implanted cardiovascular polymers: natural, synthetic and bio-inspired

Progress in Polymer Science 33 (2008) 853–874

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Progress in Polymer Science

journa l homepage: www.e lsev ier .com/ locate /ppolysc i

Implanted cardiovascular polymers: Natural, synthetic andbio-inspired

Subbu Venkatraman ∗, Freddy Boey, Luciana Lisa LaoSchool of Materials Science and Engineering, Nanyang Technological University, 50 Nanyang Avenue, Singapore 639798, Singapore

a r t i c l e i n f o

Article history:Received 14 December 2007Received in revised form 30 July 2008Accepted 30 July 2008Available online 22 August 2008

Keywords:Vascular graftEndothelial cell seedingExpanded poly(tetrafluoroethylene)PolyesterPoly(l-lactide)Biodegradable stentDrug-eluting stentVessel patency

a b s t r a c t

This review details the use of polymeric biomaterials used in implantable cardiovasculardevices. Specifically, the role of the polymer in two major types of device, the vascular graftand the cardiovascular stent, is examined critically. In these two devices, the device per-formance is critically dependent on the polymer; the material requirements are detailed,and the shortcomings of currently used polymers highlighted with a view to furtheringthe development of new materials. In each category, synthetic polymers, polymers of nat-ural origin, and polymers that mimic proteins but are synthesized, have all been evaluatedwith varying degrees of success. We find that the totally artificial graft is still the pre-ferred option when autologous vessels are not available; the development of a completelytissue-engineered graft awaits improvements in scaffold materials as well as in tissue reac-tor engineering. In the field of stents, current consensus is driving the substitution of abiodegradable, polymeric stent for the biostable metallic one. Although various biodegrad-able polymers have been evaluated, the hydrolytically degradable polyesters continue to bethe polymer of choice. New developments in biodegradable polymers are highlighted, and

Artificial proteinsCollagen and elastin their performance in terms of biocompatibility and controlled degradability are presented.

The outlook for the next decade appears hopeful, with improvements in cell-seeding and
cell growth techniques expected to enhance the performance of both types of implanteddevice.
© 2008 Elsevier Ltd. All rights reserved.

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8541.1. Scope and rationale of this review. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8541.2. Cardiovascular therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8551.3. Types of implantable polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8551.4. Outline of the review . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 855

2. The artificial blood vessel . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8562.1. The need . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8562.2. Requirements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 856

2.2.1. Viscoelasticity of blood vessels . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8562.2.2. Biocompatibility aspects . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 857

2.3. Synthetic polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8582.3.1. Synthetic polymers, fiber-based . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 858

∗ Corresponding author. Fax: +65 67909081.E-mail address: [email protected] (S. Venkatraman).

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3.7. Status and future directions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8704. Outlook for polymeric cardiovascular implants . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 871

. . . . . . . .. . . . . . . .

Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1. Introduction

Ever since the introduction of poly(methyl methacry-late) in the 1940s as an artificial lens, implanted polymershave contributed considerably to the advancement of ther-apeutic options in medicine. Although there are a fewinstances of polymers being used in implanted diagnosticdevices, their use has been more extensive in therapy. Thisreview focuses on progress in the use of implantable poly-meric biomaterials in therapeutic medicine, particularly inthe treatment of cardiovascular disease and diabetes.

The reasons to restrict this review to cardiovascular dis-ease are straightforward: (a) the need is greatest for thisdisease; (b) a large number of the world’s population suf-fers from cardiovascular problems; (c) the potential forpolymers is high for solving some of the problems asso-ciated with this disease condition; and (d) much of theresearch into new polymeric biomaterials is aimed towardsresolving problems in cardiovascular disease. This is notto deny that implanted polymers have found success else-where: they have been successfully used to treat ocularconditions [1,2]; prostate cancer [3,4]; brain cancer [5];female contraception [6]; skin burns [7]; dwarfism [8], tonote just a few examples. However, many advances in poly-meric biomaterials have occurred in the search for bettercardiovascular implants, and to a much lesser extent, for anartificial pancreas; natural, synthetic and bio-inspired poly-mers have been evaluated with varying degrees of success.It is the purpose of this review to discuss the establishedsuccesses as well as potential successes for such materials,
cardiovascular devices.
At the outset, we note that polymeric implants are mostuseful for treating conditions associated with so-called softtissues; harder tissues such as bone are better treated withmetals and ceramics. Again, the majority of disease condi-

. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 871. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 871

tions are associated with problems of soft tissues, hencethe dominance of research into implantable polymers.

1.1. Scope and rationale of this review

The review will address progress made in polymericimplants over the last decade, dealing primarily with ther-apeutic devices for cardiovascular disease. The implantsfor the cardiovascular system are wide-ranging in function,and include the following:

(a) Small-diameter blood vessel replacements.(b) Polymeric stents and polymer-coated stents.

The biomaterial requirements in these applications aredemanding and varied enough to attract thousands ofresearch papers. For example, for artificial blood vessels,there were about 900 publications in the last decade;another 55 on polymeric heart valve, and more than 500 onpolymeric stents. Correspondingly, about 35 patents havebeen granted for artificial blood vessels, 5 for heart valves,and 30 for polymeric stents, in the past 30 years. Some prod-ucts in each of these areas have been commercialized withvarying degrees of success, but none have satisfied all thefunctional requirements in their category. Such a productawaits further developments in polymer materials.

In this review, we highlight the performance require-ments of the device materials, and then critically evaluatehow the various types of polymers meet those require-ments. Such an approach helps to focus on the functional

854 S. Venkatraman et al. / Progress in Polymer Science 33 (2008) 853–874

2.3.2. Expanded PTFE (ePTFE) grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8582.3.3. Polyurethane grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8592.3.4. Combination (bio-inspired) structures . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8612.3.5. Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 861

2.4. Biodegradable polymers, synthetic . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8612.4.1. As graft . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8612.4.2. As a scaffold material . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 862

2.5. Natural materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8632.6. Bio-inspired materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8642.7. EC-seeded grafts . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8652.8. Commercial products . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8662.9. Summary . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 866

3. Polymeric, biodegradable stents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8663.1. Requirements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 867

3.1.1. Biocompatibility . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8673.1.2. Viscoelastic requirements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 868

3.2. Natural polymers . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8693.3. Bio-inspired materials . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8693.4. Surface treatments in biodegradable stents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8693.5. Bioactive agents in biodegradable stents . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8703.6. Commercial products . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 870

requirements in each application, which should truly driveinnovation in materials. This approach also highlights theexisting gap between what the device needs and whatmaterials scientists have to offer currently, thus identifyingthe gap which could spur further research.

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S. Venkatraman et al. / Progress

Secondly, we restrict the discussion in this review toardiovascular devices, and more specifically to artificiallood vessels and stents. In these two types of implantedevice, the material performance is key to the success of theevice, and enhancement in device performance is drivenrimarily by materials innovation.

Polymeric biomaterials are generally derived from threeources: natural polymers, including those of plant and ani-al origin; totally synthetic sources; and synthesis based

n materials of natural origin. The first two categoriesre self-explanatory; the third is of relatively recent vin-age. It encompasses materials synthesized to mimic aaturally occurring polymer, but not necessarily identi-al to it. The most important materials in this categoryre the man-made protein structures, which resembleatural proteins but differ from them in some detailsf the primary structure. This third class of polymersromises innovative materials that have the potential tounctionally replace diseased or unavailable cell compo-ents, such as the extra-cellular matrix, which plays atructural role in many organs and tissues. Within eachpplication, we will highlight the advances made in theevelopment of each type of polymer, and the benefits theyonfer.

.2. Cardiovascular therapy

Cardiovascular disease broadly covers a range of con-itions affecting both the heart and the blood vessels.his review focuses on arteriosclerosis, which is a generalerm describing any hardening (and loss of elasticity) ofrteries (in Greek, “sclerosis” refers to hardening). Whenrteriosclerosis occurs in the coronary arteries, we refer tohe condition as coronary artery disease (CAD). If the extentf CAD is severe enough, it can result in a heart attack.n general, arteriosclerosis is amenable to treatment withtents or with a bypass; the bypass artery may be a veinrom the patient or totally artificial. Other cardiovascularonditions such as congestive heart failure also benefit frommplanted devices, but these fall outside the scope of thiseview.

Coronary artery disease leads to narrowing of the arter-es, sometimes to complete blockage. When this happensn coronary arteries, blood supply to the muscle cells sur-ounding the heart is affected, leading to their demise. Theatient concurrently manifests all the symptoms of a “heartttack”—shortness of breath and chest pain. The causesor such narrowing are manifold, and include high bloodholesterol and hardening of the arteries due to “plaque”ormation. The treatment of choice, until the mid-1990sas bypass surgery, where the affected artery is “bypassed”

y grafting another blood vessel usually a vein from the leg.his is an unsatisfactory solution, for many reasons includ-ng the fact that veins are not good artery substitutes; for

ultiple bypasses, patients run out of vein supply; and theatency of the bypass vessel is typically not long, and is of
he order of a few years, after which surgery is again indi-ated. An artificial blood vessel that functions more likecoronary artery than a leg vein would be an ideal sub-
titute. Such a product could also be used to treat blockederipheral vessels. We will review the progress made using

er Science 33 (2008) 853–874 855

polymeric biomaterials to develop an artificial artery thatis closer to the human one in properties.

In the mid-1990s, a procedure known as balloon angio-plasty combined with the insertion of a slotted metal tube(a “stent”), revolutionized the treatment of coronary arterydisease. The main advantage was that it was carried outas an outpatient procedure, reducing costs and recoverytime significantly. With a success rate of 70–75% in termsof vessel patency, it stood to replace bypass surgery exceptfor multiple blockage cases. The main disadvantage wasthat the permanent presence of the stent the bloodstreamnecessitated systemic administration of anti-thromboticdrugs for about 6–12 months following the procedure,and aspirin intake for a lifetime. The patency rate of thestented segment, while higher than with angioplasty alone,was low mainly because the unblocked vessel sufferedreblockage in the stented segment, a process known as“restenosis”. To treat this, a drug-eluting stent (DES) wasdeveloped and approved first in 2003, with significant ini-tial success. However, later developments have cast someclouds over the use of DES. In fact there is agreement amongmedical opinion leaders [9] that the permanent presence ofa stent is detrimental; hence fully degradable, and mostlyfully polymeric stents are currently being explored. This isthe second area for our discussion.

1.3. Types of implantable polymers

Synthetic polymers have historically been the mate-rial of choice for implants. The reasons include ease ofproduction; control over properties of the polymer; readyavailability and versatility of manipulation. Conversely,polymers from natural origins such as collagen sufferedfrom source to source variability of properties; possibility ofbacterial or viral contamination; and possible antigenicity.If these organic materials are of animal origin, there is theadded complication of harvesting the polymer or proteinand purifying it.

For these reasons, synthetic polymers have dominatedthe implant landscape. Examples include PMMA for ocu-lar implants, ultra-high MW polyethylene in artificialhip joints, poly(lactide) and poly(glycolide) polymers assutures and silicone polymers as breast implants. This situ-ation may be changed drastically with the advent of thethird class of polymers, which we term “bio-inspired”.This category of polymers comprises mainly the artificialor non-naturally occurring proteins, made by biosynthe-sis. Such polymers can be synthesized under controlledconditions such that they have predictable and control-lable properties from batch to batch and there is nolikelihood of microbial contamination. Antigenicity can-not be ruled out, however. Nevertheless, these polymershold great promise for implant applications, due mainlyto the possibility of generating materials that mimic thebody’s own structural polymers in terms of structure andfunction.

1.4. Outline of the review

The discussion will address these aspects in each appli-cation:

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856 S. Venkatraman et al. / Progress

a) The material requirements of the application (the mate-rial “wish-list”).

b) The relative merits of using synthetic, natural or bio-inspired polymers for each application.

c) Polymers that have been evaluated using in vitro meth-ods.

d) Outcome of animal studies and (if available) human per-formance data.

e) Commercial success.(f) Future directions.

2. The artificial blood vessel

Bypass surgery remains the treatment of choice forheavily clogged or multiple clogged vessels. This is particu-larly true of peripheral vascular disease, where diagnosis ofblockages is usually delayed, and currently available stentsdo not perform satisfactorily. Bypass surgery uses veins(and arteries, if available) from other parts of the body.Although veins, synthetic grafts and even arterial replace-ments are not satisfactory replacements for arteries, forreasons to be detailed below, about 1.4 million bypass pro-cedures are performed in the US alone [10], with aboutthree times that number being performed worldwide.

2.1. The need

Replacements for large arteries are readily available,and are known as synthetic grafts, as they are usu-ally grafted to tissue by suturing. These are most oftenmade of Dacron®, a synthetic polyester fiber made frompoly(ethylene terephthalate) or PET. Grafts made fromexpanded poly(tetrafluoroethylene) are also available andused for the larger arteries.

2.2. Requirements

So what exactly is required of the ideal arterial substi-tute? These are listed, as critical, and “nice to have”.

Critical requirements:

(a) The grafts must be biocompatible in severalaspects: non-thrombogenic, non-inflammatory,non-immunogenic, to name the three most importantrequirements (Section 2.2.2).

(b) Grafts must be compliant and elastic; they must closelymimic the unique viscoelastic nature of an artery(Section 2.2.1); this is particularly important for small-diameter arteries.

(c) Grafts must accommodate pressure changes, and benon-disruptive to blood flow.

(d) Graft should also be able to “remodel” efficiently: i.e.,to allow for growth of a layer of endothelium in a rea-sonable period of time.

“Nice to have” requirements:

(i) Be suturable or otherwise anchorable in place.(ii) Ease of manufacture is desirable; if the item is off-the-

shelf, it is then usable in emergency situations.(iii) Easily sterilized and available in various diameters.

er Science 33 (2008) 853–874

This list is a daunting one. We can state unequivocallyat this point that no synthetic material has been foundto satisfy all of these requirements. Bioprosthetic replace-ments (i.e., arterial or vein replacements) also do not satisfyrequirements (b) and (c), as well as (ii) above. It is the aim ofthis review to identify the extent of the gap between whatis needed, and what is deliverable.

2.2.1. Viscoelasticity of blood vesselsIn order to understand the viscoelastic nature of blood

vessels, a basic understanding of their anatomy is neces-sary. Shown below are cross-sections of a vein and an artery(Fig. 1).

