (hematoxylin-eosin, original magnification × 100). (B) Miller's elastic staining confirmed the presence of elastic and collagen fibers in the constructed and native aorta (original magnification × 200) [40].
Fig12. Electron microscopy identified cellular structures in the tissue-engineered arterial wall and their orientation paralleled the blood flow (original magnification × 750) [40].
Fig13. The strain-stress curve of the tissue-engineered grafts changed and resembled the native aorta over time [40].
Other studies demonstrated that the feasibility of tissue engineering of viable, surgically implantable small caliber vascular grafts and the important effect of a `biomimetic' in vitro environment on tissue maturation and extracellular matrix formation [41].
CONCLUSION
The development of vascular grafts is one of the must important requirement in surgery. The grafts that are used until today, all have disadvantages. As yet the ideal vascular graft has not been found. Most frequently used grafts are autografts and synthetic grafts. Tissue engineering strategies are promising for more suitable and long-termed implantations. The researches must focus on this type of vascular grafts.
Refrences:
1. Use of cardiac procedures and outcomes in elderly patients with myocardial infarction in the united States and Canada, N. Engl. J.Med. 366, 1500-1505 Tu JV, Pashos CL, Naylor CD et al., (1997)
2. Blood vessels, In: Rubbin E, Faber JL (Eds.) Pathology. JP Lippincott Publishers, Philadelphia, 452-495 Benditt EP, Schwartz SM, (1988)
3. Tissue engineering of vascular grafts, Matrix Biology 19, 353-357 Ratcliffe A, (2000)
4. Post-operative alterations in the size of Dacron grafts: an ultrasonic evaluation, Ann. Surg. 189, 741 Nunn DB, FreemanMN, Hudgkins PC, (1979)
5. Albumin impragnated polyester vascular prothesis for abdominal aortic surgery: an improvement? Eur. J. Vasc. Endovasc. Surg. 12, 346-353 Chakfe N, Kretz JG, Petit H et al., (1996)
6. Experimental study of selected small calibre arterial grafts, J. Cardiovasc. Surg. 18, 155 Geeracrt AJ, Callaghan JC, (1977)
7. Endothelial and vascular smooth muscle cell function on poly(lactic-co-glycolic acid) with nano-structured surface features, Biomaterials, Volume 25, Issue 1, January 2004, Pages 53-61 Derick C. Miller, Anil Thapa, Karen M. Haberstroh and Thomas J. Webster
8. Endothelium and biomaterials: morpho-functional assessments, Biomedecine & Pharmacotherapy, Volume 50, Issue 8, 1996, Page 388 Ute Henze, M. Kaufmann, B. Klein, S. Handt and B. Klosterhalfen
9. Modified polytetrafluoroethylene: Indium 111-labeled platelet deposition on carbon-lined high porosity polytetrafluoroethylene grafts, J. Vasc. Surg. 4, 643-649 Tsuchia H, Cameron BL, Marcus CS, Wilson SE, (1992)
10. Experimantal evaluation of bleeding complications, thorombogenicity neo intimal characteristics of prosthetic patch materials used for carotid angioplasty, Cardiovasc. Surg. 4, 746-752 Rhee RY, Gloviezki P, Camria RA, Miller VM, (1996)
11. A singe staged procedure for seeding vascular grafts with autogenous endothelium, Surgery 84, 498-504 Herring MB, Gardner A, Glover J, (1978)
12. Use of freshly isolated capillary endothelial cells for the immediate establishmentof a monolayer on a vascular graft at surgary, Surgery 100, 392-399
13. Thompson MM, Budd JS, Bell PRF, (1994) Endothelial seeding of prosthetic vascular grafts, Vasc. Med. Rev. 5, 225-239 Jarrell BE, Williams SK, Stokes G et al., (1986)
14. Seeding human arterial prostheses with mechanically derived endothelium: detrimental effect of smooking, J. Vasc. Surg. 1, 279-289
15. Magometschnigg H, Kadletz M, Vodrazka M et al., (1992) Prospective clinical study with in vitro endothelial cell lining of expanded polytetrafluoroethylene grafts in crural repeat reconstruction, J. Vasc. Surg. 15, 527-535 Herring MB, Gardner A, Glover J, (1984)
16. Clinical autologous endothelization of infrainguinal ePTFE grafts in 100 patients: a 9-year experiance, Surgery 126, 847-855 Deutsch M, Meinhart J, Fischlein T, Preiss P, Zilla P, (1999)
17. Recombinant gene expression in vivo within endothelial cells of the arterial wall, Science 244, 1342-1344 Nabel EG, Plautz G, Boyce FM, Stanley JC, Nable GJ, (1989)
18. Adhesion strength of endothelial cells cultured on modified segmented polyurethane surfaces under flow shear stress, Biorheology, Volume 33, Issue 1, January-February 1996, Page 89 Y. Kawamoto, M. Kaibara and N. Wada
19. VUV modification promotes endothelial cell proliferation on PTFE vascular grafts, Nuclear Instruments and Methods in Physics Research Section B: Beam Interactions with Materials and Atoms, Volume 141, Issues 1-4, May 1998, Pages 193-196 J. L. Cezeaux, C. E. Romoser, R. S. Benson, C. K. Buck and J. E. Sackman
