The Artery

Few autogenous arteries are available for use as a conduit because of the lack of redundant arterial perfusion and available lengths of medium-sized arteries. However the patency rates of inferior mammary artery conduits in coronary artery bypass grafting is outstanding compared with vein grafts.¹⁰ The benefit is less clear when arterial harvesting is required, such as radial artery grafting.¹¹ There is also some support for this benefit in patients with challenging distal bypasses.¹² The theoretic advantage is that arteries do not need to adapt to arterial flow, and thus are relatively protected from negative remodeling. However, all bypass targets are not identical, and graft remodeling is also determined in part by the perfusion bed (myocardium versus skeletal muscle).

The Vein

Superficial and deep veins can be used as conduits for arterial bypass. Their benefit over synthetic grafts for small diameter graft bypasses is well accepted.¹³ The greater saphenous vein is the most commonly used conduit for coronary and infrainguinal bypass grafting. It is duplicated in approximately 8% of persons,¹⁴ and its course has been well described.¹⁵ The lesser saphenous vein is not easily accessible in the supine position, but it can be useful as a conduit when working posteriorly or through a separate harvest incision.^16–18 More commonly in our practice, bilateral cephalic veins, which are approximately 50 cm, can be sewn together as a composite conduit for distal bypasses.^19,20

Once implanted, veins undergo arterialization in which they develop a muscular media and elastic lamina. Adaption can be both positive and negative; therefore it is not surprising that a number of vein grafts fail or require adjunctive procedures to maintain patency (Table 16-1). The modes of failure include myointimal hyperplasia (frequently at anastomotic or valve sites or at sites of intraluminal manipulation) and progression of atherosclerosis in the inflow or outflow arteries, or both, or in the vein graft itself. Limiting trauma at time of vein harvest,²¹ proper perioperative medication (antiplatelet agents²² and a statin^23,24), proper ultrasound surveillance protocols²⁵ are critical to optimizing patency and clinical benefit of vein grafts. The patency benefit of anticoagulation of vein grafts is modest,^26,27 and the risks of bleeding are generally thought to outweigh these small benefits to vein graft patency.²⁴

TABLE 16-1 Proper Handling of Vein Grafts

Cryopreserved Vessels

Cryopreserved allografts provide an alternative to either femoral vein harvest or to antibiotic-soaked Dacron with omental flap coverage for infected aortoiliac reconstruction.³² However, their poor primary patency rates for infrainguinal bypass,^33–35 even with immunosuppressive therapy,³⁶ limit this application to patients with infection and without leg or arm vein.

Conduit Patency and Failure

The benefits of vein grafts versus synthetic grafts is well established,^13,38 but recently the PREVENT III and IV trials have demonstrated that even vein grafts have durability limitations for both infrainguinal and coronary revascularizations.^39,40 As described, primary patency rates are strikingly similar to those reported more than 30 years ago.¹³ However, secondary patency rates are better, emphasizing the importance of vein graft surveillance and prophylactic reinterventions.

A variety of mechanisms can lead to vascular graft occlusion, and they occur in a defined temporal sequence. Immediate graft failure is usually the result of technical error during the operation or the patient having a hypercoagulable status. Failure in the first month following graft placement is most likely the result of thrombosis secondary to distal flow resistance or, less commonly, restricted inflow into the graft. Small-diameter grafts are prone to early thrombosis because of their lower flow rates and higher resistance in their outflow vessels. Therefore graft thrombogenicity is of primary concern early after graft placement.

Pseudointimal hyperplasia is the most common reason for graft failure from 6 months to 3 years after graft insertion, occurring primarily in the perianastomotic region in the case of synthetic conduits and both there and the sites of valves or mechanical injuries in the case of vein grafts. Later graft failure is frequently due to the progression of distal atherosclerotic disease. Small-diameter grafts are particularly susceptible to anastomotic pseudointimal hyperplasia (IH), which is a false intima composed of smooth muscle cells, inflammatory cells, and activated endothelial cells, forms on the graft material. This is distinguished from neointimal hyperplasia, in which a new intima with confluent endothelium is generated, frequently with an increased diameter compared to the true intima. More generically, the term myointimal hyperplasia is used. Its definition encompasses both phenomena and includes the abnormal multiplication or increase in the number of normal cells, often with excessive accumulation of the extracellular matrix, in the innermost part of the blood vessel.