Both veins and arteries are seen to have three layers:the tunica interna (intima), tunica media (media) and thetunica externa (adventitia). The intimal layer is mostlyendothelial cells; the tunica media (thickest layer) consistsmainly of smooth muscle cells and the adventitial layeris mostly collagen fibers and extra-cellular matrix, with afew fibroblasts. The alignment of the muscle cells is mostlycircumferential; the orientation of the collagen fibers issometimes so. In between these layers are elastic mem-branes, which consist mostly of elastin. These features arecommon to veins and arteries. Contractility is conferred onthe blood vessels by the smooth muscle cells, while resis-tance to distension and recoil are due to the presence andorientation of collagen and elastin.

There are differences between veins and arteries (forsimilar diameter vessels):

(i) The overall wall thickness is lower for veins.(ii) The overall elastin content is lower for veins.iii) The overall collagen content is higher for veins.

(iv) The extent of alignment of smooth muscle cells ishigher for arteries.

These result in the veins having lower stiffness andsomewhat higher extensibility, compared to arteries. Thusveins that are used as bypass vessels for arteries may notaccommodate the higher pressure seen in arteries, leadingsometimes to formation of aneurysms. Arterial structure(mostly relative thicknesses of the three layers) also variesdepending on size and the relative contents of musclecells, collagen and elastin also vary. Nevertheless, arteriesexhibit, on the average, certain viscoelastic characteristicsthat are different from those exhibited by veins. For exam-ple, the stress–strain curve for arteries has a “toe” region oflow modulus, followed by a region of increasing modulus.

Conventionally, it has been shown that the toe regioncorresponds to stretching of elastin fibrils, while the stiff-ening is due to deformation of the collagen fibrils, oncethe elastin component has reached its maximum exten-sion. In an interesting comparison study of arteries andveins, Monson et al. [11] confirmed the higher modulusbehaviour of both cerebral and non-cerebral arteries com-pared to veins, although the axial extensibility appeared to
be comparable for the femoral artery and the femoral vein.Nevertheless, this difference is what is usually termed as a“compliance mismatch” when quoting the deficiencies ofsaphenous vein grafts or for that matter, synthetic grafts,compared to the vessel being bypassed.

S. Venkatraman et al. / Progress in Polymer Science 33 (2008) 853–874 857

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ig. 1. Cross-sectional schematics of an artery and a vein. (Adapted fromf the layers between artery and vein, especially of the tunica media (smorteries and veins.

Much less is known about the viscoelastic charac-eristics of arteries, such as stress relaxation and creep.ccumulated creep over time in particular can lead toneurysms. It would be very interesting to estimate theumulative effects of short pressure changes on blood ves-els caused by the pulsatile nature of blood flow, usinghe Boltzmann superposition principle, and compare thato what is observed for proposed blood vessel substi-utes. Such data do not appear to have been generated,lthough axial creep [12] measurements have been suc-essfully made for arteries and for arterial substitutes. Inhis extensive study, it was shown that over a 1200-s period,rteries (specifically, the porcine carotid) behave like vis-oelastic fluids, rather than as solids, a surprising findinghat merits further investigation over longer time scales,ith more complex model-fitting.

So what are the clinical consequences of the artificialrtery not matching the viscoelastic characteristics of theative vessel? Compliance mismatch can cause less laminarow, which is believed to lead to stenosis, either throughhrombus formation or intimal hyperplasia. The difficultiesf designing a study to isolate the effect of just complianceismatch can be appreciated. An effort was made by Abbott

t al. [13] where carotid arteries of dogs were excised, andreated with glutaraldehyde to increase the modulus. This isrimarily through crosslinking of the collagen fibrils. Threerteries were thus selected for insertion into the femoralrteries of the same dogs: a control untreated artery; onereated for 30 min, and the third treated for 1 h. Compliance

easurements were made in situ by pressurizing the vesselnd measuring the diameter increase with ultrasound tech-
iques. The radial compliance is calculated as the ratio ofhe diameter change (normalized to the starting diameter)o the pressure differential. By this measure, the ‘stiff’ arteryas five times less compliant than the native artery or
he compliant artery (at a nominal pressure differential of

ith permission from John Wiley & Sons. Note the difference in thicknesscle layer). This difference contributes to the stiffness difference between

100 mm Hg). Patency was reduced to about 37% at 12 weeksfor the stiff artery, compared to about 85% for the compli-ant one. Where patency was reduced, there was thrombusformation, thus indicating that compliance mismatch leadsto less laminar flow which in turn causes thrombosis.

What about the consequences of other mismatch of vis-coelastic properties? Only one property is of significance,that of the creep compliance. If creep compliance is muchhigher than in native arteries, the chance of a clinicallysignificant aneurysm formation increases. As alluded toabove, in uncrosslinked vessel substitutes, a cumulativecreep effect due to pressure surges may eventually resultin an aneurysm.

2.2.2. Biocompatibility aspectsThese are much more straightforward to define. The arti-

ficial blood vessel must be as follows:

(a) Non-thrombogenic.(b) Non-inflammatory.(c) Non-toxic, non-carcinogenic.(d) Non-immunogenic.

Less critical attributes are that the artificial vessel mustbe resistant to infection and that its presence should notdisturb the healing process in the vessel; more specificallyit should not hinder the formation of an endothelial celllayer on its lumen-facing surface.

Some general statements, based on past experience withbiomaterials, may be made here:

(i) Synthetic materials are usually non-immunogenic,while natural materials (which are not derived autol-ogously) tend to stimulate an immune response.Bio-inspired materials typically fall somewhere inbetween.

in Polym
(ii) Thrombus formation is caused both by exposing bloodto a foreign surface and by flow instabilities. Thus,the mechanical properties of the replacement material(which causes compliance mismatch leading to non-laminar blood flow) and it surface properties, can leadto thrombus formation.

(iii) In spite of considerable effort over the last 20 years, nothrombus-resistant material surface has been devel-oped: the closest would be a pyrolytic carbon surface.There are products (artificial graft materials) withcarbon coatings that claim to offer better thrombo-resistance [14]. Despite only modest improvements inpatency rates, such coated products are commerciallyavailable [15].

Thus, to develop a material for small-diameter graftsthat would show good patency rates, it is essential to makeits surface thrombo-resistant, and to match the arterial vis-coelasticity. Of the two requirements, the generation of athrombo-resistant surface appears to be the more difficultone. Recent advances in surface modification of polymersappear to offer some hope, however. Matching the vis-coelasticity of arteries is more feasible, as we will see inthe following sections.

2.3. Synthetic polymers

Synthetic, natural and bio-inspired materials have allbeen evaluated to match the arterial compliance. We willdiscuss each in turn.

One aspect of material design for a vascular vessel sub-stitute may be highlighted: this is the requirement fora certain level of porosity which at the same time doesnot allow blood leakage. The porosity is needed for cellu-lar ingrowth from surrounding tissue, into the substitutematerial: this constitutes the normal healing process. Thesooner the lumen-facing side of the material becomescovered with endothelial cells, the better. The time forsuch coverage can be excessively long if cell colonizationdepended mainly on circulating ECs in blood.

Based on porosity, synthetic vessel substitutes can bebroadly classified into two categories: fibrous and contigu-ous. For the fibrous type, woven or knitted have both beenevaluated, primarily made of polyester (Dacron®); suchconstructs typically need to be “pre-clotted” with a layerof fibrin to prevent leakage. The contiguous type typicallyis microporous, and does not require pre-clotting.

2.3.1. Synthetic polymers, fiber-basedEarly attempts focused on synthetic Dacron® (PET)

woven or knitted fabrics. Typically, the woven tubes aremuch less porous than the knit versions: among the latter,various knitting configurations have been tried to controlthe degree of porosity. These include weft and warp-typeknits. In addition, a bewildering number of coatings havebeen applied to these Dacron® grafts, including fluoropoly-
mer [16], collagen and heparin-bound polymer [17], fibrin[18], fibroblast growth factor [19], silicone elastomer [20],gelatin [21], to mention just some of the more widely stud-ied coatings. Although some animal studies had indicateddifferences in acute thrombosis [22] or in time of healing
er Science 33 (2008) 853–874

[21], no differentiation was seen in clinical trials [23–25],even though one of them [25] was a study of a larger ves-sel (the aorta). (In some of these trials, the control wasnot always an uncoated Dacron® prosthesis, but an ePTFEprosthesis, which is currently the preferred material forperipheral vessel replacements; nevertheless, the lack of animprovement with the coatings is a reasonable conclusionfrom these trials.)

A word about the mode of “failure” of Dacron®-basedgrafts: initially there is protein adsorption on to theDacron® surface, followed by platelet adhesion, inflam-matory cell penetration and EC/SMC migration. On thelumen-facing side, a platelet-containing fibrin coagulumis formed, whereas a dense layer of foreign body cells isformed on the outer surface. The loss of patency is even-tually due to a combination of thrombus formation andintimal hyperplasia. While it is difficult to estimate howmuch patency is lost due to compliance mismatch andhow much is due to thrombogenicity, it can be assumedthat patency loss occurring over a longer time frame (6months upwards) is likely to be caused by compliancemismatch.

2.3.2. Expanded PTFE (ePTFE) graftsIf saphenous vein is not available, this is the mate-

rial of choice for small-diameter vascular grafts. Since thelate 1970s, expanded PTFE has been produced by a pasteextrusion process, followed by solvent evaporation, biaxialstretching and final high temperature sintering. It has beenshown that the pore size can be manipulated from 30 �mto about 100 �m.

Mechanically, ePTFE is much stiffer than arteries orveins, measured moduli being reported to be around3–6 MPa [26,27] in both the circumferential and longi-tudinal directions. For comparison, the arterial modulus(canine) at 300 kPa stress is about 600 kPa, while canineveins exhibit a slightly higher modulus of 900 kPa. Also forcomparison, Dacron® grafts register vastly different moduliin the circumferential (12 MPa) as compared to the longitu-dinal (0.7 MPa) direction. In spite of this obvious modulusmismatch, ePTFE has gained popularity for peripheral ves-sel bypass, particularly for the femoropopliteal artery.Grafts from ePTFE have a nice “feel”, low coefficient offriction, are mechanically relatively stable, and are eas-ily manufactured in many sizes, and shapes, including thecrimped variety.

Whether ePTFE grafts are superior to Dacron® grafts inthe clinic is open to debate. Four studies, two involving knit-ted Dacron® [28,29] one involving heparinized Dacron®

[24] and the fourth involving a gelatin-coated, doublevelour Dacron® graft [30] were compared against ePTFEGoretex® grafts, in a multi-centred studies. For the knittedDacron® studies, one was non-velour, plain Dacron® graft[29], while the other involved a double velour graft withcollagen impregnation [28]. In the non-velour case, overallpatency was actually higher for the knitted Dacron® at 5
years (48% vs. 27%) whereas the double velour study didnot see any statistically significant difference in patencyrates at 5 years (45% vs. 43%). In the gelatin-coated Dacron®
study, 427 patients received randomized femoropoplitealbypasses (above the knee or ABK). At 2 years, patency rates

in Polymer Science 33 (2008) 853–874 859

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ere actually higher for the Dacron® group (70% vs. 57%)ompared to the ePTFE group. For the prospective studynvolving heparinized Dacron®, for patients with both ATKnd below the knee (BTK) femoropopliteal bypasses, pri-ary patency rate at 3 years was higher (46%) for theacron® group versus 35% for the ePTFE group, although

his difference disappeared at 5 years.The overall conclusion from the studies involving ATK

emoropopliteal bypass indicates no substantial clinical dif-erence between the two types of grafts; patency rates foroth are still inferior to venous grafts. The events lead-

ng to the low patency are predominantly thrombotic; inurn it appears that none of these grafts ever developsfull coverage of endothelial cells on their lumen-facing

ide. This lack of coverage is due entirely to the lack oforosity needed for transmural (i.e., EC migration fromhe tissue side) endothelialization, as the EC concentra-ion in circulating blood is quite low. As mentioned above,ncreasing porosity beyond a certain point is unadvisablen order to prevent extravasation (i.e., leakage) of bloodlements.

Attempts to address more rapid endothelialization withhe use of higher porosity ePTFE have been reported. In

study of grafts inserted into the carotid and femoralrteries of dogs [31], it was found that ePTFE grafts withmean pore size (internodal distance) of 90 �m showed

igher EC coverage at 18 weeks (75%) versus the stan-ard ePTFE graft (30 �m internodal distance; EC coverage3%). With 60 �m ePTFE, however, the higher porosityid not show higher patency rate compared to the 30 �mPTFE, in another study involving aortoiliac grafting in dogs32]. A human study involving 60 �m and 30 �m ePTFErafts however showed virtually no endothelialization at 3onths [33]. The endothelialization was assessed only indi-

ectly by radioactive indium infusion and platelet imaging.istology of two patients confirmed very little ingrowthf cells from the tissue to the lumen side. This lack ofrowth was attributed, at least partially, to the presencef a reinforcing wrap on the tissue-contacting side of theraft. Disappointingly, there appears to be no clinical studyf the 90 �m ePTFE graft to date.

Other modifications to ePTFE to improve patency ratesave been similarly unsuccessful: these include carbon-oating or carbon impregnation [14,32,34]; fibrin glue withrowth factors [35,36]. None of these treatments have beenested in humans, except for the carbon impregnation,here no improvement in primary or secondary patencyas found. Animal studies with FGF-releasing fibrin [35]

reatment appeared to enhance endothelialization. How-ver, another animal study involving VEGF-releasing fibrinreatment did not see any improvement in the pig carotidrtery [36], indicating that no firm conclusion can be yetrawn. More biologically oriented treatments (such as cel-

ular attachment) are now gaining attention, as detailedelow.

.3.3. Polyurethane graftsThe compliance mismatch concern has prompted the

tudy of a more elastomeric (erroneously termed “morelastic” in literature) material, the most popular beingolyurethane.

Fig. 2. Typical stress–strain curve for an artery and a vein. In general, veinsare more “collapsible” than arteries, due primarily to thickness differencesin the three layers, particularly of the medial layer. The stiffening in thelatter part of both curves is usually attributed to the collagen.

2.3.3.1. Mechanical aspects. In the literature, “compliance-matching” usually relates to matching the compliance ofthe first part of the stress–strain curve only (see Fig. 2),i.e., the compliance usually attributed to the elastin com-ponent. The secondary stiffening (attributed to the collagenfibers) has been discussed less, if at all. In this light, it isinstructive to tabulate the moduli of these synthetic graftmaterials, for reference (Table 1).

To make the stress dependence more explicit, we havereplotted these data in a bar graph, for the modulus in thecircumferential direction (Fig. 3).

