Materials in the Construction of Stent-Grafts

eryone's time is valuable and limited. Again, by working with experts in the business community, a better, more professional and well-funded event will result. Any event, article or advertising will reflect upon the radiologist as a professional. Therefore, quality marketing will only enhance the reputation of the physician for thorough medical procedures. An additional and important component in planning and designing any program is to include the participation of other related medical specialists. Not only does this obtain their buy-in, but it allows everyone to gain a greater appreciation for the unique benefits of various techniques in treating a common disease. Failing to involve other specialists in the initial stages of planning an outreach program could invite direct conflict in which the interventional radiologist is doomed to lose as he or she is a minority in any medical community. For example, surgeons, cardiologists and other medical specialists frequently control hospital committees which deal with such critical issues as credentialing and capital acquisition. Direct conflict, therefore, is detrimental to the long-term growth of an interventional service. Finally, no matter what marketing campaign you select or how carefully you have planned your campaign, significant credibility will be lost if planning has not included systems for handling response to your campaign. It is essential that telephones be staffed, messages responded to promptly, and that your clinical area is able to schedule additional procedures.

Conclusion Interventional radiology is still an unexplored asset in many medical communities. Carefully planned education and marketing efforts to targeted groups are essential to increase visibility. Marketing efforts should include resources already in place within the local medical community as well as those available at regional or national levels. Communication and cooperation with the target groups is essential for any marketing effort to be successful. If the interventional radiologist initiates, organizes, and oversees the marketing efforts, they have accomplished their primary goals: increasing their practice by building visibility, credibility and knowledge within the medical community.

Symposium - Stent Grafts Saturday, February 28, 1998 12:00-5:00 pm 12:00 pm

Natural History and Surgical Repair of Aortic Aneurysms Robert Thompson, MD 12:25 pm

Materials in the Construction of Stent-Grafts

28 Joseph Bonn, MD

THE burgeoning field of endovascular stent-grafts forces us to view two developing technologies, expandable metallic stents, and synthetic vascular grafts, from a unique perspective: neither device is being used in its original manifestation, but both are interacting in an entirely new application. Stents are no longer simply propping open the elastic recoil or dissection of a failed angioplasty but are now anchoring and scaffolding a vascular graft in occlusive or aneurysmal disease, bridging the interface of diseased vessel and synthetic material, or pulsating against excluded new-formed thrombus in an aneurysm. Likewise, vascular grafts are no longer bypassing through perivascular surgical tunnels, but are being stretched beyond their typical diameters and pressed by metallic struts against long segments of endothelial cell- and plaque-lined luminal wall where cellular interactions have yet to be fully elucidated. While vascular stents are a subject clearly more familiar to most interventional radiologists, vascular grafts occupy a niche only of passing familiarity. The modem synthetic vascular graft came into its own with the postwar growth of the plastiCS industry, beginning with Voorhees and colleagues using the parachute synthetic fabric Vinyon-N at Columbia in 1954. By 1957, the Society for Vascular Surgery had outlined the qualities of the ideal vascular graft: porosity, flexibility, durability, compliance, and proper tissue reactivity (or biocompatability). In the ideal graft, many of these features are a compromise: porosity contributes to incorporation and anchoring of the graft material to its environment, but too much porosity leads to difficulties achieving hemostasis, while too little allows a peri-graft fluid or pseudomembranous lining to form, freely disseminating any infection or hemorrhage. Compliance propagates the pulsatile waveform of arterial flow: too little creates a mismatch at the graft-vessel junction which is thought to contribute to vibrations and turbulence leading to neointimal hyperplasia, while too much can lead to graft weakening and aneurysm formation. Current synthetic vascular grafts can be grouped into two categories: textiles or non-textiles. Textiles grafts are made of fibers woven or knitted into cloth; the original textile vascular grafts were constructed of Nylon, Orlon, Ivalon, and Teflon; however all of these polymers failed in vivo tensile stresses. Textile vascular grafts in clinical practice are mostly made of Dacron, the trade name for polyethylene terephthalate (PET). A graft woven of PET fibers possesses minimal compliance, porosity and ease of handling while demonstrating maximal tensile strength and resistance to leakage. Its knitted form, made of a continuous chain of thread, has more optimal compliance and ease of handling, but the more porous looped fibers of knitted fabric require either preclotting to form a hemostatic thrombus lining, or in more recent versions, a collagen, gelatin or albumin luminal impregnated coating. Two additional processes are sometimes applied to graft materials to achieve specific goals. Velour, which

