Fisher John P. e.a. (ed.) Tissue Engineering

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25-22 Tissue Engineering

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Tissue Engineered

Vascular Grafts

Rachael H. Schmedlen

Wafa M. Elbjeirami

Andrea S. Gobin

Jennifer L. West

Rice University

26.1 Introduction.............................................. 26-1

Current Vascular Grafts • Tissue Engineering

26.2 Cell Sources for Vascular Tissue Engineering .......... 26-3

26.3 Genetic Modiﬁcation of Vascular Cells................. 26-4

26.4 Scaffolds for Vascular Tissue Engineering.............. 26-5

Natural Scaffold Materials • Synthetic Polymer Scaffolds

26.5 Bioreactors for Mechanical Conditioning.............. 26-8

26.6 Conclusions .............................................. 26-8

References ....................................................... 26-9

26.1 Introduction

Cardiovascular disease remains the leading cause of death in the United States, claiming more lives each

year than the next ﬁve leading causes of death combined (American Heart Association National Center

for Health Statistics). Coronary heart disease caused more than 1 in every 5 American deaths in 2000 and

required approximately 500,000 coronary artery bypass graft surgeries (CABGs) that year. Bypass grafting

is also used in the treatment of aneurysmal disease or trauma. At present, surgeons use autologous tissue

and synthetic biomaterials as vascular grafts. Transplantation of autologous tissue has the best outcome

in small diameter applications such as CABG because synthetic grafts lack long-term patency for small

diameter applications (<6 mm). However, autologous tissue is limited in supply. Recent advances in tissue

engineering provide hope that new blood vessel substitutes may one day be fabricated for small diameter

applications, such as CABG, where treatment options are often severely limited.

26.1.1 Current Vascular Grafts

Currently, occluded vessels with diameters <6 mm are bypassed with autologous native blood vessels such

as the saphenous vein as a treatment. Autologous blood vessels were ﬁrst employed in the beginning of the

20th century when Goyanes reported the use of a graft to replace an excised segment of artery by bridging

the defect with the patient’s own popliteal vein [Goyanes, 1906]. Since then, venous grafts have retained

the best long-term patency rate, although limitations such as deterioration when exposed to increased ﬂow

and pressure still remain [Crawford et al., 1981; Szilagyi et al., 1973]. More recently, the internal mammary

artery from the chest has shown to be a superior conduit to the saphenous vein with better long-term

26-1

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26-2 Tissue Engineering

patency and has gained increasing use [Kock et al., 1996]. These native vessels generally display 50 to 70%

graft patency over 10 years (VA study). Moreover, the superior patency of native vessels is attributed to

their compliant natural tissue characteristics and to their antithrombogenic luminal endothelial lining

[Richardson et al., 1980]. Unfortunately, ∼60% of patients who require vascular bypass surgery do not

have a suitable vessel for grafting [Moneta and Porter, 1995], either due to earlier procedures or advanced

peripheral vascular disease. As a result, alternative conduits constructed with synthetic materials have

been investigated.

The use of synthetic vascular prostheses for replacement of natural blood vessels has been explored since

the 1950s [Vorhees et al., 1952]. The most widely employed materials include Dacron™ (polyethylene

terepthalate; PET), Gore-Tex™ (expanded polytetraﬂuoroethylene; ePTFE), and compliant polyurethanes.

A strong, yet ﬂexible polymer, PET may be fabricated in woven, velour, or knitted ﬁber conﬁgurations

[Coury et al., 1996]. ePTFE possesses a smooth surface and exhibits good durability and biocompatibility.

Polyurethane has been increasingly studied because it has been developed to more closely match the

compliance of native vessels than Gore-Tex™ or Dacron, which may reduce some clinical complications

[Giudiceandrea et al., 1998]. These materials are readily available, relatively inexpensive, and have seen

success clinically in applications with vessel diameters greater than 6 mm. However, synthetic grafts with

a small diameter (<6 mm, e.g., coronary artery bypass and femoral-crural bypass) exhibit unsatisfactory

long-term patency due to graft thrombogencity and neointima formation.