Notably, the large ratio between the modulus at lowand high stresses for the artery is not seen in any of thegraft materials. Notwithstanding this, the PEEU polymermatches the artery reasonably well. The stiffening of thePEEU at high stress is not due to a change in material prop-erties, but rather to the use of a knitted construction. As thecrimps straighten, it leads to a higher modulus. This sort ofbehaviour is inelastic, and does not permit recovery to theoriginal state, which of course is what arteries and veinsdo. In the following discussion, the matching compliancebehaviour of PU is restricted only to the first part of thestress–strain curve, and not at all to viscoelasticity.

2.3.3.2. Polyester-based PU (PEU). While initial animal tri-als were very promising, the hydrolytic instability ofpolyester-PUs proved to be a drawback. Earlier versions ofpolyester urethanes, such as Estane®, proved to be unstableto hydrolysis or to enzymatic attack or both [37]. They havesince not been used beyond animal studies and will not bediscussed further.

2.3.3.3. Polyether PU (PEEU). These PUs use predominantlypoly(tetramethylene oxide) (PTMO) as the soft segmentand have been extensively studied for its in vitro degra-dation characteristics [37]. Hydrolytically, it is more stablethan polyester urethanes under both acidic and alkaline
conditions. However, it degrades under enzymatic attack,oxidative environments and under stress [38]. In fact, thein vivo performance of devices made with PEEU appears tohave been affected by “environmental stress cracking” [39],in which the enzymes are part of the environment.

860 S. Venkatraman et al. / Progress in Polymer Science 33 (2008) 853–874

Table 1List of measured moduli for various graft materials, compared to values obtained for the iliac artery and iliac vein

Polymer Modulus, circumferential(MPa) (applied stress, kPa inparenthesis)

Modulus, longitudinal (MPa)(applied stress, kPa inparenthesis)

Remarks

Dacron® (PET)a woven 12 MPa (180 kPa) 0.87 MPa (87 kPa)15 MPa (360 kPa) 1.1 MPa (170 kPa)

Dacron® (PET)a knitted 12 (115) 0.64 (54)14 (230) 0.64 (110)

ePTFEa 5.9 (60) 2.0 (27) GoreTex®

7.9 (120) 3.6 (54)Polyether urethane (PEEU)a (Mitrathane®) 1.9 (97) 0.64 (46) Knitted prototype, 45◦ warp angle

4.3 (195) 1.1 (92)Polycarbonate urethane (PCU) 3.1 (?)b 3.1 (?)b Based on Corvita(R); No stress dependenceIliac arterya 1.3 (135) 0.40 (64)

4.8 (270) 0.77 (130)Iliac veina 4.8 (200) 1.8 (97)

6.3 (400) 5.3 (190)

The figure in parenthesis is the value of the imposed stress corresponding to each modulus value. Circumferential modulus is also sometimes referred toas the “radial” modulus, and is a measure of stiffness when the tubular construct is expanded isotropically outward.

of choice. FDA approval for the PCU-based Vascugraft®

was granted only for the more limited indication ofvascular access in hemodialysis. A similar fate seemsto have befallen the Thoratec polyetherurethane urea

a Taken from [27].b Taken from [140].

One of the earliest grafts developed, a PEEU withthe brand name of Mitrathane®, from the US Company,Mitral [40], was produced by phase inversion from solu-tions in dimethyl acetamide to generate microporosity.It yielded relatively poor performance in animal trials[41], apparently because the pores are closed, and donot interconnect. Histological studies confirm that theocclusion when Mitrathane®-based graft is used, is acuteand attributed to a variety of factors, including poorattachment at the anastomotic site [42]. Other compa-nies, such as Newtec Vascular Products (U.K.) developedthe PEEU Pulse-Tec® graft, while Thoratec Laboratories(U.S.A.) presented the Vectra® graft. Pulse-Tec® died anearly death, for reasons not entirely clear, whereas Vectra®

actually was approved by the FDA in 2000, but only forvascular access during hemodialysis, which is a short-term application [43]. Evidently, the promise held out byPEEU was not realized in a long-term, blood-contactingimplant application, due primarily to thrombogenicity butalso due to its instability under enzymatic and oxidativeattack.

2.3.3.4. Polycarbonate urethanes (PCUs). PCUs were devel-oped to resolve the instability issues present in PEU andPEEU. The first patent was issued to L. Pinchuk of CorvitaCorporation in 1992 [44]. However, the first graft proto-type using PCU (trade-named Vascugraft®) was developedby Braun-Melsungen AG and tested as early as 1983 in ani-mals [45] without an explicit reference to the presence ofa polycarbonate segment. Zhang et al. [46] subsequentlyconcluded that the Vascugraft® prosthesis did containa polycarbonate soft segment. Based on this and otherreports, we will consider Vascugraft® as incorporating (pre-
dominantly) a PCU.
The Vascugraft® is claimed to be made microporous, byusing a spray technique. A low viscosity solution is sprayedthrough a nozzle fine enough to generate fiber formationaround a rotating mandrel. Many sprayings are used to gen-

erate a reasonably thick structure, with inter-connectedpores [46]. The Vascugraft® also contains an “overwrap”formed with an extruded, harder PCU.

One study in rats [47] found faster endothelializationrates, as well as lower amount of neo-intima forma-tion, in Vascugaft® compared to ePTFE. However, this wasonly reported well after the company (Braun-MelsungenAG) stopped the development of Vascugraft® follow-ing disappointing clinical results in a 15-patient study[48]. Interestingly, the early occlusion observed in theclinical study was not related to the instability of thePCU, but to thrombus formation attributable to poorrun-off or distal blockage [49]. The primary patencyrate in the clinical study was comparable to ePTFE, butfive occlusions were found. Without a compelling rea-son to change, ePTFE grafts remain the synthetic graft

Fig. 3. The circumferential modulus for selected graft materials, at twodifferent imposed stresses (based mostly on data in [27]). The differentialin modulus at low and high stresses is marked for the artery.

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roduct, Vectra®, following a successful 142-patient trial50].

As of this writing, it appears as though that a biostableU does not offer any benefits over ePTFE as a small-iameter arterial prosthesis material. The closer fit of at

east part of the compliance curve with arteries for the PUs,oes not translate into higher patency. Does this mean thatcompliance matching” is irrelevant? Not necessarily, forhe following reasons:

1. “Compliance matching” could imply matching the entirestress–strain curve.

. “Compliance matching” may be a necessary but not suf-ficient requirement for enhanced patency.

. The PU prostheses failed due to factors other than thehemodynamic disturbances prompted by “compliancemismatch”. One such factor is acute thrombus forma-tion due to the interaction with the material surface.Fundamentally, PU is not any different from ePTFE inthis regard. In spite of other efforts to modify surfacesof PU, including hirudin attachment [51], improvementsin early occlusion rates have been hard to come by.

Without further study, the precise reason for PU fail-re cannot be pinpointed, making it unlikely that PUs wille successfully used for a small-diameter prosthesis in theear future.

.3.4. Combination (bio-inspired) structuresA notable effort to match the entire stress–strain curve

f an artery has been made by Gupta and Kasyanov [52],sing a combination of PU and Dacron® in two differ-nt ways to match the entire stress–strain curve for theuman common carotid artery (CCA). One of the graftssed Dacron® fibers as the warp (running longitudinally),ith a combination of Dacron® and stretched PU (typeot specified) used as the weft (running circumferentially).his construction gave crimping in the circumferentialirection. A second construction used the combination ofacron® and stretched PU fibers in both the warp and theeft, allowing for “stretchability” in both directions. Bencheasurements showed the first construction to mimic the

CA stress–strain curves more closely, although the matchas by no means exact. Both constructs were more closelyatched to the CCA than a Dacron®-only graft.In an accompanying implantation study, only the first

onstruct (no controls) was implanted in the femoral andarotid arteries of dogs for 1 year, with the grafts explantednd evaluated mechanically and histologically. No occlu-ions were found and neo-intimal formation, at least of ECormation on the surface was evident. The grafts appearedo have been well-tolerated. A 3-month explanted grafthowed some ‘stiffening’ in the stress–strain curves at ear-ier strain levels than before implantation, also testifying tocertain degree of tissue ingrowth. Surprisingly, there areo follow-up reports of these constructs in further animal
r human experimentation.
.3.5. SummaryIn summary, the synthetic polymer graft still dominates

he marketplace for grafting in the peripheral arteries, in

er Science 33 (2008) 853–874 861

spite of relatively low patency rates compared to venousreplacements. The low patency rate appears to be dueboth to “compliance mismatch” and thrombogenicity atthe polymer surface as well as failures at the anastomoticsite. The following sections describe other approaches toimprove the patency rates.

2.4. Biodegradable polymers, synthetic

Biodegradable polymers have been used in two ways:

(a) As an implantable graft, which then degrades while tis-sue remodeling occurs around it and inside it.

(b) As a scaffold for growing vascular tissue ex vivo, fol-lowed by implantation of cell-containing polymericstructure.

2.4.1. As graftEarly investigations focused on fast-degrading poly-

mers, such polyglactin 910, which is a copolymer ofpoly(lactide) and poly(glycolide) (90% glycolide, alsoknown as PLGA), which is marketed also as a biodegradablesuture [53]. Greisler and coworkers have explored variousimprovements [54–56], with some promising results.

It was thought that a faster the rate of erosion (dis-appearance) of the mesh or graft would enable easieringression and growth of tissues into the mesh. In addition,some authors [57,58] believe that the degrading polymeractually stimulates endothelialization via an intermediary,the macrophages. Presumably, internalization of degradingpieces of polymer would activate macrophages into secret-ing growth factors aimed at endothelial cell proliferation.

However, a fast-eroding graft, while showing good cellgrowth in the interior of the graft, suffers from poor burststrength. Without a supporting mesh, the growing cell bodyis unable to withstand blood pressure surges, and showsextensive dilation or even developing aneurysms. Attemptshave been made to slow down the degradation rate usingpoly(dioxanone) or PDS polymer [55]. Poly(dioxanone) is apolyether-ester, and was one of the earliest polymers usedin sutures. Its absorption rate in the body is longer thanVicryl® or polyglactin 910 (90/10 PGLA) and is quoted tobe about 6 months compared to around 3–4 months forpolyglactin 910. Correspondingly, the thickness of the so-called inner capsule (IC), which measures the thickness ofnew cell formation, is about 500 �m for polyglactin 910and about 230 �m for PDS. Evidently, slower degradationrates also slow down cell growth within the mesh. Conse-quently, dilation effects are expected to be less significantand for a longer period with the presence of PDS than withthe PGLA. However, the cell ingrowth rate into PDS is alsoslower.

So can a compromise be achieved? Bi-component fibermatrices were tried, consisting of 74% PGLA and 26% PDS[59], which do indeed show an intermediate cell ingrowthrate (∼300 �m at 1 month), in rabbits. Myofibroblasts
and collagen were detectable beneath what appeared tobe endothelial cells. Maintaining good burst strength iscrucial for the tissue/mesh structure. PDS-based matriceswithstood about 1200 mm Hg (at 3 months), which is anacceptable burst strength. The corresponding value for the

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hybrid material graft is not reported, but is expected to belower.

In spite of these very promising results with completelydegradable grafts, there appears to have been issues withthe idea of “premature” degradation of these structures,especially before complete cell coverage has taken place;given the nature of the degradation process in these mate-rials, which is the so-called bulk degradation mechanism,there is a finite possibility that pieces of the graft couldbreak off and float in the bloodstream. This is much moreof an issue with grafts than with stents, as we will see inSection 3.1.1.

As of this writing, there are no reported human trialswith a fully or partially biodegradable graft.

2.4.2. As a scaffold materialWhen arterial cells are grown on a biodegradable

substrate and then implanted back into the patient, itis essential that autologous cells be used for to avoidhost rejection. This necessitates a longer processing time,enhances likelihood of infections and a planned surgicalprocedure, and is clearly not indicated for emergency situ-ations.

There are excellent reviews [60–62] of the many notableattempts to produce a tissue-engineered graft using asynthetic biodegradable scaffold. We will not attempt to
exhaustively review this work here, but only point toselected examples to highlight the advantages and the lim-itations of this approach (Fig. 4).
In what follows, we will not also describe the verysignificant advances made with respect to tissue reactor

Fig. 4. Schematic representation of use of collagen gel as a mediumfor growing blood cells in vitro. (Adapted from Nerem and Seliktar[60]; Reprinted, with permission from the Annual Review of BiomedicalEngineering, vol. 3© 2001 by Annual Reviews www.annualreviews.org).Generally, the approach is to use varying ratios of collagen and elastin inthe outermost (adventitial) and middle (medial) layers. Fibroblast cells aregrown in the outer layer, while smooth muscle cells (SMCs) are grown inthe medial layer. The innermost layer of endothelial cells (ECs) is usuallydeposited by a seeding process.

er Science 33 (2008) 853–874

engineering; rather, we will focus on the role of the poly-meric biomaterial in the process. Considerable innovationhas gone into trying to shorten the time required to growa functioning artery ex vivo; even greater advances areexpected in the future, but these details are outside thescope of this review.

Following a report [63] of enhanced attachment ofendothelial and smooth muscle cells in vitro on to a PGAmesh that had been treated with alkali (1N NaOH for 1 min),Niklason et al. [64] demonstrated the feasibility of engi-neering an artery ex vivo on these surface-treated PGAmeshes. Using a “biomimetic” system, which allowed forapplication of a pulsatile radial stress on the tubular mesh,bovine SMCs were seeded and grown first, followed byendothelial cell growth on the same mesh. The meshesthat were seeded and had cells grow on them under theradial stress, exhibited good mechanical properties (burstpressure of ∼2000 mm Hg at 8 weeks), the requisite phar-macological responses and showed evidence for collagenproduction.

Four such vessels were implanted in pigs: one was avessel grown from bovine cells, but seeded with porcineautologous ECs; another was an autologous vessel grownunder pulsatile stress; the last two were autologous ves-sels, but grown without pulsatile stress. After 3 weeks,the two autologous vessels grown without stress, showedthrombosis while the other two (including the xenograft)were patent. Although these results indicate success in thegrowth of cells with the correct alignment, the failure ofthe non-pulsed vessel was attributed to denudation of ECsrather than due to the lack of alignment of SMCs or ECswithin the vessel.

In a similar vein (no pun intended), PGA mesh coatedwith poly(4-hydroxy butyrate) (P4HB), was also seededwith myofibroblasts, and cultured under pulsatile stress forseveral days [65]. These authors confirmed the enhance-ment of mechanical properties of vessels grown underradial stress, by a factor of 6.