20. RGD-Peptides: A New Technique of Endothelial Cell Seeding Artifical Surfaces, Cardiovascular Surgery, Volume 3, September 1995, Page 41 K. P. WalluscheckG. SteinhoffA. Haverich
21. Binding and release of basic fibroblast growth factor from heparinized collagen matrices, Biomaterials, Volume 22, Issue 16, 15 August 2001, Pages 2291-2299 M. J. B. Wissink, R. Beernink, J. S. Pieper, A. A. Poot, G. H. M. Engbers, T. Beugeling, W. G. van Aken and J. Feijen
22. Improved endothelialization of vascular grafts by local release of growth factor from heparinized collagen matrices, Journal of Controlled Release, Volume 64, Issues 1-3, 14 February 2000, Pages 103-114 M. J. B. Wissink, R. Beernink, A. A. Poot, G. H. M. Engbers, T. Beugeling, W. G. van Aken and J. Feijen
23. Endothelial cell seeding of (heparinized) collagen matrices: effects of bFGF pre-loading on proliferation (after low density seeding) and pro-coagulant factors, Journal of Controlled Release, Volume 67, Issues 2-3, 3 July 2000, Pages 141-155
M. J. B. Wissink, R. Beernink, N. M. Scharenborg, A. A. Poot, G. H. M. Engbers, T. Beugeling, W. G. van Aken and J. Feijen
24. Optimization of Recombinant t-PA Secretion from Seeded Vascular Grafts, Journal of Surgical Research, Volume 57, Issue 4, October 1994, Pages 495-504 Vafa Shayani, Kurt D. Newman and David A. Dichek
25. Gene therapy for arterial thrombosis, Cardiovascular Research, Volume 35, Issue 3, September 1997, Pages 459-469 Giuseppe Vassalli and David A. Dichek
26. Accelerated endothelialization of interpositional 1-mm vascular grafts1, Journal of Surgical Research, Volume 113, Issue 2, August 2003, Pages 234-242 Kameha R. Kidd, Vangie B. Patula and Stuart K. Williams
27. Elastic response of filament wound arterial prostheses under internal pressure, Biomaterials, Volume 15, Issue 14, November 1994, Pages 1115-1121 G. Hellener, D. Cohn and G. Marom
28. Selecting valid in vitro biocompatibility tests that predict the in vivo healing response of synthetic vascular prostheses*1, Biomaterials, Volume 17, Issue 19, October 1996, Pages 1835-1842 Yves Marois, Robert Guidoin, Raynald Roy, Tamas Vidovsky, Barbara Jakubiec, Marie-Françoise Sigot-Luizard, Julian Braybrook, Yahye Mehri, Gaétan Laroche and Martin King
29. Healing Comparison of Small Intestine Submucosa and ePTFE Grafts in the Canine Carotid Artery, Journal of Surgical Research, Volume 58, Issue 4, April 1995, Pages 415-420 G. E. Sandusky, G. C. Lantz and S. F. Badylak
30. Identification of dendritic cells in ePTFE grafts explanted from humans, Cardiovascular Surgery, Volume 8, Issue 4, June 2000, Pages 265-273 Stephanie J. Inder, Yuri V. Bobryshev, Sanjay M. Cherian, Reginald S. A. Lord and Andy Y. Wang
31. Arterial homografts, Eur. J. Vasc. Endovasc. Surg. 12, 272-281 Callow AD, (1996)
32. Maintanence of vaible arterial allografts by cryopreservation, Surgery 83, 382-391 Borne CH, Roon AJ, Moore WS, (1977)
33. A decade of experiance with the glutaraldehyte-tanned human umbilical vein graft for revascularization of the lower limb, J. Vasc. Surg. 7, 336-346 Dardik H, Miller N, Darchk A, (1988)
34. Inravenous digital subtraction angiography and duplex scanning in the detection of late human umbilical vein degeneration, Br. J. Surg. 75, 668-670 Nevelsteen A, Smet G, Wilms G, Marchal G, Suy R, (1988)
35. Aneurysm formation in modified human umbilical vein grafts, Eur. J. Vasc. Endovasc. Surg. 11, 417-420 Strobel R, Boontje AH, Van Den Dungen JJAM, (1996)
36. A five-year experiance with the bovine heterograft for the vascular access, Arch. Surg. 121, 941-999 Brems J, Castaneda M, Garvin PJ, (1986)
37. Requirements for growing tissue-engineered vascular grafts, Cardiovascular Pathology, Volume 12, Issue 2, March-April 2003, Pages 59-64 Shannon L. Mitchell and Laura E. Niklason
38. Tissue Engineering of Small Diameter Vascular Grafts, European Journal of Vascular and Endovascular Surgery, Volume 23, Issue 6, June 2002, Pages 475-485 Omke E. Teebken and Axel Haverich
39. Tissue Engineering of Vascular Grafts: Human Cell Seeding of Decellularised Porcine Matrix, European Journal of Vascular and Endovascular Surgery, Volume 19, Issue 4, April 2000, Pages 381-386 O. E. Teebken, A. Bader, G. Steinhoff and A. Haverich
40. Tissue engineering of autologous aorta using a new biodegradable polymer, The Annals of Thoracic Surgery, Volume 68, Issue 6, December 1999, Pages 2298-2304 Dominique Shum-Tim, Ulrich Stock, Jeff Hrkach, Toshiharu Shinoka, Jamie Lien, Marsha A. Moses, Andrea Stamp, George Taylor, Adrian M. Moran, William Landis et al.
41. Tissue engineering of small caliber vascular grafts*1, European Journal of Cardio-Thoracic Surgery, Volume 20, Issue 1, July 2001, Pages 164-169 Simon P. Hoerstrup, Gregor Zünd, Ralf Sodian, Andrea M. Schnell, Jürg Grünenfelder and Marko I. Turina
