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TISSUE ENGINEERING OF BLOOD VESSELS

INTRODUCTION

Tissue Engineering

End- Stage organ failure or tissue loss is one of the most devastating and costly problems in medicine. Organ or tissue transplantation is the common therapy for failure or loss of organs or tissues. Unfortunately, organ and tissue transplantation have some limitations such as donor shortages. Additionally, the recipients must follow lifelong immunosupression regimens, which may increase risk of infection, tumor development and unwanted side effects. Functional replacement of organs with artificial materials is limited by increased risk of infection, thromboembolism and finite durability. In order to overcome the above problems the field of tissue engineering was born as a means to replace diseased tissue with living tissue that is designed and constructed to meet the needs of each individual patient. Tissue engineering is an interdisciplinary field hat incorporates principles of engineering and polymer chemistry in to the biological sciences. Tissue engineering is a rapidly expanding field, in which the techniques are being developed for culturing a variety of tissues in both in vitro and in vivo using polymer scaffolds to support tissue growth. The goals in tissue engineering include the replacement of damaged, injured or missing body tissues with biological compatible substitutes such as bioengineered tissues and restore function through delivery of living elements which become integrated into the patient. Tissue engineering concept was used when researchers encased mouse tumor cells in a polymer membrane and placed them into abdominal cavity of a pig in 1933 (Fuchs et al., 2001).

Soft tissue engineering offers the potential of using the patient’s own cells to create the necessary tissues for surgical repair and functional organs for replacement. Many short and long term areas of applications have been targeted. Short-term targets include tissues such as skin and blood vessel substitutes for conditions such as coronary by-pass and genitourinary systems.  

It is necessary for tissue regeneration to increase the number of cells that constitute the tissue. Cells are seeded onto a scaffold, which is either synthetic or neutral at appropriate cultivation time, followed by implantation into the host. The function of the matrix is to guide the development of the new tissue and provides structural support.

Tissue Engineering of Blood Vessels

            Tissue engineered blood vessels have a long history, beginning with early developments in polymeric vascular graft technology and in cell and tissue transplantation. More recent biologic manipulation is geared toward modulating the tissue reaction to implants and to thereby impacting on the common failure models of thrombosis and intimal hyperplasia.

            The pathology that affects the small and medium sized blood vessels is one of the primary causes of death in society. Atherosclerosis is the major disease of blood vessels. The atherosclerotic lesion consists of a lipid core surrounded by extracellular matrix and smooth muscle cells and covered by a fibrous cap. As it increases in size it clots the vessel and restrict blood flow. These vessels are usually replaced by autologous veins, or sometimes with autologous arteries. However, up to 30% of the patients who require arterial bypass surgery do not posses suitable or sufficient saphenous veins of the leg, which remains the standard conduit for coronary bypass surgery (Soker et al., 2002). Besides, arterial conduits have restricted dimensions and are limited in supply and venous conduits lack vasomotor tone and may have varicose degenerative alterations that can lead to aneurysum formation in the higher-pressure arterial circulation. Allografts are problematic because of high rate of rejection. As an alternative, use of synthetic vascular grafts such as ePTFE (expanded polytetrafluoroethylene) and, Dacron has been successful in treating the pathology of large diameter (>6 mm internal diameter) arteries. However, it has been difficult to develop grafts smaller than 5 mm internal diameter because of biological reactions at the blood material and the tissue material interface. Synthetic materials are excessively thrombotic when used to bypass arteries less than 6 mm in diameter, with thrombosis rates higher than 40% after 6 months (Niklason et al., 1999).

There is therefore, a massive clinical need for an alternative supply of vessels to replace diseased arteries. Tissue engineering offers the potential of providing vessels that can be used to replace diseased and damaged native blood cells (Teebken et al., 2000). Consequently, significant efforts in the past 15 years have focused on the development of a small-diameter blood vessel equivalent using tissue engineering. Before going on to strategies used for tissue engineering of blood vessels we should have a look at structure and properties of blood vessels briefly.