refers to the textile process of extending the woven or knitted threads perpendicular to the inner or outer surface of the graft fabric, effectively increases the contact area of the graft material, more firmly anchoring the newly formed luminal thrombus lining, or better incorporating the outer graft into the surrounding tissues. Crimping, or folding of the graft material into corrugations improves the kink resistance of the graft for an equal thickness of material, but in doing so decreases the graft flexibility. Both velour and crimping create a graft that is thicker and less favorable for introducing through a sheath in an endovascular application. In one series of large stent-grafts for thoracic aortic aneurysms, the graft material crimping was ironed flat to reduce the working device diameter (Dake). Non-textile graft material includes primarily expanded polytetrafluroethylene (ePTFE) a polymer of carbon atoms enveloped in parallel coatings of large fluorine atoms, invented in 1969 by Robert W. Gore and made into a vascular graft in 1975. In the process of producing the expanded form, PTFE particles first are extruded under pressure, acquiring from the shear stress a strength longitudinal to the direction of the extrusion. These extruded tubes are then heated and stretched to create solid "nodes" oriented perpendicular to the length of the tube, and "fibrils" which are the longitudinal interconnections between nodes. The fibril lengths, also known as the internodal distance (IND) and typically on the order of 30 microns, determine the porosity of this material, .,0% of which is air. Due to its fluorine coating, PTFE is inert, stable, electronegative and hydrophobic; these features contribute to its relative non-thrombogenicity and, despite its high level of porosity and handling ease, its relative impetviousness to blood or need for preclotting. These positive qualities of the polymer, however, are offset by a lack of compliance of the ePTFE graft, especially in comparison to typical arterial viscoelastic properties. The node and fibril design of expanded PTFE lends porosity and flexibility at the expense of tensile strength: femoropopliteal grafts of the Gore-Tex brand of ePTFE placed in the early clinical experience were noted to easily form aneurysms. This was corrected by the addition of a thin outer coating of non-porous, sheet-like ePTFE for the needed tensile strength and ability to hold a suture. This coating must be removed in order to balloon dilate any Gore-Tex ePTFE that is part of a covered stem or stent-graft. A PTFE graft competitive manufacturer, Impra, relies on a thicker wall instead of the outer coating for the additional tensile strength. Another series of non-textile vascular grafts have been formed of polyurethanes, a family of chemical compounds all of which are hydrophilic, unlike the hydrophobic PTFE. Polyurethane microfiber grafts are formed by one of several manufacturing processes: casting, electrostatic and wet spinning of fibers and monofilaments, extrusion, dip coating or spraying of mandrils with polymer/additive solutions. These processes form a

synthetic conduit with a macroporous, randomly oriented fiber outer coating, unlike the regular nodes and fibrils of expanded PTFE. It preferentially binds albumin on its luminal surface when inserted as a vascular graft, making it less thrombogenic, and its compliance can be manipulated by adjusting the wall density. Some early forms of polyurethanes were not adequately durable as vascular grafts, although they have been successful as venous access catheters (Zhang). The properties of textile and non-textile synthetic vascular grafts described above pertain primarily to their use as traditional surgical bypasses. When utilized as part of a stent-graft, however, several other issues need to be considered. The first is the effect of balloon dilation of non-textile grafts such as ePTFE when used in smaller vessels such as the aortoiliac system either as a covering on a Palmaz balloon-expandable stent, or as the anchor of a hybrid stem-graft which is sutured into place distally (Cynamon). Little is known about the acute biocompatability or long-term durability effects of stretching ePTFE conduits to 3-4 times their original intended diameter. One initial experiment has demonstrated a decrease in graft wall thickness and both longitudinal and radial tensile strengths, and an increase in porosity, although without a change in the ePTFE internodal distance (F. Palmaz). It is unclear what long-term clinical significance a reduction in tensile strength may have for instance when such a material is used in a stent-graft to exclude an aneurysm. Stent-grafts for larger vessels such as the thoracic and abdominal aortas have been fabricated primarily of textile graft material, with variations of woven and knitted PET grafts supported mostly internally by self-expanding stents. Smaller diameter textile stent-grafts, with PET covering a self-expanding or balloon-expandable stent endoskeleton, have also been used in long-segment infra-inguinal occlusive disease. The second issue which must be considered, and which remains to be investigated, is the long-term impact of metallic stents pulsating on vascular g~aft material, and whether durability or biocompatability is affected by these repetitive stresses on fabric designed not with metal friction in mind. Vascular stem-grafts are a new frontier in the treatment of occlusive and aneurysmal vascular disease, and the materials currently being tested will soon be considered crude reminders of a pioneering era. In their place will come synthetic and biological graft materials and stents tailored and mated ;pecifically for endovascular use, with coatings and impregnations to optimize device function in the endoluminal environment. Research is already ongoing to replace in the near future all of these synthetic facsimiles with biological and gene engineering processes designed to stimulate the body to build in place its own new blood vessel.

References Sawyer PN, Kaplitt M], ed. Vascular Grafts. AppletonCentury-Crofts, New York, 1978.

