Proceedings of the Institution of Mechanical Engineers, Part H: Journal of Engineering in Medicine http://pih.sagepub.com/

Engineering Design of Vascular Prostheses T V How, R Guidoin and S K Young Proceedings of the Institution of Mechanical Engineers, Part H: Journal of Engineering in Medicine 1992 206: 61 DOI: 10.1243/PIME_PROC_1992_206_269_02 The online version of this article can be found at: http://pih.sagepub.com/content/206/2/61

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61

Engineering design of vascular prostheses T V How, BSc, PhD, CEng, MIEE, MBES Department of Clinical Engineering, University of Liverpool

R Guidoin, PhD, DSc Laboratoire de Chirurgie ExPCrimentale, Universite Laval, Quebec, Canada

S K Young, MB, ChB Department of Clinical Engineering, University of Liverpool The replacement and bypass of arteries of diameter greater than 6 mm with textile vascular prostheses has proved very successful since they were jirst introduced forty years ago. Although manufacturers continue to improve their products and make them of consistent quality for increased safety and performance and to facilitate their use by surgeons, most of the research work in this area is concerned with the development of small-diameter prostheses. Current expanded PTFE and textile prostheses do not perform satisfactorily when their diameters are reduced to less than 6 mm. For the small-diameter prostheses it will he necessary to develop less thromhogenic materials and to design the structure of the prostheses more closely to match the mechanical properties of the natural arteries. The purpose of this paper is to discuss the design requirements and to review the development of large- and small-diameter vascular prostheses. 1 INTRODUCTION

Surgical reconstruction of occluded arteries is a routine and generally successful operation for restoring adequate blood supply to organs or limbs. This usually takes the form of a vascular prosthesis that is joined to the artery upstream and downstream of the obstruction to restore the blood supply to the distal ischaemic tissues. Textile arterial prostheses made of polyethylene terephthalate (Dacron@) and polytetrafluoroethylene (Teflon@)have been used routinely for this purpose for the past four decades. These prostheses perform satisfactorily when they are used for the replacement or bypass of the large arteries with high blood flow velocity. When the internal diameter of the arteries is less than 6 mm and the blood flow velocity is low these prostheses achieve much lower success rates (1). The causes of failure of vascular prostheses are multifactorial and complex and the exact mechanisms are not well understood. Most failures occur by thrombosis resulting in occlusion, and this may take place early (within the first month of implantation) or much later. The major factors thought to be responsible for these failures can be divided into those related to the material (such as the surface characteristics and the physical, mechanical and haemodynamic properties of the graft) and those specific to the host (such as abnormal coagulation, the conditions of the inflow and outflow vessels, and the progression of atherosclerosis) (2). Autologous veins are considered to be the best grafts for the reconstruction of small peripheral arteries such as those of the lower limbs (3). In the case of coronary arteries, the internal mammary artery graft is preferred by many surgeons because of its superior long-term patency compared with the reversed saphenous vein graft (4). In about 20-30 per cent of patients requiring lower limb arterial bypass, the autologous veins are not suitable as they may be too small or diseased or may have been used in previous operations (5). None of the The M S was received o n 22 M a y 1992 and was accepted for publication on 6 August 1992. H02592 @lMechE 1992

current alternatives, for example knitted Dacron, expanded polytetrafluoroethylene (PTFE) prostheses and umbilical vein grafts, have achieved satisfactory long-term results (6, 7). There has been considerable effort in the development of better small-diameter vascular prostheses in the past 15 years but none, as yet, has reached the stage of clinical acceptance. Much of the work has centred on the use of polyurethanes because of their generally lower thrombogenicity and more desirable mechanical properties. However, doubts have been expressed about the biostability of some of the polyurethanes following long periods of implantation in oioo (8-11), although attempts are being made to overcome this problem by chemical means (12). The purpose of this paper is to review the design considerations, the construction and mechanical properties of Dacron, PTFE and several elastomeric smalldiameter vascular prostheses under development. 2 DESIGN CONSIDERATIONS FOR VASCULAR

PROSTHESES The first vascular prostheses used in the early 1950s cannot be said to have been designed in the engineering sense. These prostheses were seamed conduits made by sewing together two pieces of flat woven material (13). Although the technology for producing seamless woven tubes was available at the time, it was only after the principle of a porous textile tube as an arterial replacement became accepted that this was employed to manufacture arterial prostheses. Most of the new designs of textile prostheses can be considered as adaptive designs in which relatively minor modifications were made to the original. These include various knitted constructions, which have been devised to provide better strength, velour surfaces to improve healing, crimping and external support to prevent buckling and compression of the prosthesis. However, the design of small-diameter prostheses has required a different approach as it was realized at an early stage that merely reducing the scale of the implants was insuff-

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cient. This approach may be termed ‘development design’ whereby the existing successful prostheses are used as a basis for developing new devices. The technical work involved in the development is considerable and the final product may be quite different from the original. Many small-diameter vascular prostheses that have been proposed to date (1419) fall into this category. In order to design small-diameter arterial prostheses, the complete design specifications should be determined and considered at the outset. Based on the features of successful large-diameter vascular prostheses and autologous vein grafts, several authors have listed the principal requirements that must be met (1, 20). The material should be relatively inert so that it does not provoke an excessive fibrous reaction. The surface should be non-thrombogenic. The prosthesis should be porous to help the growth of a new intima. It should be strong enough to withstand the maximum blood pressures encountered in the body. Since the environment in which the prosthesis is required to perform is aggressive, the material should be chemically stable (11) and its strength should not deteriorate with time by biochemical degradation, fatigue or creep. The prosthesis should also have a smooth flow surface and, unlike polyester and PTFE prostheses, it should have compliant walls. The prosthesis should be designed for maintenance-free operation since further surgical intervention after implantation is costly and undesirable for the patient. 2.1 Porosity

The notion of a porous conduit made from a synthetic fabric material for use as an arterial prosthesis was first demonstrated by Voorhees et al. (21). According to this concept the fabric structure acts as a skeleton within which blood clots and becomes entrapped. The thrombus is then transformed to a compacted fibrin lining (the pseudointima) which is firmly supported by the fabric structure, providing a less thrombogenic surface in contact with blood than the base material. The ultimate non-thrombogenic surface is one that is covered with endothelial cells (that is the luminal surface of arteries and veins). Although, when implanted in experimental animals, current vascular prostheses become completely endothelialized within a few months, in man the endothelial layer extends only a few millimetres from the suture lines. All successful vascular prostheses that have been developed for clinical use are based on the concept of a porous conduit. Although a solid walled tube may work, this would rely on the use of a completely nonthrombogenic material (20). Since this material is not yet available, most of the effort in the development of small-diameter prostheses has been concerned with methods of producing tubular conduits with porous walls. Porosity is related to the void space within the prosthesis wall and can be determined by measuring the void fraction, defined as the ratio of the volume of the void space to the total volume of the porous material. The void fraction for fabric prostheses may vary from about 0.5 for woven to 0.8 for knitted constructions (22). In the early 1960s, a simple test was devised by

surgeons to estimate porosity indirectly, that is by measuring the volume flow of water across a sample of the prosthesis of known area subjected to a transmural pressure of 120 mmHg (16 kPa) (23, 24). The water permeability, usually given in terms of millilitres per square centimetre per minute, is higher for knitted than woven textile prostheses. This parameter therefore provides an indication as to whether or not a particular prosthesis should be preclotted before implantation. In a study of 34 textile prostheses it was found that water permeability does not correlate well with the void fraction and is therefore not a satisfactory parameter for measuring porosity (22). Although the void fraction is a useful indicator of the healing characteristics of the prosthesis it does not give any indication of the distribution of pore sues, which is known to play an important role in the tissue response to the prosthesis (25, 26). When the pore size is greater than 45 pm the prosthesis becomes infiltrated predominantly with fibrous tissue (type I collagen). When it is less than 45 pm the wall is penetrated by fibrohystiocytes which produce type 111 collagen. Type I1 collagen is associated with organs exhibiting elastic behaviour such as the dermis and the cardiovascular system and is therefore more elastic than type I collagen which is the constituent of fibrous tissue. Moreover, in prostheses with pore sizes less than 30 pm the tissue ingrowth is limited to a depth of less than 50 pm, leaving the bulk of the wall free of fibrous tissue (27). Thus it is possible to control the tissue response, and consequently the compliance of the prosthesis, by selecting the structure having the appropriate pore size distribution. This may be important for small-diameter prostheses as it would help to preserve the vessel compliance. 2.2 Flow surface