The first clinical application using a syntheticbiodegradable scaffold was reported in 2002 by Japaneseresearchers [66], for the reconstruction of a larger (10-mmdiameter) pulmonary artery in a 4-year-old. Autologouscells, cultured for a month, were used to seed a porousPCL/PLA copolymer (50:50) reinforced by a PGA mesh,the whole being designed to completely erode in 2months. Since this culturing involved xenoserum, theseresearchers switched to bone marrow cells (BMCs) in2001. They reported successful implantation in 20 patientsby 2002 [67]. The grafts used were varied as follows:three of the grafts used autologous cells seeded ontoa PGA/PCL–PLA copolymer mesh (1 month procedureas above). The rest of the grafts used bone marrowcells aspirated from the patient, seeded on to PCL–PLAcopolymer reinforced with a woven PLA fabric, andincubated for a few hours at 37 ◦C. The rationale for thelatter procedure was that BMCs have been reported to
grow into ECs under lab conditions. A follow-up of someof these patients [68] showed the tubular grafts to bestill patent at 32 months. If the good results persist,the BMC procedure is a time-saving (and consequentlycost-saving) and far simpler procedure that works with
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S. Venkatraman et al. / Progress in Polymer Science 33 (2008) 853–874 863

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low-degrading polymer tubular grafts. Further results arewaited.

.5. Natural materials

To date, the only “natural” polymers that have been eval-ated at length are collagen and elastin. Although thereere some early attempts to use these polymers in a graft,ost of the work involved the use of collagen as a geledium for growing vascular cells in vitro, into a com-

lete blood vessel The earliest attempt by Weinberg andell [69] used certain binding sites on collagen to seed androw the three major types of vascular cells: endothelialells, smooth muscle cells and fibroblasts.

Briefly, their methodology was as follows: collagen isissolved in acidic water, mixed with smooth muscle cellsnd cell culture medium, and poured into an annular mold,here the mixture gels at 37 ◦C. In about a week, the mix-

ure contracts over the central mandrel to form the media.hen a non-woven, Dacron® open mesh fabric is placed onhe cell layer. Another mixture of collagen solution, fibrob-asts and cell culture medium is poured into the mold toorm the outer layer (adventitia). After another 2 weeks,hen the outer layer has contracted, the whole assemblage

s pulled off the mandrel and mechanically tested. It waslso seeded on one side with endothelial cells and testedor physiologically relevant responses. Physiologically, theells behaved as they would in a healthy blood vessel; how-ver, the mechanical strength with this construct was poor40–50 mm Hg burst strength). Addition of two more layersf Dacron® and media/adventitia brought the strength upo 320 mm Hg, still short of arterial strength. Nevertheless,
his study showed the potential for complete constructionf an artery ex vivo, leading to a significant increase inesearch activity in the following decade.
Early failures to create a mechanically strong vessel wereue mostly to the lack of alignment of the smooth muscle

er of a tissue-engineered matrix reproduced from [62], with permissionor the cells, but must be able to transmit an externally imposed strain to

cells in the medial layer and perhaps of the collagen as well.It is also possible that the amount of collagen overall wasnot regulated well enough (Fig. 5).

A key step in improving alignment came with the recog-nition that cells grown in stressed conditions, tend toalign more than when grown under static or nearly staticconditions [64,70,71]. Kanda [71] gave details of how thegrowing cells were mechanically stressed in a simple butelegant apparatus. One SMC-containing collagen gel waskept stress-free, another subjected to a static stress, anda third gel was dynamically stressed on an inflatable sili-cone balloon, using a respirator. Histological results of the“vessel” after 5 days, clearly showed that stressed sampleshad circumferential alignment of both SMCs and collagen,and the extent of alignment was higher in the dynami-cally stressed sample. This appears to have been the firstattempt to demonstrate alignment of cellular elements andof collagen parallel to the direction of the stress.

More extensive confirmation of this came in 1999 [72].In this study, the collagen-gel with SMCs was directlypoured on to a silicone tube mounted over a glass man-drel, then grown for 4–8 days in medium until shrinkagewas complete. The silicone tube with the cell layer on top,was then dynamically stressed (5% and 10% distension) forthe same length of time as a control, unstressed sample inmedium. This study showed more extensive cellular align-ment at 10% strain; although values of ultimate strengthsof 60 kPa were reported, the modulus after 8 days of condi-tioned growth was about 130 kPa, 1/10th of that of an artery(Table 1). Nevertheless, these studies clearly delineated theimportance of mechanical stresses in regulating the growthof aligned cells and collagen fibers.

Interestingly, this work did not use a collagen sub-strate. This could be because the silicone was able totransmit radial stresses more efficiently to the gel matrixthan a collagen/elastin tube. The recent use of electro-spun collagen and elastin fibers [73] could result in a

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composite graft/cell hybrid that is made entirely from nat-urally occurring polymers. These researchers were able toelectrospin collagen I and retain or generate the quater-nary structure associated with the regular bands seen byelectron microscopy [74], using a hexafluoro-propanol sol-vent. It is still unclear if these quaternary structures arebeneficial for blood-contacting applications, with reportsindicating the need to remove this structure to preventthrombosis [75].

A three-layered construct resembling vascular tissuehas in fact been constructed with electrospun elastin andcollagen [73], in a three-step process. The adventitia waslaid down on a substrate of collagen (80%) and elastin (20%)fibers. This was followed by a SMC growth on a 30/70 colla-gen/elastin tube representing the medial layer, and finallyECs were seeded on the inside. The entire process required15 days. The resulting structure could potentially resolvethe tissue-engineered graft problem (Fig. 6).

The challenge now is to weave these fibers into aconventional-type graft while retaining circumferentialalignment of the fibers, and approximating the native ves-sel composition and mechanical properties. To date, thereare no reports of such studies, or indeed of the use of elec-trospun fiber matrices in animals.

Another approach to obtaining a collagen/elastin matrix
is by simply removing cells from cardiac tissue [76]. In afour-step process, Courtman et al. decellularized bovinepericardial tissue to yield a mixture of collagen (94%) andelastin (4%), along with a proteoglycan sugar residue. Thisdecellularized matrix exhibited a stress–strain curve sim- b
Fig. 6. Schematic representation of the protein biosynthesis method (adapted frosteps involved are the identification of the primary sequence, and then translatin

er Science 33 (2008) 853–874

ilar to that of a typical artery wall (Fig. 2), with stressrelaxation also similar to that of the tissue. Moreover thequaternary structure of collagen is retained. The matrix,which took 4 days to generate, could be a promisingsubstrate for embedding and growing a collagen–cell gelmatrix. More studies along these lines are awaited.

In summary, although natural extra-cellular matrix pro-teins could be potentially used for either a graft or a scaffoldfor tissue-engineered constructs, very few studies of theiruse have been reported. This could change soon as theuse of electrospun matrices becomes more wide-rangingto accommodate the weaving or knitting into tubular graftconstructs.

2.6. Bio-inspired materials

Bio-inspiration may come from two sources: one, wherethe graft construct is inspired by the native vessel structureand two, where the component materials themselves areinspired by those in the extra-cellular matrix. An exampleof the former was discussed in Section 2.3.4. This sectionnow discusses the synthesis and use of the so-called “artifi-cial” proteins that mimic those occurring in cardiovascularextra-cellular matrix.

Artificial or synthetic proteins are better candidates
than tissue-derived proteins for the following reasons:
a. They have better batch-to-batch reproducibility of struc-ture.

. They are not contaminated by microbials.

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c. They do not need complicated extraction proceduresthat may destroy some structural features.

. They are potentially less immunogenic.

. In the particular case of elastin, native elastin occurs ina crosslinked form and hence is not readily processible.

The biosynthetic route to proteins is a powerful one,ith exceptional control over structure, and hence overroperties. It is shown schematically below [77].

First, the desired amino-acid sequence is selected, andhe complementary nucleotide sequence identified. Typ-cally, the DNA is synthesized by solid-phase techniqueshough at times it may be cloned from an organismhich produces the desired protein. The DNA is inserted

nto a plasmid to transform a bacterial host to pro-uce the desired protein. The molecular weight is strictlyontrolled by controlling the length of the nucleotide “mul-imer”, leading to a protein with a monodisperse moleculareight. Frequently, the synthesized protein is coaxed

nto spontaneous self-assembly by varying the externalonditions.

Perhaps the first person to demonstrate the power ofNA-based methods for protein synthesis was Urry [78],

ocusing on elastin. Elastin is composed of four main aminocids: glycine, valine, alanine and praline, with some lysines well. Crosslinking occurs via the linking of lysine units,o form a six-membered ring, sometimes referred to asdesmosine”. The more hydrophobic AAs are responsibleor the so-called inverse gel transition in elastin, i.e., as theemperature is raised, elastin aggregates. The most exten-ively studied elastin-like polypeptide has the sequenceVPGVG). In this protein, the transition temperature cane fine-tuned, for example, by substituting for the sec-nd valine in the sequence with more hydrophobic or lessydrophobic groups. This is only possible via DNA-directedynthesis.

For use in vascular grafts, the elastin-like polypeptideELPP) has been modified for enhanced cellular interac-ion. Examples of modification include the incorporationf a RGD sequence [79] or REDV sequence [80]. The REDVequence is also found in fibronectin and is claimed to bepecific for endothelial cell attachment/binding [81]. Stud-es of EC attachment on ELPPs with the REDV sequencehow enhanced attachment of HUVECs [82]. The adher-nt ECs are also somewhat stable to flowing shear stressexposure only for 2 min). However, selectivity to ECs hasot been evaluated so far. Nevertheless, this sort of studyerves to highlight the versatility of the engineered proteinpproach to functionalize materials for a specific applica-ion.

While this approach to bio-inspired materials is promis-ng, the following strategic issues need to be addressed to

ove ahead:

Are these elastin-like polypeptides better used as a fab-ricated tubular graft vessel? If so, would spinning theminto fibers or making them into a blend be preferred?How fast can endothelialization be achieved by thesegrafts in vivo? In situ endothelialization is very slow, duemainly to the low number of circulating endothelial cells.

er Science 33 (2008) 853–874 865

Nevertheless, controlled protein synthesis opens upnew possibilities for an artificial vascular graft. The materi-als thus produced bridge the gap between purely syntheticmaterials such as Dacron® and ePTFE on the one hand, andcompletely biological materials such as decellularized tis-sue on the other.

2.7. EC-seeded grafts

There are two main approaches to endothelialize animplanted graft:

(a) Seed the graft with autologous ECs ex vivo, then implantthe graft.

(b) Create surface features on the graft which induces insitu endothelialization.

The latter includes techniques to improve the transmu-ral (i.e., from surrounding tissue) endothelial cell migrationto the graft surface, as well as attachment of mitogenssuch as fibroblast growth factor (FBF) and endothelial cellgrowth factor (ECGF) to the graft surface via heparin and/oralbumin coatings [83,36,84]. This latter approach has notreally shown enough promise (in the case of grafts, butperhaps not in the case of stents, see below) to be pur-sued seriously. The first approach has gone to clinicaltrials.

Early animal studies using the single-stage seeding ofDacron® gave sufficient positive results in animals to havegone on to human trials [85]. An inoculum of autologousECs was added to the clotting mix for the graft, followed byimmediate graft implantation. This process unfortunatelyyielded incomplete coverage on the lumen facing surfaceby the ECs. Budd et al. [86] used an overnight incubationtechnique that gave confluent surface coverage on an ePTFEgraft. Patency in dogs improved considerably when com-pared to unseeded ePTFE grafts [87].

The next improvement came with the recognition thatcoating the graft with components of the extra-cellularmatrix, such as fibronectin or plasma proteins, enhancedendothelialization rates [88]. This was a key step inadopting a practical solution to seeded graft fabrication.Presumably the RGD sequence (Arginine-Glycine-Asparticacid tripeptide sequence) in the fibronectin had helpedensure stable cell attachment on to a surface containingfibronectin [81]. The same researchers also demonstratedthat another four-peptide sequence of Arg-Glu-Asp-Val(REDV) is in fact more selective to EC attachment (and not toplatelet and SMC attachment), because ECs possess a recep-tor that binds to this sequence. This sequence is also foundin fibronectin. Such studies laid the mechanistic frameworkto obtaining good EC seeded grafts in humans, as describedbelow. It is to be noted that ex vivo seeding is mandatedin view of the relative non-selectivity of coatings such asfibronectin or fibrin glue for cells.

The first human trial for seeded ePTFE grafts in 1987
[89] gave disappointing results. Subsequently, a fibronectincoating [98] and fibrin glue [90] were used with EC takenfrom the jugular or cephalic vein with a seeding densityof 3000 cells/cm2, requiring 9–10 days after seeding toachieve full surface. Meinhart et al. [91] reported an 8-year

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follow-up trial, showing encouraging patency rates. Thehuman fibronectin-coated, EC-seeded grafts in 43 patientsshowed 70% patency rates (comparable to saphenous veingrafts) and somewhat lower patency rates for fibrin glue-coated grafts. Overall the entire cohort of 108 patients hadan impressive patency rate of 66% at 7 years. Evidently, bothfibronectin and fibrin glue contain the RGD domain thatpromotes EC adhesion. The 9-year data [92] continued toconfirm the 7-year trend.

A more efficient endothelialization method, calledelectrostatic seeding [93,94], placed the graft within a con-ducting hollow rod, with a guide wire passed through itas the second electrode. The graft is then immersed inEC-containing medium and a voltage applied to create atemporary positive charge in the graft to attract negativelycharged ECs. In vivo studies showed the persistence of theseeded ECs after 1 week in a canine artery [95] and nothrombosis at 6 weeks [96]. However, longer-term data forpatency using a control graft are not yet available for usto compare this technique of seeding against others. Nev-ertheless, the short time required to seed the graft in thismethod (less than 2 h) is indeed attractive.

Most data for seeded-EC patient trials is from Zilla andcoworkers. Nevertheless, even the limited data is sufficientto affirm the advantages of seeded ePTFE grafts comparedto the unseeded graft. Seeding time (using the Zilla method)is not excessive (10 days overall) and offers an improvedalternative for patients as they await the holy grail of thecompletely tissue-engineered graft.

2.8. Commercial products

A survey of vascular grafts suppliers shows the followingproduct categories being sold:

1. Dacron® grafts, woven and knitted (Boston Scientific,C.R.Bard).

2. ePTFE grafts (C.R.Bard, Boston Scientific, Edwards LifeSciences).

