Vascular Graft Surface Modifications –
Endothelialization of Vessel Implants
Julie Y. Ji
BE553
Tissue Engineering
Spring, 2001
University of Pennsylvania
Table of Contents
1. Introduction ................................................................................................... 1
Vessel Biology ........................................................................................ 1
In Vivo Hemodynamic Environment ..................................................... 1
Motivation............................................................................................... 2
2. Surface Modifications ................................................................................... 4
Biomolecular Based Surface Coatings ................................................... 4
Chemical Based Surface Coatings.......................................................... 6
3. Shear Stress ................................................................................................... 7
4. Surface Materials .......................................................................................... 9
5. General Considerations ................................................................................. 10
6. Tissue Engineering – More Novel Approaches ............................................ 10
7. Conclusions ................................................................................................... 11
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1. Introduction
Vessel Biology
Blood vessels serve to distribute and transport blood to and from every organ and tissue in the
body. The two main branched networks of blood vessels are the arteries and veins. Blood is carried away
from the heart by arteries to the lungs to be oxygenated or to the various tissues in capillaries, and returns
to the heart through veins. Depending their function and location, blood vessels can vary greatly in size,
structural and mechanical properties, and biological and biochemical contents. The basic structure of the
vessel is a tube of three layers: intima, media, and adventitia. The structure is most distinct in larger
vessels. The intima is the layer closest to the lumen where blood flows. It is made up mainly of a layer of
endothelial cells (EC) attached to a basement membrane and matrix molecules. The next layer outside is
the internal elastic lamina, a band of elastic fibers that separates intima and media. The media contains
mainly smooth muscle cells plus some fibers of elastic tissue. The external elastic lamina then separates
the media from the adventitia, which is made of a collagenous extracellular matrix (ECM) with
fibroblasts, blood vessels, and nerves. The adventitia serves as structural support to the vessel and gives
it shape and rigidity.
The ECM outside vascular cells provides the mechanical properties of the vessel through its
major molecular components. Collagen, type I, III, and IV, gives the vessel its tensile stiffness. Elastin
fibers provide elastic properties, and proteoglycans (versican, decorin and biglycan, lumican and perlican)
provide compressibility. Together they contribute to the viscoelastic nature of blood vessels, and allow
them to function properly.
In Vivo Hemodynamic Environment
The physiological environment that vascular cells experience also includes mechanical stresses
imposed by the hemodynamics of blood flow through vessels. This relationship is most important for
vascular endothelium cells because they are actively involved in the biology of blood vessels by
modulating and participating in the release of biochemical regulators. The endothelium also serves as a
signal transduction interface between blood flow and the vessel wall.
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Blood flow through a cylindrical vessel imposes hemodynamic stress on the endothelium. The
stress has two components: a normal component, pressure, and a tangential component, shear stress. If
assuming blood to be Newtonian, shear stress (w) is directly related to wall shear rate Sw, or velocity
gradient at the wall, through blood viscosity ():
 u 
 w   μS w   
 y  w
(1)
Flow in blood vessels is mostly laminar, but due to the complex geometry of the vascular system,
shear stress is both time varying and spatially varying. The average mean shear stress in a straight section
of an artery is in the range of 10 to 20 dynes/cm2, with extreme variations existing at branching points
where shear stress be as high as 50 dynes/cm2 or over 100 dynes/cm2 with pulsatile flow [10].
For a Newtonian fluid in cylindrical tube under laminar flow, shear stress at the vessel wall is
calculated as follows:
w 
4Q
r 3
(1)
where  is viscosity (dynes/s-cm2), Q is the flow rate (cm3/s), and r is radius in cm.
Motivation
The pathology affecting small and medium-sized blood vessel is the primary cause of death in
American and the Western societies. Atherosclerosis is the major disease of the blood vessel, and
atherosclerotic lesions are raised focal plaque within the intima region consisting of a lipid core
surrounded by smooth muscle cells and ECM, and covered by a fibrous cap. As it increases in size, it
restricts blood flow, and once its surface ruptures, blood clot forms and the vessel is blocked. In cardiac
and peripheral bypass surgery, atherosclerotic regions of vessels are replaced by autologous veins or
arteries obtained from another part of the body, with the anticipated side effect of donor-site morbidity.
Some patients, therefore, do not have appropriate vessel replacement due to prior removal, disease, or
inadequate size. This has prompted the need for vascular grafts as replacements. Effective vascular
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grafts, however, must maintain the biological structure of native vessel while still maintaining its
mechanical properties.
The first successful application of synthetic vascular graft in the arterial wall, pioneered by
Voorhees et al in 1952, was made of Vinyon N [20]. At the time, the blood vessel was considered as
simply a conduit for blood flow. Since then, however, the extraordinary expansion of the general
knowledge of blood vessel biology has led to much more sophisticated approaches to vessel replacement.