Pathophysiologically, denuded intima of a vein graft or exposed luminal area of a synthetic graft may lead to thrombosis via platelet deposition and activation of the coagulation cascade, which over time promotes pathologic SMC migration, proliferation, and ECM deposition, leading to IH. IH in turn narrows the vessel lumen (restenosis), decreasing blood flow to the point that it can promote local thrombosis or lead to symptomatic ischemia in the relevant distal end organs, such as the brain (cardiovascular accidents), heart (myocardial infarctions), and extremities (critical limb ischemia; Figure 16-1).

FIGURE 16-1 Myointimal hyperplasia: arteriographic and histologic findings.

A, Arteriogram showing a focal vein graft stenosis after a femoropopliteal bypass (left), and histology of this segment of vein graft (right). The arrow encompasses the diameter of myointimal hyperplasia. Deep to the arrow, a well-organized muscular vessel media is identified. B, Summary of the waves of myointimal hyperplasia in experimental models. Of note, the first wave occurs in the media. The second wave occurs in the media to myointimal region, and the third and fourth waves occur in the myointima. Strategies to limit this problem must understand cellular and bioactive proteins in a temporal and transmural fashion.

(Reprinted with permission from Elsevier. Brewster LP, Brey EM, Greisler HG et al: Advanced Drug Delivery Reviews, 58 (9) 604–629, 2006.)

Thrombosis and intimal hyperplasia represent the most common causes of graft failure and are both largely mediated at the luminal interface of the vessel or graft; therefore the inner lining of grafts has been the subject of much investigation. This is discussed below in two parts: the body’s response to grafts and the design of desirable blood-biomaterial interfaces.

Cellular and Molecular Mediators of Graft Outcome

A confluent quiescent endothelial lining of injured vessels and synthetic grafts is desirable for clinical applications.⁴¹ Re-endothelialization of injured vessels or endothelialization of synthetic grafts can occur through transanastomotic endothelial cell migration (usually limited in humans to 1 to 2 cm), transmural tissue ingrowth accompanied by microvessels from the perigraft tissue, and from deposition of circulating endothelial cells, endothelial progenitor cells (EPCs), or stem cells. Homing can be promoted through affixation of EC attractant antibodies to the grafts in a fashion similar to that used in coronary artery stents.⁴²

However, ECs growing on prosthetic graft surfaces are not necessarily the same as their normal quiescent counterparts in uninjured vessels. These ECs are often “activated,” secreting bioactive substances (e.g., platelet-derived growth factor [PDGF]) that actually promote thrombogenesis and changes in SMCs to a synthetic phenotype. This has been seen in the perianastomotic region, which is the most frequent site of interventional failure after operation.⁴³ SMCs found within the myointima of prosthetic grafts are also functionally altered. They produce significantly higher amounts of PDGF, as well as various extracellular matrix proteins, than those of the adjacent vessel, which contribute to the development of intimal hyperplasia along with the body’s inflammatory reactions to prosthetic material.⁴⁴

The inflammatory response to vascular interventions is complex. Both inflammatory cells and proteins mediate it in a cooperative manner. Potent chemoattractants like complement 5a (C5a) and leukotriene B₄ recruit neutrophils to the graft surface, where they localize in the fibrin coagulum of the graft’s inner and outer capsule via β2 integrins. In addition, immunoglobulin G binds to the neutrophils’ Fcγ receptors activating the neutrophils’ proinflammatory response while inhibiting normal clearance of bacteria. Neutrophils also interact with various other deposited proteins including C3bi and factor X, and they adhere to the endothelial cells in the perianastomotic region through selectin- and integrin-mediated mechanisms. L-selectin is thought to modulate neutrophil and endothelial cell interactions by presenting neutrophil ligands to the both E- and P-selectin on the vascular endothelium. In addition, selectin-carbohydrate bonds are important for the initial cellular contact, whereas the integrin-peptide bonds are responsible for strengthening of this adhesion as well as the transmigration of neutrophils. Both intercellular adhesion molecular 1 (ICAM1) and vascular cell adhesion molecule 1 (VCAM1) on the EC surface bind these integrins as well, and ECs upregulate ICAM1 expression and express VCAM1 when stimulated by inflammatory agonists such as interleukin (IL)-1, tumor necrosis factor (TNF)-α, lipopolysaccharide, and thrombin. Furthermore, activated neutrophils release oxygen-free radicals and various proteases, which result in matrix degradation that may inhibit both endothelialization and tissue incorporation of the vascular graft.