Structure and functions of blood vessels

The function of blood vessels is to carry blood from heart to the tissues and organs and from tissues and organs to the heart. They form a branched system of arteries and veins through out the body, that vary in size, mechanical properties, biochemical and cellular content, and ultrastructural organization depending on their location and specific function. The blood pumped from the heart is transported through the largest artery referred to as aorta. The blood from the large arteries is delivered to tissues and organs by small diameter muscular arteries. These then branch into smaller arteriols and capillaries, which function to distribute blood within the tissue and organs. Blood is returned to heart through venules, which combine to form veins.

The normal blood vessel is a complex tubular structure comprised of three concentric tubes or tunics. The large and medium sized arteries have distinct structural features, intima, media and adventitia. However, these are less obvious in small arteriols and do not exist in capillaries. The intima, most intimately positioned at the lumen that supports blood flow, is lined by a basement membrane below which resides a sparse layer of vascular smooth muscle cells. The endothelial cells provide a mechanical barrier to the solutes and solvents in the plasma, sense alterations in blood flow, plasma constituents and noxius elements, and secrete a variety of powerful chemical mediators that regulate blood cell trafficking, vasomotor tone, growth and vascular remodeling. The smooth muscle cells serve as responsive elements that set vascular tone, though they can assume many of the other vasoregulatory biochemical processes in a far less efficient manner than endothelial cells. Intima is adjacent to the internal elastic lamina, which is a band of elastic tissue, found most prominently in large arteries and which separates the intima from the media. The bulk of the smooth muscle cells reside in the media, which are organized concentrically, with bands or fibers of elastic tissue. The contraction and relaxation of these units allow the artery to constrict or dilate, thereby regulating blood flow. The external elastic lamina separates the media from the outer tube the adventitia. Adventitia is a loose fibrous network of fibroblasts through which course the vessels that nourish the blood vessel wall and the nerves that supply neural control and functions to add rigidity and form to the blood vessel (Edelman et al., 2000)

 The extra cellular matrix sorrounding the vascular cells is complex and combines to provide the biomechanical properties of tissue. The molecular network consists primarily of collagen, elastin in the form of fibers proteoglycans, hyaluronan, glycoproteins. The mechanical properties critical to blood vessel function include the tensile stiffness, elasticity, compressibility, and viscoelasticity. The collagen,s provide the tensile stiffness, the elastin the elastic properties, the proteoglycans contribute to the compressibility, and combined with the collagen, the elastin they are responsible for the elastic properties (Ratcliffe 2000).

Strategies for Tissue Engineering of Blood Vessels

             The challenges faced by approach of tissue engineering for replacing blood vessels are substantial. They include providing a conduit that will have sufficient strength not to burst with changes in blood pressure, a vessel wall that is elastic and can withstand cyclic loading, matching compliance of the grafts with the adjacent host vessel and aligning of the lumen that is antithrombotic.  Many tissue engineering approaches can rely on remodeling of the tissue in vivo to approach functionality in time, however the tissue engineered vascular grafts must function immediately on implantation. It is considered that tissue engineering offers the opportunity to use cell and tissue growth, by a material selection and scaffold fabrication, and cell type and phenotypic regulation, to manufacture a blood vessel that can function long term in vivo (Ratcliffe 2000).

The use of synthetic materials to fabricate a blood vessel substitute again in the 1950’ s and has led to vascular prostheses made from a wide variety of materials, including nylon, Teflon, Orlon, Dacron, polyethylene and polyurethane. Expended poly (tetrafluoroethylene) (e-PTFE) was introduced about 20 years ago, and Dacron and PTFE remain the most widely used synthetic materials. Synthetic grafts are most commonly made from Dacron or PTFE. Dacron grafts are manufactured in either a woven or knitted form. Woven grafts have smaller pores and do not leak as much blood. PTFE is a velour graft. Its smooth surface is less thrombogenic than Dacron. Its smooth wall is prone to kinking as it passes around joints necessitating it to be extermly supported. Suitable vein is not always available and in this situation PTFE should be used. It can be used in conjuction with vein as a composite graft.  

