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EARLY HISTORY OF VASCULAR GRAFTS

            The development of peripheral vascular reconstructive surgery has been closely associated with the development of prosthetic vascular grafts. The quest for an ideal vascular conduit began soon after Carrel and Guthrie demonstrated in a canine model that homologous and herelogous artery and vein segments could be used as arterial substitutes. Although, living, nonrejectable arterial and venous autografts appear to be near-ideal vessel conduits, problems of procurement and size restrictions have prompted efforts to develop stable prosthetic materials. The characteristics of the ideal graft have long been defined. It must be readily available in multiple lenghts and sizes. It must be durable, withstanding after implantation the dual threats of biodegradation and mechanical fatigue. The ideal graft should have and maintain the same compliance as a normal artery. It should be flexible, maintaining its contour and bending without partial occlusion as it crosses joints. Its handling characteristics should include an ease of suturing and a maintenance of integrity. The graft must not harm the host in anyway. Its luminal surface must interact with blood elements in a minimally traumatic, nonthrombogenic fashion. It should be resistant to infection. It must be capable of sterilization without graft alteration. Finally the ideal graft should have an optimal porosity, allowing for good incorporation without causing unmanagable bleeding following implantation.

            During the evolution of vascular surgery, a variety of potential vascular grafts have been tested and have ultimately been rejected. Early graft development focused on the use of impervious nonbiologic tubes. Attempts to implant solid-wall inert arterial prosthesis of glass, metal, and paraffin met with failure.These grafts functioned adequately as short-term passive conduits, but they were never incorporated by the host and were ultimately subject to suture line disruption. Most solid wall plastic substances sutured into arterial defects fared only slightly better. Hufnagel, however, had fair success with a solid-wall tube of methylmethacrylate if the tubes was highly polished, hemodynamically contoured, and attached to the artery by fixation rings rather than anastomotic suturing. Amodification of this intubation technique employing a low-porosity woven polyester tube attached to a rigid, grooved polypropylene support ring at each end is still successfully applied in the surgical management of select acute aortic dissections and aortic aneurysms.

            The explosive growth of vascular surgery began with the introduction of pliable, porous fabric prostheses. Vinyon-N, a polymer of vinyl chloride and acrylonitrile introduced in 1952, was the result of the pioneering efforts of Vooheers et al. In their laboratory they had observed that a single silk suture projecting into the right ventricle of a dog became coated, over time, by a glistening tissue film devoid of clot. Based on this observation, they reasoned that if a prosthetic graft were constructed of a mesh type cloth, the leakage of blood throug it would be quickly halted by the formation on its inner surface of a fibrin layer. This layer might also serve as  acceptable blood-graft interface. Ultimately, fibroblasts might grow through the interstices. This fibroblastic encapsulation might serve as a viable basis for endothelial proliferation, or undergo sufficient adaptation to form a functional intima. Vinyon-N was first implanted clinically in the femoropopliteal position. Unfortunately, it proved to fatigue early, progressively losing tensile strength, and was relegated to a position of historic importance only. The inability to retain tensile

strength also proved true of Ivalon, a nontextile vinyl chloride polymer produced in sponge form; Fortisan, a regenerative cellulose product; and conduits made of fine-caliber, multifilament strands of stainless steel.

            Braided nylon tubes introduced by Edwards and Tapp were also found to rapidly lose strength and were soon discarded. These prostheses have additional historic importance in that they were the first grafts crimped in an effort to impart elasticity and shape retention with bending. In point of fact, we now know that most of the elasticity is lost during implantation and that the only real advantage of crimping is the facilitation of anastomoses. Furthermore, unless care is taken to “stretch out” the crimps during implantation, surface thrombogenicity, effective internal diameter, flow dynamics, and rate of healing are all adversely affected.

            The early conduits were all constructed with a longitudinal seam, with the materials typically fraying, it was necessary to cuff the ends of these conduits when sewing an end-to-end anastomosis with a host artery. Orlon, an additional polymer of acrylonitrile, was the first prosthetic graft manufactured as seamless straight and Y-shaped tubes. Unfortunately, Orlon, too, proved unable to retain tensile strength following implantation.

            Prosthetic grafts, once implanted, are not as compliant as indigenous arteries; with subsequent graft incorporation further hampering compliance. It is thought that compliance mismatch between graft and artery and the resulting hemodynamic effects at the anastomosis might predispose to intimal hyperplasia, graft thrombosis, or even false aneurysm formation. Synthetic elastomer, introduced in 1957, were designed as one solution to problems of compliance mismatch. These conduits simulate the elastic modulus of the artery, giving an elastic response to the graft in both the longitudinal and lateral directions. One present elastomer configuration has a microporous form, obtained by incorporating sea urchin spicules in polyurethane solution and then dissolving the spicules with hydochloric acid after graft formation. Equally intriguing is a biodegradable hydrophilic polyurethane prosthesis. Unfortunately, even in current forms, early graft thrombosis remains a significant problem.