The surface of a woven or knitted fabric prosthesis cannot be said to be smooth. Where the yarns cross over each other, projections of the order of the width of the yarn (0.3-0.5 mm) are formed, producing a grossly uneven flow surface. This effect is worsened in crimped prostheses where the height of the crimps may be significant in relation to the prosthesis diameter (1). A rough surface may lead to increased flow resistance, particularly when the flow is turbulent (28). Both the relative height and the shape of the roughness play a role in determining the flow resistance (29). Savvides and Gerrard (30)have gttempted a numerical analysis of pulsatile blood flow through corrugated tubes with dimensions similar to those of crimped Dacron arterial prostheses. They found that for a simulated femoral artery waveform the effect of the corrugations on the total resistance to flow was small. Although local flow separation occurred in the furrows, no stagnant regions were present and, compared with uniform tubes, increased mixing at the wall was predicted by the numerical model. This would imply that mural thrombus should not be more likely to form in crimped than in uniform prostheses. The evidence in uiuo, however, indicates that this is not the case and there are significantly greater deposits of fibrin and cellular elements in crimped prostheses (31). When a prosthesis with an uneven surface is

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ENGINEERING DESIGN OF VASCULAR PROSTHESES

implanted, fibrin and cellular materials are deposited to a thickness greater than the height of the roughness (32). In crimped prostheses, this can lead to a significant reduction in the luminal cross-section, particularly when the vessel diameter is small. In addition, because the thickness of fibrin and blood cell deposits in the furrows is greater, healing is delayed and the risk of thrombosis is increased (31, 33). These are not seen in non-crimped prostheses which have a thinner and more uniform lining. Thus, it is an essential requirement that new vascular prostheses should be designed with a ‘smooth’ luminal surface, and this is especially important for small-diameter prostheses. 2.3 Mechanical properties

Mechanical properties are of major importance when selecting a material for the fabrication of vascular prostheses. They are also important for the safety of the patient because they have an influence on the performance of the implant. Not only should vascular prostheses have sufficient strength to withstand the highest blood pressures in the body, but the strength should be maintained over the lifetime of the prosthesis. Mechanical stability and durability of the material may be affected by biochemical degradation which may lead to changes in molecular weight, surface cracking, alteration in prosthesis dimensions or failure. All current vascular prostheses have tensile strengths much greater than those of arterial tissue (Table 1). This safety margin is essential, at least in the case of Dacron prostheses, since a reduction in strength has been noted after implantation as a result’of dilatation (38, 39) and the slow rate of degradation (40). The AAMI standard has published standard test methods for measuring tensile strength (41). These include tests on ring, sheet and tubular samples. However, no minimum value is given because of the diversity of materials and structures, which would be different for each material. Since a vascular prosthesis undergoes cyclic deformations (approximately 5 x lo7 every year), the fatigue properties are important. Fatigue life may be affected by the type of material, the microstructure of the prosthesis wall, manufacturing flaws, material damage caused by improper handling during surgery and the environment (that is blood and tissue fluids). Measurements made using a ‘life tester’ over a period of 24 weeks (2.32 x lo7 cycles) have demonstrated the mechanical durability of Teflon prostheses (42). In the case of Dacron prostheses, there was an unrecoverable increase in diameter and a drop in tensile strength which was attributed to the damage of fibres caused by abrasion between yarns. Table 1 Tensile strength of Dacron fabric and polyurethane vascular prostheses compared with that of the canine aorta

2.3.1 Compliance

The viscoelastic properties of arteries play an important role in the performance of the cardiovascular system and it is well known that compliance of the central blood vessels is essential for efficient propagation of the pressure pulse to the peripheral vessels. Since the vascular prosthesis materials used clinically are much less compliant than arterial tissue, it is thought that failure of vascular prostheses of small diameter is associated with the mismatch in elastic properties between the host artery and the prosthesis (the stress-strain curves of knitted and woven Dacron prostheses, and those for the abdominal and femoral arteries are shown in Fig. 1 for comparison). Although some authors have consistently identified a definite link between compliance matching and performance of the implant (44-47), others have failed to observe a significant difference between compliant and stiff prostheses (48, 49). Nevertheless, most surgeons now believe that new small-diameter vascular prostheses should be designed with compliance matched as closely as possible to the host artery, since these ought to perform better haemodynamically. The haemodynamic changes brought about by the implantation of a rigid prosthesis into a naturally compliant vessel can, in general, be anticipated by the application of fluid dynamic principles. A discontinuity in elastic properties is known to cause reflection of the blood pressure wave, leading to loss of energy. Wave reflection depends on the input impedance upstream and downstream of the anastomoses and can be minimized by matching the impedances. Under pulsatile pressure, the mismatch in elastic properties will also cause a discontinuity in cross-sectional area at the anastomosis, which has implications for the local haemodynamics. The effect is similar to that of a stenosis and is characterized by the presence of flow separation and turbulence (50). Flow disturbances are associated with low and high shear stresses and there is evidence that regions exposed to abnormally low (51) and high (52) shear stresses are predisposed to intimal hyperplasia. Intimal hyperplasia is a major cause of long-term failure of small-diameter prostheses. It leads to a gradual

06

1 I I-

Longitudinal strength

Material

M Pa

MPa

Reference

Dacron fabric Polyurethane Thoracic aorta Abdominal aorta

128 5.62-1 0.25 0.73 1.47

139 11.96-8.05 0.32 5.29

34 35 36 37

Knitted Dacron

/ 0.2 Femoral artery 00.00

Circumferential strength

63

I

0.25

I

0.50

I

I

0.75 Strain

1.00

I

1.25

I

1.50

Fig. 1 Stress-strain curves of strip specimens of woven and knitted Dacron vascular prostheses compared with those of the distal abdominal aorta and the common femoral artery [Data from Hasegawa and Azuma (4311 Proc Instn Mech Engrs Vol 206