Both of these materials have been very successful in applications requiring large diameter (greater than 5 – 6 mm) vascular substitutes in areas of high blood flow. However, in low flow or smaller diameter applications such as peripheral vascular repair below the nee or coronary bypass, these grafts are not effective as mentioned before. One strategy to overcome these problems, which are mentioned before, is seeding of these materials with endothelial cells. Engineering of vascular grafts by transplanting autologous cells on specially designed scaffolds represents a new experimental concept to alleviate the disadvantages of allografts, autografts, nonliving xenografts or prostheses. These disadvantages include thromboembolism and thrombosis, anticoagulant related haemorrhage, neointima, hyperplesia, and eurysm formation, as well as the inability to grow with a young patient (Teebken et al., 2000)

Small vessel vascular tissue engineering can be categorized into three groups. These include using non-biodegradable grafts seeded with cells, using natural materials such as collagen, as grafts with or without cells, and using biodegradable polymer matrices seeded with cells. Another new strategy is using endothelial cells and smooth muscle cells without a use of scaffold. Several researchers have also investigated hybrid vascular grafts, consistinfg of synthetic materials such as Dacron or polytetrafluoroethylene seeded with cultered endothelial cells before implantation.

Polymers, Natural Materials and Cells Used in Tissue Engineering of Blood Vessels

            There is a pressing clinical need for better small caliber vascular grafts. A tissue engineered conduit, with the potential to remodel and to grow, made of autologous cells and biocompatible scaffold, would represent a significant advance. Scaffold guides the development of the new tissue and provides structural support. 

            Polymer scaffolds can be produced from natural or synthetic materials. Natural materials may closely mimic the native cellular environment as they are often extracellular matrix components and include collagen, hydroxyapatite, matrigel and alginate, among others. However, collagen gels seeded with cells have been limited by the fact that the resultant cell densities per unit volume that can be achieved are much lower than those observed in vivo, while prostheses utilizing non degradable materials suffer from risks of infection. (Niklason et al., 19997)

 Cell attachment can be improved by modifying the polymer chemically or by coating it. Growth factors can also be incorporated into the matrix. Defined shapes and sizes can be fabricated readily and reproducibly. Ideally these polymers must be biocompatible and bioabsorbable, nonimunogenic, support cell growth, and be able to induce angiogenesis to supply the newly formed tissue. The most widely used polymers in tissue engineering fulfilling these criteria include the poly (hydroxy acids) of the aliphatic polyesters, which include the polyglycolic acid (PGA), polylactic acid (PLA), and the copolymers (PLGA) of these materials. An electron microscopic visual of can be seen in figure 1. Polyglycolic acid was first developed as Dexon, a synthetic absorbable suture. These polymers can be produced in a number of shapes and physical forms, including meshes, sponges, and films with biodegradation rate that can be changed through copolymer proportions. The degradation of these polymers is via passive hydrolysis at the ester linkages and degradation times can be     “ tuned” from weeks to years by varying the proportions of the monomer units. These polymers may be easily functionalized if they are co-polymerized with an amino acid such as lysine to allow covalent attachment of various peptide sequences. Biodegradable polyesters have been used extensively for tissue engineering purposes. Copolymers of lactic and glycolic acid (PLGA) have been processed into three-dimensional sponges with varying porosities an flexibilities to serve as scaffolds for cell growth. These polymer sponges have been fashioned into tubes to serve as scaffolds for tissue such as blood vessels.

When vascular prostheses are constructed of either polyglycolic acid (PGA) or of Dacron to nearly identical weave characteristics, including wall thickness, porosity and elastic modulus, the tissue in growth following implantation of these to materials differs greatly. According to Greisler and his coworkers, the PGA material was resorbed between four weeks and three months following implantation. By three months the PGA was fully resorbed. Slower resorbtion produced by an increase in the ratio of lactid: glycolid rings resulted in a smilar but more slowly tissue in-growth. This altered chemistry was used in production of polydioxanone (PDS) grafts (Greisler et al., 1995).