                                 [1]                                                                        [1]

                Dacron was introduced into America in 1946 as a polyester polymer of ethylene glycol and terephthalic acid. Polytetrafluoroethylene (Teflon) first became available as a multifilament yarn in 1954. Neither of these two conduits tended to deteriorate after implantation. [2] Teflon and Dacron are excellent materials because they are strong nonreactive, noncarcinogenic and tend to retain tensile strength indefinitely. Of the numerous synthetic materials introduced, only versions of these two are in current clinical use. [3]

The Developmental Progression Of Biomaterials in Peripheral Vascular Surgery

Carrel and Guthrie

Technical factors in vascular reconstructions, 1906

Goyanes

The first autologous vein bypass, 1906

Tuffier

Solid-wall inert arterial prostheses, 1915

Murray and Best

Heparin use for anticoagulation, 1940

Hufnagel

Methyl methacrylate tubes, 1947

Voorhees

Vinyon-N, the first fabric graft, 1952

Edwards and Tapp

Crimped nylon prostheses, 1955

Sanger

Seamless fabric grafts, 1956

DeBakey

Dacron, 1957

Edwards and Lyons

Teflon fabric grafts, 1958

Soyer

ePTFE, 1972

Aortic and Small Diameter Vessel Replacements

            Vascular prostheses can be classified based on the materials of construction, method of construction and according to the diameter (Silver and Doillon, 1989). The distinction between large and small diameter vessel replacements is arbitrarily based on the degree of blood compatibility. Large diameter (12-38 mm in diameter) vascular replacement with polyethyleneterephthalate, PET), commercial products contaianing Dacron, is the accepted clinical practise. Medium diameter vessel replacement with polytetrafluoroethylene(PTFE), tradenamed Gore-Tex, in addition to Dacron and biologicals dominate the market (5-10 mm in diameter). In small diameter applications (less than 4 mm in diameter) biologicals and materials currently under development are the hope for cerebral and coronary arteries.

Vessel Replacement Using Synthetics

            The commercial market is split almost evenly between large and medium diameter vascular prostheses. The companies involved include Meadox Medical, Bard, W.L. Gore & Associates, Pacesetters, Impra and Golaski. Nearly all large diameter vascular prostheses are made of Dacron; Bard and Meadox are the leaders in this area. Three-quarters of these prostheses are bifurcated while one-forth are straight. The bifurcated are used to repace the lower part of the aorta where it branches. About 70 % of the medium diameter market is held by Gore-Tex while Dacron and biologicals account for 25 % and 5 %, respectively. Currently only 1 % of the small diameter market is held by synthetic grafts due to lack of blood compatible materials.

Manufacturers Of Vascular Grafts

Construction of Large and Medium Diameter Vascular Prostheses

            According to Sauvage et al. (1986), the modern era of effective reconstructive arterial surgery began in 1952 when Voorhees, Jaretski and Blakemore introduced fabric prostheses for replacement of the abdominal aorta. High rates of success are reported for tightly woven, crimped, nonsupported Dacron in the abdominal aortoiliac area; and for knitted, noncrimped, supported Dacron prostheses for axilofemoral and femoropopliteal bypass. [4]

Graft Preparation:

            Vascular grafts composed of Dacron are made using a variety of processes including fiber extrusion and fabric formation using knitting and weaving techniques (Silver and Doillon, 1989). Fabrics composed of woven and non-woven yarns are made using standard textile manufacturing processes. Large diameter vascular grafts are supplied in both porous and non-porous forms. The porosity of these fabrics is related to the number of fibers present, their diameter and the presence of a type I collagen or albumin coating to seal the pores between the fibers. [5]

            On a microscobic level the pores between the fibers in vascular grafts promote tissue ingrowth from the tissue implant interface which leads to adhesion to the surrounding tissue. Fabrics with pores 10-45 mm in diameter promote optimal tissue ingrowth and rapid wound healing. The liquid permeability of vascular grafts is expressed in ml/ cm² per minute and relates the rate of  water movement through a fabric at a pressure of 120 mm Hg. Porosity is typically measured by pressurizing tubes of the material to be tested. Most prostheses display porosities of 2300-5300 ml/ cm²; however, with the recent introduction of coated grafts the rate of liquid penetration through the graft is significantly decreased without adversely affecting tissue ingrowth. Traditionally, a balance of blood loss due to the porous nature of grafts was required to support rapid tissue ingrowth. Other parameters of critical importance are biocompatibility, blood compatibility, graft stiffness, fatigue lifetime and handling characteristics. All of these properties depend on the type of fiber used and the geometry of fibers in the finished textile product.

            The properties of the textile product reflect the individual yarns used to make a fabric. Yarn properties in turn reflect the number and twist of individual fibers as well as the material properties of the polymer used. Continuous fiber is termed monofilament, which differs in physical properties from yarn composed of similar polymeric materials. Physical properties of a fabric are also influenced by the presence of crimp, degree of twist in the yarn, the direction of the fiber axis, and the yarn diameter.

            Fibers can be heat processed to form a regular zig-zag (crimp) that results in a low modulus region to the stress-strain curve. In addition, they can be twisted together to form yarns that resist deformation in proportion to the twist. Yarn can be processed into fabrics by either weaving or knitting. The direction of the fiber with respect to the axis of the prosthesis can be either parallel (warp direction), circumferential (weft) or wrapped at an angle (bias) to the axis. Increased degrees of extension can be obtained along the prosthesis axis when the yarn is incorporated in the bias direction.

            Woven structures are constructed by criss-crossing layers of fibers that are orthogonal to each other. The yarns pass over and under each other creating interfiber friction which increases the bending stiffness of a fabric. The ability of a fabric to drape or lie smoothly over a surface is related to the diameter or denier and the fiber stiffness. In contrast, knitted structures are prepared by looping the yarn with a moving needle and then interconnecting the loops to form a continuous structure. As a result knitted fabrics tend to be highly porous and have lower ultimate tensile strengths compared to woven ones.