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Stanley]C, ed. Biologic and Synthetic Vascular Prostheses. Grune & Stratton, New York, 1982. Greisler HP. New Biologic and Synthetic Vascular Prostheses. KG. Landes Company, Austin, 1991. Kogel HC, ed. The Prosthetic Substitution of Blood Vessels. Quintessenz-Verlags, Munich, 1991. Hastings GW, ed. Cardiovascular Biomaterials. Springer-Verlag, New York, 1992. Voorhees AB, ]aretzki A, Blakemore AH. The use of tubes constructed from Vinyon "N" cloth in bridging arterial defects. Annals of Surgery 1952; 135: 332-336. Dake MD, Miller DC Semba CP et al. Transluminal placement of endovascular stent-grafts for the treatment of descending thoracic aortic aneurysms. N Engl ] Med 1994; 331:1729-1734. Zhang Z, Marois Y, Guidoin RG, et al. Vascugraft polyurethane arterial prosthesis as femoro-popliteal and femoro-peroneal bypasses in humans: pathological, structural, and chemical analyses of four excised grafts. Biomater 1997; 18:113-124. Cynamon], Marin ML, Veith F], et al. Stent-graft repair of aorto-iliac occlusive disease coexisting with common femoral artery disease.]VIR 1997; 8:19-26. Palmaz F, Palmaz ]C, Sprague EA. Physical properties of PTFE bypass material after radial dilation. ]V1R 1996; 7:657-663. 12:45 pm

Mechanics and Biology of Stent Grafts Julio C. Palmaz, MD THE polymeric materials under consideration in this brief review are polyethylene terephthalate (PED polytetrafluoroethylene (PTFE) and polyurethanes (PUs). PET and PTFE have both been extensively applied as vascular prosthetic implants. PU vascular prosthetics have been used in a limited fashion and can be considered under evaluation at the present time. The advantages and limitations of these materials must be evaluated in regard to the relevance of their biological and biomechanical properties in vascular applications. An ideal material for arterial implant should be thromboresistant, its mechanical propenies should not degrade with time and it should be incorporated by tissue but not elicit excessive proliferation, inflamatory or degenerative response. These properties are not equally relevant in regard to diameter. For example, thromboresistance and tissue reaction are more important to the fate of small conduits and mechanical endurance is more relevant to large ones. Metals and alloys, panicularly as intravascular stents, have been employed with increasing frequency in the past 10 years. Much of the experience learned with intravascular stents is relevant to the research and development of endovascular devices since, almost invariably, these are composed of metallic and polymeric pans. For the purposes of this review, only polymers will 30 be considered. The emphasis will be on their physico-

chemical propenies and the biological response they elicit in blood and vascular tissues when used as implants.

Chemical Structure of Vascular Polymers PET is composed of long chains of alternating units of glycol and terephthalic acid. The tensile strength of PET is derived from the high dissociation energies of the covalent bonds along the polymer chain. Weaker hydrogen and van der Waals' forces determine the spatial conformation of the chains which are characteristically disposed as alternating segments of tangled and parallel arrangement at 10 nm intervals. This gives PET fibers the microscopic appearance of concentric dark and light bands surrounding nuclei or spherulites (1). PTFE is chemically composed of carbon chains saturated with fluorine. The chain twists 180 degrees every 13 carbon atoms. Because of the large size of fluorine atoms in comparison to carbon atoms the former arrange themselves as a sheath completely covering the carbon chain (2). This arrangement provides conformational rigidity and chemical stability that may explain some of the characteristics of this polymer such as low friction coefficient and high melting point (3). These physical aspects of PTFE correlate with some of the biological properties such as low thrombogenicity and modest tissue reaction. Microscopically, PTFE shows bands with width and length depending on the molecular weight and conditions of crystallization. These bands have perpendicular striations resulting from repeating crystalline and non crystalline zones. In the crystalline portions or striae, the PTFE molecules are arranged in a parallel fashion, perpendicular to the long axis of the bands and chain folding occurs in the non-crystalline areas (4). PUs are a chemically complex and diverse group of related compounds. PUs can be rigid or soft, depending on their composition. Current vascular PUs are elastomeric and, like other related compounds, have polymeric chains with alternating aliphatic soft segments and urethane derivative or hard segments. While the urethane ponion of the chain tends to form hydrogen bonds with adjacent similar moieties forming crystals , the aliphatic segments are randomly oriented (5). PUs have been tried in the past, as vascular substitutes, but have failed because of biodegradation. Early PUs such as polyether and polyester polyurethanes were susceptible to hydrolytic degradation at the ester linkages in the soft segments (6). In recently developed PUs, such as polycarbonate polyurethane (Corethane, Corvita, Inc, Miami, Fl; ChronoFlex, PolyMedica, Woburn MA) the polymer chain sites deemed to be at risk of cleavage were replaced for chemical groups resilient to degradation. These modifications in the chemical structure are claimed to have solved the biodegradation problem of the implanted materials (6).