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T V HOW, R GUIDOIN AND S K YOUNG

reduction of the luminal diameter and may eventually result in occlusion of the lumen. When a rigid prosthesis is implanted in a compliant pulsatile flow system, stress concentration at the suture line may occur because of the restraint on the movement of the arterial wall imposed by the prosthesis. This may be responsible for cases of anastomotic aneurysms (53) and intimal hyperplasia (52) which have been observed with the use of stiff prostheses. The severity of these effects depends on the degree of compliance mismatch and the length of rigid prosthesis used. In the case of superficial femoral artery bypass, the prosthesis is between 30 and 60cm long and its compliance, assuming a 6 mm diameter expanded PTFE prosthesis, is about 6 times lower than that of a normal femoral artery (Table 2). Such a degree of compliance mismatch can cause measurable flow disturbance in pulsatile blood flow (54).It has also been shown that the rigid prosthesis acts as a low-pass filter and attenuates the high-frequency components of the flow waveforms (55).It is often pointed out that as the artery being bypassed is diseased and often calcified, attempts to match compliance would be futile since these vessels do not have the same compliance as normal healthy arteries. Langewouters et a!. (56)have compared the compliance of mildly diseased aortas (those with fatty streaks and spots, and fibrous plaques) with severely atherosclerotic aortas (those with atheroma, ulceration and calcium deposits) and found that there were no differences over the physiological pressure range. Compliance is defined as the fractional change in volume of a vessel segment per unit change in transmural pressure (AvlvAp). It can be shown that if the change in internal diameter di is small and second-order terms in Adi are neglected, then

C is normally expressed as a percentage and the constant 2 is usually omitted (37).The inside diameter and wall thickness are difficult to measure, especially in vivo, and therefore compliance is defined by most authors in terms of outside diameter do:

Arteries and arterial prostheses are anisotropic and exhibit non-linear elastic behaviour. Their compliances are dependent on the conditions of measurement (that is longitudinal extension, mean transmural pressure) and in order to allow a comparison to be made with other data, these conditions should be specified. The values of C for arteries and vascular prostheses reported in the literature are listed in Table 2. Since C has been determined from volume or diameter measurements, the values have been recalculated according to equation (2) in order that a comparison can be made between different studies. The in nitro values are normally determined under static or quasi-static conditions at a mean pressure of 13.3 kPa (100 mmHg). When measurements are carried out in oivo, Ado is taken as the maximum change in diameter between systole and diastole, do is assumed to be the diameter in diastole and Ap is the pulse pres-

Table 2 Compliance of vascular prostheses, saphenous vein graft and femoral artery Diameter

C

~~

Vessel

mm

%/mmHg

Reference

Woven Dacron Knitted Dacron Warp-Knit, Ext. Velour Expanded YTFE Polyurethane Human umbilical vein Human saphenous vein Human femoral artery

10 8

0.0008 0.02

57 58

6

0.012 0.07-0.19 0.037

59 60 61 61 61

4

0.044 0.059

sure. Since the compliance of arteries and arterial prostheses is frequency dependent (62),the measured in vivo value will depend on the frequency content of the pressure and diameter signals. It is therefore more appropriate to determine the frequency components of the diameter and pressure signals by Fourier analysis before calculating the compliance values at each frequency.

2.4 Geometry Arterial prostheses are usually cylindrical in shape. The length of a prosthesis used depends on the length of the arterial segment to be replaced or bypassed. Since the resistance to flow through a tube is inversely proportional to the fourth power of the diameter in steady flow and to the power of 2.25 in pulsatile flow (30),it is important that a prosthesis of adequate size should be selected. Allowance should also be made for the pseudointimal layer which forms on the luminal surface of the prosthesis. This is particularly important in smalldiameter prostheses where thc thickness of the pseudointima may be significant in relation to the internal diameter. Selection of prosthesis diameter is also determined by the diameter of the host artery. In order to avoid flow disturbances and energy losses due to sudden changes in the flow cross-sectional area, the diameter of the prosthesis should be similar to that of the artery. In the case of bypass of the arteries of the lower limbs, this is not possible because of the disparity in size between the proximal vessel (the common femoral artery with a diameter of 6-8 mm) and the distal vessels (the popliteal and tibia1 arteries with diameters ranging from less than 2 to 4 mm). This may be overcome by using prostheses with a tapered lumen. Tapered prostheses also have haemodynamic advantages over cylindrical prostheses. Converging flow in a tapered conduit is known to be more stable than flow in a cylindrical conduit because of the more favourable pressure gradient. Moreover, flow disturbances, such as those produced at an anastomosis, are attenuated in tapered prostheses of angle of taper as low as 0.5" (63). The tapered geometry also has the effect of amplifying wall shear rates. This is due to the increase in flow velocity as the cross-sectional area decreases and to the more uniform velocity profiles seen in tapered prostheses. Both a stable flow and moderately high shear rates are desirable in arterial prostheses since there is evidence that there is a greater tendency for intimal hyperplasia (51)to form and mural thrombus (64)to accumulate on artificial surfaces when the shear stresses are low and the blood flow is disturbed or turbulent. The effect of blood flow velocity (and wall shear

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ENGINEERING DESIGN OF VASCULAR PROSTHESES

stress) on thrombus formation in arterial prostheses has been studied by Sauvage et al. (65). They found that for a given prosthesis material, the amount of thrombus deposited on the walls was dependent on the flow velocity. They defined a thrombotic threshold velocity (TTV) below which more than 50 per cent of the luminal surface was covered with thrombus. Above this threshold velocity, less thrombus accumulated on the wall, presumably by the action of shearing forces. Measurements of TTV were made on externally supported knitted Dacron, expanded PTFE and crimped knitted Dacron prostheses and were found to be 2.3, 5.2 and 8.8 cmfs respectively. Thus to minimize the risk of thrombosis it would be advantageous to use a prosthesis of small diameter in order to provide flow velocities well above the TTV for the material. However, as mentioned previously, proper consideration must also be given to the flow resistance as it increases with decreasing diameter. Increased flow velocity in a vascular prosthesis (usually for femorodistal bypass) may also be obtained by creating an arteriovenous shunt at the distal anastomosis, thereby providing an additional low-resistance channel for the prosthesis flow (66). Although the side effects of an arteriovenous shunt such as ‘steal’ from the distal artery and increase in blood volume and cardiac output do not appear to be serious, the benefit, in terms of improved patency of the bypass, has yet to be demonstrated. In the case of large-diameter prostheses, precise matching of their size to the host vessels does not appear to be important. This is illustrated in the case of bifurcation prostheses. Analysis of the transmission of pulsatile energy through bifurcations shows that reflection of the pulse waves is smallest when the ratio of the sum of the cross-sectional areas of the branches to that of the trunk is approximately 1.15 (55). The pulse waves passing through bifurcations having these relative dimensions would be transmitted with the least energy loss. Most textile Dacron bifurcation prostheses have an area ratio of about 0.5 and therefore there would be substantial reflections. Despite these haemodynamic disadvantages, the performance of these prostheses, measured in terms of gross patency, has been good (67). 3 VASCULAR PROSTHESES

3.1 Fabric prostheses Fabric prostheses are manufactured in two basic constructions : woven or knitted. The woven prosthesis consists of longitudinal and circumferential yarns running alternately over and under each other at right angles, as shown in Fig. 2. This construction imparts a low porosity to the fabric and the woven prosthesis can therefore be implanted without preclotting. For this reason they are widely used in emergency situations such as resection of aortic aneurysms where blood loss must be minimized. Woven prostheses are strong and stiff, both radially and longitudinally, and have a tendency to fray at cut ends. Knitted prostheses are produced by looping yarns around needles. The number and size of needles and the yarn size determine the spacing between the yarns and therefore the porosity of the fabric. In the prosthesis shown in Fig. 3, the yarns run predominantly in the

65

( h ) W o v c n douhlc Vclour (Meadox)