            During surgery, porous vascular grafts are pre-clotted with the patient’s own blood to decrease the leakage, once the graft is sutured in place. During pre-clotting, proteins such as albumin and fibrin are integrated along with blood cells into the wall of the prosthesis. The material is replaced by connective tissue.

Medium and Small Diameter Arterial Replacement

            In contrast with the success that has been achieved using knitted, crimped unsupported (external support) Dacron prostheses for abdominal aorta and iliac artery replacement, reduced success rates have been measured for these grafts as bypasses for axillofemoral and above-knee femoral popliteal, below-knee femoropopliteal and below-knee femorotibial sites (less than 50 % patent). These grafts were modified to reduce kinking by adding external support to non-crimped, knitted Dacron prostheses. Increased numbers of patent grafts were reported after these modifications; however, today about three-quarters of the grafts used to repair diseased arteries in the leg are made of expanded polytetrafluoroetylene, PTFE.

            Microporous expanded PTFE is currently the most successful graft material for small diameter arterial reconstruction. After evaluating various modifications of expanded porous PTFE, it was determined that a pore size of 20-30 mm was optimal with external support. This provided the best handling and clinical success rates. Above the knee these grafts were as effective as autografts up to 30 months; below the knee, they were less effective.

Endothelial Cell Seeded Vascular Grafts

            Research and clinical studies have led to the conclusion that saphenous vein autografts are preferred blood vessel substitutes for peripheral vascular reconstructions. However, because this graft is a biological product, person-to-person variation in length and varying availability due to the presence of diseased segments or prior surgical intervention, limit its use in patients. Small diameter vascular prostheses composed of synthetic materials have had limited success not only because of the problem of blood compatibility but also because of the technical problems encountered during graft insertion into vessels with low flow rates. Several research groups have reported improved small diameter vascular graft performance as a result of seeding prostheses of Dacron and PTFE with autologous endothelial cells. A report on PTFE grafts with pore sizes between 28-52 mm, seeded with autologous endothelial cells, show higher mean flow rates than did non-seeded grafts. The study

also concluded that endothelial cell seeding of small diameter PTFE vascular grafts improved patency and the amount of thrombus-free surface areas were greatest when the pores were 40 mm in diameter; however, neither endothelial cell seeding nor pore size affected the performance of PTFE grafts under conditions of reduced flow.

            Another approach to fabrication of small diameter vascular replacements involves use of a variety of polyurethanes. Consisting of a fibrous conduit spun on a rotating mandril. However, animal and clinical results have been disappointing and a recent report indicates that occlusion of polyurethane grafts is a result of a hyperplastic response. [4]

DACRON PROSTHESES

            DeBakey and co-workers are credited with the introduction of Dacron grafts. These conduits are currently fabricated in one of two ways. Multifilamented Dacron yarn is either woven or knitted. Woven grafts are fabricated using a simple over-and-under pattern of threads. This produces a graft with very limited porosity, preventing a need for preclotting with non-heparinized blood prior to implantation. Unfortunately, these grafts have little stretch and are minimally pliable. For this reason some surgeons find woven Dacron prostheses difficult to work with. In knitted Dacron grafts, threads are looped to form a continuous, interconnecting chain. Knitted fabrics have a certain amount of stretch in all directions and more versatile in degree of porosity. Velour is a variant of knitted cloth in which a soft plush surface is created by extending loops of yarn upward at right angles to the surface of the fabric. The velour finish may be placed on the exterior, interior, or both sides of the conduit. Additional elasticity is imparted and porosity varied. Dacron prostheses may be classified according to their porosity as either microporous (woven) or macroporous (woven, knitted, or velour). It is important for grafts to be at least modestly porous to allow incorporation of the outer wall and organization of the luminal surface. Porosity is dependent on graft pore size and can be expressed as an absolute measure in microns, or as an amount of water flowing through a centimeter of graft per unit of time at standard pressure.

            The most desirable graft fabrications are those that combine maximal porosity with good handling qualities. Although an “ideal” graft porosity has been cited as being 10,000 ml/ min, the practical upper limit of graft implantation porosity is less than 4000 ml/min. Most conduits in clinical use have porosities in the range of 1200 to 1900 ml/ min. The higher the porosity, the greater the risk of hemorhage. Bleeding from grafts with high implantation porosities is difficult to control. Attempts have been made to create grafts having “large” pores filled with a biodegradable material that is slowly broken down by fibroblasts, but no clinically useful large-pore “coated” conduit has been developed. The large pores were found to reduce the strength of these grafts, predisposing them to aneurysm formation.

            More recently, collagen-coated knitted conduits of “normal” pore size have been successfully introduced. These treated fabrics have handling qualities indistinguishable from nonimpregnated grafts and have been proven nearly impervious to blood. The need to preclot these knitted prostheses prior to implantation

is eliminated. Knitted Dacron grafts, similarly coated with albumin, proved only relatively impervious to blood, required rehydration prior to implantation, and were somewhat stiffer than nonimpregnated conduits.

Host Responses

            Shortly after a Dacron graft is exposed to arterial blood flow, a predictable sequence of events occurs. Fibrin is absorbed onto the inner graft surface. In large-  calibre prostheses subjected to high volume flow this fibrin layer usually remains thin, typically less than 1 mm. In smaller-calibre Dacron prostheses in low flow environments, the fibrin layer may increase in thickness, ultimately promoting graft occlusion.