Circulating monocytes and macrophages are also attracted to areas of injured or regenerating endothelium, especially in response to IL-1 and TNF-α. There are many plasma monocyte recruitment and activating factors, including LTB₄, platelet factor 4, and PDGF. This process is propagated in the presence of these plasma activating factors driving monocytes to differentiate into macrophages that direct the host’s chronic inflammatory response via the release of proteases and oxygen-free radicals.

A variety of cytokines are released from inflammatory cells activated by vascular grafts. Cultured monocytes and macrophages incubated with Dacron and expanded polytetrafluoroethylene (ePTFE) have been demonstrated to produce different amounts of IL-1β, IL-6, and TNF-α that are biomaterial specific.⁴⁵ Because the inflammatory reaction elicits a cascade of growth processes, it has been proposed that approaches attenuating the initial inflammatory reaction may improve long-term graft patency. Alternatively, directing the inflammatory response to promote favorable cellular and protein responses may promote intimal generation and tissue incorporation.

Creation of a More Hospitable Blood Biomaterial Interface

The long-term patency of vascular grafts depends upon the intrinsic properties of the graft itself and the hemodynamic environment in which the graft is placed, as well as patient variables (e.g., diabetes, renal failure, nutritional status), and may or may not be improved by prior or concomitant interventions such as proximal or distal angioplasty. It is now clear that tissue incorporation is important for graft function; therefore grafts that have this ability are considered desirable for medium- and small-caliber vessel replacement. The compliance mismatch between arteries and grafts causes flow disruption in vivo, which could contribute to anastomotic pseudointimal hyperplasia.⁴⁶ For this reason, a number of surgeons have suggested interposing a segment of vein between the synthetic graft and artery, creating a composite graft at the distal anastomosis. This has also led some investigators to design more compliant grafts using more flexible materials or changing the parameters of graft construction to improve graft compliance. Although animal experiments have suggested concept validity, the clinical benefit of this approach remains controversial. Many factors may contribute to this confusion, including longitudinal variability in the diameter and compliance of the arterial tree and the effect of activated endothelium on intimal hyperplasia. Furthermore, there is a robust fibrotic response after implantation that leads compliant grafts to become more noncompliant after implantation; therefore, even if a compliance match were attained initially, this benefit would unlikely persist. Synthesis of a conduit designed to match the compliance of the arterial wall is simplistic, as the compliance of the diseased arteries to which the grafts would be anastomosed is widely variable. Furthermore, in the para-anastomotic region there are dynamic changes in compliance that vary over time. First a para-anastomotic hypercompliant zone exhibits a 50% gain in compliance, then later its compliance is lessened 60% from baseline.⁴⁷ It is likely that this bimodal adaptation limits the practical value of graft designs focused on initial mismatch. Addressing later mismatch may be more valuable, but synthetic grafts are intrinsically limited in this ability to adapt over time.

Anticoagulant Affixation

Early platelet deposition on vascular grafts is mediated by von Willebrand factor and platelet membrane glycoproteins. After adherence to a graft, platelets degranulate, releasing many bioactive substances, including serotonin, epinephrine, ADP, and thromboxane A2; these substances in turn activate additional platelets and promote a prothrombogenic reaction. Activated platelets also release growth factors, such as PDGF, epidermal growth factor (EGF), and transforming growth factor-β, which promote SMC migration and proliferation as well as ECM degradation and ECM protein synthesis. In addition, platelets release monocyte chemoattractants such as platelet factor 4 and β-thromboglobulin, which mediate the recruitment of macrophages to the graft. Platelet deposition and activation continues chronically after graft implantation as evidenced by increased thromboxane levels and decreased systemic platelet counts 1 year after Dacron graft implantation in a canine model,⁴⁸ and human studies have confirmed platelet adhesion to grafts up to 1 year after implantation.