Fig. 2 Scanning electron micrograph of the internal surfaces of woven Dacron prostheses. Magnification: x 60

longitudinal direction. This is known as a warp knit. In prostheses where the yarns run in the circumferential direction, the term used is weft knitting. Knitted prostheses are more conformable and are easier to handle than woven prostheses. They are less susceptible to fraying and are therefore better at retaining sutures. However, being much more porous, knitted prostheses must be preclotted with the patient’s blood prior to implantation. Recently, knitted prostheses impregnated with proteins such as albumin, collagen or gelatin have been introduced (68-70). Since they are initially impervious to blood, they require no preclotting. Woven and knitted velour prostheses were introduced in the 1970s in order to promote better healing and incorporation of the prosthesis into the surrounding tissue (71). They are constructed in such a way that loops of yarn extend into the lumen or the external wall or both. High porosity and low tensile strength are potential drawbacks of velour fabrics (72). 3.1.1 Crimps and external support

When used as a bypass, an arterial prosthesis is

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T V HOW. R GUIDOIN AND S K YOUNG

( a ) V P S0K Triaxial (Vascutck)

(h) Cooky 11 (Meadox)

Fig. 3 Scanning electron micrograph of the internal surfaces of two knitted Dacron prostheses, both warp knit structures. Magnification: x 60

Stephens et a]. (75) considered the problem of tubes of finite length subjected to internal pressure. Because of the interaction between the two failure modes and the increasing influence of the short-wave axial buckling with reduced length, short tubes were shown to undergo much lower radial deflection before collapse than long tubes. In pressurized tubes, cross-sectional flattening was constrained when the tube was bent and, consequently, buckling was found to be delayed compared to the unpressurized tube. To overcome the problem of buckling, most woven and knitted fabric prostheses have Circumferential crimps which are usually circular but sometimes helical. Crimping provides radial stiffness and reduces the tensile and compressive stresses on the outer and inner walls, and therefore delays or eliminates cross-sectional flattening when the prosthesis is bent. It also has an influence on the longitudinal extensibility and lateral flexibility of the prosthesis, thereby improving its handling. Concern has been expressed about the damage to Dacron fibres caused by crimping and the flow disturbances that may be present when blood flows along the corrugated surfaces (1). The effect on the flow becomes more pronounced in small-diameter prostheses where the height of the corrugations becomes more significant with respect to the conduit diameter. To overcome these problems manufacturers have introduced non-crimped prostheses with a helical polypropylene coil fused to the external surface of the prosthesis. This external coil provides the necessary radial stiffness to resist cross-sectional flattening and buckling and therefore achieves a similar effect to that of crimping. Externally supported Dacron prostheses have been used for femoropopliteal bypass (76). Angiographic examinations of prostheses placed below the knee appear to show that the external reinforcement prevents buckling and compression of the prosthesis during knee flexion. Although their circumferential compliance is very much reduced by the rigid polypropylene coil, externally supported Dacron prostheses appear to have better patency results than crimped prostheses. 3.2 Expanded PTFE prostbeses

implanted with a certain degree of curvature depending on the site. Prostheses placed in below-knee femoropopliteal bypass are subjected to bends of small radii of curvature during knee flexion. It is important that arterial prostheses should bend freely without buckling since this obstructs blood flow and may lead to thrombosis. The phenomenon of buckling can be appreciated qualitatively by considering a circular tube subjected to pure bending. The longitudinal tension in the outer wall and the compression on the inner wall have components directed towards the tube centre, causing the cross-section to become flattened to an oval shape (with the short axis in the plane of the curvature). Crosssectional flattening leads to reduced flexural stiffness and, if bending is increased further, buckling occurs. For a long thin tube of radius r, Brazier (73) was first to show that buckling occurs when the radial deflection on the minor axis (or ‘flattening’) reaches 2r/9. In the case of a tube of finite length, a second collapse mode, known as short-wave axial buckling, can occur when the compressed region of the tube wrinkles axially (74).

Prostheses made of expanded PTFE were introduced in 1975 (77). They are manufactured by paste-forming PTFE resin mixed with a lubricant and extruded under high pressure into tubes. The lubricant is removed and the tube is then stretched at high temperature, producing a strong expanded material of high porosity (78). The microstructure is characterized by cicumferentially oriented nodes about 5-10 pm wide by 5-100 pm long, and by interconnecting longitudinal fibrils of less than 0.5 pm diameter (Fig. 4). The mean internode distance, which is an indicator of the degree of porosity, is between 20 and 30 pm. The prosthesis does not require preclotting before implantation. The inner and outer surfaces are smooth to the naked eye. Because of its particular microstructure, expanded PTFE has very much greater tensile than compressive stiffness in the longitudinal direction. In addition, the circumferentially oriented nodes provide some radial stiffness, and consequently the prosthesis is relatively resistant to buckling when bent. Although early designs of expanded PTFE were susceptible to creep and dilatation, this problem @IMechE 1992

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ENGINEERING DESIGN OF VASCULAR PROSTHESES

Table 3 Mechanical properties of proprietary polyurethanes and other elastomers Elastic modulus (at % elongation)

(a) Internal wrlace of Gore-Tex showing circumferentially oricntc'tl node\ ahout S pi11 wide hq 15-XO prn long w i t h intcrconiicciing longitudinal f'ibrils of than 0 . 5 prn diameter

(h) External surlace olGoi-e-Tcr. \bowing layct- 01' low poro\iiy rcinlorcciiicni

Fig. 4 Scanning electron micrograph of expanded PTFE prostheses. Magnification: x loo0

has largely been overcome either by incorporating a thin reinforcing layer of fibrous expanded PTFE of low porosity on the external surface (Gore-Tex@, W. L. Gore and Associates) or by using thicker walls (Impra, Impra Inc.). Expanded PTFE prostheses with external helical support coils are also available. Because of their improved kink resistance they are often used for belowknee femoropopliteal bypass. In the absence of a suitable autologous vein, the expanded PTFE has become the most commonly used prosthesis material for peripheral arterial bypass. It is also widely used as a bridge graft for secondary vascular access in patients undergoing haemodialysis and is favoured by surgeons because of the ease with which it can be thrombectomized (79).

3 3 Small-diameter vascular prostheses Over the past fifteen years several small-diameter vascular prostheses have been developed. Most of the designs

Tensile

Elongation

strength at break -~

Elastomer*

MPa

MPa

%

Reference

Avcothane-51 Biomer Pellethane Texin MD85A Toyobo TM5 Silastic Q7-4735 Silastic 47-4765

4.81(50) 5.00(50) 5.03(100) 7.55(50) 6.18(50) 1.03(200) 3.03(200)

19.81 43.16 53.98 38.95 42.08 9.03 7.58

1265 1231 284 860 1263 1310

80 80 81 80 80 82 82

900

*Avcothane-51 Block copolymer or 90% polyether urethane and 10% polydimethyl siloxanc. Now sold as Cardiothane-51, Kontron Instruments, Everett, USA Biomer Segmented polyether urethane urca, Ethicon Inc., USA Segmented polyether urethane, Dow Chemical, USA Pellethane Texin MD85A Segmented polyether urethane, Sumitomo Bakelite Company, Japan Tohoyo TM5 Segmented polyether urethane, Toboyo Company, Japan Silastic Q7-4735 Medical grade elastomer. Dow Coming Corporation, USA and Q7-4765