Platelet adhesion and aggregation on the graft surface often results in reocclusion of the vessel. This

is caused by the adsorption of proteins that mediate platelet adhesion and coagulation processes on the

relatively hydrophobic polymeric materials. The need for a nonthrombogenic surface has lead to the invest-

igation of endothelial cell (EC) seeding on the lumen of the graft. Although EC seeding improves synthetic

graft patency, EC retention under physiological shear conditions remains problematic [Thompson et al.,

1994; Deutsch et al., 1999]. Furthermore, differences in elasticity between a synthetic material and the

adjacent tissue create discrepancies in strain at the anastomoses, referred to as compliance mismatch,

contributing to intimal hyperplasia [Geary et al., 1993; Greisler et al., 1993]. Moreover, synthetic grafts

have been shown to carry an increased risk of infection [Clowes, 1993], further limiting their clinical

efﬁcacy.

26.1.2 Tissue Engineering

Most tissue engineering strategies attempt to create small caliber vascular grafts by closely mimicking

the structure, function, and physiologic environment of native vessels. Normal arteries possess three

distinct tissue layers (Figure 26.1): the intima, media, and adventitia. The intima consists of an EC

Media

Adventitia

Intima

FIGURE 26.1 The arterial wall is composed of three distinct layers: the intima, media, and adventitia. The intima,

composed of endothelial cells, provides a nonthrombogenic surface. In the medial layer, smooth muscle cells and

elastin ﬁbers align circumferentially and provide mechanical integrity and contractility. The outer adventitia is a

supportive connective tissue.

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Tissue Engineered Vascular Grafts 26-3

monolayer, which prevents platelet aggregation and regulates vessel permeability, vascular smooth muscle

cell behavior, and homeostasis. Within the medial layer, smooth muscle cells (SMCs) and elastin ﬁbers

are aligned circumferentially, contributing the majority of the vessel’s mechanical strength [Wight, 1996].

Dilation and constriction of the vessel are due to medial SMCs’ response to external stimuli. Finally, the

adventitial layer contains ﬁbroblasts, connective tissue, the microvascular supply, and a neural network

that regulates the vasotone of the blood vessel. The recreation of some or all of the vessel layers and their

properties may result in the development of a patent, functional vascular graft. In all likelihood, an intima

and media will be required to achieve any degree of success.

The general concept for creating a tissue-engineered vascular graft (TEVG) usually involves the har-

vest of desired cells, cell expansion in culture, cell seeding onto a scaffold, construct culture in an

environment that induces tissue formation, and implantation of the construct back into the patient.

Many options exist at each step of this process and each must be carefully considered. First, cell source

and culture conditions must be determined. Due to concerns regarding immunogenicity, autologous

cells are likely to be required, but these may be differentiated SMCs and ECs or a variety of types of

vascular progenitors. Expansion in vitro is usually necessary, but in vitro culture time must also be min-

imized to avoid cell dedifferentiation. Genetic modiﬁcation of cells may also be a means to improve

TEVG performance. In addition to the cellular component, some type of material, or scaffolding, is

generally required to provide mechanical support and integrity. Many varieties of scaffolds are used in

tissue engineering, primarily composed of extracellular matrix (ECM) proteins and synthetic polymeric

materials. Once cells are seeded onto scaffolds, another in vitro culture period is usually needed to allow

for new tissue formation and development of appropriate mechanical and functional characteristics.

During the culture period, the construct should receive appropriate chemical and mechanical signals

for cells to synthesize proteins, remodel the tissue, and organize their environment such that the con-

struct will develop into a functional graft with mechanical properties similar to native vessels. Therefore,

a tissue-engineered construct may require several weeks of preparation before it can be implanted into the

patient.

While efforts to create TEVGs remain in an early developmental stage, several problems potentially

exist. Graft patency is still threatened by thrombosis, in all likelihood due to issues with retention of ECs

after implantation or with alterations in EC function after culture in vitro. Also, the possibility of burst

failure after implantation in the physiological ﬂow environment raises concern since the consequences

would be catastrophic. Mechanical properties of TEVGs are generally observed to be lower than those of

native arteries, and thus a number of approaches have begun to address these issues. All of the strategies

discussed above — cell source, genetic modiﬁcation, scaffold materials, and culture conditions — will

likely impact the fabrication of an optimal, clinically useful TEVG.