Material for vascular grafts, for example, has evolved from being biologically inert to biocompatible, or
conductive toward functional interactions with surrounding tissues.
In larger blood vessels, with diameter greater than 6 mm, synthetic vascular grafts have worked
reasonably well. Clinical application of small diameter, less than 6 mm, synthetic prostheses continue to
meet with difficulties: thrombus formation in the short term and intimal hyperplasia in the long term.
Given the anticoagulant and vasodilatory properties of ECs, these problems may be overcome partially if
grafts are seeded with an adherent monolayer of ECs prior to implantation.
The concept of seeding ECs on the lumen side vascular grafts was first introduced in 1979 by
Herring and colleagues [6]. Advantages of using EC-seeded vascular grafts include presence of
nonthrombogenic surface at time of implantation that would not only prevent immediate clotting, but also
act on distal surrounding tissues to prevent constriction proliferation of smooth muscle cells. Practical
success of this method has been limited by autologous EC harvesting and viability and poor cell retention
after exposure to flow conditions.
Efforts to overcome these difficulties have concentrated mainly in two different approaches.
First, the biocompatibility approach aims at improving graft material and the efficiency of luminal
seeding of an endothelial interface layer to increase graft patency. This includes use of coatings of
biocompatible substances to improve EC attachment. The second, more novel tissue-based approach
attempts to recreate vessel constructs based on co-culturing of endothelial and smooth muscle cells [23].
However, tissue engineering of multi-layered cell-based vessel is still new in its development, and its
practicality in clinical applications is limited by time and expense. This paper offers an overview of the
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different approaches taken to modify and enhance vascular graft surfaces for biocompatibility and
improved cell-substrate interactions, with the immediate goal being the proper adhesion of endothelial
cells on lumen surface of tubular grafts and the formation of a vasoactive endothelium monolayer.
2. Surface Modifications
A major problem with endothelial seeding procedure is that most of the currently used polymerbased graft materials often are resistant to cell attachment. Surface modifications that have been used to
enhance EC attachment on vascular prosthesis can be categorized under two different approaches:
1. Realization of a coating layer made of biological molecules: immobilization of fibronectin,
laminin, collagen, and peptides
2. Chemical modifications of the material surfaces: synthetic modification of polymer side chains,
for example, the use of plasma treatments to modify the surface of vascular prostheses without
changing graft dynamics and physical properties
Biomolecular Based Surface Coatings
Clinically, expanded poly(tetrafluoro ethylene) (ePTFE) is still the most common construction
material for small-bore vascular grafts. The material is stable, non-adherent, and chemically inert. The
unique through-pore microporous wall structure and highly flexible mechanical properties of ePTFE
grafts hold promise for enhanced healing and compliance as an implantable graft. However, the highly
hydrophobic surfaces of ePTFE limit endothelia surface adhesion. They can be coated with a protein
substrate to render the surface more favorable for endothelia cell attachment. Possible seeding substrates
that have been tested include whole blood and ECM proteins such as fibronectin, laminin, and collagen.
Following in vitro study of EC attachment on ePTFE material, attachment is found to be poor on
untreated surface, but can be greatly improved by preclotting surfaces with blood or pre-coating with
either fibronectin, laminin, and type IV collagen [18]. Preclotted ePTFE is obtained by incubating the
graft in whole blood for 2 to 3 hr for clotting to occur. Graft segment is removed and blood clot is broken
up by vigorous shaking. The graft is then washed in phosphate buffer saline till macroscopically free of
red clots. This procedure has shown to encourage cell attachment to ePTFE, resulting in virtually
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confluent endothelial monolayer after just 1-hour incubation in vitro. Preclotting grafts tends to yield a
more consistent and even spread of fibrin across the surface than protein coating. Though cell adhesion
and spreading on protein-coated surfaces might occur at a faster rate, gaps exist in areas of confluent ECs
due to patchy surface protein binding. In the end, preclotted surfaces give a consistent and complete layer
of endothelial covering.
Synthetic cell-binding peptides based on receptor-binding domains of cell adhesion protein can be
covalently grafted on to material surfaces to provide substrates for which cells have adhesion receptors
directly. Examples include the cell-binding peptide of many cell adhesion proteins, Arg-Gly-Asp (RGD)
and Tyr-Ile-Gly-Ser-Arg (YIGSR), the cell-binding peptide of laminin. Covalently grafting PTFE
surfaces with RGD and YIGSR containing peptides improves the adhesion of Human EC [8].
Advantages of using peptide grafting are: the process can be easily controlled and the resulting surfaces
are temporally stable. Protein adsorption on the other hand may be difficult to control and proteins are
temporally dynamic.