have been based on the use of polyurethanes because of their low thrombogenicity. Polyurethanes are available in a wide range of chemical composition and mechanical properties (see Table 3). Those that have been used for vascular prostheses have lower tensile moduli than Dacron or PTFE while having adequate strength and good fatigue properties. Moreover, they lend themselves to a wide variety of manufacturing techniques. Some polyurethanes are thermo-formable and others can be solvent cast or precipitated in a non-solvent such as water. Fibres can be drawn or extruded to produce a wide variety of microporous structures. Lyman et al. (14) produced 4 mm internal diameter prostheses from copolyether-urethane urea by a coagulation process. A 4 mm diameter glass rod is dipped first in a solution of the polyurethane and then in water to precipitate the polymer. After drying, a Dacron mesh is slipped over the coated rod and the dipping procedure is repeated. The resulting Dacron-reinforced tubular conduit is then removed from the rod. The prosthesis is porous and has a controlled void structure similar to a spongy material. The mechanical properties of this prosthesis have been studied by Stewart and Lyman (83) using a linear finite elastic model. Their data show that the prosthesis is highly anisotropic with elastic moduli of 2.24 and 6.97 MPa in the circumferential and longitudinal directions respectively. The high elastic moduli compared with those of the natural arteries C0.69 MPa at 100 mmHg for the femoral artery (W)]and the anisotropy are due to the Dacron mesh which is incorporated into the prosthesis wall. Hayashi et al. (85) have developed a prosthesis having a similar microstructure to that of Lyman's. The polymer used was a polyurethane to which was added polydimethylsiloxane (silicone rubber) to improve the blood compatibility of the base material. Porous tubes of internal diameter 3 mm were produced by precipitation. In this case a reinforcing polyester mesh was placed on the outside of the prosthesis. The prosthesis was generally more porous on the inner surface and the ratio of the area of the pores to the area of the prosthesis surface ranged from 0.35 to 0.41. The stress-strain relationship was effectively linear up to a strain of 100 per cent. The elastic modulus, which varied from 0.056 to 0.08 MPa, was much lower than that reported for the

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femoral artery (84). A wall thickness of 0.8 mm was needed to obtain the same compliance as the femoral artery. The external reinforcing polyester mesh was necessary since the tensile strength was only between 0.38 and 0.43 MPa. To overcome the problem of the reduction in compliance that accompanies perigraft fibrosis, a new prosthesis which is thought to provide volume compliance by wall compression has been developed (19). It is claimed that the internal compression of the wall can accommodate changes in volume without significant external wall movement. However, it is unclear how wall compresssion can occur after the prosthesis has healed and the internal and external capsules have formed. The prostheses consist of two concentric layers bonded together. The thicker porous outer layer is produced by a precipitation process while the inner layer of approximately 15 pm is made by solvent casting. Prostheses of internal diameter 4 mm and 6 cm long have been implanted in sheep with a cumulative patency of 78 per cent at 7 months (86). The relatively thick wall (1 mm) provided increased radial stiffness and resistance to buckling when subjected to bending loads. White and co-workers (87, 88) developed a novel process, termed replamineform, for fabricating microporous prostheses. The process is based on replicating the microporous calcite structure of sea urchin spines. The spines are treated in sodium hypochlorite solution to remove all organic matter. They are then machined into tubes of the required dimensions, typically 4-6 mm internal diameter and 1 mm wall thickness. The poIyrner solution, polyurethane or silicone rubber, is injected into the calcite structure and, after curing, the calcite is dissolved, leaving a microporous tube with a pore size of 18-25 pm. Replamineform prostheses made from polyurethane had a compliance of 0.083 per cent/ mmHg at implantation, but this decreased by half 8 weeks later (89).This technique allows prostheses of different pore sizes to be fabricated and has proved useful in studying the effect of pore size on the healing characteristics of vascular prostheses. However, the maximum length of conduit that can be fabricated by this technique is limited by the length of the sea urchin spines. Various techniques of fabricating non-woven microfibrous vascular prostheses have been reported. The fibres are formed from the polymer solution by extrusion through a fine orifice (17), by a spraying method (16) or by electrostatic spinning (15). The fibres are wound helically on to a rotating mandrel with the direction of the helix being reversed with each successive layer of fibres. When the required wall thickness is attained the process is stopped and the fibrous tube is removed from the mandrel. The diameter of fibres produced by extrusion through a 200 pm orifice ranges from 10 to 30 pm (17). The angle, &a, of the helical fibres to the mandrel axis can be controlled. This is an important process parameter since it affects the anisotropic properties and the kink resistance: the prosthesis becomes more resistant to buckling as a increases. The mechanical properties of these prostheses have been extensively studied (90, 91). The cylindrical properties were examined using a twodimensional incremental model and an expression relating compliance with internal and external radii, circumferential elastic modulus and Poisson's ratios.

Prostheses made with various winding angles were examined and it was found that those with a of 45" had incremental moduli and Poisson's ratios closely matched to those of the iliac arteries at physiological pressures. Kowligi et al. (92) recently described a method (flotation-precipitation) that combines spraying and precipitation techniques for producing fibrous grafts. The polyurethane solution is sprayed through a nozzle on to flowing water in a water tank. The precipitated polyurethane fibres that form on the water surface are transferred to a rotating mandrel where layers of fibrous polymer accumulate to form a tubular conduit. The mandrel is then dried in a vacuum oven at 50-60°C to remove all excess solvents. The prosthesis is removed from the mandrel after soaking in water for one hour. Prostheses fabricated by this technique have a wide distribution of pore sizes, ranging from 10 pm to 1.5 mm. The compliance of prostheses of internal diameter 3 and 4 mm varied from 0.272 to 0.374 per cent/kPa (0.0360.050 per cent/mmHg). The fibres produced by electrostatic spinning of a proprietary polyether urethane urea (Biomer@)are at least one order of magnitude smaller (1-2 pm) than those produced by extrusion. To the naked eye the inner and outer surfaces of the electrostatically spun prostheses are smooth. The small-diameter fibres have advantages in terms of healing since tissue attachment is more delicate and closer to the fibres. The spacing between the fibres is on average less than 20 pm and this helps to limit the penetration of connective tissue cells to a depth of 50 pm and to maintain the compliance of the prosthesis (27). The microfibrous prostheses have been shown to be highly anisotropic, and control of their anisotropy can be achieved by modifying the manufacturing process parameters (60).Because of the symmetry in their manufacture, the prostheses behave macroscopically as multi-layered cross-ply tubes. Thus they can be modelled as orthotropic cylinders and when they are subjected to physiological loads (internal pressure and axial load), only normal stresses need to be considered (93). When designing microfibrous vascular prostheses with known mechanical properties it is important to know the effect that each of the electrostatic spinning process parameters has on these properties. This investigation has been carried out previously and it was shown that the electrostatic spinning process is stable and reproducible and prostheses with a wide range of mechanical properties can be produced. Assuming the prostheses to be orthotropic, incompressible and homogeneous, the non-linear elastic properties were analysed (60, 94). A model based on a polynomial form of strain energy density function was used (93). It was shown that the same constitutive equation was applicable to prostheses of different diameters and wall thicknesses, provided that they were manufactured using the same polymer and that the electrostatic spinning conditions were kept constant. From the model, the effect of diameter and wall thickness of electrostatically spun prostheses on compliance (at a mean pressure of 100 mmHg) were calculated (Fig. 5). All of the small-diameter prostheses described above can be considered to be hydraulically smooth. High resistance to buckling is an important requirement for these prostheses since they may be used for below-knee

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Internal diameter mm a 3.0 b 4.0 c 5.0 d 6.0

30 24 I

13 X

c,

18

30

Wall thickness mm

24 I

X

2

c,

2

2 .

2 12 .

6?

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12

G?