26.2 Cell Sources for Vascular Tissue Engineering

The development of a functional TEVG is likely to require the construction of an intima and media

composed of ECs and SMCs. Limitations imposed by immunogenicity will probably require the use of

autologous cells, so the majority of studies to date have utilized differentiated SMCs and ECs isolated from

harvested blood vessels. Issues with donor site morbidity and the performance of these cells types in the

engineered tissues have led to the consideration of alternative cell sources. Recent advances in stem cell

biology offer hope for suitable progenitors that can be effectively differentiated into ECs and SMCs for

use in vascular tissue engineering.

Circulating endothelial progenitor cells [Wijelath et al., 2004; Griese et al., 2003; He et al., 2003; Shirota

et al., 2003] and smooth muscle progenitor cells [Simper et al., 2002] can be isolated from blood, offering a

potential source of autologous cells for tissue engineering if they can be appropriately expanded in culture.

Blood-derived endothelial progenitors have already been utilized as linings on synthetic vascular grafts in

several studies [He et al., 2003; Shirota et al., 2003]. Grafts lined with these endothelial progenitors have

been implanted in a canine carotid model, and after more than 30 days, 11 out of 12 grafts remained patent,

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26-4 Tissue Engineering

and cells lining the surface appeared to be ECs [He et al., 2003]. Umbilical cord blood [Murga et al., 2004]

and bone marrow [Hamilton et al., 2004] may be additional sources of autologous vascular progenitor

cells.

26.3 Genetic Modiﬁcation of Vascular Cells

The leading cause of vascular graft failure has been attributed to thrombosis. Genetic engineering of

vascular cells ex vivo may provide an effective strategy to improve graft properties for tissue engineering

applications. Investigators have reported transduction of vascular cells such as ECs [Wilson et al., 1989;

Dunn et al., 1996], ﬁbroblasts [Scharfmann et al., 1991], and SMCs [Lynch et al., 1992]. The seeding of

small-diameter vascular graft constructs with either ECs or SMCs genetically engineered to secrete anti-

thrombotic factors exists as a potential method of improving graft patency rates. As an example, baboon

ECs have been genetically modiﬁed with a retroviral vector encoding antithrombotic factors, namely

tissue plasminogen activator (tPA) and glycosylphosphatidylinositol-anchored urokinase-plasminogen

activator (uPA) [Dichek et al., 1996]. The modiﬁed cells were seeded on the luminal surface of collagen-

coated Dacron surfaces and introduced in arteriovenous shunts in baboons. A signiﬁcant reduction in

platelet and ﬁbrin accumulation was observed in grafts containing modiﬁed ECs expressing either tPA or

uPA. In another approach, the use of a platelet aggregation inhibitor, nitric oxide (NO), has shown to be

promising. Using an ex vivo approach, bovine SMCs liposomally transfected with nitric oxide synthase III

(NOS III) and GTP cyclohydrolase, which produces a cofactor essential for NOS activity, were grown as

monolayers on plastic slides or biomaterials of interest and then placed in a parallel-plate ﬂow chamber.

Whole blood was introduced into the ﬂow chamber to assess platelet adherence to the cell monolayers. The

number of platelets that adhered to the NOS-transduced SMCs was signiﬁcantly lower than those bound

to mock-transduced SMCs and similar to the numbers adhered to cultured ECs. Moreover, the NOS-

expressing SMCs exhibited decreased proliferative activity, which may reduce the incidence of intimal

hyperplasia [Scott-Burden et al., 1996]. These studies demonstrated that an ex vivo genetic engineering

approach may be useful for altering thrombogenicity of TEVG surfaces.

An alternative gene therapy approach to reducing the thrombogenicity of TEVG surfaces would utilize

transfected VEGF. VEGF has shown to be mitogenic for ECs in vitro and stimulates angiogenesis in vivo

[reviewed in Ahrendt et al., 1998]. More importantly, the mitogenic response associated with VEGF is

restricted to ECs, which allows VEGF to be administered without concerns of SMC intimal hyperplasia.

VEGF production in the TEVG may encourage proliferation of ECs seeded on the lumenal surface as well

as stimulate endogenous host EC migration from the anastomoses. Transfecting SMCs with VEGF may

allow for localized and prolonged treatment. Additionally, VEGF-producing SMCs have been found to

promote EC proliferation and migration using in vitro models [Elbjeirami et al., 2004].