The deposition and activation of platelets elicit various pathologic cascades; therefore the thrombogenic nature of the synthetic graft surface can lead to both early and late graft failure. A myriad of approaches have been studied to attenuate platelet deposition, aggregation, and degranulation. Antiplatelet agents directly targeting platelet or graft-binding molecules such as platelet surface GPIIb/IIIa and different functional domains of thrombin have been shown to at least transiently decrease the accumulation of platelets on Dacron grafts.⁴⁹ The surface thrombogenicity of grafts can be altered experimentally through disruption of the blood system’s coagulation cascade on thrombogenic surfaces. We have reported improved patency, thromboresistance, and the absence of intraluminal graft thrombus in heparin-analog–coated ePTFE grafts using a canine bilateral aortoiliac artery model.⁵⁰ Similarly, combined data tables of clinical data showed a decrease in subacute thrombosis with heparin-coated coronary artery stents.⁵¹ Genetic approaches to increase thromboresistance have been used by multiple groups through the overexpression of thrombotic inhibitors, but because quiescent ECs themselves are antithrombotic, there may be limited benefit of this approach compared with a functioning endothelium.

A heparin-bonded Dacron graft from Maquet (Wayne, NJ) is currently available on the market. The heparin is bound primarily through Van der Waals bonds to the polyester fiber pretreated with a cationic agent, tridodecil-methyl-ammonium chloride. The external third of the graft wall is coated with collagen to prevent blood extravasation.⁵² In a comparative clinical trial involving 209 patients undergoing femoropopliteal bypass grafts, the heparin-bonded Dacron graft exhibited a slightly better but significant (p = 0.04) patency with 1, 2, and 3 years of 70%, 63%, and 55%, compared to untreated ePTFE graft of 56%, 46%, and 42%, respectively.⁵³

A heparin-bonded ePTFE graft developed by W.L. Gore is currently available worldwide. Heparin binding is through covalent end-point attachment of the heparin to the pretreated bioactive surface of the graft. An animal study showed improved patency of the graft compared with the standard ePTFE graft using a canine carotid interposition model up to 6 months.⁵⁴ The surface heparin activity measured by antithrombin III uptake per unit area was 24.7 ± 7.9 pmol/cm² at 2 weeks and remained at 15.3 ± 3.7 pmol/cm² by 12 weeks after implantation. Clinical data to date suggests decent early results⁵⁵ and 5-year results.⁵⁶

Whether the anticoagulation works through continuous release of heparin from the material establishing an effective concentration at the interface between blood and the graft surface or through nonconsumptive mechanisms of active function of the heparin immobilized on the material surface is unclear. A major concern with heparinization of the graft surface is the duration of heparin function. Premature release or disturbance of functional heparin or the presence of a physical barrier owing to adherent blood components implies a theoretical inefficacy of the approach. Although theoretically probable,⁵⁷ there have been discordant case reports on whether these grafts cause heparin-induced thrombocytopenia.^58,59

Prevalent Grafts

Polyethylene terephthalate (PET or Dacron) and ePTFE are the predominant materials currently used in prosthetic vascular grafts, but both Dacron and ePTFE react with blood components and perigraft tissues in both clinically desirable and undesirable manners. All grafts, regardless of their composition and structure, evoke complex but predictable host responses that begin immediately upon restoration of perfusion, and this improved understanding of the cellular and molecular components of biomaterial and tissue interactions has led to more intelligent designs of grafts that maximize beneficial ingrowth while minimizing the chronic inflammatory changes that lead to graft dilation or occlusion. These approaches include protein adsorptive grafts (e.g., growth factors, anticoagulants, antibiotics) and improved graft skeletal construction via synthetic polymers or biologically derived structural proteins that can be bonded to various bioactive cytokines and growth factors to induce a more favorable host response.

Dacron

PET was first introduced in 1939. DuPont (Wilmington, DE) made further developments and patented its well-known Dacron fiber in 1950.⁶⁰ Vascular grafts made from Dacron were first implanted by Julian in 1956⁶¹ and DeBakey in 1958.⁶²

Clinically available Dacron grafts are fabricated in either woven or knitted forms. In woven grafts, the multifilament Dacron threads are fabricated in an over-and-under pattern in both lengthwise (warp) and circumferential (weft) directions. This structure results in limited porosity and the best dimensional stability of the finished grafts. Accordingly, woven grafts have less bleeding through interstices and less likelihood of structural deformation after implantation. The disadvantages of such grafts are less desirable handling features, reduced compliance, and a tendency to fray at cut edges.