 

Fig 1. Scanning electron microscopy of a fiber-based polymer scaffold consisting of a nonwoven mesh of polyglycolic acid (PGA) fibers.

Polyhydroxyalkanoates (RHA) as are another class of natural polymers with thermoplastic properties. They are biocompatible, resorbable, extremely flexible, and induce minimal inflammatory response. Because of these properties as well as their tensile strength, they are currently being used in the tissue engineering of heart valves and blood vessels (Fuchs et al., 2001).

               To study the feasibility of culturing an arterial prosthesis using degradable polyesters as substrates, Niklason and her coworkers, first studied bovine aortic smooth muscle and endothelial cells seeded onto PGA mesh and PLGA films, and grown under static conditions. They found that PLGA Films easily supported confluent monolayers of both cell types. To improve the results they have exposed the grafts to physical forces such as pulsatile stretch, to the cell polymer constructs during the culture periods, which can be seen in figure 2 (Niklason et al., 1997). 

Figure 2.  Perfusion system providing pulsatile flow to four bioreactors, each of which contains one tissue engineered vessel. . Gas exchange to perfusate occurs via the reservoir bag and perfusate pressure is monitored continuously.

The use of seeded acellularized vessels from animal or human origin may provide good functional results as an alternative strategy. Scanning electron microscopy of decellularized porcine aorta can be seen in figure3 (Teebken et al., 2000). Different from biodegradable polymer scaffolds, availability and physiological design of implants are not a problem, especially concerning small vessels. A vascular graft produced using this this approach can be seen in figure 4 (Teebken et al., 2001).

Figure 3. Scanning electron micrograph of a decellularised porcine aorta 60 min after seeding with human endothelial cells. Note the attached and round cells and the collagen fibres. Magnification 1 ´1000.

 

Figure 4. Explanted carotid segments. Compared to autografts (a), perivascular tissue reaction is increased in seeded acellular matrix grafts (b), and PDS (c) 4 months after implantation. Bar indicates 1 cm.

Endothelial Cell Seeding of vascular Grafts

Cells used in tissue engineering can be derived from numerous sources, including primary tissues and cell lines. Cells may be allogenic, xenogenic, syngeneic, or autologous. Ideally, the cells should be nonimmunogenic, highly proliferative, easy to harvest, and have the ability to differentiate into a variety of cell types with specialized functions (Fuchs et al., 2001).

In the face of the many limitations of synthetic vascular grafts, investigators have turned their attention towards seeding vascular grafts with endothelial cells as a means to improve long-term patency. By employing the endothelial cell’s ability to inhibit the full range of the vascular response to injury, investigators hoped to develop the first true cellular delivery therapy for vascular disease. Endothelial cells seeded vascular grafts were developed in order to inhibit thrombosis.

 Endothelial cells were mechanically or enzymatically harvested from autologous venous structures and mixed with whole blood or tissue culture media prior to incubation with the synthetic graft material. The primary determinants of seeding success appear to include the source of the endothelial cell, the duration of incubation prior to implantation or exposure to fluid flow, the precise seeding mode, and the substrates upon which the cells are seeded. Sources of endothelial cells are autologous venous endothelial cells whole vein homogenates, microvascular endothelial cells from adipose tissue.

The potential use of autologous endothelial cells produces a unique set of problems. Adequate time for cell harvest, isolation, identification and selective growth, and then graft preparation, seeding, electroincubation, characterization and implantation is needed. The incubation time is critical to cell seeding efficiency. One study has demonstrated that endothelial cells are adhered to grafts for 20 min incubations, but far more so if the time was extended 4-5 fold. Several studies have also confirmed the increased adherence efficiency of EC to grafts. Therefore, it appears that longer incubations for cell seeding result in enhanced cell retention upon the graft. Several groups have demonstrated the feasibility of pre-incubating seeded vascular grafts to generate a confluent, endoluminal endothelial cell monolayer prior to graft implantation. The time required to achieve a confluent endothelial lining after in vitro graft culture ranges from 3 days to 2 weeks. In both animals and humans, these grafts perform as well or better than single procedure harvest / seeding grafts (Edelman et al., 2000)