Improving the mechanical properties of TEVGs has also been a goal of numerous research efforts, and

genetic modiﬁcation of cells used to seed the TEVG may prove beneﬁcial. The mechanical properties

of a TEVG will be related in large part to the ECM that forms, both its composition and structure.

ECM crosslinking can result from the enzymatic activity of lysyl oxidase (LO) or tissue transglutaminase

[Aeschlimann et al., 1991] and may be a means to improve the mechanical properties of the TEVG. LO,

a copper-dependent amine oxidase, forms lysine-derived crosslinks in connective tissue, particularly in

collagen and elastin [Rucker et al., 1998]. Desmosine is produced in LO-mediated crosslinking of elastin

and is commonly used as a biochemical marker of ECM crosslinking [Venturi et al., 1996]. The LO-

catalyzed crosslinks are present in various connective tissues within the body including bone, cartilage,

skin, and lung and are believed to be a major source of mechanical strength in tissues. Additionally,

the LO-mediated enzymatic reaction renders crosslinked ﬁbers less susceptible to proteolytic degradation

[Vater et al., 1979]. A gene therapy strategy has demonstrated the enhancement of mechanical properties

of tissue-engineered collagen constructs using vascular SMCs transfected with LO [Elbjeirami et al., 2003].

The elastic modulus and ultimate tensile strength of collagen gels seeded with LO-transfected SMCs nearly

doubled as compared to gels seeded with mock-transfected SMCs. These enhanced mechanical properties

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Tissue Engineered Vascular Grafts 26-5

resulted from increased ECM crosslinking rather than increased amounts of ECM, changes in the ECM

composition, or increased cellularity. LO-mediated crosslinking, tissue transglutaminase, or glycation,

may potentially be used in combination with mechanical conditioning, or biochemical factors, such

as TGF-β, that increase the synthesis of ECM proteins to achieve synergistic effects. This strategy may

ultimately enhance mechanical characteristics of TEVGs and minimize in vitro culture times prior to

implantation.

Another effort in the use of genetic modiﬁcation of vascular cells to improve TEVG performance has

focused on regulation of SMC phenotype [Stegemann et al., 2004]. SMCs that were stably transfected with

the gene for cyclic guanosine monophosphate-dependent protein kinase (PKG) were seeded into type I

collagen constructs. PKG is an important regulator of SMC phenotype, and PKG-transfected cells showed

substantially increased expression of smooth muscle α-actin, indicating a reduction in dedifferentiation

during culture. Thus, TEVGs formed with such genetically modiﬁed cells may function more similarly to

native tissues.

26.4 Scaffolds for Vascular Tissue Engineering

26.4.1 Natural Scaffold Materials

26.4.1.1 Collagen

One of the ﬁrst TEVGs developed was based on a natural collagen type I scaffold supported by knitted

Dacron [Weinberg and Bell, 1986]. This study provided early evidence of the feasibility of tissue engineer-

ing a blood vessel substitute by creating layers similar to those of a blood vessel using ECM components

and vascular cells. Cultured bovine ECs, SMCs, and ﬁbroblasts embedded within denatured bovine col-

lagen were used to construct a multilayered vascular graft. Endothelium formation was conﬁrmed by

production of biosynthetic markers such as prostacylcin and von Willebrand’s factor. The burst strength

of the engineered tissue was proportional to the collagen content and cell density. However, the resultant

engineered graft did not achieve sufﬁcient mechanical strength to withstand physiological conditions

unless a supporting Dacron mesh sleeve was added. In fact, maximum burst strengths seen in this model

(≈325 mmHg) were signiﬁcantly lower than that of native coronary artery (≈ 5000 mmHg) or saphenous

vein (≈2000 mmHg).

Glycation, the nonenzymatic crosslinking of ECM proteins by reducing sugars, has been proposed as a

strategy to enhance the stiffness and strength of collagen gel constructs seeded with SMCs [Girton et al.,

1999, 2000]. These constructs were cultured for 10 weeks under high concentrations of glucose or ribose

and then assessed for mechanical properties. The circumferential tensile stiffness of constructs incubated

in medium containing 30 mM ribose showed a 16-fold increase over normally cultured constructs while

tensile strength increased by 4-fold. Although the compliance of the construct compared favorably, the

tensile strength and burst pressure still fell signiﬁcantly below that of arteries.