            Immediately following implantation the outer lining of the graft is also completely encapsulated with fibrin. Organization of the outer fibrinous capsule begins within 2 days. This capsule is comprised of an outer layer containing nutrient vessels, a middle layer of dense collagen, and an inner layer, in contact with the graft and consisting of organized tissue and foreign body cellular infiltrate. This encasement of the conduit causes a loss of preexisting elasticity or compliance mismatch. Platelets adhere to the intraluminal fibrin layer and are one of the major constituents of thrombus formed on the surface of a prosthetic conduit. Using indium 111-labeled platelets, numerous investigators have shown platelet deposition and activation with Dacron grafts. Moreover, Dacron activates both the complement and coagulation cascades and also triggers the release of superoxides and thromboxanes by leukocytes in an in vitro setting.

Indications  

            The crimped, knitted Dacron graft is the most common conduit used for infrarenal aortic replacements. In situations where bleeding is a significant possibility, woven grafts are usually chosen. Such situations may include patients with coagulopathy or those with ruptured aneurysms who have already sustained a large blood loss. Most surgeons also use woven grafts preferentially when performing suprarenal aortic reconstructions. While some published reports show reasonable long-term patency when Dacron is used in the above-knee, femoropopliteal position, early thrombosis in bypasses across the knee joint seems to be the rule.

            Knitted porous prostheses function satisfactorily in axillofemoral and femorofemoral locations. For the purposes of extra-anatomic arterial reconstructions, the graft is available with a supporting continuous coil densely incorporated into its weave. Dacron has also proved useful for patch angioplasty.

Complications 

            Infection occurs in approximately 2 % of synthetic graft implantations. Grafts may become infected by direct inoculation with microbes during the operative procedure, as well as by hematogenous seeding from transient bacteriemias months or years after implantations. Dacron should rarely, if ever, be implanted in infected or

potentially infected locations. Recent reports, however, have suggested that acceptable results can be obtained when placing Dacron prosthetic graft in the anatomic bed of a primarily infected aortic aneurysm or when using Dacron prostheses for the in situ replacement of infected abdominal aortic grafts. This policy has not been widely embraced by all vascular surgeons, however, and is subject to closer scrutiny.

            Structural failures, including aneurysm formation, are rare occurences and are most commonly due to mechanical failures in fabrication. Aneurysm formation should not be confused with graft dilatation, since a 10 to 20 % increase in graft diameter is common following implantation. Additional complications do occur and include the occasional formation of seroma around some grafts, and the development of anastomotic pseudoaneurysms. These occur at a rate of 1 to 4 % and are multifactorial in etiology. [1]

New Bioactivation Mode For Vascular Prostheses Made of Dacron Polyester

         The performance of vascular prostheses, especially those of small calibre, still presents several problems. The use of polyesters such as the well-known polyethylene terephthalate and the more recent use of polytetrafluoroethylene or polyurethanes, provided a marked improvement without, however, completely resolving the problems concerning patency and, in particular, bioactivation. These synthetic materials, knitted or woven, are porous and must be coated by the attachment of proteins such as albumin, type I collagen or fibrin, which also improve the biocompatibility. However, with regard to the small calibre, the use such procedures does not allow the device to be haemocompatible or even bioactive.

            With a view to more closely reproducing the in vivo situation, an original and attractive solution would consist of interacting these polymers with the components of a vascular subendothelium. For other purposes, a new artificial connective matrix closely resembling a subendothelium layer has been developed. This biomaterial results from the particular interaction of elastin or, better still, elastin-solubilized peptides (ESPs) with type I+III collagens under very precise conditions. These biological conditions and the addition of some connective proteins and glycosaminoglycans (laminin, fibronectin, type IV collagen and heparan sulphate) conferred on the material produced: i) a composition very close to a subendothelium, ii) a well-organized thin-layered structure, iii) the qualities of endothelium tissue, i.e. haemocompatibility and the capability to allow the growth of endothelial cells.

This subendothelium-like structure is fragile. With a view to improving its strength, this matrix could be linked to a Vicryl lattice using the acyl-azide chemical process as shown above.

            Nevertheless, the biodegradibility of Vicryl and its watertightness due to its hydrophilic character prompted us to choose Dacron polyester which is already used in vascular prostheses. A recent publication reports the ability of the Dacron polyester to be covalently modified by proteins under mild chemical conditions.

            Dacron is covalently bonded irreversibly with a subendothelium-like structure and the subsequent development of a vascular prosthesis.

            Chemical activation of Dacron polyester has three stages: activation of the polyester Dacron, chemical bonding of ESP, and adsorption of collagen together with other components such as laminin, fibronectin and heparan sulphate.

            The well-known and currently employed acyl-azide method of polyester activation was performed to modify the Dacron. As shown in the figure, this chemical process involves the hydrazinolysis of ester bonds and their conversion to azide, followed by reaction with the ESPs. The main aim was to monitor this reaction to determine the resulting bioactivity and strength.

            Several strategies to improve the blood compatibility of polymers are currently being developed. The different techniques used can be classified by three general approaches.

            First, methods that increase the hydrophobicity of polymer surfaces may reduce the interaction with blood components. Second, providing a biomaterial surface with a biologically active compound offers numerous possibilities to prevent thrombus formation. Third, to prevent polymer surfaces being recognized as foreign surfaces, many researchers try to disguise them. An example of this technique is material endothelization, the endothelium being considered as a perfect blood compatible material.

            There is a possibility of linking this artificial matrix to a Vicryl lattice, using a mild chemical method of polyester activation. This technique was applied to a Dacron coating, and ESP immobilization was obtained at the same biochemical conditions of temperature and concentration. However, Dacron is less sensitive to the chemical process, the hydrazine concentration and hydrazinolysis time being greater.