Another factor important in endothelial cell adherence to vascular grafts is graft surface. Multiple alternative graft materials have been considered since Herring’s initial reports of endothelial cell seeding on Dacron grafts, but considerations other than biocompatibility have come to the fore. Vascular grafts made of ePTFE are used as well as Dacron, but while the latter was easily seeded, only with alteration of porosity of early designs was endothelial cell seeding upon ePTFE possible. Moreover, though different bare polymer substrates affect cell adherence, the most significant advances in vascular graft endothelial seeding have resulted from coating the synthetic compounds with biocompatible proteins. Herring’s studies utilized pre-clotting of Dacron with blood, which presumably laid down a fibrin-rich substrate upon the Dacron polymer. Specific application of fibronectin protein greatly increased graft surface retention of endothelial cells, even after exposure to fluid flow conditions. Fibronectin not only enhanced cellular attachment, but also accelerated endothelial cell proliferation however; fibronectin coatings also mildly enhanced platelet deposition, although without generally compromising patency. As a result, a host of other protein-based coatings have been studied.

Several researches have demonstrated the improvement of EC (Endothelial Cell) adhesion by coating grafts with fibronectin or other allogenic human materials such as fibrin glue prior to the cell seeding of grafts. Unfortunately, such extracellular matrix protein (ECM) coatings also decrease the production of antithrombotic factors (i.e. t-PA and prostacyclin)(Siphea et al., 1999).

Another approach is based on tissue engineering. Biodegradable polymers such as polyglactin 910,poly (L-lactic acid) (PLLA), poly (glycolic acid)

(PGA), or poly (DL-lactic-co-glycolic acid) (PLGA) are shaped into vascular grafts, serving as a scaffold for tissue ingrowth. In time, these biologically resorbable scaffolds are dissolved, leaving behind a regenerated neo-vascular conduit with arterial-like mechanical properties. In vivo, the formation of fibrovascular tissue in the intima of implanted grafts exerts a compressional force on the forming tissue. Although scaffolds made of PGA fiberbers and coated with PLLA may withstand such a force, the ingrowth and organization of fibrovascular tissue inside the scaffold will lead to intimal hyperplasia, resulting in the occlusion of the regenerated vascular tissue. Furthermore, since an endothelial cell lining fails to develop on the luminal surface of the regenerated tissue, there are no regulatory mechanisms for minimizing this occlusion.

            Sipheia and his co-workers have demonstrated affect of the fibronectin coating and surface modification of PLLA in endothelial cell adherence to grafts, which can be seen in figure 5.

In addition to coating of grafts with proteins intended to increase endothelial cell adhesion, several investigators have attempted to enhance endothelial cell proliferation and inhibit smooth muscle cell proliferation by impregnating grafts with endothelial cell growth factors. Greisler et al. showed that precoating ePTFE with a fibrin glue containing fibroblast.

Studies in grafts coated with fibrin glue that releases vascular endothelial growth factor (vEGF) along with heparin have shown stimulation of endothelial cell proliferation and suppression of vascular smooth muscle cell proliferation in vitro. Similarly, endothelial cells seeded upon albumin-heparin coated grafts with immobilized basic FGF (bFGF) have been shown to proliferate more rapidly. Such carefully formulated endotethelial growth factor laden grafts may accelerate re- endothelialization of seeded grafts while limiting other maladaptive cellular responses. (Edelman et al., 2000)

Figure 5. Photomicrographs of HUVEC (Human umblival vein endothelial cells) grown on various PLLA substrates. HUVEC were plated on various PLLA substrates at a density of 2.5x104 cells/cm2. After 7 days, samples were fixed with 4% glutaraldehyde and stained with 0.1% toluidine blue. (a) Control PLLA; (b) Fn-coated control PLLA; (c) modified PLLA; (d) Fn-coated modifed PLLA (U100).