6

6

0

0

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0.45

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,

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(a) Wall thickness for various internal diameters

I

I

1

4.0 4.5 5.0 Internal diameter mm

I

I

5.5

6.0

(b) Internal diameter for various wall thicknesses

Fig. 5 Variation in compliance (at a mean pressure of 13.3 kPa) of electro-

statically spun polyurethane prostheses femoropopliteal bypass. The kink resistance can be enhanced by increasing the wall thickness or by providing a reinforcing coil within the wall or on the external surface. The drawback is that this reduces the compliance, thereby nullifying an important advantage of elastomeric prostheses. With prostheses of fibrous construction it is possible to improve the kink resistance by modifying the winding angle, the distribution of fibre diameters and the inter-fibre spacing, since these have an effect on the radial and longitudinal stiffnesses. Despite the low thrombogenicity and the more desirable mechanical properties of polyurethanes compared with Dacron and PTFE, small-diameter vascular prostheses made from polyurethane have not been entirely satisfactory. One cause for concern is the potential for long-term degradation of polyurethanes (B), leading to alterations in their chemical, physical, mechanical and biochemical properties. This may result in loss of strength and release of potentially toxic breakdown products (9). Some polyurethanes are more resistant to degradation than others and it is more important therefore to pay special attention to the material selection (12). The techniques of manufacturing microporous prostheses that have been described are generally versatile and they may be adapted for new materials with improved stability and non-thrombogenicity. 4 CONCLUSION

If the performance of an arterial prosthesis is measured in terms of patency it can be concluded that the major problems in the design and development of largediameter prostheses have largely been overcome, although further improvements will continue to be made. In the case of peripheral arterial bypass, the autologous vein remains the graft of choice. Alternative prostheses currently used in the absence of a suitable vein do not perform satisfactorily and there is a real need for more successful small-diameter prostheses. In the development of these prostheses the availability of a material with desirable properties (that is nonthrombogenicity, long-term stability in the body and good physical and mechanical properties) is of primary importance. This material is not yet available, but work

in recent years has gathered pace. Major directions for this research are the development of new polymeric materials, the incorporation of anti-platelet drugs in, and surface modifications of, existing materials, and seeding of endothelial cells on to currently available prostheses. As well as the problems associated with the choice of material, there is still uncertainty regarding the detailed specification for the design of small-diameter prostheses. It has yet to be determined which requirements are essential and which are of secondary importance. In this respect, more information is required on the role of the wall mechanics and haemodynamics, particularly at the anastomoses, on the in viuo performance of smalldiameter prostheses. This is important since these factors can be influenced by the design of the prosthesis and the configuration of the bypass. ACKNOWLEDGEMENTS

The first author would like to thank Laval University and the Faculty of Medicine, Laval University, for their generous support and the Laboratoire de Chirurgie Experimentale for their hospitality during the period when this work was carried out. REFERENCES 1 Charlesworth, D. Arterial replacements. In Recent advances in surgery (Ed. S . Taylor), Vol. 10, 1980, pp. 93-111 (Churchill Livingstone). 2 Haimovici, H. Ideal arterial graft: an unmet challenge-scope and limitations. Surgery, 1980,92, 117-1 19. 3 Harris, P. L., How, T. V. and Jones, D. R. Prospectively randomized clinical trial to compare in siru and reversed saphenous vein grafts for femoropopliteal bypass. Br. J . Surg., 1987,74, 252-255. 4 Lytle, B. W., Cosgrove, D. M., Loop, F. D., Borsh, J., Goormastic, M. and Taylor, P. C. Perioperative risk of bilateral internal mammary artery grafting: analysis of 500 cases from 1971 to 1984. Circulation, 1986, 74(Suppl. III), 111-37-111-41. 5 Myhre, H. 0.Peripheral reconstruction when adequate saphenous vein is lacking. Acta Chir. Scand. Suppl., 1985,529,73-75. 6 Bergan, J. J., Yao, J. S. T., Flinn, W. R. and Graham, L. M. Prosthetic grafts for the treatment of lower limb ischaemia: present status. Br. J. Surg., 1982, 69(Suppl.), S34-S37. 7 Brewster, D. C, LaSalle, A. J, Robison, J. G., Strayhorn, E. C. and Darling, R. C. Factors affecting patency of femoropopliteal bypass grafts. Surg. Gynecol. Obsfet.,1983, 157,437-442.

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8 Paynter, R. W., Askill, I. N., Glick, S. H.and Guidoin, R. The hydrolytic stability of Mitrathane (a polyurethane u r e a b a n X-ray photoelectron spectroscopy study. J. Biomed. Mater. Res., 1988, 22,687-698. 9 Lemm, W. and Bucherl, E. S. The degradation of polymeric materials. In Polymers: their properties and blood compatibility (Ed. S. Dawidsj, 1989, pp. 247-256 (Kluwer Academic Publishers). 10 Stokes, P, Urbanski, P. and Upton, J. The in uiuo auto-oxidation of polyether polyurethane by metal ions. J . Biornater. Sci., Poljmer Edn, 1990,1,207-230. 11 Griesser, H. J. Degradation of polyurethanes in biomedical applications-a review. Polymer Degradation and Stability, 1991, 33, 329--354. 12 Pinchuk, L., Esquivel, M. C., Martin, J. B. and Wilson, G. J. CorethaneTM:a new replacement for polyether urethanes for longterm implant applications. Trans. Seventeenth Annual Meeting of Society for Biomaterials, 1-5 May 1991, Arizona, p. 98. 13 Edwards, W. S. Arterial grafts. Past, present, and future. Arch. Surg., 1978,113,1225-1233. 14 Lyman, D. J, A h , D., Jackson, R. and Knutson, K. Development of small diameter vascular prostheses. Trans. Am. Sac. Art$ Intern. Organs, 1977,23,253-261. 15 Annis, D., Bornat, A., Edwards, R. O., Higham, A,, Loveday, B. and Wilson, J. An elastomeric vascular prosthesis. Trans. Am. SOC. Art$ Intern. Organs, 1978,24,209-214. 16 Plank, H. and Ehler, P. Blood vessel prostheses. UK Pat., GB 2,015,118B, 1982. 17 Leidner, J., Wong, E. W. C, MacGregor, D. C. and Wilson, G. J. A novel process for the manufacturing of porous grafts: process description and product evaluation. J. Biomed. Mater. Res., 1983, 17,229-247. 18 Kowligi, R. R. and Calcotte, R. W. Vascular prostheses from polyurethanes: methods for fabrication and evaluation. In High performance biomatmiuls (Ed. M. Szycher), 1991, pp. 425-442 (Technomic Press, Lancaster, Pa.). 19 Charlesworth, D., White, E. T. and Kent, S. UK Pat., GB 2,130, 521B, 1 1 December 1985. 20 Sauvage, L. R, Berger, K. E., Mansfield, P. B, Wood, S. J., Smith, J. C. and Overton, L. B. Future directions in the development of arterial prostheses for small and medium caliber arteries. Surg. Clin. N. Am., 1974,54,213-228. 21 Voorhees, A. B., Jaretzki, A. and Blakemore, A. H. The use of tubes constructed from Vinyon ‘N’ cloth in bridging arterial defects. Ann. Surg., 1952,135,332-336. 22 Guidoin, R., King, M., Marceau, D., Cardou, A., de la Faye, D., Legendre, J. M. and Blais, P. Textile arterial prostheses: is water permeability equivalent to porosity? J , Biomed. Mater. Res., 1987, 21,6547. 23 Wesolowski, S. A., Fries, C. C., Karlson, K. E, De Bakey, M. and Sawyer, P. N. Porosity: primary determinant of ultimate fate of synthetic vascular grafts. Surgery, 1961,50,91-96. 24 Wesolowski, S. A., Fries, C. C., Hennigar, G., Fox, L. M., Sawyer, P. N. and Sauvage, L. R. Factors contributing to long-term failure in human vascular prostheses. J . Cardiovasc. Surg., 1964, 5, 544567. 25 Long, J, Tan, E., DeSantis, S, Shors, E., Klein, S., Goldberg, L., Nelson, R, Uitto, J. and White, R. Implant microstructure and collagen synthesis. Trans. Am. SOC.Art$ Intern. Organs, 1982, 28, 195-199. 26 White, R.A. Vascular prostheses. Present status and future development. In Blood compatibility (Ed. U. F. Williams), Vol. 2, 1987, pp. 47-61 (CRC Press, B o a Raton, Fla.). 27 de Cossart, L., How, T. V. and Annis, D. A two year study of the performance of a small diameter polyurethane (Biomer@)arterial prosthesis. J. Cardiouasc. Surg., 1989,30, 38&394. 28 Schlichting, H. Boundary layer theory, 7th edition, 1979 (McGrawHill). 29 Prandtl, L. and Tietjens, 0. C. Applied hydro- and aeromechanics, 1934 (Dover Publications, New York). 30 Savvides, C. N. and Gerrard, J. H. Numerical analysis of the flow through a corrugated tube with application to arterial prostheses. J . Fluid Mech., 1984,138, 129-160. 31 Herring, M., Dilley, R., Gardner, A. and Glover, J. The effects of crimping on the healing of prosthetic arterial grafts. 1. Cardiouasc. Surg., 1980,21,596-603. 32 Hamlin, G. W., Rajah, F. M., Crow, M. J. and Kester, R. C. Evaluation of the thromhgenic potential of three types of arterial grafts studied in an artificial circulation. Br. J. Surg., 1978, 65, 272-276.