            By scanning electron microscopy, it can be detected that bound ESP is still able to interact with type I+III collagens, this interaction depending on the conditions of reaction, namely, collagen concentration, absence or presence of proteoglycans and connective proteins. However, the coating distribution on the polymer seems to be irregular: certain parts are covered by a thin layer of the artificial connective matrix, while in other parts, fibers are totally embedded in a thick meshwork of material. This situation probably results from the difficulty of reaction between a solid phase and a viscous solution. Modifications in the conditions of ESP and collagen fixation (concentration, temperature, incubation type) and using polyester thread in place of samples of prostheses could provide a method of improving this distribution. The further development of tubular devices could be made without extensive difficulties. [6]

POLYTETRAFLUOROETYLENE PROSTHESES

            Although the early woven Teflon prostheses did not deteriorate following implantation, the limited porosity, poor handling, characteristics, and the tendency to fray at suture lines led to the cessation of its use. A heating and mechanical stretching process was developed in the late 1960s creating a nonfabric modification: expanded reinforced polytetrafluoroethylene (ePTFE). These conduits enjoy all the advantages of fabric Teflon conduits without the aforementioned disadvantages. Being polymers of carbon and fluorine, these grafts are highly electronegative and hydrophobic. They consist of transversely oriented, plate-like portions of PTFE nodules interconnected by multiple PTFE fibrils running longitudinally. Void spaces surrounding the node-fibril structure occupy 85 % of the volume of the graft wall. The resulting graft porosity is essentially blood-tight at implantation, but does allow cellular ingrowth. Pore size is thought to be the primary determinant of tissue ingrowth and is determined by fibril length variations. In most ePTFE conduits, pore size averages 20 to 30 mm. ePTFE is durable, does not dilate, holds suture, and does not undergo biologic deterioration. These grafts can be sterilized with steam or gas without structural or mechanical compromise. The graft is available with reinforcements of plastic external supporting rings to prevent compression of the grafts when crossing joints or pressure areas. Unfortunately, the grafts are non-compliant. EPTFE was introduced clinically in 1971 when Norton and Eiseman used these grafts to replace the portal vein during pancreatectomies. They were first applied as arterial substitutes in 1976 when Campbell et al. Used ePTFE for femoropopliteal and femorotibial bypass in patients without ipsilateral saphenous vein and limbthreatening ischemia.

Host Responses

                Healing events following implantation are quiet similar to those occuring following the implantation of Dacron conduits. These events are modified by both graft porosity and wall thickness; the thickness of tha intraluminal fibrin lining appears to vary inversely with porosity and wall thickness. As with all human prosthetic grafts, endothelial cells do not line its luminal surface except at or near anastomotic sites where ingrowth from native vessels occurs. This graft is less attractive to platelets and less perturbing to the coagulation and complement pathways than Dacron. When used in intermediate-sized vessels (6 to 10 mm), it is commonly believed that ePTFE provides equal if not superior patency vs. Dacron. A recently published retrospective review that demonstrates improved patency of Dacron grafts vs. ePTFE when used for femoropopliteal reconstructions casts doubt on this unfounded assumption. When used in small-diameter vessel reconstruction, ePTFE grafts are consistently hampered by the development of intimal hyperplasia leading to graft failure. Investigators have demonstrated that endothelium and smooth muscle cells in the perianastomotic locale continue to proliferate despite complete endothelial coverage. It has been postulated that the implanted graft functions as a continuing site of tissue perturbation, triggering local platelet aggregation and the release of platelet-derived growth factors (PDGF) and PDGF-like mitogens. It has been suggested that intimal hyperplasia adjacent to perianastomotic regions could be reduced by antiplatelet therapy, however, the value of postoperative treatment anti-platelet agents is unproven.

Indications 

            ePTFE grafts have been used most commonly in the femoropopliteal position. The graft seems to be a reasonable alternative conduit in patients who require femoropopliteal bypasses above the knee. It also seems to be reasonable choise when patients are found to be without autologous vein, and have life expectancies of 2 years or less. Extension of these grafts below the knee or insertion into a popliteal segment having poor distal runoff are considered contraindications to the use of ePTFE. Nevertheless, ePTFE has been used in these circumstances with mixed results. While arterial reconstructions in the infrainguinal tree have been the most common indication for the use of this conduit, a bifurcated large-diameter ePTFE prosthesis has recently become available allowing the use of ePTFE for aortic reconstruction. Additionally, ePTFE has been used satisfactorily for patch angioplasty, for visceral vessel reconstruction, and for brachiocephalic arterial reconstruction. The graft is possibly a preferable conduit vs. autologous vein in subclavian-carotid bypasses. Extra-anatomic bypasses in the axillofemoral, axillobifemoral, and femoral-femoral positions have routinely been performed with ePTFE.

            Patency rates are equivalent to those achieved with Dacron grafts. EPTFE has become the graft material of choice for the creation of prosthetic arteriovenous fistula for hemodialysis. Small-caliber ePTFE has been used in the correction of congenital cardiac defects including atretic aortic lesions and the creation of systemic to pulmonary shunts. Only on rare occasions has ePTFE been used for aortocoronary bypasses in patients with unsuitable autologous conduits. Large-diameter ePTFE grafts have been used successfully in the venous circulation to replace the vena cava, iliac, and visceral veins. In fact, ePTFE grafts have been used with good success in the performance of portosystemic shunts for the control of portal hypertension. In the lower extremity, small-calibre ePTFE has been used for veno-venous bypass (usually with adjunctive arterial-venous fistula).