3. Carbon-impregnated ePTFE graft (C.R.Bard).4. Collagen-impregnated Dacron® (Boston Scientific).5. PU grafts (none, although CardioTech is in clinicals

with a PCU-based graft; note that a vascular accessgraft is being sold as Vectra® VAG by Thoratec, usingPEEU).

2.9. Summary

The following list summarizes the status of artificialarterial grafts:

(a) Saphenous (or other autologous) veins are preferredfor bypasses—their long-term (5 years and beyond)patency rates remain the best.

(b) Where sapheneous veins are not available, ePTFE-basedgrafts are used, although they offer lower patency rates.

(c) Seeding ePTFE graft with autologous ECs can give ade-quate EC coverage; clinical data showed patency ratecomparable to that of sapheneous veins beyond 5years.

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(d) Pre-clinical work on alternatives has shown some inter-esting conclusions:(i) Matching the distensibility of a synthetic graft to an

artery does not improve patency rates and throm-bus formation is not eliminated.

(ii) Currently, it takes too long (60–90 days) to developa completely tissue-engineered artery ex vivo to beclinically useful.

(iii) Tissue-engineered vessels still require a scaffoldto strengthen it. Biodegradable scaffolds would bepreferred.

(iv) Scaffolds using acellular, extra-cellular matrixhave shown good results but immunogenicity ofxenogenic matrices is still a concern.

(v) The development of artificial proteins with control-lable sequences of amino acids is highly promisingfor the next generation of grafts.

The near term outlook is to use EC-seeded, fully syn-thetic grafts. A variation on this theme would be seedingby bone marrow cells, as noted above. Over the mediumterm, we anticipate the use of EC-seeded, fully biodegrad-able grafts (with long biodegradation times). The long-termsolution is a complete vessel grown with autologous cells exvivo, with growth rates surpassing current rates by a factorof 2 or more.

Most grafting methods require invasive surgery, thoughnon-surgical methods (endovascular insertion) have beenproposed using procedures similar to that used in per-cutaneous transluminal angioplasty (PTA) procedures, i.e.,insertion of the graft through an artery in the leg andguiding it to the desired site. This involves a differentof material requirements, including self-expandability orballoon expandability, and good mechanical strength; how-ever, the porosity need not be low, or for that matter,the graft need not be contiguous. This feature facilitatestransmural endothelialization—something that is beingalready used to advantage in coronary stenting. Generally,a surgeon would prefer using a bypass graft while an inter-ventional cardiologist would prefer stenting. The choiceis also dictated by the extent of the blockage, and by thenature of the plaque itself. This leads us to the related fieldof stent materials.

3. Polymeric, biodegradable stents

Stenting following balloon angioplasty caused a revo-lution in interventional cardiology. Since it was approvedin 1994, an incredible array of stent designs has emergedfor both coronary artery and, to a lesser extent, peripheralartery blockages. The stent material has traditionally beenmetallic, and usually is stainless steel, cobalt–chromium orNitinol®. The last-named is used only in peripheral appli-cations.

Metal stents can result in reblockage of the same stentedsegment of the artery as before, a process known as
restenosis, for about 30–40% of the patients. By addinga drug-eluting coating, this rate has been reduced toabout 5% or less. Longer-term follow-up data has high-lighted problems with metallic drug-coated stents [97,98],significantly reducing their usage after 2005. The key prob-

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em is that the drug release delays endothelialization ofhe metal surface, although other issues have also beenetailed in an earlier review by the authors [99]. These

ssues in both bare and coated metal stents have causedoncerns in deploying a stent permanently in the body,specially when clinical data has shown that the stent is noteeded 6–9 months after deployment [100]. Consequently,

ncreased attention is now being placed on fully biodegrad-ble polymeric stents, which we will address in the nextection.

The concept of a fully polymeric, biostable stent doesot require much discussion, except to relate it to the prob-

ems seen with grafts. If a polymeric biomaterial can beound that has low thrombogenicity and can be deployedndovascularly, biostable stents are a reasonable option; ifo such material can be found, then given the redundancyf stents beyond 9 months, a biodegradable stent would belearly preferred.

.1. Requirements

Fully biodegradable coronary or peripheral stents mustatisfy the following requirements:

a) Both the stent material and its degradation productsmust be biocompatible.

b) Structural integrity must be maintained for 6 monthsand full degradation should occur in about 12–18months.

c) The degradation of the stent must not result inpieces of degraded material being released into thelumen.

d) The stent must be deployable either by balloon expan-sion or by self-expansion.

e) The stent must be safely anchored after deployment andnot migrate. It must also be able to structurally with-stand the blood vessel contractions.

f) The stent must be radio-opaque to enable its safe andfacile deployment.

Requirements (a)–(c) are applicable to both grafts (Sec-ion 2.4.1) and stents. Although reports have shown thataster graft scaffold degradation facilitates faster in situessel growth, the rapid degradation will also increasehe chances of pieces breaking off into the lumen. This isvoided with slower degradation, as cell or tissue over-rowth (endothelialization, in particular) would preventhe dislodgment of the pieces.

Requirements (d)–(f) are specific to endovasculareployment. For PTA, the crimped stent together withhe deflated balloon must have an initial diameter smallnough to enable the stent and balloon to navigate to thelocked coronary or peripheral artery site. A guide wire issed to guide the stent to the blockage, whereupon thealloon is inflated to expand the stent against the vesselall, so that it will be firmly anchored; this procedure is

alled balloon-expansion. Other devices use a sheath whichs then released at the site to allow the stent to self-expandgainst the vessel wall; this procedure goes by the name ofelf-expansion. Both methods dictate different viscoelasticroperties of the stent materials.

er Science 33 (2008) 853–874 867

In order for the cardiologist to see the stent to ensuresafe deployment and proper coverage, the stent mustbe visible to X-ray fluoroscopy employed in PTA pro-cedures. The polymeric biomaterial must be inherentlyradio-opaque or made to be radio-opaque by the use of anadditive.

3.1.1. BiocompatibilityIf we exclude thrombogenicity, there are many mate-

rials that meet the biocompatibility requirements for abiodegradable stent. The pioneering work on biodegrad-able stents was carried out by Zidar and coworkers at DukeUniversity. They performed animal tests using polymer-coated and fully polymeric stents [101]. One PLLA stentwhich was a filament-based, diamond-braided design wassuccessfully deployed in a canine femoral artery for 18months, with minimal neointima formation, thromboticresponse and inflammation. While the canine model mightnot have been ideal, the study paved the way for seriousdevelopment of biodegradable stents in general and PLLA-based stents in particular.

Polymers used in coated stents deployed in coronaryarteries were found to be less than satisfactory in termsof provoking both neointimal and inflammatory responses[102]. Biodegradable and biostable polymers were usedto coat a metal stent, then inserted in porcine coro-nary arteries. The biodegradable polymers included PLGA(85%lactide); polyorthoester, POE; caprolactone or PCL:and poly(hydroxybutyrate/hydroxyl valerate) copolymer;while the set of biostable polymers included polyurethane,poly(dimethyl siloxane) and poly(ethylene terephthalate).In this extensive, multicenter study, the major finding wasthat all the polymers showed greater neointimal thicken-ing than metal stents at 4 weeks; in addition, severe tomoderate inflammation was observed. In contrast to thiswork, Zidar et al.’s work [101] on the PLLA filament-basedstent, although carried out in a canine model, showed nosignificant inflammatory or proliferative response at 18months.

For PLLA, there appears to be an effect of molecu-lar weight on tissue reactions [103]: this study reportedcellular response to a stent coated with low-MW PLLA(80,000 Da) and high-MW PLLA (320,000 Da). Both thepolymers were spray-coated on to metal stents from achloroform solution. The stent coated with high-MW PLLAexhibited much less neo-intimal inflammatory response; itis likely that the higher-MW PLLA degraded slower, lower-ing the local pH much more slowly than the PLLA of lowerMW.

Other polymers, for example, a blend of poly(methylmethacrylate) and poly(2-hydroxyethyl methacrylate)[104], have not shown similar inflammatory responses inanimal studies. Conversely, PET [105] and polyurethane orPU [106] have shown inflammatory responses in pigs. ThePU was solution-cast on to a metal stent and dried beforeinsertion into porcine coronary artery. Both thrombotic
occlusion and extensive hyperplasia were detected at 48 hand at 28 days, respectively. In contrast, a study of stain-less steel stents coated with PEEU [107] found very littleocclusion or neo-intimal proliferation in porcine coronaryarteries. The woven poly(ethylene terephthalate) mesh

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stents [116,117]. One of them used PLLA of MW 600,000 fab-ricated into a laser cut tube [116] with a definite slanted-slitpattern. The stent did not recoil on balloon expansion, prob-ably helped by the slanted slit design. Tensile tests howevershowed 30–60% recovery. Although this seems large, the


used [105], presumably of Dacron® fibers, was also foundto cause severe neo-intimal proliferation at 4–6 weeks.

A newer class of materials, based on the amino acid tyro-sine, has been reported [108]. One extensive study done tocompare histological response to degrading polycarbonatesand polyadipates [109], at equivalent degrees of degrada-tion, showed that there were two distinct stages in thelocal tissue response for PLLA. When MW dropped by 25%,the thickness of the capsule around the implant, consist-ing of collagen enveloping macrophages and fibroblasts,measured 49 �m. As the MW dropped to 50%, the capsulethickness actually decreased to 20 �m, but increased backto 50 �m as MW dropped to 15%.

This two-stage behaviour was not noted in the othertwo polymers. The polycarbonate (initial MW = 44,000)had very long degradation time compared to the PLLA(initial MW = 80,000), whereas the polyadipate (initialMW = 72,000) degraded faster than the other two. Therationalization of the observed difference in histologicalresponse is that PLLA suffers substantial mass loss aroundthe time that the second capsule diameter increase isobserved; this mass loss may lead to localized pH decreasein the tissue, and/or the released “particles” or oligomersmay elicit a macrophage response. The polyadipate doesnot exhibit this mass loss prior to complete disintegration,enabling it to have lower inflammatory response in com-parison. It degrades by a bulk degradation process in vitro,without substantial mass loss until complete disintegra-tion occurs, differentiating it from poly(l-lactic acid) andpoly(glycolic acid) polymers.

Poly(ethylene carbonate) is an exciting newer polymerthat exhibits surface erosion in vivo, while retaining a higherosion rate [110], unlike the polyanhydrides, which areextremely slow-degrading polymers when formulated tosurface-erode. In addition, the eroding polymer does notappear to elicit any detrimental inflammatory response.This material has been touted as a drug-containing coatingon metal stents [111].

Unfavourable animal studies not withstanding, a PLLAstent was used in human coronary arteries for the firsttime in 2000 [112], without any undue inflammatoryreactions reported at 6 months. Guidant Inc. (now AbbottVascular) conducted clinical trials using a PLLA-basedstent known by the acronym BVS [113], using an entirelydifferent design. A third “self-reinforced” type of PLLAstent has been in ureteric stent clinicals [114]. This groupin Finland used extruded PLLA and PLGA stents, and haspaved the way to develop these materials for non-urethralapplications. The tyrosine-derived carbonate material hasbeen fabricated into a stent as well, and is reported to bein clinical trials [113,115]. None of the other biodegradablepolymers have been reported to be in clinical trials with ablood-contacting stent.

3.1.2. Viscoelastic requirementsThe viscoelastic requirements for a polymeric stent are

demanding. Specifically:

(a) The material must withstand a minimum collapsestrength of 1 bar (comparable to a slotted-tube metalstent).

er Science 33 (2008) 853–874

(b) If the stent is to be balloon expanded, it must have plas-tic flow in the strain range corresponding to the balloonexpansion to avoid recoil.

(c) Alternatively, if the stent is to be self-expanded, it needsa viscoelastic memory effect at 37 ◦C. This method canuse either a sheath or balloon crimping.

(d) If a method, such as a slide-and-lock mechanism, isused to lock the stent in place after balloon expansion,a more rigid polymer may be required.

Clearly, the modes of deployment for polymer stents dif-fer vastly from metal stents, due mainly to the polymer’scapacity to recoil, a consequence of the viscoelastic natureof polymers. Metallic stents are easily balloon-expandedwith only minimal recoil (less than 5%) precisely becausethe balloon expansion produces strains in the metal thatare in its plastic range of deformation. On the other hand,polymers with Tg and Tm higher than the expansion tem-perature, will deform elastically. Once the balloon deflates,the polymer’s elastic strain will be recovered, causingrecoil. There are three ways to overcome recoil in polymericstents, discussed below.

3.1.2.1. Plastic deformation. When a stent is balloonexpanded from its initial diameter to twice that value,the overall strain is about 100%. In this range, most poly-mers with Tg, Tm above 37 ◦C, will exhibit mostly plasticdeformation; but the elastic part of the deformation isnot zero. In particular, it depends on factors such as therate of strain, duration of stress imposition as well as thetemperature. Shown below are typical stress–strain curvesfor a PLLA (Tg ∼ 70 ◦C) and a PLGA polymer (Tg ∼ 45 ◦C),at 25 ◦C (Fig. 7).

As the temperature of deformation approaches Tg, therelative amount of plastic deformation increases, for a givenstrain. So one approach is to plasticize the polymer suffi-ciently to minimize elastic recovery; although not explicitlystated, this has been the approach for two biodegradable

Fig. 7. Stress–strain curves for a typical PLLA (Tg ∼ 70 ◦C) and PLGA(Tg ∼ 45 ◦C). Data were taken on a tensile testing machine, using dog-bone-shaped samples. Rate of extension was 5 mm/min.

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tent was successfully deployed in a porcine carotid artery118]. The tensile test, done using flat strips, was probablyot representative of the creep or recovery behaviour ofhe tube in radial expansion. Measuring the tube recoil as itas expanded for short periods (2–5 min) would have beenore instructive. The authors recently blended poly(4-

ydroxy butyrate) (P4HB) with PLLA in a 22/78 ratio to giveore rapid expansion with minimal recoil [119], compro-ising on a lower modulus at ambient, compared to pure

LLA. The stent used was a slotted-tube design, cut using aO2 laser.

In the second approach [117], a PdlLA material was plas-icized with supercritical CO2, then degassed partially tobtain a porous sent material. The stent is then deformednto a double intertwining helicoidal shape joined at thends. A special balloon catheter is used to expand the stentn a study in porcine arteries. Presumably, the expansiononditions (balloon shape, pressure and duration of pres-ure application) are set to maximize plastic deformation,ut these details are not given.