Finally, flow and shear stresses reduce the endothelial cell retention on the grafts.  The introduction of fluid flow and concomitant shear stress to a freshly seeded graft decreased endothelial cell retention by nearly one-third from static flow to high flow states when cells were seeded upon ePTFE grafts by a pre-clot technique. Similar findings noted that endothelial cell retention improved when seeded grafts were not immediately exposed to pulsatile shear stress (Edelman et al., 2000).

The seeding of genetically modified endothelial cells onto vascular grafts is also being studied. Recombinant endothelial cells which over express tissue plasminogen activator seeded on a vascular graft has been studied in a sheep. However, t-PA reduced the adhesion of EC because of its proteolytic activity. This result pushed the investigators to pay close attention to the nature of the recombinant gene products being delivered by seeded endothelial cells.  Further studies have shown that long-term gene expression may be more difficult than it was thought before. Nevertheless, enthusiasm for using genetically modified endothelial cells remains high.

History of Tissue Engineering of Blood Vessels

 

            Initial attempts at tissue engineering a blood vessel substitute involved seeding of a lumen of a synthetic graft with endothelial cells. Modifying the prosthetic graft surface with a monolayer of of autologous cells was initially suggested by Herring et al.(1987) as a means to provide a more biocompatible surface with the potential to decrease thrombosis and intimal hyperplassia.

            Weinberg and Bell were the first to construct a tissue engineered blood vessel composed of natural materials. Weinberg and Bell (1986) demonstrated in vitro development of a model blood vessel with three layers, corresponding to an intima, media, and adventitia. A confluent layer of endothelial cells was grown in culture onto the lumen of a tubular collagen construct consisting of an outer layer of fibroblasts and a middle layer of smooth muscle cells. In this study, the cells and matrix were cast in an annular mold and external Dacron mesh was used to provide the additional mechanical support. Electron microscopy revealed endothelial cells lining the lumen, and staining revealed the presence of von Willebrand factor. However, mechanical strength of this model was not sufficient to attain adequate burst strengths for in vivo applications. Also a major difference from the normal artery was the lack of elastin. Matsuda and Miwa (1995) also created a hybrid construct using an artificial scaffold of polyurethane seeded with smooth muscle cells and endothelial cells. This construct was shown to remodel in vivo such that the endothelial cells on the lumen became oriented in parallel to the direction of blood flow, and the smooth muscle cells in the medial layer redistributed and became circumferentially oriented. Implants were successful in the canine model for 1 year.

            The constructs of Weinberg and Matsuda used a composite approach to reinforce the strength of the cellular layers. A key challenge in the development of  a cellular model blood vessel was to create a construct with the required mechanical properties.  L’Heureux and colleagues (1998) improved on the mechanical strength of these constructs by alterations in culture conditions. Human vascular smooth muscle cells were cultured with ascorbic acid and produced a cellular sheet, which was placed around a tubular support to produce the media of the vessel. A sheet of fibroblasts was then wrapped around the media to serve as the adventitia, and endothelial cells were seeded in the lumen after a period of culturing. The tissue-engineered vessel exhibited a well defined, three layered organization as well as extracellular matrix proteins such as elastin on histologic analysis. The vessel construct had a burst strength of more than 2,000 mm Hg, which is comparable to human vessels. These structures were implanted into the femoral arteries of mongrel dogs and remained patent in 50% at 1 week.

Girton et al. (1999) have generated blood vessel medial layer equivalents with smooth muscle cells and collagen and have shown that the mechanical strength can be enhanced by nonenzymatic cross-linking of proteins with reducing sugars such as glucose and ribose. This glycation of the media equivalents significantly increased both stiffness and tensile strength of the constructs.

Bovine vessels containing vascular biopsy-derived smooth muscle and endothelial cells have now been cultured on tubes of partially hydrolyzed poly(glycolic acid) under pulsatile conditions (Niklasson et al., 1999). These vessels also have a burst strength greater than 2000 mmHg and exhibited the beginnings of vascular contractile responses.  Teebken and colleagues created decellularized matrix tubes by enzymatic cell extraction of porcine aortas and then seeded these natural grafts with human endothelial cells and myofibroblasts. The grafts were exposed to pulsatile flow conditions, and on histologic analysis, they resembled native vessels and had an intact endothelial cell monolayer (fuchs et al., 2001).