Complications

            Similar to that found when implanting Dacron grafts, one of the major risks of ePTFE implantation is infection. Although it has been demonstrated in canines that ePTFE may be more susceptible to bacterial infection than Dacron, many clinicians suspect otherwise. Additionally, in the presence of infection, ePTFE anastomoses have a lower incidence of disruption. Some have theorized that this phemomenon is secondary to an enhanced healing capacity of ePTFE relative to other prosthetic conduits.

            Pseodoaneurysm formation can result from many factors including the breakdown of an infected graft to artery suture line or may follow partial or complete graft laceration during hemodialysis. While aneurysms were documented arising from the body of early ePTFE grafts, they are uncommon occurences in presently available conduits reinforced with either an additional outer circumference of ePTFE or by the thickening of graft walls. Though ePTFE is less thrombogenic than other prosthetic grafts, both early and late graft thrombosis does occur. Technical factors are usually responsible for early graft failure. In a number of instances contributing factors like

poor inflow, poor outflow, and postoperative hypotension negatively influence graft patency. Late graft thromboses are often due to pseudointimal hyperplasia at the anastomotic site. An additional cause of late failure is the progression of distal arteriosclerotic disease. [1]

Lipid Uptake In Synthetic Vascular Prostheses Explanted From Humans

            EPTFE vascular prostheses are widely acknowledged as the most reliable synthetic arterial substitutes in peripheral vascular surgery when the autologous saphenous vein is not available. However, the absence of endothelialization is a serious limitation, as thrombosis is the most frequent complication to occur with these prostheses. Following implantation of a synthetic vascular prosthesis, endothelial coverage of the luminal surface is the expected outcome. However, several retrieval studies have analyzed prostheses explanted following late complications and have shown that a çomplex sequence of events does occur, beginning with plasma protein adsorption and plasma blood cell deposition., and culminating with the total or partial encapsulation of the prostheses. Despite the tremendous efforts made to promote or enhance endothelial coverage and neo-intima formation on the luminal surface of synthetic vascular prostheses implanted in humans, endothelialization is rare and, when observed, is limited to within only a few millimeters of the anastomosis.

            It is common knowledge that atherosclerosis in humans is an evolutive process that begins to develop after the first 10 years of life. Moreover, atherosclerosis is generally considered as a human disease that affects large, medium, and small-size arteries, with a predilection for the critical arterial beds such as the coronary, cerebral, and peripheral arteries, where the intima thickens by progressive accumulation of both cellular and non-cellular elements. In healthy vessels, lipids have an important role in the metabolism of the arterial wall as they diffuse across the endothelium from the intima to the media, returning to the blood circulation by flowing through the vasa-vasorum. In atherosclerosed vessels, lipids and other blood elements, including smooth muscle cells, are known to be the most important components of the onset of atherosclerotic plaques. In this context, the conditions dictated by atherosclerotic arteries adjacent to synthetic vascular prostheses must be considered in the healing and neo-intima formation processes of these prostheses. Therefore, lipid deposits, present in all atherosclerosed arteries, may have a major role also in the progression of an atherosclerotic process within synthetic vascular prostheses.

            As far back as 1958, different studies have been conducted to investigate the effects of lipids on synthetic arterial prostheses. It has been reported that chronic lipid hypercholesterolemia produced changes in Nylon and Orlon arterial prostheses implanted as canine thoracic aorta bypasses. In 1986, Walton et al. Investigated the healing response of 32 Dacron and seven Teflon synthetic arterial prostheses explanted from humans by immunologic and histologic chemical methods. They observed that over 2 years of implantation the reactive healing tissue invariably contained bonded lipid in a distribution resembling that seen in atherosclerotic arteries. The authors concluded that an insudative atherosclerotic-like process, indistinguishable from true atherosclerosis, occurs in synthetic prostheses when

implanted in humans. Some years later, in 1991, Chignier et al. Used histological methods to show that lipid uptake was more important in synthetic arterial prostheses implanted as canine aortic bypasses than in actual canine arteries. Moreover, in our preliminary study, Fourier transform infrared spectroscopy was used to investigate lipid uptake in 104 ePTFE arterial prostheses explanted from humans, and it was discovered that unsaturated fatty acids strongly bonded to the polymer structure were present in all of the retrieved ePTFE devices. 

In general, the luminal surface was found to be poorly healed with only a thin fibrin layer and showing virtually no evidence of collagen infiltration into the microporous structure. In particular, only 1.4 % of the explants showed the presence of endothelial-like cells, whereas 35 % contained cholesterol deposits and 81 % show evidence of lipid deposition. These findings, although qualitative, have therefore clearly demonstrated that ePTFE arterial prostheses are exposed to lipid uptake when implanted in humans.

Among the parameters investigated, only the duration of implantation, the sex of the patient and the diameter of the prosthesis have been shown to have a significant effect on the lipid concentration measured on the surfaces of the explanted prostheses.

The lipid concentration in the prosthesis increases with the duration of implantation. In addition , it is known that lipid adsorption is rapid because lipids have been observed on the prosthesis surface as early as one day after implantation. On this basis, the strong affinity of lipids for the high hydrophobic PTFE polymer is believed to promote lipid deposition on surface and stimulate lipid transport across the wall of the prosthesis.

Men are more prone to than women to developing atherosclerotic plaques in their native arteries. Coronary arterial diseases in the USA appear 10 years later in women than they do in men.

ePTFE prostheses with a 6 mm diameter are more prone to lipid uptake than those with an 8 mm dimeter. This may be related to the diminished patency rates reported to be substantially lower for small-diameter devices compared to large-diameter synthetic prostheses. In fact, the primary failure mode of the small-diameter synthetic prostheses is thrombus formation, and thrombosis has been known for many years to be a contributing factor in the pathogenesis of atherosclerosis.