33 Takebayashi, J, Kamatani, M., Katagami, Y., Hayashi, K., Yanagi, Z., Nakaue, T., Kise, L,Tokura, K. and Asada, K. A comparative study on the patency of crimped and noncrimped vascular prostheses, with emphasis on the earliest morphological changes. J . Surg. Rex, 1975,19,209-218. 34 Harrison, J. H. Synthetic materials as vascular prostheses. Am. J . Surg., 1958,95, 16-24. 35 How, T. V., Clarke, R. M. and Annis, D. Uniaxial tensile properties of the Liverpool artificial artery. In Biomedical polymers, 1982, pp. 171-179 (Biological Engineering Society). 36 Cohen, J., Litwin, S. B, Aaron, A. and Fine, S. The rupture force and tensile strength of canine aortic tissue. J. Surg. Res., 1972, 13, 321-333. 37 Baird, R. N., Kidson, I. G., L’ltalien, G. J. and Abbott, W. M. Dynamic compliance of arterial grafts. Am. J. Physiol., 1977, 233, H568-HS72. 38 Nunn, D. B., Freeman, M. H. and Hndgins, P. C. Postoperative alteration in size of Dacron@ aortic grafts. Ann. Surg., 1978, 189, 741-745. 39 Nunn, D. B., Carter, M. M., Donohue, M. T. and Hudgins, P. C. Postoperative dilation of knitted Dacron aortic bifurcation graft. J. Vasc. Surg., 1990,12,291-297. 40 Maarek, J. My Guidoin, R., Anbin, M. and Prud’homme, R. E. Molecular weight characterization of virgin and explanted polyester arterial prostheses. J. Biomed. Mater. Res., 1984,18,881-894. 41 ANSIiAAMI VP20-1986 American national standard for vascular graft prostheses (Association for the Advancement of Medical Instrumenlation, Arlington, Mass.). 42 Edwards, W. S, Snyder, R. W., Botzko, K. and Larking, J. Comparison of durability of tensile strength of Teflon and Dacron grafts. In Grafi materials in vusculur surgery (Ed. H.Dardik), 1978, pp. 169-182 (Symposia Specialists, Miami, Fla.). 43 Hasegawa, M. and Azuma, T. Mechanical properties of synthetic arterial grafts. J. Biomechanics, 1979, 12, 509-51 7. 44 Seifert, K. B., Albo, D., Knowlton, H. and Lyman, D. J. Effect of elasticity of prosthetic wall on patency of small-diameter arterial prostheses. Surg. Forum, 1979,30,206-208. 45 Lyman, D. J., Lyman, E. C. and Eichwald, E The effect of graft compliance on the patency of small-diameter polyurethane vascular grafts. Third World Biomaterials Congress, Kyoto, Japan, 21-25 April 1988, p. 55. 46 Kidson, 1. G. and Abbott, W. A. Low compliance and arterial graft occlusion. Circulation, 1978,58(Suppl. l), I- 1-1-4. 47 Abbott, W. A., Megerman, J, Hasson, J. E., L’ltalien, G. and Warnock, D. F. Effect of compliance mismatch on vascular graft patency. J. Vasc. Surg., 1987,5,376-382. 48 Fisher, A. C, How, T. V., de Cossart, L. and An&, D. The longer term patency of a compliant small diameter arterial prosthesis: the effects of the withdrawing of aspirin and dipyridamole therapy: the effect of reduced compliance. Trans. Am. SOC. Art$ Intern. Organs, 1985,31,324-328. 49 Uchida, N., Emoto, H., Kamhic, H., Harasaki, H., Chen, J. F., Hsu, S. H, Murabayashi, S. and Nose, Y. Compliance effect on patency of small diameter vascular grafts. Trans. Am. Soc. Art$ Intern. Organs, 1989,35,556-558. 50 Kinley, C. E. and Marble, A.E. Compliance: a continuing problem with vascular grafts. J . Cardiovasc. Surg., 1980, 21, 163-170. 51 Rittgers, S. E., Karayannacos, P. E., Guy, G. F, Nerem, R M.. Shaw, G. M., Hostetler, J. R. and Vasco, J. S. Velocity distribution and intimal proliferation in autologous vein grafts in dogs. Circ. Res., 1978,42,792-801. 52 Sottinrai, V. S., Yao, J. S. T, Flinn, W. R. and Batson, R. C. Intimal hyperplasia and neointima: an ultrastructural analysis of thrombosed grafts in humans. Surgery, 1983,93, 809-817. 53 Mehigan, D. G., Fitzpatrick, B, Browne, H. I. and BouchierHayes, D. J. Is compliance mismatch the major cause of anastomotic arterial aneurysms? Analysis of 42 cases. J. Cardiouasc. Surg., 1985,26,147-150. 54 Miyawaki, F., How, T. V. and Annis, D. Effect of compliance mismatch on flow disturbances in a model of an arterial graft replacement. Med. Bid. Engng Comput., 1990,28,457-464. 55 Womersley, J. R. Oscillatory flow in arteries. TI: the reflection of the pulse wave at junctions and rigid inserts in the arterial system. Phys. Med. B i d , 1958,2,313-323. 56 Langewouters, G. J., Wesseling, K. H. and Goedhard, W. J. A. Age-related changes in viscoelasticily of normal and arteriosclerotic human aortas. In Biomechanics: current interdisciplinary research (Eds S. M. Perren and E. Schneiderj, 1985, pp. 245-250 (Martinus Nijhof).