 An alternative to generating a cellularized vascular construct in vitro is to construct a prosthesis that is remodeled in vivo. Zarge et al. (1997) reviewed progress in the development of bioresorbable synthetic grafts that induce tissue ingrowth and remodeling after implantation to form a “neoarterry”. The clinical efficacy of thse grafts depend on a balance between rapid graft resorption during tissue ingrowth and slower degradation rates for maintanance of structural stability.

There have also been efforts to develop resorbable grafts composed of naturally occurring materials. Collagen and collagen-based materials have been increasingly used ina variety of medical products since the mid 1960s. Huynh and colleagues constructed a 4-mm diameter graft from small intestinal submucosa and type I bovine collagen. Small intestinal submucosa is a biomaterial, composed primarily of type I collagen, that has shown good patency as a large diameter graft in the canine aorta. Their results revealed excellent patency up to 3 months; histologically, the grafts were remodeled into cellularized vessels that responded appropriately to vasoactive agents such as norepinephrine, serotonin, and bradykinin  (Fuchs et al., 2001).

However, in general, the collagen constructs were designed to be similar to the synthetic polymer prostheses in terms of their ability to persist. Consequently, purification and cross-linking methods were developed to optimize mechanical strength and enhance the stability of the collagen in vivo.  In these cases, the material is seen as a foreign body and is essentially encapsulated by inflammatory cells and organizd scar tissue rather than being integrated into surrounding tissue.

Researches currently going on involve investigation of the ability of different polymer scaffolds to support attachment and growth of vascular endothelial and smooth muscle cells, comparison of different dynamic seeding methods in the construction of the vascular conduit, blood vessels [35], the use of gene therapy in vivo to modify grafts (transfection with adenovirus expressing tissue plasminogen activator), and the addition of peptide sequences to constructs to improve endothelial cell attachment (Fuchs et al., 2001).

An Example to the Tissue Engineering of Blood Vessels :

Development of a hybrid cardiovascular graft using a tissue engineering approach

Because of the the high failure rates of prosthetic grafts in peripheral vascular bypass surgery there is a great need for the development of tissue engineered graft whose compliance mirrors that of an artery, and possess a tissue-engineered luminal surface which make it less thrombogenic.

Seifalian and his co-workers have developed a compliant graft (MyoLinK) based on poly(carbonate-Urea) urethane chemistry that has compliance similar to human artery. MyoLink graft has a honeycomb structure that allows it to maintain compliance and pulsatile flow in vivo. It achieves this property by compressing its wall that accommodates increases in volume without dilating its external region.

In tissue engineering of blood vessels endothelial cell coating of the lumen of the graft is employed in order to reduce thrombogenic effects. For this purpose usually expanded polytetrafluoroethylen (ePTFE) is used. However Seifaian and his coworkers have shown that MyoLinK graft has better cell retention, needs less incubation time and therefore can be used instead of ePTFE. However, there is still need for a substance to facilitate adhesion of EC (Endothelial Cell) to the graft. Arginine-glycine- aspartate (RGD) residues of fibronectin have been used for this purpose. The whole fibronectin have not been used because of its vulnerability to hydrolysis. Heparin was also used to provide additional anthithrombogenic effect.

Seifalian and his coworkers have tried a new method to covalently bind RGD and heparin to surface in order to prevent desorption of these molecules thereby providing highly nontithrombogenic surface. They have used a new grafting technique using the spacer arms that reduces the steric hinderance caused by the proximity of the ligand to the rigid surface of the polymer backbone.

Results and Conclusion

As a result they have seen that RGD/Heparin moiety attached to the polymer surface at a concentration rate >14%. No degradation of the MyoLinK polymer was detected postbonding and the bioactivity of each peptide covalently bonded vas very high. Significantly higher metabolic activity was recorded on RGD/Hep-bonded grafts than with MyoLinK-seeded graft indicating more cells were present on the RGD7Hep graft after exposure to shear stress.