Guidant’s (now Abbot Vascular) PLLA-type BVS stent113] is reported to be similar to that described in [116]. TheVS patent [120] alludes to a laser-cut tube that is biaxiallytretched to form the stent. The material is then cooled andrimped on to a balloon. The exact mode of stent fabricationnd deployment is not clear at present.

.1.2.2. Elastic memory. A totally different approach to min-mize recoil is to incorporate elastic memory into the stent.emperature conditioning is one way to incorporate elas-ic memory, although other methods are possible, suchs solvent removal and reimbibition. Patents using thesepproaches have been granted [121]. These patents refer toevelopment of crosslinked polymers with transition tem-eratures in the appropriate range; however, these are notiodegradable per se.

Thermal activation [122] is achieved by setting the tubet the expanded diameter at T1. The stent is then cooled andet to the lower deployment diameter, at a lower temper-ture T2. The deployment temperature (Td) is in between2 and T1. When heated to Td, the body temperature, theaterial’s molecular mobility increases enabling the tube

o expand to the larger diameter.The Igaki-Tamai stent [112] uses such a temperature-

riggered memory effect to self-expand their filament-spunLLA, woven into a zig-zag pattern, assisted by balloonxpansion. Its main disadvantage is that it had to be heatedo 80 ◦C to trigger the self-expansion. The stent howeveras able to function otherwise, and is now being pursued

or stenting peripheral vessels.In our laboratory, we have attempted [123,124] to incor-

orate elastic memory into uncrosslinked, well-acceptediodegradable polymers such as PLLA and PLGA. Theesign is a multilayered structure, using poly(lactide)r poly(lactide-co-glycolide) layers. The outer layer washosen to provide significant memory effects at body tem-
eratures. The stresses and duration of stress impositionlso play a role in the extent of recoverable memory [122].he authors showed that a multilayered uncrosslinkedolymer system was able to function effectively as a self-xpanding stent at 37 ◦C.
er Science 33 (2008) 853–874 869

Other polymers used in stents include poly-(capralactone), or PCL, which is self expandable butneeds heating to 60 ◦C [125]. For true self-expansion(100% recovery of elastic memory), we need a crosslinkedpolymer with a transition at 37 ◦C or slightly above; such apolymer remains to be synthesized.

3.1.2.3. Other designs. A novel slide-and-lock mechanismhas been used to lock the stent in its expanded state [126],using tyrosine-derived polycarbonates (see Section 3.1.1)[108,109]. By incorporating iodine into the polymer, thepolymer was made radio-opaque without using externalagents, such as barium sulphate. The slide-and-lock mech-anism ensures a good fit against the vessel wall, althoughthe anchoring efficiency has not been reported. The com-pany is focusing on coronary stenting with this concept. Theresults of the clinical trials are eagerly awaited, especiallyin comparison to those involving PLLA stents.

Peripheral stents, unlike coronary stents, are located invessels which are exposed directly to external impact andpressures. They must have high elastic resilience to fullyregain their original shape after impact. This is why onlyhighly resilient Nitinol stents are preferred over stainlesssteel stents in the peripheral vasculature. Highly plasticstents and stents with a slide and lock mechanism wouldbe unsuitable for the periphery.

It remains to be seen which of these approaches wouldbe both most “cardiologist-friendly”, and have minimallong-term effects. The BVS stent, which has a drug-coatedlayer, is already in advanced clinical trials, while the othersare in pre-clinical studies.

3.2. Natural polymers

There appears to be only one report of a stent made fromnatural polymers: a self-expanding chitosan stent [127].This was a helicoidal stent made from strips of chitosanfilm, cast from an acidic (de-acetylated) chitosan solution.The strip is then wound around a mandrel, while stretch-ing it at the same time, at ambient conditions. Although notused in cardiovascular stenting, this was tested in the vasdeferens of rats, and found to self-expand after insertion, atbody temperatures. Collagen has not been studied to dateas a stent material.

3.3. Bio-inspired materials

To date, the use of artificial proteins or a bio-inspiredstructure has not been proposed for stenting. Nevertheless,the capability to functionalize the artificial proteins wouldprove useful in future considerations.

3.4. Surface treatments in biodegradable stents

The polymers used for biodegradable stents are allthrombogenic. For this reason, it is essential to have a layer
of endothelial cells covering the lumen-facing side of thestent, as quickly as possible after implantation. Bare metalstents have been reported to completely endothelializeafter 3–6 months in humans, although not verified by anydefinitive studies. This period may differ for biodegradable

in Polymer Science 33 (2008) 853–874

Fig. 8. An example of how the drug release of paclitaxel from fullybiodegradable stents can be manipulated relative to what is achievablefrom a coated metal stent. FR, TAXUS is the data for the fast-release versionof Taxus® stent, which has paclitaxel in a biostable polymer coated ontoa metal stent; MR, TAXUS is the slower-release formulation in the same


stents in both animals and humans. For the biodegradablestents, there are no reports of the time taken to endothe-lialize, even in animal studies; we can only assume that theprocess should be comparable in rate to bare metal stents.This is in contrast to the graft situation where endothe-lialization takes longer, because of the lack of access oftransmural ECs to the lumen side of the graft.

Another reason that biodegradable stents should beendothelialized quickly, is to eliminate the chance of piecesof the degrading stent breaking off into the lumen. This alsodictates that the degradation time of the stent be at leastlonger than the endothelialization time. The biodegradabil-ity also opens up the possibility of ‘revascularization’ at thesame stented site, should a lesion redevelop in the vicinity.

The group in NTU has employed physical methods suchas pore creation [128] and lithography to create nano-sized surface features; both have been shown to enhanceendothelialization rates in vitro. We have also used chem-ical modification [129] to attach ligands and thus directlyinfluence adherence of ECs. However, the effectiveness ofthese modifications to accelerate endothelialization ratesin vivo remains to be demonstrated.

Another elegant technique is to anchor an antibody thatcaptures floating endothelial progenitor cells (EPCs) fromthe bloodstream [130]. In clinical trials, the Genous® stent,which is not drug-eluting, was reported to have fared aswell as the paclitaxel-eluting stent. Follow-up data is nowneeded to address the concern that this stent works wellonly for a subsection of patients with a high amount offloating EPCs.

3.5. Bioactive agents in biodegradable stents

Bioactive agents are incorporated into the body of thestents, in order to achieve localized release with all itsaccompanying benefits, such as lack of a systemic effect,higher efficacy and fewer side effects Agents evaluatedin our laboratories include heparin [131]; drugs such assirolimus and paclitaxel [132–134]; plasmid DNA [135].Other groups have reported on paclitaxel incorporation inPdlLA [117] as well as rapamycin in PLLA [118].

A remarkable degree of control of drug release is achiev-able with drug incorporation within the stent body. Thepatent by Venkatraman and Boey [136] also describes theuse of multilayers to enable multidrug release, in differentdirections and at different rates. The degradation rates ofbiodegradable polymers can also be manipulated by using ebeam radiation to cause varying but well controlled degreeof chain scission in the polymer backbone, thus changingthe degradation rate [137].

In general, if the stent material has to be fabricated usingmelt extrusion, the only option is to coat the drug layeronto the stent. This has been reported for sirolimus [118].The Guidant/Abbott Vascular BVS system is also coatedwith a drug-containing layer. If solution casting is used,more bioactive agents may be incorporated into the body
of the stent, with better control over the release profiles.The issue for the need for a range of release profiles hasbeen addressed in an earlier recent review [99]. The authorsbelieve a short-term release of an anti-proliferative agentfrom a fully biodegradable stent that endothelializes in
stent. PLGA/PLLA DL refers to a double-layered stent developed in our lab-oratory, in which the paclitaxel is contained in the PLGA layer, whereasPLGA SL refers to a single-layer stent with paclitaxel incorporated intoPLGA. Redrawn based on data for the Taxus® stent release as given in [138].

3–6 months, optimally resolves the two major issues inthe cardiovascular stent world: restenosis and late-stagethrombosis (Fig. 8).

3.6. Commercial products

There are presently no commercial biodegradable stentsavailable. Two biodegradable stents are being clinicallytested: BVS (Abbott Vascular) and tyrosine–polycarbonate(REVA Medical/Boston Scientific). Detailed data on clinicalperformance are not available as of now, but conferencepresentations have presented a very encouraging picture.

3.7. Status and future directions

Somewhat in contrast to vascular grafts, the success ofpolymer stents depends both on stent design and the mate-rial. For stents, there is no ex vivo tissue-engineered optionavailable or needed: in situ cell coverage is sufficient tosolve most problems. The key appeal for stenting is thatit does not involve surgery, but is only an outpatient pro-cedure. However, the act of stenting itself (which involvesforcibly pushing the plaque to the vessel wall) may result invessel trauma, and hence restenosis. Nevertheless, stentingwill remain a vital therapeutic tool in interventional cardi-ology, and with earlier diagnosis of incipient blockages incoronary and peripheral vessels, may take over from graftscompletely at some point in the future. If and when it doesso, the most attractive stent option is a biodegradable one,but the following issues remain unsolved:

(a) Stents made with materials that degrade by bulkdegradation may not be the ideal choice, with theassociated concerns regarding floating debris in thebloodstream.

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b) Surprisingly, surface-eroding materials have not beenevaluated for stents: this could be partly due to the factthat endothelialization of such a stent surface is onlytemporary, as the eroding surface may take ECs with it.

(c) Faster endothelialization is the key to a successfulbiodegradable stent. The faster complete coverageoccurs, the better. Drugs that are non-selective willretard the rate of EC growth, hence a rapid release ofsuch drugs may be desirable.

d) Mechanically, the biodegradable stent must satisfy therequirements of adequate mechanical strength that ispreferred for direct stenting, i.e., combining angioplastyand stenting in one step. Current stents do not appearto match metals in this regard.

e) Deployment is still an issue, and although many elegantapproaches have been tried, a completely acceptablesolution (which does not involve storage of stent at lowtemperatures, for instance) is yet to be found; in thisrespect, newer materials may offer some hope.

. Outlook for polymeric cardiovascular implants

In the opinion of these authors, both grafts and fullyiodegradable stents can benefit immensely from the intro-uction of new polymers. A truly biomimetic graft material,hat fully satisfies the viscoelastic properties of bloodessels (not just match the compliance) including creep,ecovery and the entire stress–strain curve, is yet to beeveloped. This material is not entirely beyond the reach ofynthetic or biosynthetic chemists. Recent developments inrtificial proteins may well pave the way to mimicking theomposition and configurational aspects of a blood vessel.

The key problem to resolve for biodegradable stents ishat of fast endothelialization, and to a lesser extent, ofnflammatory reactions as the stent degrades. Work in ouraboratories and elsewhere is addressing these two keyoncerns in various ways. While it certainly appears thateveloping an intrinsically thrombo-resistant material is

mpossible, we feel that surface modifications will consti-ute an important part of generating favourable biologicalnterfaces in cardiovascular implants.

Since stents and grafts share a common interface (i.e.,lood), can learning from one field be used in the other? Inur opinion, this cross-device learning has not been exten-ive enough. Specifically, we feel that the techniques forx vivo endothelialization should be adopted in the stentorld; similarly, the learnings from the biodegradable stent

esearch should be translated into developing a biodegrad-ble graft material or a scaffold.

We end on an optimistic note. Although substantialmprovements in the graft world have been slow in beingommercially successful since ePTFE, we are poised to makeome inroads via EC seeding, as well as via tissue reac-or engineering. Research into the reasons for thrombusormation, and other aspects of blood incompatibility willave the way to directing specific functionalization of artifi-
ial proteins. Biodegradable stents are in advanced clinicals,nd with expected improvements in materials and bioac-ive agent delivery, should replace permanent stents in theext 10–15 years. Taken together, better management ofardiovascular disease is at hand.
er Science 33 (2008) 853–874 871

Acknowledgements

We thank Serene Kok Su Ling for redrawing Fig. 1, andXia Yun for redrawing Fig. 6.

References

[1] Allan B. Intra-ocular lens implants. Br Med J 2000;320:73–4.[2] Dhillon B, Kamal A, Leen C. Intravitreal sustained-release ganci-

clovir implantation to control cytomegalovirus retinitis in AIDS. IntJ STD AIDS 1998;9:227–30.

[3] Dlugi AM, Miller JD, Knittle J. Lupron®depot (leuprolide acetatefor depot suspension) in the treatment of endometriosis: arandomized, placebo-controlled, double-blind study. Fertil Steril1990;54:419–27.

[4] Berges R. Eligard®: pharmacokinetics, effect on testosterone andPSA levels and tolerability. Eur Urol Suppl 2005;4:20–5.

[5] Westphal M, Hilt DC, Bortey E, Delavault P, Olivares R, Warnke PC,et al. A phase 3 trial of local chemotherapy with biodegradable car-mustine (BCNU) wafers (Gliadel® wafers) in patients with primarymalignant glioma. Neuro-Oncology 2003;5:79–88.

[6] Westhoff C. Depot-medroxyprogesterone acetate injection (Depo-Provera®): a highly effective contraceptive option with proven long-term safety. Contraception 2003;68:75–87.

[7] Eaglstein WH, Falanga V. Tissue engineering and the developmentof Apligraf® , a human skin equivalent. Clin Ther 1997;19:894–905.

[8] Cook DM, Biller BMK, Vance ML, Hoffman AR, Phillips LS, FordKM, et al. The pharmacokinetic and pharmacodynamic character-istics of a long-acting growth hormone (GH) preparation (NutropinDepot) in GH-deficient adults. J Clin Endocrinol Metab 2002;87:4508–14.

[9] Columbo A, Karvouni E. Biodegradable stents: ‘fulfilling the missionand stepping away’. Circulation 2000;102:371–3.

[10] Campbell GR, Campbell JH. Development of tissue-engineeredgrafts. Curr Pharm Biotechnol 2007;8:43–50.

[11] Monson KL, Goldsmith W, Barbaro NM, Manley GT. Axial mechanicalproperties of fresh human cerebral blood vessels. J Biomech Eng –Trans ASME 2003;125:288–94.

[12] Berglund JD, Nerem RM, Sambanis A. Viscoelastic testing method-ologies for tissue engineered blood vessels. J Biomech Eng – TransASME 2005;127:1176–84.

[13] Abbott WA, Megerman J, Hasson JE, L’Italien G, Warnock DF. Effectof compliance mismatch on vascular graft patency. J Vasc Surg1987;5:376–81.

[14] Bacourt F. Prospective randomized study of carbon-impregnatedpolytetrafluoroethylene grafts for below the knee popliteal and dis-tal bypass: results at 2 years. Ann Vasc Surg 1997;11:596–603.