Expanding upon work with collagen scaffolds, others have further explored the potential utility of

collagen scaffolds by evaluating decellularized collagen matrices from porcine tissue [Hiles et al., 1993;

Huynhet al., 1999]. Small intestine submucosa (SIS) has been isolated and chemically treated to remove the

cellular component, leaving an intact matrix containing mainly collagen. The mechanical properties and

remodeling ability of SIS have been previously investigated for numerous tissue-engineering applications

[Sacks et al., 1999; Gloeckner et al., 2000]. For vascular grafts, SIS possesses acceptable mechanical

properties and exhibits better compliance than currently employed vein grafts [Roeder et al., 1999].

In one effort, Huynh et al. prepared SIS tubes and minimally chemically crosslinked the collagen layers

to provide mechanical strength while maintaining biocompatibility. The inner surface of the grafts were

additionally treated with a heparin complex to inhibit thrombosis and then implanted into rabbits. For

time periods up to 13 weeks, grafts remained patent, and SMCs and ECs were found to have migrated

into the graft from the surrounding tissue.

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26-6 Tissue Engineering

26.4.1.2 Decellularized Vascular Matrix

Xenogenic acellular matrix conduits have also been investigated for the development of TEVG. In this

approach, vessels are treated with reagents such as trypsin and EDTA [Bader et al., 2000] or sodium dodecyl

sulfate [Schaner et al., 2004] to remove the cells, leaving an intact matrix that can be used as a scaffold.

Decellularized porcine arteries have been seeded with human ECs and SMCs isolated from the saphenous

vein [Teebken et al., 2000]. After exposure to pulsatile ﬂow, the cell-seeded porcine matrix developed

an endothelial monolayer. This study demonstrated the feasibility of generating vascular grafts in vitro

from acellular, animal-derived, vascular matrices and human cells. Attempts to improve biocompatibility

of xenograft matrices have included proteolysis of bovine or porcine carotid arteries [Moazami et al.,

1998; Teebken et al., 2000]. This decellularized form provided an ECM with properly arranged collagen

and elastin ﬁbers that had minimal immune or inﬂammatory response [Teebken et al., 2000]. However,

ﬁbrosis eventually rendered these grafts unsuccessful.

26.4.1.3 Cell Sheets

L’Heureux and his collegues have applied a unique approach in the fabrication of a TEVG using biological

materials, essentially using cells to generate the scaffold for the tissue-engineering construct [L’Heureux

et al., 1998]. Umbilical vein SMCs and human skin ﬁbroblasts were cultured for 5 weeks to form sheets

containing superconﬂuent cells and cell-synthesized ECM. The cell sheets were sequentially rolled around

a mandrel to form ﬁrst the medial layer, then the adventitial layer. For 8 weeks, the construct was cultured

in a bioreactor designed to provide perfusion of the culture medium and mechanical support. Following

culture, the mandrel was removed, and the lumen was seeded with ECs, thus forming the three layers

representative of a native artery. The burst strength of these grafts well exceeded that of native veins.

Based on histological analysis, it appeared that SMCs were aligned circumferentially and produced a

signiﬁcant amount of new ECM. When implanted in vivo, these vascular substitutes had a 50% patency

rate at one week after implantation. The construct consists entirely of human cells and human ECM

proteins while exhibiting impressive mechanical strength, which makes this strategy attractive. However,

much of the constructs’ strength was attributed to the adventitial layer instead of the medial layer, as is

normally observed in blood vessels. Some thrombus formation was also observed at the graft site following

implantation. Furthermore, a 12-week time span was required to prepare a construct.

26.4.2 Synthetic Polymer Scaffolds

While the natural scaffold materials discussed above have achieved some success, concerns with disease

transmission, difﬁculties with material processing, and often poor mechanical properties have led some

groups to concentrate on the development of synthetic biomaterials as scaffolds. Bioabsorbable synthetic

polymers may be designed to provide a transitional environment by providing a supporting structure

to developing tissue. The degradation rate of these constructs may often be tailored to match the rate

at which new tissue is formed, so that space is created for cell growth and matrix deposition. These

materials serve as guides for tissue regeneration in three dimensions and offer the possibility to control

structural variables and scaffold properties, such as the molecular structure, molecular weight, degradation

properties, porosity, and mechanical properties [Ma et al., 1995].

