Initially designed for use in the textile industry, Dacron is designed to be made into small strands that are either woven into a mesh or knitted into a porous fabric. Dacron was developed by the British chemists J.R. Whinfield and J.T. Dickinson in 1941. When fabricated in a woven fashion, this prosthesis has small pores. Knitted grafts have greater porosity (larger pores), which allows for greater tissue in growth of surrounding tissue and incorporation of the prosthetic material. Knitted fabric was introduced in the late 1950s and is composed of yarn interloped around each other in either a longitudinal (warp-knit) or transverse (weft-knit) fashion. Clinically, the difference is noted by the tendency for the weft-knit fabric to loosen. The knitted fabric offers greater ease of handling but has a tendency to elongate over time. Knitted grafts initially required preclotting prior to implantation due to the high porosity. An inner albumin or gelatin coating was subsequently added to decrease seepage of blood after implantation, without compromising the tissue in growth advantage seen with knitted grafts.

This material is biocompatible, resilient, flexible, durable, and resistant to biodegradation. Although not biodegradable, Dacron grafts tend to dilate over time. The fabric has been manufactured with a crimping design, which allows the graft to have greater ease of flexibility without causing kinking. This improvement in design for ease of use may have some detrimental effect by creating an uneven internal lumen, which theoretically could lead to greater thrombotic tendency. Dr. DeBakey used a sewing machine in 1952 to create a tube of Dacron for aortic reconstruction. Since that time, Dacron has become the most utilized prosthetic for large vessel (>6 mm) reconstruction.

In 1952, Vorhees used the first vascular conduit derived from a polymer known as Vinyon “N.” This material underwent significant deterioration during the sterilization process. Dr. Roy Plunkett developed PTFE in 1938 while working for the DuPont Corporation. It was first marketed under the Teflon trademark in 1945, but researchers at the W.L. Gore Corporation subsequently developed expanded PTFE, which possesses properties that made it more desirable for vascular reconstruction. These properties included greater porosity and compliance. Since the first report of use in the lower extremity with ePTFE by Campbell in 1974, it has become the most widely utilized prosthetic conduit for revascularization below the groin level. ePTFE is nonbiodegradable, and has an electronegative charge that is felt to aid in resisting thrombosis. The size of the pores in the fabric has been studied to optimize the endothelialization of the graft. Currently 30 μm appears to be the optimal pore size for maximizing endothelialization, without excessive oozing from the conduit. External supporting rings have been added to ePTFE grafts to minimize kinking and compression of the material. An external wrap was added to mitigate the tendency for degradation and prevent aneurysmal deterioration. In addition, some grafts have been lined with an antithrombogenic carbon coating in an effort to enhance patency rates. Most recently, heparin has been bound to the lining of the ePTFE graft and utilized clinically. Other materials have been used to coat ePTFE grafts including polypropylene sulfide–polyethylene glycol (PEG), hydrophilic acrylic polymers, which release salicylic acid, and an elastomer polycitrate using a spin-shearing technique. Although the results of these materials are preliminary, there does appear to be some promise to these and other polymeric coatings.

The advantage of polyurethane can be found in its compliance. The production of polyurethane allows for greater radial compliance, which should lead to a decrease in neointimal hyperplasia. The material flexibility can be modified while maintaining tensile strength. Although biocompatible, the initial use found that a significant amount of biodegradation occurs. In addition, the material tends to elongate over time, causing significant dilatation of the conduit. Currently, carbonate-based polyurethane has shown some promise in increasing the stability of the graft with prevention of aneurysmal degeneration. Other substances have been used to increase stabilization, but without success. Unfortunately, some polyurethanes have been shown to have carcinogenic effects in laboratory animals due to breakdown products during degradation.

Polypropylene is a biostable, inert, and crystalline material in nature. Initially designed and utilized as suture material, experimental evidence suggests promise for use as an arterial conduit. Like polyurethane, polypropylene has been found to elongate over time. It has been evaluated using composite polymers of polyglactin and polydioxanone (PDS) with some encouraging initial patency results. In light of the disadvantages seen in polyurethane and polypropylene grafts, several hybrid compounds are currently under investigation.

Biodegradable polymers degrade in the human body as a result of chemical reaction with the surrounding tissues and result in smaller fragments. The rate of degradation is dependent upon the pore size. Smaller pores illicit an inflammatory reaction and larger pores result in seepage of blood through the material. Conceptually, these
polymers could be placed as a scaffold for ingrowth of human cells that would slowly replace the prosthetic polymer as it deteriorated and function as an arterial conduit. Several single and multiple compound polymers have been studied with varying results. The major advantage to compound polymers (biodegradable and nonbiodegradable) appears to come with prevention of aneurysmal degeneration.

Acellular vascular graft development has been investigated with xenogeneic materials. These materials are treated to remove the cellular components with preservation of the acellular scaffold. This process has proven to be difficult, but theoretical preservation of normal arterial mechanical properties (i.e., compliance) may be beneficial to patency rates of these grafts.

Prosthetic graft seeding with endothelial or smooth muscle cells within the inner lumen of the conduit may further improve patency rates. The major limitation appears to be the low adherence rate of these cells to the prosthetic, particularly in the pressurized arterial system. Several substances have been used to increase this adherence including fibrin–gelatin, granulocyte-stimulating factor, and fibronectin coating. Carbon deposition, photo discharge, chemical vapor deposition, and plasma discharge technology have also been used, but to date, the formation of the neointima in the prosthetic graft has had limited ability to mimic natural endothelium.