[15] Kapfer X, Meichelboeck W, Groegler F-M. Comparison of carbon-impregnated and standard ePTFE prostheses in extra-anatomicalanterior tibial artery bypass: a prospective randomized multicenterstudy. Eur J Vasc Endovasc Surg 2006;32:155–68.

[16] Eiberg JP, Roder O, Stahl-Madsen M, Eldrup N, Qvarfordt P, LaursenA, et al. Fluropolymer-coated Dacron® graft versus PTFE graftsfor femorofemoral crossover by pass. Eur J Vasc Endovasc Surg2006;32:431–8.

[17] Parsson H, Jundzill W, Johansson K, Jonung T, Norgren L.Healing characteristics of polymer-coated or collagen-treatedDacron® grafts: an experimental porcine study. Cardiovasc Surg1994;2:242–8.

[18] Sreerekha PR, Krishnan LK. Cultivation of endothelial pro-genitor cells on fibrin matrix and layering on Dacron®/polytetrafluoroethylene vascular grafts. Artif Organs 2006;30:242–9.

[19] van der Bas JM, Quax PH, Van Der Berg C, Visser MJ, van der Linden E,van Bockel JH. Ingrowth of aorta wall into stent grafts impregnatedwith basic fibroblast growth factor: a porcine in vivo study of bloodvessel prosthesis healing. J Vasc Surg 2004;39:850–8.

[20] Granke K, Ochsner JL, McClugage SG, Zdrahal P. Analysis of grafthealing in a new elastomer-coated vascular prosthesis. CardiovascSurg 1993;1:254–61.

[21] Hake U, Gabbert H, Iversen S, Jakob H, Schmiedt W, Oelert H. Eval-uation of the healing of precoated vascular Dacron® prostheses.Langenbecks Arch Chir 1991;376:323–9.

[22] Hake U, Gabbert H, Iversen S, Jakob H, Schmiedt W, Oelert H. Healingparameters of a new albumin-coated knitted Dacron® graft. ThoracCardiovasc Surg 1991;39:208–13.

in Polym
[23] Prager MR, Hoblaj T, Nanobashvili J, Sporn E, Polterauer P, WagnerO, et al. Collagen- versus gelatine-coated Dacron® versus stretchPTFE bifurcation grafts for aortoiliac occlusive disease: long-termresults of a prospective randomized multicenter trial. Surgery2003;134:80–5.

[24] Devine C, McCollum C. Heparin-bonded Dacron® or polytetraflu-orethylene for femoropopliteal bypass: five-year results of aprospective randomized multicenter clinical trial. J Vasc Surg2004;40:924–31.

[25] Branchereau A, Rudondy P, Gournier J-P, Espinoza H. The albumin-coated Dacron® aorticprosthesis: a clinical study. Ann Vasc Surg1990;4:138–42.

[26] Grigioni AM, Daniele C, D’Avenio G, Barbaro V. Biomechanics andhemodynamics of grafting. In: Tura A, editor. Vascular grafts: exper-iment and modelling. Boston: WIT Press; 2003. p. 41–82.

[27] Lee JM, Wilson GJ. Anisotropic tensile viscoelastic propertiesof vascular graft materials at low strain rates. Biomaterials1986;7:423–31.

[28] Green RM, Abbott WM, Matsumoto T, Wheeler JR, Miller N, VeithFJ, et al. Prosthetic above-knee femoropopliteal bypass grafting: fiveyear results of a randomized trial. J Vasc Surg 2000;31:417–25.

[29] Pevec WC, Darling RC, L’Italien GJ, Abbott WM. Femoropoplitealreconstruction with knitted nonvelour Dacron® versus expandedpolytetrafluoroethylene. J Vasc Surg 1992;16:60–5.

[30] Jensen LP, Lepäntalo M, Fossdal JE, Røder OC, Jensen BS, Madsen MS,et al. Dacron® or PTFE for above-knee femoropopliteal bypass. Amulticenter randomised study. Eur J Endovasc Surg 2007;34:44–9.

[31] Cameron BL, Suchida HT, Connall TP, Nagae T, Furukawa K, Wil-son SE. High porosity PTFE improves endothelialization of arterialgrafts without increasing early thrombogenicity. J Cardiovasc Surg1993;34:281–5.

[32] Akers DL, Du YH, Kempscinski RF. The effect of carbon coatingand porosity on early patency of expanded polytetrafluoroethylenegrafts: an experimental study. J Vasc Surg 1993;18:10–5.

[33] Kohler TR, Stratton JR, Kirkman TR, Johansen KH, Zierler BK, ClovesAW. Conventional high porosity polytetrafluoroethylene grafts:clinical evaluation. Surgery 1992;112:901–7.

[34] Kapfer X, Meichelboeck W, Groegler FM. Comparison of carbon-impregnated and standard ePTFE prostheses in extra-anatomicalanterior tibial artery bypass: a prospective randomized multicenterstudy. Eur J Vasc Endovasc Surg 2006;32:155–68.

[35] Greisler HP, Cziperle DJ, Kim DU, Garfield JD, Petsikas D, MurchanPM, et al. Enhanced endothelialization of expanded polytetrafluo-roethylene grafts by fibroblast growth factor type 1 pretreatment.Surgery 1992;112:244–55.

[36] Walpoth BH, Zammaretti P, Cikirikcioglu M, Khabiri E, Djebaili MK,Pache J-C, et al. Enhanced thickening of expanded polytetraflu-oroethylene grafts coated with fibrin or fibrin-releasing vascularendothelial growth factor in the pig carotid artery interpositionmodel. J Thorac Cardiovasc Surg 2007;133:1163–70.

[37] Pinchuk L. A review of the biostability and carcinogenicity ofpolyurethanes in medicine and the new generation of ‘biostable’polyurethanes. J Biomater Sci – Polym Ed 1994;6:225–67.

[38] Tanzi MC, Fare S, Petrini P. in vitro stability of polyether and poly-carbonate urethanes. J Biomater Appl 2000;14:325–48.

[39] Zhao Q, Agger MP, Fitzpatrick M, Anderson JM, Hiltner A, Stokes K, etal. Cellular interactions with biomaterials: in vivo cracking of pre-stressed Pellethane 2363-80A. J Biomed Mater Res 1990;24:621–37.

[40] Gilding DK, Reed AM, Askill IN, Briana SG. Biocompatible microp-orous polymeric materials and methods of making the same. USpatent 4,704,310; 1987.

[41] Martz H, Paynter R, Slimane SB, Beaudoin G, Guidoin R. Hydrophilicmicroporous polyurethane versus expanded PTFE grafts as substi-tutes in the carotid arteries of dogs. A limited study. J Biomed MaterRes 1988;22:63–9.

[42] Teijeira FJ, Lamoureaux G, Tetreault J-P, Bauset R, Guidoin R, MaroisY, et al. Hydrophilic polyurethane versus autologous femoral veinas substitutes in the femoral arteries of dogs: quantification ofplatelets and fibrin deposits. Biomaterials 1989;10:80–4.

[43] Kakkos SK, Haddad R, Haddad GK, Reddy DJ, Nypaver TJ, Lin JC, et al.Results of aggressive graft surveillance and endovascular treatmenton secondary patency rates of Vectra® Vascular Access Grafts. J VascSurg 2007;45:974–80.

[44] Pinchuk L. Crack-resistant polycarbonate urethane polymer pros-thesis. US patent 5,133,742; 1992.

[45] Hess F, Jerusalem C, Braun B. Endothelialization process of a fibrouspolyurethane microvascular prosthesis after implantation in theabdominal aorta of the rat—a scanning electron microscopic study.J Cardiovasc Surg 1983;24:516–24.

er Science 33 (2008) 853–874

[46] Zhang Z, King MW, Guidoin R, Therrien M, Pezolet M, Adnot A, et al.Morphological: physical and chemical evaluation of the Vascugraft®

arterial prosthesis: comparison of a novel polyurethane device withother microporous structures. Biomaterials 1994;15:483–501.

[47] Jeschke MG, Hermanutz V, Wolf SE, Koveker GB. Polyurethanevascular prosthesis decreases neointimal formation comparedwith expanded polytetrafluoroethylene. J Vasc Surg 1999;29:168–76.

[48] Bull PG, Denck H, Guidoin R, Gruber H. Preliminary clinical expe-rience with polyurethane vascular prosthesis in femoro-poplitealreconstruction. Eur J Vasc Surg 1992;6:217–24.

[49] Zhang Z, Marois Y, Guidoin RG, Bull P, Marois M, How T, et al.Vascugraft® polyurethane arterial prosthesis as femoro-poplitealand femoro-peroneal bypasses in humans: pathological, struc-tural and chemical analyses of four excized grafts. Biomaterials1997;18:113–24.

[50] Glickman MH, Stokes GK, Ross JR, Schuman ED, Sternbergh III WC,Lindsberg JS, et al. Multicenter evaluation of a polyurethaneureavascular access graft as compared with the expanded polytetrafluo-roethylene vascular access graft in hemodialysis applications. J VascSurg 2001;34:465–73.

[51] Phaneuf MD, Dempsey DJ, Bide MJ, Szycher M, Quist WC, LoGerfoFW. Bioengineering of a novel small diameter polyurethane vas-cular graft with covalently bound recombinant hirudin. ASAIO J1998;44:M653–8.

[52] Gupta BS, Kasyanov VA. Biomechanics of human common carotidartery and design of novel hybrid textile compliant vascular grafts.J Biomed Mater Res 1997;34:341–9.

[53] Gomez-Alonso A, Garcia-Criado FJ, Parreno-Manchado FC, Garcia-Sanchez JE, Garcia-Sanchez E, Parreno-Manchado A, et al. Studyof the efficacy of coated Vicryl Plus® antibacterial suture (coatedpolyglactin 910 suture with triclosan) in two animal models of gen-eral surgery. J Infect 2007;54:82–8.

[54] Greisler HP. Arterial regeneration over absorbable prostheses. ArchSurg 1982;117:1425–31.

[55] Greisler HP, Ellinger J, Schwarcz TH, Golan J, Raymond RM, Kim DU.Arterial regeneration over polydioxanone prostheses in the rabbit.Arch Surg 1987;122:715–21.

[56] Greisler PH, Tatersall CW, Klosak JJ, Cabusao EA, Garfield JD,Kim DU. Partially resorbable vascular grafts in dogs. Surgery1991;110:645–55.

[57] Greisler HP, Dennis JW, Endean ED, Ellinger J, Friesel R, Burgess W.Macrophage/biomaterial interactions: the stimulation of endothe-lialization. J Vasc Surg 1989;9:588–93.

[58] Zweep HP, Satoh S, van der Lei B, Hinrichs WL, Feijen J, Wildevuur CR.Degradation of a supporting prosthesis can optimize arterializationof autologous veins. Ann Thorac Surg 1993;56:1117–22.

[59] Greisler HP, Endean ED, Klosak JJ, Ellinger J, Dennis JW, Buttle K, etal. Polyglactin 910/polydioxanone bicomponent totally resorbablevascular prostheses. J Vasc Surg 1988;7:697–705.

[60] Nerem RD, Seliktar D. Vascular tissue engineering. Ann Rev BiomedEng 2001;3:225–43.

[61] Teebken OE, Haverich A. Tissue engineering of small diameter vas-cular grafts. Eur J Vasc Endovasc Surg 2002;23:475–85.

[62] Isenberg BC, Williams C, Tranquillo RT. Small-diameter artificialarteries engineered in vitro. Circ Res 2006;98:25–35.

[63] Gao J, Niklason L, Langer R. Surface hydrolysis of poly (glycolic acid)meshes increases the seeding density of vascular smooth musclecells. J Biomed Mater Res 1998;42:417–24.

[64] Niklason LE, Gao J, Abbott WM, Hirschi KK, Houser S, Marini R, etal. Functional arteries grown in vitro. Science 1999;284:489–93.

[65] Hoerstrup SP, Zund G, Sodian R, Schnell AM, Grunenfelder J, TurinaMI. Tissue engineering of small caliber vascular grafts. Eur J Cardio-Thorac Surg 2001;20:164–9.

[66] Hibino N, Imai Y, Shin-Oka T, Aoki M, Watanabe M, Kosaka Y, etal. First successful clinical application of tissue-engineered bloodvessel, Kyobu geka. Jpn J Thorac Surg 2002;55:368–73.

[67] Matsumura G, Hibino N, Ikada Y, Kurosawa H, Shin’oka T. Successfulapplication of tissue-engineered vascular autografts: clinical expe-rience. Biomaterials 2003;24:2303–8.

[68] Shin’oka T, Matsumara G, Hibino N, Naito Y, Watanabe M, Kon-uma T, et al. Midterm clinical result of tissue-engineered vascularautografts seeded with autologous bone marrow cells. J Thorac Car-
diovasc Surg 2005;129:1329–37.
[69] Weinberg CB, Bell E. A blood vessel model constructed from collagenand cultured vascular cells. Science 1986;231:397–400.

[70] Ziegler T, Nerem RM. Tissue engineering of a blood vessel: regu-lation of vascular biology by mechanical stresses. J Cell Biochem1994;56:204–9.

in Polym
[71] Kanda K, Matsuda T. In vitro reconstruction of hybrid arterialmedia with molecular and cellular orientations. Cell Transplant1994;3:537–45.

[72] Seliktar D, Black RA, Vito RP, Nerem RM. Dynamic mechanical condi-tioning of collagen-gel blood vessel constructs induces remodelingin vitro. Ann Biomed Eng 2000;28:351–62.

[73] Boland ED, Matthews JA, Pawlowski KJ, Simpson DG, Wnek GE,Bowlin GL. Electrospinning collagen and elastin: preliminary vas-cular tissue engineering. Front Biosci 2004;9:1422–32.

[74] Matthews JA, Wnek GE, Simpson DG, Bowlin GL. Electrospinning ofcollagen nanofibers. Biomacromolecules 2002;3:232–8.

[75] Yannas I. Natural materials. In: Ratner B, Hoffman AS, Schoen FJ,Lemmons JE, editors. Biomaterials science: an introduction to mate-rials in medicine. San Diego: Elsevier; 2004. p. 127–36.

[76] Courtman DW, Pereira CA, Kashef V, McComb D, Lee JM, WilsonGJ. Development of a pericardial acellular matrix biomaterial: bio-chemical and mechanical effects of cell extraction. J Biomed MaterRes 1994;28:655–66.

[77] van Hest JCM, Tirrell DA. Protein-based materials, toward a newlevel of structural control. Chem Commun 2001;19:1897–904.