Knitted grafts use a textile technique in which the Dacron threads are looped to form a continuous interlocking chain. Most knitted grafts are manufactured with the threads predominantly oriented in a longitudinal direction (warp knitting) as opposed to a circumferential direction (weft knitting), which has been proved to be structurally unstable. Warp knits have good handling characteristics and reasonable dimensional stability. The loop structure creates greater porosity and radial distension. The velour technique that extends the loops of yarn on the surfaces of the fabrics has been used in the attempt to increase tissue incorporation. An external velour surface permits more extensive and firmer incorporation of the graft into surrounding tissue, but the function of an internal velour structure remains controversial, with the suggestion that it may enhance firm anchorage of the fibrin–platelet pseudointima.^63,64

Crimping technique is used to increase flexibility, distension, and kink resistance of textile grafts. It is recognized, however, that much of the initial elasticity is lost with the stretching during implantation and with later tissue incorporation. Crimping also reduces the effective internal diameter of the graft and creates an uneven luminal surface. The latter can potentially interfere with laminar blood flow that leads to increased thrombogenicity of the graft. Although such considerations may not be critical in large-diameter grafts, they are important in small-diameter situations. As a result, prosthetic rings or coils are often applied to the graft external surface of grafts to provide external support to resist kinking and mechanical compression.

The high porosity of the knitted graft necessitates preclotting to prevent transmural blood extravasation. Gelatin (Vascutek, U.K.) and collagen (Maquet, Wayne, N.J.) are predominantly used to seal knitted Dacron graft pores. The gelatin and collagen in the Vascutek and Maquet grafts are cross-linked by low concentrations of formaldehyde, a method resulting in a weak linkage that allows the gelatin or collagen to be degraded in the body in less than 2 weeks.^65,66

Dacron grafts, especially knitted grafts, have been demonstrated to be prone to dilate when implanted into the arterial environment. A 10% to 20% increase in graft diameter size after 1 year following restoration of blood flow is in the expected range. A follow-up study in 95 patients from 2 weeks to 138 months postoperatively (mean, 33 months) demonstrated a mean increase of 17.6% (0% to 84%) compared with initial graft size.⁶⁷ Woven grafts are structurally more stable. A direct relationship between uncomplicated graft dilatation and structural failure has not been established. Consequently, there is no recommendation on a specific degree of dilation that constitutes a significant hazard and warrants graft replacement.

Structural failure has been reported sporadically in the literature despite advances in the design and manufacturing of grafts (Figure 16-2).

FIGURE 16-2 Contained rupture of prior aortic aneurysm repair.

Computed tomographic scan images demonstrate rupture of the Dacron graft likely near the proximal anastomosis. Arteriogram at time of reoperation (bottom right) demonstrates the dilation of the graft compared to native aorta.

In 1997, an inquiry to the U.S. Food and Drug Administration by Wilson et al.⁶⁸ revealed 68 cases of Dacron graft failure owing to structure defect, with an average time to failure of 7.4 years (range, 4 to 18 years). The true incidence of structure failure is difficult to estimate because most patients with grafts are not followed-up for long periods or are not being evaluated periodically for graft integrity. A number of factors may contribute to the graft degradation including the design of the textile structure, manufacturing process, surgical handling or application of clamps during implantation, mechanical fatiguing due to cyclic stresses of pulsatile blood flow, and chemical or physical alterations associated with biodegradation. The first generation of double-velour knitted Dacron grafts introduced in the mid 1970s, such as the Microvel and Cooley Double Velour grafts (Boston Scientific Meadox, Oakland, N.J.), showed excessive dilation and rupture after implantation.^68,69 It is believed that structural failure of these grafts was related to the fabrication techniques such as the use of trilobal filaments, insufficient strength of guideline yarns, and the weakness of connecting lines where two knitted bands join together to form a tube.^67,70 The trilobal filaments were replaced with more durable cylindrical filaments more than 30 years ago.