Although the incidence and severity of atherosclerosis appears to be age related, it is not actually. Lipid deposition on ePTFE vascular prostheses does not relate to the patient’s age.

The effect of different implantation sites on lipid uptake was also investigated. No relationship was found between the different sites of implantation and the amount of lipid uptake.

It can be demonstrated that lipid deposition occurs in ePTFE vascular prostheses, and that the level of deposition is dependent on the duration of implantation, the sex of the patient, and the diameter of the prosthesis. On one hand, prostheses implanted for longer period of time clearly retained more lipids onto their surface. On the other hand, A higher lipid concentration was observed for prostheses implanted in women with respect to those that were retrieved from men. No relationship was established between lipid uptake and the age of the patients, the site of implantation, or reason for explantation. Due to the strong affinity between ePTFE and lipids, the role of lipids has to be considered when attempting to improve the healing response of these devices. [7]

Vascugraft Polyurethane Arterial Prosthesis as Femoro-popliteal and Femoro-peroneal Bypasses In Humans: Pathological, Structural and Chemical Analyses of Four Excised Grafts

Polyurethane application in vascular surgery as small or medium-caliber blood conduits has been limited by low patency rates and uncertain long-term biostability in animals. Nevertheless, because of their outstanding physical and chemical characteristics, polyurethanes continue to be regarded as suitable candidates for use as small-diameter prostheses and remain the subject of numerous investigations.

The Vascugraft prosthesis is made from a novel polyurethane material with a microporous fibrous structure. The Vascugraft prosthesis is made from a unique polycarbonate urethane. The fibres are formed by a spraying technique from a solution of polyurethane projected towards a rotating mandril onto which they are deposited. A fibrous tubing is therefore created with the communicating pores in the wall  structure. The resulting prosthesis is an elastic tube with a smooth, white appearances. It is easy to suture and cut cleanly without flossy edges. The Vascugraft prostheses were retrieved from four patients. The reason for implantation in all cases was limb salvage. The Vascugraft was implanted as a femoro-popliteal bypass in three cases, and as a femoro-peroneal bypass in one case. Implantation ranged from 21 to 358 days, and in all cases the grafts were surgically excised due to thrombosis.

Virgin Vascugraft prostheses have a non-woven microfibrous structure. On the internal surface, some of the microfibres were fused together forming thick bundles. These bundles were less common on the external surface. Some discrete patches were also observed on both surfaces, probably caused by large droplets of polyurethane accumulation on the mandril during graft manufacture. Removal of the monofilament from the external surface, resulting in peeling of the texture. The specimen explanted one year after implantation showed significant damage on the external surface. In addition to the broken microfibres, there was evidence of material loss causing parts of the external surface to appear grossly uneven. There was a tendency for the damage to be concentrated in areas where non-fibrous patches were present. Moreover, in the vicinity of the non-fibrous patches, the microfibres were broken at regular intervals.

The molecular weight of the explanted and cleaned Vascugraft prostheses did not vary significantly, even after one year of implantation.

All of the grafts failed due to thrombosis occurring distally as the result of graft kinking in one case, graft-artery mismatch in one case and poor run-off in two cases. The pathological observations carried out on the Vascugraft prostheses were similar to those with other synthetic vascular grafts such as PTFE and Dacron, when implanted as lower-limb substitutes. Inadequate inflow or poor run-off is possibly the most frequent cause of failure and is consecutive to the presence or development of distal obstructive disease. The luminal surface is always covered by thrombotic material, and complete healing is never achieved. Soft tubing, such as polyurethane and PTFE grafts, easily kinks and traumatizes, and therefore promotes thrombus formation in the vicinity of the graft flexion.

Short-term implantation did not affect the surface morphology or the basic chemical structure of the Vascugraft prostheses. No damage to the microfibres was found on either the external or luminal surface of the prostheses up to 46 days of implantation. Furthermore, the molecular weight was found to be within the normal range in all four explanted Vascugraft prostheses.

In the prosthesis implanted for 358 days, the deterioration of the microfibrous structure observed on the external surface casts doubt on the long-term stability of this vascular prosthesis. Similar surface changes in the same prosthesis were also found after one year in a canine thoraco-abdominal model. Although these deteriorations were detected only on the surface, the possibility of any further deterioration deeper in the graft wall may compromise the long-term performance of the graft.

Small-diameter polyurethane grafts implanted as femoro-popliteal and femoro-peroneal bypasses were investigated after explantation. The grafts were retrieved after thrombotic occlusion in all cases due to kinking, graft-artery mismatch and poor run-off. In contrast to animal implants, no internal capsule had developed: however, endothelial-like cells were observed at the anastomotic site of one harvested graft.

The polyurethane vascular prostheses were stable for short-term implantation. Chemical and morphological characteristics of the prostheses were maintained up to 46 days. Although the bulk properties of the polyurethane material remained stable 358 days after implantation, a deterioration of the surface microfibres and a significant reduction of carbonate groups were observed on the outermost surface of the grafts.