26.4.2.1 Polyester Scaffolds

Polyglycolic acid (PGA), a biodegradable polyester, has demonstrated relatively good biocompatibility and

has been extensively studied for numerous tissue-engineering applications [Freed et al., 1993; Kim et al.,

1998; Mooney et al., 1996]. These well-characterized materials have been the ﬁrst to illustrate the feasibility

of polymeric scaffolds in vascular tissue engineering [Niklason et al., 1999]. TEVGs were fabricated using

PGA mesh scaffolds, seeded with bovine aortic SMCs, then cultured in pulsatile ﬂow bioreactors. After

8 weeks, these cultured vessels exhibited increased concentrations of collagen and enhanced mechanical

properties. Burst pressure of these grafts compared well to that of typical vein grafts (human saphenous

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Tissue Engineered Vascular Grafts 26-7

vein = 1680 ± 307 mmHg vs. PGA graft = 2150 ± 700 mmHg). When these grafts were seeded with

autologous ECs on their luminal surface, continuously perfused for 3 days, and then implanted into pigs,

they remained patent for up to 4 weeks.

26.4.2.2 Hydrogel Scaffolds

Derivatives of polyethylene glycol (PEG) currently are being studied as hydrogel scaffolds for blood

vessel substitutes. These materials are hydrophilic, biocompatible, and intrinsically resistant to protein

adsorption and cell adhesion [Merrill et al., 1983; Gombotz et al., 1991]. Thus, PEG essentially provides

a “blank slate,” devoid of biological interactions, upon which the desired biofunctionality can be built.

Aqueous solutions of acrylated PEG can be rapidly photopolymerized in direct contact with cells and

tissues [Sawhney et al., 1994; Hill-West et al., 1994]. Furthermore, PEG-based materials can be rendered

bioactive by inclusion of proteolytically degradable peptides into the polymer backbone [West et al.,

1999] and by grafting adhesion peptides [Hern et al., 1998] or growth factors [Mann et al., 2001a] into the

hydrogel network during the photopolymerization process. Recently, PEG hydrogels that largely mimic

the properties of collagen have been developed. PEG hydrogels grafted with a synthetic adhesive peptide

RGDS and the collagenase-sensitive peptide sequence GGGLGPAGGK permitted cell migration [Gobin

and West 2002; Mann et al., 2001b]. Following 7 days of incubation, approximately 70% as many cells

migrated through collagenase-sensitive hydrogels as through collagen gels, with no statistically signiﬁcant

difference between the two groups. The elastin-derived peptide VAPG has shown to be speciﬁc for SMC

adhesion, and PEG hydrogels modiﬁed with this adhesive peptide rather than RGDS supported adhesion

and growth of vascular SMCs but not ﬁbroblasts or platelets [Gobin et al., 2003]. Moreover, bioactive

molecules like TGF-β may be covalently incorporated into scaffolds to induce protein synthesis by vascular

SMCs. TGF-β has been reported to stimulate expression of several matrix components, including elastin,

collagen, ﬁbronectin, and proteoglycans [Amento et al., 1991; Lawrence et al., 1994; Tajima, 1996]. TGF-β

covalently immobilized to PEG-based hydrogels signiﬁcantly increased collagen production of vascular

SMCs seeded within these scaffold materials [Mann et al., 2001]. Mechanical testing of these engineered

tissues also determined that the elastic modulus was higher in TGF-β-tethered PEG scaffolds than PEG

scaffolds without TGF-β, indicating that material properties for TEVGs may be improved using this

technology. A cell-seeded graft formed from this biomimetic hydrogel scaffold is shown in Figure 26.2.

These types of bioactive materials may allow one to capture the advantages of a natural scaffold, such as

speciﬁc cell–material interactions and proteolytic remodeling in response to tissue formation, while also

having the beneﬁts of a synthetic material, namely the ease of processing and the ability to manipulate

mechanical properties.

FIGURE 26.2 A PEG-based scaffold seeded with smooth muscle cells and endothelial cells ready for insertion into a

bioreactor for in vitro culture of a TEVG (left, standing upright; right, laying on side). The cell-seeded scaffold is formed

via photopolymerization, so the dimensions are easily tailored for a given application and cells are homogeneously

seeded throughout the material. The scaffold is designed to degrade in response to cellular proteolytic activity during

tissue formation.