After implantation, a coagulum containing fibrin, platelets, and blood cells builds up during the first few hours to days and stabilizes over a period of 6 to 18 months, forming a compacted layer. The histologic characteristics observed within Dacron grafts is a compact fibrin layer on the blood contacting surface and densely packed foreign body giant cells between the outer layer of graft wall and surrounding connective tissue capsule. The fibrin layer within the midgraft region beyond 1 to 2 cm from either anastomosis remains acellular regardless of whether the grafts are woven or knitted. Protein impregnation changes the surface properties of Dacron grafts and may induce more inflammatory reaction, but does not change the clinical patency rates of these grafts.^65,71,72

ePTFE

PTFE was patented by DuPont in 1937 as Teflon. In the 1960s, ePTFE was developed and is the material used in vascular grafts. The expanded polymer is manufactured by heating, stretching, and extruding processes producing a microporous material that is more supportive of firm tissue adhesion.

The PTFE molecule is biostable, and the graft made from it does not undergo biological deterioration within the body. The surface of the graft is electronegative, which minimizes its reaction with blood components. ePTFE grafts are manufactured by stretching a melt-extruded solid polymer tube that then cracks into a nontextile porous tube. The characteristic structure of ePTFE is a node-fibril structure in which solid nodes connect through fine fibrils with an average internodal distance in the range of 30 µm for standard grafts.

The initial host response to ePTFE grafts is similar to that of Dacron grafts.^63,64 A fibrin coagulum or amorphous platelet-rich material develops over a time sequence similar in both materials. Lack of luminal surface cellular coverage can be found at the midgraft region for years following human implants.^73–75 In outer wrap–reinforced grafts, the wrap delays the infiltration of the cells from perigraft tissue.⁶⁴

In addition to the standard ePTFE graft, a number of modified grafts have been developed to offer specific properties for various applications. Thin-wall grafts, which have a wall thickness of 0.2 to 0.3 mm versus 0.4 to 0.6 mm for the standard grafts, possess better handling characteristics and improved compliance. The grafts failed to show a better patency for hemodialysis arteriovenous access compared with the standard graft⁷⁶ and are prone to more postpuncture blood loss. Extended stretch ePTFE graft has longitudinal extensibility that is achieved by a microcrimping of the fibrils. When the stretch graft is extended, the fibrils are extended to their full length, and the microstructure of the graft is then identical to that of the standard ePTFE graft. Stretch ePTFE grafts have somewhat better patency rates when used for atrioventricular access. Unfortunately, no data support such an advantage when used for peripheral reconstructions. Rings or coils are often applied to the abluminal surface of the grafts to provide external support in order to increase the graft resistance to external compression and to reduce kinking.

Still, the rate and extent of tissue ingrowth can be improved by optimizing graft permeability. Clowes et al.⁷⁷ have demonstrated enhanced tissue ingrowth and complete reendothelialization of 60 µm or 90 µm internodal distance ePTFE grafts in a baboon model.⁷⁷ Transinterstitial capillary ingrowth was not seen with the more commonly used 30-µm internodal distance ePTFE. Human trials using ePTFE with these expanded internodal distances failed to show any advantage in platelet deposition (considered an indicator of endothelial coverage) compared with standard 30-µm internodal distance ePTFE grafts.⁷⁸

The delivery of potent angiogens or genes that promote EC-specific mitogenesis or chemotaxis upon injured vessels or prosthetic surfaces may be used to regenerate a rapid and complete endothelium after vascular intervention. We have demonstrated the feasibility of this approach in endarterectomized canine carotid arteries (Figure 16-3).

FIGURE 16-3 Endothelial healing with localized delivery of FGF-1 in fibrin glue.

A, Scanning electron micrograph (×400) illustrating the differences in re-endothelialization between normal saline untreated arteries and treatment arteries at 30 days after balloon angioplasty injury. The top micrograph is an untreated artery, and the bottom micrograph is an HBGAM/FGF-1 treated artery. B, Light microscopy micrograph (×10) comparing mitotic indices of treatment groups to untreated and fibrin glue treated control arteries at 30 days after injury as measured by bromodeoxyuridine (BrdU) incorporation (brown staining). The control arteries have significantly less BrdU incorporation than the treatment arteries (p < 0.017).

(From Brewster L, Brey EM, Addis M, et al: Improving endothelial healing with novel chimeric mitogens. Am J Surg 192:589–593, 2006.)