As a conclusion Seifalian and his coworkers have developed a new bonding chemistry by surface analysis that enables uniform layers of peptide and heparin to be attached to the surface that allows the inherent bioactivity of each moiety to be retained.  This strategy reduces washing off the cells when exposed to blood flow.

Suggestion of an Alternative Approach

I think a different strategy using genetically engineered animal cells should be used. The animal cells can be genetically modified so as to cease proliferation when a repressor substance is given which is normally not found in the cells of that animal. Artery samples from that geneticaly-modified animals can be taken and partially decellularized in order to remove the endothelial cell line and outer fibroblasts in order to avoid tissue rejection when implanted to the patient. The remaining smooth muscle cells in the media should be replaced by patient’s own smooth cells by seeding them on it and inhibiting the prolifration of animal smooth cells by giving the inhibitory substance to media. Therefore, after a time animal cells are replaced by proliferating patient’s smooth muscular cells. The lumen of this graft should be seeded with endothelial cells to reduce thrombosis and can be oriented linearly under pulsatile flow conditions. The outer part of the graft should be seeded with fibroblasts and collagen or another synthetic or natural materil to give the final form and rigidity to the graft. Elastin should also be incorporated in order to provide elasticity. Therefore, a three-layered blood vessel may be created using this procedure.

Conclusions

Although several approaches to the tissue engineering of a small diameter vascular graft have begun to show feasibility, demonstrating potential clinical efficacy will require much more data from long-term implant studies. The requirement for a construct with sufficient mechanical strength, a nonthrombogenic surface, long term resistance to hyperplessia and aneurysm formation, as well as compatibility with the host tissue, remains a major challenge for tissue engineering.      

 

 

 

 

 

 

 

 

 

 

 

 

 

 

References

  1. Fuchs, Julie R, Nasseri, B.A., and Vacanti J.P. Tissue engineering: A 21st century solution to surgical reconstruction. Ann Thorac surg 2001; 72:577-91

  2. Edelman, E.R., and Parikh, S.A. Endothelial cell delivery for cardiovascular therapy. Advanced Drug Delivery Reviews 42 (2000) 139-161

  3. Greisler, H.P., Gosselin, C., Ren, D., Kang, S.S., and Kim, D.U. Biointeractive polymers and tissue engineered blood vessels. Biomaterials 17  (1996) 329-336

  4. Niklason, L.E., Langer, R.S. Advances in tissue engineering of blood vessels and other tissues. Transplant immunology 197; 5: 303-306 (1997)

  5. Niklason, L.E.,gao, J., Abbot, W.M., Hirschi, K.K., Houser, S., Marini, R., Langer, R. Functional arteries grown in vitro. Science 284, 489-493 (1999)

  6. Ratcliffe, A. Tissue engineering of vascular grafts. Matrix Biology 19(2000) 353-357

  7. Seifalian, A.M., Tiwari, A., Salacinski, H.J., Punshon, G., Hamilton, G. Development of a hybrid cardiovascular graft using a tissue engineering approach. The FASEB J., 2002, vol. 16, 791-796.

  8. Sipehia, R., Chu, C.F.L., Lu, A., Lizkowski, M. Enhanced growth of animal and human endothelial cells on biodegradable polymers. Biochemica et biophysica Acta 1472 (1999) 479-485

  9.  Soker, S., Nomi, M., Anthony A., De Coppi, P. Principles of neovascularization for tissue engineering. Molecular Aspects of Medicine 23 (2002) 463-483

  10. Teebken, O.E., Bader, A., Steinhoff, G., Haverich, A. Tissue engineering of vascular grafts : human cell seeding of decellularized porcine matrix. Eur J Vasc Enovasc surg 19, 381-386 (2000)

  11. Teebken, O.E., Pichlmaier, A.M., Haverich, A. Cell seeded decellularized allogenic matrix grafts and biodegradable polydioxanone-prostheses compared with arteria allografts in a porcine model.  Eur J Vasc Enovasc surg 22, 139-145  (2001)

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