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57 Hokanson, D. E.and Strandness, D. E. Stress-strain characteristics of various arterial grafts. Surg. Gynecol. Obstet., 1968, 127,57-60. 58 Usui, Y., Goff, S., Sauvage, L. R., Wu, H. D., Robel, S. B. and Walker, M. Effect of healing on compliance of porous Dacron grafts. Ann. Vasc. Surg., 1988, 2, 120-126. 59 H a d , K. C., McCabe, C., Abbott, W. M, Fallon, J. and Megerman, J. Current PTFE grafts. A biomechanical, scanning electron, and light microscopic evaluation. Ann. Surg., 1982,195,456-463. 60 How, T. V. and Clarke, R. M. The elastic properties of a polyurethane arterial prosthesis. J . Biomechanics, 1984, 17,597-608. 61 Walden, R., L’Italien, G. J, Megennan, J. and Abbott, W. M. Matched elastic properties and successful arterial grafting. Arch. Surg., 1980,115, 1166-1169. 62 How, T. V. and Annis, D. Viscoelastic behaviour of polyurethane vascular prostheses. J . Biomed. Mater. Res., 1987,21, 1093-1108. 63 Black, R. A. and How, T. V. Attenuation of flow disturbances in tapered vascular grafts. Trans. ASME, J . Biomech. Engng, 1989, 111,303-310. 64 Shultz, J. S. and Lindenaner, S. M. Ex vivo evaluation of blood surface interfaces. In Biologic and synthetic vascular prostheses (Eds J. C . Stanley et at.), 1982, pp. 113-129 (Grune and Stratton, New York). 65 Sauvage, L. R, Walker, M. W., Berger, K., Robel, S. B., Lischko, M. M., Yates, S. G. and Logan, G. A. Current arterial prostheses: experimental evaluation by implantation in the carotid and circumflex coronary arteries of the dog. Arch. Surg., 1979, 114, 687691. 66 Harris, P. L. and Campbell, H. Adjuvant distal arteriovenous shunt with femorotibial bypass for critical ischaemia. Br. J . Sury., 1983,70,377-380. 67 Newman, D. L. and Bowden, N. L. R. An in viuo assessment of aorto-iliac bifurcation grafts. J . Cardiovasc. Surg., 14,298-302. 68 Guidoin, R, Snyder, R., Martin, L., Botzko, K, Marois, M., Awad, J., King, M. W, Domurado, D., Bedos, M. and Gosselin, C. Albumin coating of a knitted polyester arterial prosthesis: an alternative to preclotting. Ann. Thorac. Swg., 1984,37,457-465. 69 Quinones-Baldrich, W. J., Moore, W. S., Siomek, S. and Chvapil, M. Development of a leak-proof knitted Dacron vascular prosthesis. J. Vasc., Surg., 1986, 3, 895-903. 70 Guidoin, R, Marceau, D, Rao, T. J., King, M. W., Merhi, Y., Roy, P. E., Martin, L. and Duval, M. In uitro and in viuo characterization of an impervious polyester arterial prosthesis: the Gelseal Triaxial graft. Biomaterials, 1987, 8, 433-41 1. 71 King, M. W. Designing fabrics for blood vessel replacement. Can. Text. J., 1991,108,24-30. 72 Guidoin, R., King, M. W, Therrien, M, Douville, Y, Debille, E., Boyer, D., Simoneau, S. and Tremblay, L. Woven velour polyester arterial grafts with polypropylene wrap: a cosmetic change or improved design? In High performance biomuterials (Ed. M. Szycher), 1991, pp. 449-474 (Technomic Press, Lancaster, Pa.). 73 Brazier, L. G. On the flexure of thin cylindrical shells and other thin sections. Proc. R . Soc., 1927,116, 104-114. 74 Fabian, 0. Collapse of cylindrical, elastic tubes under combined bending, pressure and axial loads. Int. J. Solids Structs, 1977, 13, 1257-1 270. 75 Stephens, W. B., Starnes, J. H. and Almroth, B. 0. Collapse of long cylindrical shells under combined bending and pressure loads. Am. Inst. Aeronaut. Astronaut. J., 1975, 13,20-25.

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76 Kenney, D. A., Sauvage, L. R., Wood, S. J., Berger, K., Davis, C. C., Smith, J. C., Rittenhouse, E. A., Hall, D. G. and Mansfield, P. B. Comparison of noncrimped, externally supported (EXS) and crimped, nonsupported Dacron prostheses for axillofemoral and above-knee femoropopliteal bypass. Surgery, 1982,92,931- 946. 77 Boyce, B. Physical characteristics of expanded polytetrafluoroethylene grafts. In Biologic and synthetic vascular prostheses (Eds J. C. Stanley et al.), 1982, pp. 553-560 (Grune and Stratton, New York). 78 Gore, R. W. Very highly stretched polytetrafluoroethylene and process thereof. US Pat. 3,962,153, 8 June 1976. 79 Tilney, N. L. and Whittemore, A. D. Dialysis access in difficult patients. In Vascular surgery (Eds P. R. F. Bell and N. L. Tilney), 1984, pp. 175-192 (Butterworth, London). 80 Hayashi, K., Takano, H., Matsuda, T. and Umezn, M. Mechanical stability of elastomeric polymers for blood pump applications. J. Biomed. Mater. Rex, 1985, 19, 179-193. 81 McMillin, C. R. Physical testing of elastomers for cardiovascular applications. Art$ Organs, 1983,7,78-91. 82 Dow Corning Corporation. Silastic medical grade ETR elastomers (Q7-4735, Q7-4750, Q7-4765), product information, 1982. 83 Stewart, S. F. C. and Lyman, D. J. Finite elasticity modelling of the biaxial and uniaxial properties of compliant vascular grafts. Trans. ASME, J . Biomech. Engng, 1988,110,344-348. 84 Bergel, D. H. The static elastic properties of the arterial wall. J . Physiol., 1961, 156, 445-457. 85 Hayashi, K., Takamizawa, K., Saito, T., Kira, K., Hiramatsu, K. and Kondo, K. Elastic properties and strength of a novel smalldiameter, compliant polyurethane vascular graft. 1. Biomed. Mater. Rex, 1989,23,229-244. 86 Underwood, C. J., Tait, W. F. and Charlesworth, D. Design considerations for a small bore vascular prosthesis. Int. J . Art$ Organs, 1988, 11, 272-276. 87 White, R. A., White, E. W., Hamon, E. L., Rohner, R. F. and Webb, W. R. Preliminary report: evaluation of tissue ingrowth into experimental replamineform vascular prostheses. Surgery, 1976,79,229-232. 88 Hiratzka, L. F., Goeken, J. A., White, R. A. and Wright, C. B. I n vivo comparison of replamineform, silastic, and bioelectric polyurethane arterial grafts. Arch. Sury., 1979, 114, 698-702. 89 White, R., Goldberg, L., Hirose, F, Klein, S., BOSCO, P., Miranda, R., Long, J., Nelson, R. and Shors, E. Effect of healing on small internal diameter arterial graft compliance. Biomnter. Med. Dev. Art$ Organs, 1983,11,21-29. 90 Lee, J. M. and Wilson, G. T. Anisotropic viscoelastic properties of vascular graft materials tested at low strain rates. Biomaterials, 1986,7,423-431. 91 Nahon, D., Lee, J. M. and Wilson, G. J. A two-dimensional incremental study of the static mechanical properties of vascular grafts. Clin. Mater., 1986, 1, 177-197. 92 Kowligi, R. R., von Maltzahn, W. W. and Eberhart, R. C. Synthetic vascular graft fabrication by precipitation-flotation method. Trans. Am. Soc. Art$ Intern. Organs, 1988,34,800-804. 93 Vaishnav, R. N., Young, J. T., Janicki, J. S. and Patel. D. J. Nonlinear anisotropic elastic properties of canine aorta. Biophys. J., 1972,12,1008-1027. 94 How, T. V. Elastic deformation of a tapered vascular prosthesis. J . Mater. Sci.: Mater. Med., 1991, 2,94-100.

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Engineering design of vascular prostheses.

The replacement and bypass of arteries of diameter greater than 6 mm with textile vascular prostheses has proved very successful since they were first...
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