It is therefore important that further improvements be made before this prosthesis can be reintroduced as a blood conduit for human implantation, which is very unlikely. [8]

An Albumin-Coated Polyester Arterial Graft: In Vivo Assessment of Biocompatibility and Healing Characteristics

Textile polyester vascular grafts sealed with bioresorbable coatings provide attractive advantages for surgeons and patients, because they are ready to anastomose, have the ability to limit blood loss and thus reduce the need for blood transfusion and haemostasis. Experimental collagen- and gelatin-coated grafts were first introduced in the early 1960s but only became commercially available for general clinical use in the 1980s. Albumin-coated grafts (ACGs) have also been the subject of development since the mid 1970s. In recent years, glutaraldehyde cross-linked human albumin impregnated in polyester knitted grafts has demonstrated short-term antithrombogenic and haemocompatible properties that were superior to those of preclotted polyester grafts. Moreover, the ACG has been extensively investigated in various animal models as a thoracic aortic graft, an aortiliac bypass graft and an infrarenal aortic substitute.

The performance of the ACG was assessed in terms of its ability to avoid blood loss, its patency, morphology, histopathology, prostoglandin secretion and surface thrombogenicity, as well as the rate of resorption of the cross-linked albumin. The impact of the albumin on the healing performance was determined by including a second parallel series of implantations over the same 6-month time frame using the uncoated , preclotted, polyester DeBakey Vasculour II (V-II) prosthesis. The other part of the present investigation involved an evaluation of the biocompatibility of the same two prostheses using a murine intraperitoneal implant model. Of particular interest was the host’s immune response to the bioresorbable coating, and changes in peripheral T cell populations in response to the intraperitoneal implants were evaluated by cytofluorometry. Histopathological analyses of the tissue response and acid phosphatase activity at the implantation sites were also conducted.

The Bard DeBakey V-II vascular prosthesis served as the non-coated control graft. The V-II has a warp-knit velour structure of texturized Dacron polyester yarns. The graft has acceptable bursting strength when compared to other commercially available warp-knitted prostheses. The V-II prosthesis as received had been sterilized by ethylene oxide.

DeBakey albumin-coated graft (ACG) is a V-II graft impregnated with human albumin and cross-linked with glutaraldehyde. The albumin coating fully impregnates and seals the walls of the textile prosthesis. The graft is then packaged in saline and sterilized by gamma irradiation. The graft is non-pyrogenic.

The albumin-coated vascular graft (ACG) and its uncoated polyester substrate, the Vascular II (V_II), were evaluated in terms of biocompatibility and biofunctionality using two in vivo animal studies. Biocompatibility and immunoreactivity were assessed by implanting intraperitoneally in the rat small segments of the ACG and the V-II graft and harvesting them with their surrounding tissue 3 days, 1, 2 and 4 weeks later. Cytofluorometric determination of total T cells has revealed that no significant difference in any of the T cell populations was found between the ACG and the V-II graft. The cellular reactivity of the ACG in terms of acid phosphatase activity at the implant side was significantly greater at 3 days but not

at longer periods. Biofunctionality was evaluated by implanting both grafts as a thoracoabdominal vascular bypass in dogs for 11 different periods ranging from 4 hours to 6 months. The rate of albumin resorption was such that traces were still present at 1 month, but no longer observable at 2 months. Tissue incorporation into the graft wall was earlier for the V-II (2 weeks) than for the ACG (4 weeks), which showed complete encapsulation, tissue incorporation and endothelialization after 2 months in vivo. Only small differences were observed between both grafts in terms of platelet and fibrin uptake on the luminal surface.

            Vascular grafts, which have been impregnated with bioresorbable proteins and so eliminate the need for preclotting, are of particular interest to vascular surgeons because of their ability to reduce blood loss. This is of considerable importance in cases of emergency or where patients are under general heparinization. There was no blood loss from any of the ACG prostheses evaluated in this study. The ACG demonstrated good handling and suturing characteristics, and proved to be a flexible and effective leak-proof prosthesis. There was no indication of any T cell subset modifications in the circulating blood. Once the albumin is resorbed, the healing is able to progress in a similar fashion to the preclotted Dacron prostheses. [9] 

OMNIFLOW II VASCULAR PRODUCTS

The Omniflow II Vascular prosthesis is a unique combination of ovine collagen with an integral endoskeleton of polyester mesh. The prosthesis is chemically stabilised, which, along with the integration of mesh and collagen, imparts superior mural strength. The haemocompatible flow surface is comprised of stable biological material.

 The Omniflow II prosthesis is available in straight and curved configurations, however it can  be manufactured in a number of configurations to order depending on the application.

                                                         [10]

FUTURE DIRECTIONS

            Although neither of the currently used prosthetic grafts fulfills criteria for an ideal arterial substitute, both Dacron and ePTFE have acceptable patency when used for large- (more than 10 mm) and intermediate-sized vessel (6 to 10 mm) reconstructions. When used to reconstruct smaller-diameter arteries in peripheral systems where low flow, high resistance, and intimal proliferation prevail, these grafts generally fail. Central to the further advancement of vascular surgery is the development of a successful small-diameter synthetic prosthesis. This graft of the future must be a better mimic of the blood vessel it replaces. Biologic manipulations must be found to interfere with the untoward biologic reactions that result in increased thrombogenicity and proliferative activity. A plethora of coatings have been incorporated on existing graft materials, with mixed degrees of success, in an effort to improve biocompatibility. These include antibiotics, heparin, plasma TFE (the gaseous form of Teflon), and growth factors. The graft of our future should be seeded with endothelium. A greater understanding in cell physiology with the resulting regulation of the seeded endothelial cells is also required. Current work offers hope that a prosthetic material will be found that is readily available, has a better resistance to thrombosis and infection, and is suitable for the replacement of small-calibre blood vessels.