[78] Urry DW. Physical chemistry of biological free energy transductionas demonstrated by elastic protein-based polymers. J Phys Chem B1997;101:11007–28.

[79] Nicol A, Gowda DC, Parker TM, Urry DW. Cell adhesive propertiesof bioelastic materials containing cell attachment sequences. In:Gebelein CC, Carraher CE, editors. Biotechnology of bioactive poly-mers. New York: Plenum Press; 1994. p. 95.

[80] Panitch A, Yamaoka T, Fournier MJT, Mason L, Tirrell DA. Design andbiosynthesis of elastin-like artificial extracellular matrix proteinscontaining periodically spaced fibronectin CS5 domains. Macro-molecules 1999;389:167–70.

[81] Hubbell JA, Massia SP, Desai NP, Drumheller PD. Endothelial cell-selective materials for tissue engineering in the vascular graft via anew receptor. Nat Biotechnol 1991;9:568–72.

[82] Heilshorn SC, DiZio KA, Walsh ER, Tirrell DA. Endothelial cell adhe-sion to the fibronectin CS5 domain in artificial extracellular matrixproteins. Biomaterials 2003;24:4245–52.

[83] Gray JL, Kang SS, Zenni GC, Kim DU, Kim PI, Burgess WH, et al. FGF-1 affixation stimulates ePTFE endothelialization without intimalhyperplasia. J Surg Res 1994:596–612.

[84] Doi K, Matsuda T. Enhanced vascularization in a microporouspolyurethane graft impregnated with basic fibroblast growth factorand heparin. J Biomed Mater Res 1997;34:361–70.

[85] Herring M, Baughman S, Glover J. Endothelium develops onseeded human arterial prosthesis: a brief clinical note. J Vasc Surg1985;2:727–30.

[86] Budd JS, Allen KE, Bell PRF. Effects of 2 methods of endothelial-cell seeding on cell retention during blood-flow. Br J Surg 1991;78:878–82.

[87] Budd JS, Allen KE, Hartley G, Bell PRF. The effect of preformedconfluent endothelial cell monolayers on the patency and throm-bogenicity of small caliber vascular grafts. Eur J Vasc Surg 1991;5:397–405.

[88] Williams SK, Jarrell BE, Friend L, Radomski JS, Carabasi RA, KoolpeE, et al. Adult human endothelial cell compatibility with prostheticgraft material. J Surg Res 1985;38:618–29;Zilla P, Fasol R, Deutsch M, Fischlein T, Minar E, Hammerle A, et al.Endothelial cell seeding of polytetrafluoroethylene vascular graftsin humans: a preliminary report. J Vasc Surg 1987;6:535–41.

[89] Kadletz M, Moser R, Preiss P, Deutsch M, Zilla P, Fasol R. In vitrolining of fibronectin coated PTFE grafts with cryopreserved saphe-nous vein endothelial cells. Thorac Cardiovasc Surg 1987;35(S2):143–7.

[90] Zilla P, Fasol R, Preiss P, Kadletz M, Deutsch M, Schima H, et al. Useof fibrin glue as a substrate for in vitro endothelialization of PTFEvascular grafts. Surgery 1989;105:515–22.

[91] Meinhart J, Deustch M, Zilla P. Eight years of clinical endothelial celltransplantation: closing the gap between prosthetic grafts and veingrafts. ASAIO J 1997;43:M515–21.

[92] Deutsch M, Meinhart J, Fischlein T, Preiss P, Zilla P. Clinical autolo-gous in vitro endothelialization of infrainguinal ePTFE grafts in 100patients: a 9-year experience. Surgery 1999;126:847–55.

[93] Bowlin GL, Rittgers SE, Milsted A, Schmidt SP. In vitro evaluation
of electrostatic endothelial cell transplantation onto 4 mm inte-rior diameter expanded polytetrafluoroethylene grafts. J Vasc Surg1998;27:504–11.
[94] Fields C, Cassano A, Allen C, Meyer A, Pawloski KJ, Bowlin GL, et al.Endothelial cell seeding of a 4-mm I.D. polyurethane vascular graft.J Biomater Appl 2002;17:45–70.

er Science 33 (2008) 853–874 873

[95] Bowlin GL, Meyer A, Fields C, Cassano A, Makhoul RG, Allen C, et al.The persistence of electrostatically seeded endothelial cell lining asmall diameter expanded polytetrafluoroethylene graft. J BiomaterAppl 2001;16:157–73.

[96] Pawlowski KJ, Rittgers SE, Schmidt SP, Bowlin GL. Endothelial cellseeding of polymeric vascular grafts. Front Biosci 2004;9:1412–21.

[97] McFadden EP, Stabile E, Regar ECheneau E, Ong ATL, Kinnaird T, Sud-dath WO, et al. Late thrombosis in drug-eluting coronary stents afterdiscontinuation of antiplatelet therapy. Lancet 2004;364:1519–21.

[98] Virmani R, Guagliumi G, Farb A, Musumeci G, Grieco N, Motta T, et al.Localized hypersensitivity and late coronary thrombosis secondaryto a sirolimus-eluting stent: should we be cautious? Circulation2004;109:701–5.

[99] Venkatraman S, Boey F. Release profiles in drug-eluting stents:issues and uncertainties. J Control Release 2007;120:149–60.

[100] Kimura T, Yokoi H, Nakagawa Y. Three-year follow-up afterimplantation of metallic coronary artery stents. N Engl J Med1996;334:561–6.

[101] Zidar J, Lincoff A, Stack R. Biodegradable stents. In: Topol EJ, editor.Textbook of interventional cardiology. New York: Saunders; 1994.p. 787–802.

[102] van der Giessen WJ, Lincoff AM, Schwartz RS, van Beusekom HMM,Serruys PW, Holmes Jr DR, et al. Marked inflammatory sequelae toimplantation of biodegradable and nonbiodegradable polymers inporcine coronary arteries. Circulation 1996;94:1690–7.

[103] Lincoff MA, Furst JG, Ellis SG, Tuch RJ, Topol EJ. Sustained localdelivery of Dexmethasone by a novel intravascular eluting stent toprevent restenosis in the porcine coronary injury model. J Am CollCardiol 1997;29:808–16.

[104] Bar FW, van der Veen FH, Benzina A, Habets J, Koole LH. Newbiocompatible polymer surface coating for stents results in a lowneo-intimal response. J Biomed Mater Res 2000;52:193–8.

[105] Murphy JG, Schwartz RS, Edwards WD, Camrud AR, Vliestra RE,Holmes Jr DR. Percutaneous polymeric stents in porcine coronaryarteries: initial experience with polyethylene terephthalate stents.Circulation 1992;82:1596–604.

[106] Holmes DR, Camrud AR, Jorgenson MA, Edwards WD, Schwartz RS.Polymeric stenting in the porcine coronary artery model; differ-ential outcome of exogenous fibrin sleeves versus polyurethane-coated stents. J Am Coll Cardiol 1994;24:525–31.

[107] de Scheerder IK, Wilczek KL, Verbeken EV, Vandorpe J, Lan PN,Schacht E, et al. Biocompatibility of polymer coated oversizedmetallic stents implanted in normal porcine coronary arteries.Atherosclerosis 1995;114:105–14.

[108] Pulapura S, Kohn J. Tyrosine-derived polycarbonates: backbone-modified “pseudo”-poly (amino acids) designed for biomedicalapplications. Biopolymers 1992;32:411–7.

[109] Hooper KA, Macon ND, Kohn J. Comparative histological evalua-tion of new tyrosine-derived polymers and poly (l-lactic acid) asa function of polymer degradation. J Biomed Mater Res 1998;41:443–54.

[110] Dadsetan M, Christenson EM, Unger F, Ausborn M, Kissel T, Hiltner A,et al. In vivo biocompatibility and biodegradation of poly (ethylenecarbonate). J Control Release 2003;93:259–70.

[111] Unger F, Wedstedt U, Hanefield P, Wombacher R, Zimmermann S,Greiner A, et al. Poly (ethylene carbonate): a thermoplastic andbiodegradable polymer for drug eluting stent coatings? J ControlRelease 2007;117:312–21.

[112] Tamai H, Igaki K, Kyo E, Kosuga K, Kawashima A, Matsui S, et al. Initialand 6-month results of biodegradable poly-l-lactic acid coronarystents in humans. Circulation 2000;102:399–404.

[113] Waksman R. Update on bioabsorbable stents: from bench to clinical.J Interv Cardiol 2006;19:414–21.

[114] Isotalo T, Talja M, Valimaa T, Tormala P, Tammela TLJ. A bioab-sorbable self-expandable self-reinforced poly l-lactic acid urethralstent for recurrent urethral strictures: long term results. J Endourol2002;16:759–62.

[115] Reva Medical. Reva Medical announces enrollment of firstpatients in the RESORB clinical trial. http://www.teamreva.com/company news RESORB.html 2007.

[116] Grabow N, Schlun M, Sternberg K, Hakkanson N, Kramer S, SchmitzK-P. Mechanical properties of laser-cut poly (lactide) microspeci-mens: implications for stent design, manufacture and sterilization.
J Biomech Eng – Trans ASME 2005;127:25–31.
[117] Vogt F, Stein A, Rettemeir G, Krott N, Hoffmann R, vom Dahl J, et al.Long-term assessment of a novel biodegradable paclitaxel-elutingcoronary polylactide stent. Eur Heart J 2005;25:1330–40.

[118] Bunger CM, Grabow N, Sternberg K, Kroger C, Ketner L, SchmitzK-P, et al. Sirolimus-eluting biodegradable poly-l-lactide stent for

http://www.teamreva.com/company_news_RESORB.html%202007

http://www.teamreva.com/company_news_RESORB.html%202007

in Polym
peripheral vascular application: a preliminary study in porcinecarotid arteries. J Surg Res 2007;139:77–82.

[119] Grabow N, Bunger CM, Schultze C, Schmol K, Martin DP, WilliamsSF, et al. A biodegradable slotted tube stent based on poly (L-lactide)and poly (4-hydroxy butyrate) for rapid balloon expansion. AnnBiomed Eng 2007;35:2031–8.

[120] Huang B, Schaible S, Kaufman R, Kennedy J, Gale DC, Amin S, et al.Method of fabricating an implantable device with biaxially orientedpolymers. US patent application US 2006/0020330A1; 2006.

[121] Froix M. Expandable polymeric stent with memory and deliveryapparatus and method. US patent 6,248,129; 2001.

[122] Wong YS, Ying X, Boey F, Venkatraman S. Shape memory inuncrosslinked biodegradable polymers. J Biomater Sci – Polym Ed2008;19:175–91.

[123] Venkatraman S, Tan LP, Joso JF, Boey F, Wang X. Biodegradable stentswith elastic memory. Biomaterials 2006;27:1573–8.

[124] Tan LP, Venkatraman S, Joso JFD, Boey F. Collapse pressures of bilay-ered biodegradable stents. J Biomed Mater Res B 2006;79:102–7.

[125] Beck A. About a new balloon-expandable plastic stent—preliminaryreport on the heat-labile thermo-stent. Radiologe 1990;30:347–50.

[126] Reva Medical. Unique stent geometry—slide & lock design.http://www.teamreva.com/heart stents about.html 2007.

[127] Lauto A, Ohebshalom M, Esposito M, Mingin J, Li PS, Felsen D, et al.Self-expandable chitosan stent: design and preparation. Biomate-rials 2001;22:1869–74.

[128] Deepa N, Venkatraman S. Effect of pore size and interpore distanceon endothelial cell growth on polymers. J Biomed Mater Res A;doi:10.1002/jbm.a.31749, in press.

[129] Xia Y, Pan J, Chaw ST, Boey F, Venkatraman S. Bio-molecule immo-
bilization to enhance endothelial cell adhesion and growth on poly(l-lactide) surface. J Biomed Mater Res A, under review.
[130] Cardiovascular News International. First do no harm – theGenous stent. http://www.orbusneich.com/news/media/0701-04%20CardioVascular%20News%20Issue%204%20Jan%202007.pdf2007.

er Science 33 (2008) 853–874

[131] Tan LP, Venkatraman S, Sung PF, Wang XT. Effect of plasticiza-tion on heparin release from biodegradable matrices. Int J Pharm2004;283:89–96.

[132] Alexis F, Rath SK, Boey F. In vitro study of release mechanisms ofpaclitaxel and rapamycin from drug-incorporated biodegradablestent matrices. J Control Release 2004;98:67–74.

[133] Wang X, Venkatraman S, Boey F, Loo SC, Tan LP. Effects of con-trolled release of sirolimus from polymer matrices on humanartery coronary smooth muscle cells. J Biomater Sci – Polym Ed2007;18:1401–14.

[134] Lao LL, Venkatraman S. Paclitaxel release from single anddouble-layered poly (dl-lactide-co-glycolide)/poly (l-lactide) filmfor biodegradable stent application. J Biomed Mater Res A2008;87A:1–7.

[135] Yamini R, Mondal D, Godbey WT, Venkatraman S. Sustained releaseof complexed and naked pDNA from polymer films. J Biomed MaterRes B 2007;85B:496–503.

[136] Venkatraman S, Boey F. Polymeric stent and method of manufacture.US and PCT application, PCT/SG2004/000180. US 10/867,617 July2004.

[137] Chia NK, Venkatraman S, Boey F, Cadart S, Loo JSC. Controlleddegradation of multilayered poly(lactide-co-glycolide) films usingelectron beam irradiation. J Biomed Mater Res A 2008;84A:980–7.

[138] Ranade SV, Miller KM, Richard RE, Chan AK, Allen MJ, Helmus MN.Physical characterization of controlled release of apclitaxel fromthe TAXUS® ExpressTM drug-eluting stent. J Biomed Mater Res A2004;71:625–34.

[139] Tortora G, Grabowski SR. Principles of anatomy and physiology. 7th
ed. New York, New York: Harper Collins College Publishers; 1993,625 pp., [reprinted with permission from John Wiley & Sons, Inc.].
[140] Eberhart A, Zhang Z, Guidoin R, Laroche G, Guay L, de la FayeD, et al. A new generation of polyurethane vascular prosthe-ses: rara avis or ignis fatuus? J Biomed Mater Res B 1999;48:546–58.

http://www.teamreva.com/heart_stents_about.html%202007

http://dx.doi.org/10.1002/jbm.a.31749

http://www.orbusneich.com/news/media/0701-04%2520CardioVascular%2520News%2520Issue%25204%2520Jan%25202007.pdf%202007

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