Since the discovery that novel adhesion receptors are present on the human EC for the peptide
sequence Arg-Glu-Asp-Cal (REDV) present in the III-CS domain of human plasma fibronectin, substrates
containing synthetic REDV peptides have been tested for selective binding of ECs. Thought they are
otherwise cell nonadhesive substrates, endothelial cells are able to bind and spread in the presence of
REDV peptides while fibroblasts, vascular smooth muscle cells, and platelet fail to attach. The resulting
endothelial monolayer on REDV is also nonthrombogenic, while on surfaces with other receptor-binding
peptides, ECs attach and spread with less specificity and demonstrates more thrombogenicity [7].
In a study that further evaluated EC attachment on PTFE grafts with known matrix components:
collagen type I and/or III, laminin, fibronectin, gelatin, and RGD-containing peptides, it was found,
however, that laminin and gelatin-coated surfaces yielded the most improved EC attachment [16]. In a
similar study, ECs are seeded on ePTFE grafts coated with RGD-containing peptides are also exposed to
physiological shear stress in an artificial flow circuit and cell attachment and retention are both assessed.
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Grafts coated with RGD synthetic peptides result in the highest attachment and retention rate, compared
to even fibronectin coated grafts [21].
Naturally produced ECM has also been shown to be an excellent substrate for EC attachment,
proliferation, and resistance to shear stress on PTFE grafts [17]. To coat PTFE grafts with ECM, ECs are
first seeded on PTFE grafts precoated with fibronectin, and allowed to grow for 12 days after cells have
reached confluency. The attached ECs are then lysed, without disrupting the graft itself, leaving behind
the underlying ECM intact and attached to the surface. After the remaining nuclei and cytoskeletal
elements are removed, these ECM-precoated grafts are used again for EC seeding. The strength of
subsequent EC attachment is evaluated under constant physiological shear of 15 dynes/cm2 for 1 hr.
Vascular grafts precoated with fibronectin and a naturally produced ECM yielded significantly increased
adhesion strength, indicated by high retention rate of ECs, as compared to cultures on surfaces coated
with fibronectin alone. In further in vivo studies, PTFE or canine EC seeded on ECM precoated grafts
were implanted in the common carotid arteries of dogs. ECM coated PTFE implants seeded with
autologous ECs revealed significant increase in EC overage and reduced thrombus formation. The ECM
coating covers the graft as a uniform layer and contains attachment proteins and growth promoting factors
responsible for better EC attachment and proliferation than other substances such as fibronectin, matrigel,
or collagen. Thus, the presence of adhesive macromolecules and potent EC growth promoting factors
found on naturally formed ECM might render ECM as a promising substrate for vascular prostheses.
Such ECM grafts would be highly stable and can be easily stored and distributed.
Chemical Based Surface Modifications
Plasma treatment provides a means of altering surface characteristics of vascular grafts without
changing its physical properties. Low-pressure plasma techniques are becoming popular as a versatile
tool for modifying surface chemistries that improve the wettability, adhesion, and biocompatibility of
polymer surfaces by introducing functional groups or depositing thin films of defined composition on
them. Plasma treatment is often accomplished using the radio frequency glow discharger (RFGD).
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Nitrogen-rich plasma provides amide groups that are considered the main promoter for surface
cell adhesion. Immobilization of nitrogen-containing functional groups on ePTFE vascular grafts is done
using amide (-CONH2) and amine (-CH2NH2) plasma (butylamine) treatments. ECs are seeded on treated
surfaces, and placed in an artificial circulatory system for 5 days to test for the impact of constant or
pulsatile flow in the presence of a normal force against the vessel wall (cyclic stretch). Plasma coating of
amide and functional groups enhances EC adhesion to vascular grafts under both constant and pulsatile
flow conditions, and the formation of an EC monolayer on plasma-coated surface is observed [19].
Polyethyleneterephthalate (PET) surfaces are also plasma-treated with oxygen and ammonia in
the presence of a gas mixture (NH3, NH3/H2, O2/H2, and O2/H2O) to verify the effect of functional groups
grafting onto the endothelial cell growth. There is no toxic effect on ECs by the treated PET samples, and
the presence of well spread and flattened ECs, as well as increase in EC growth with incubation time can
be observed [13]. The RFGD treatment with H2 and NH3 involves the increase of amino groups with
respect to other nitrogen moieties. Amino groups are thought to become positively charged at
physiological pH and enhance graft surface interactions with cells and biomolecules carrying negative
charges. Further studies using plasma-treatment of graft surfaces are needed still, but this procedure does
hold potential in generating biocompatible vascular prostheses.
3. Shear Stress
ECs in vivo are highly adherent and can resist disruption by hemodynamics shear stress at levels
that far exceed physiological conditions. They are also highly differentiated in vivo, with organized
cytoskeleton, Weibel-Palade bodies, and basal stress fibers as well as focal adhesion points. In culture,
however, ECs rapidly lose many of their differentiated features and, on artificial graft surfaces, do not
sufficiently differentiate and adhere to resist physiological shear stress. Exposing EC to chronic shear
stress in vitro, applied in a stepwise fashion over several days, can induce cells to adhere more tightly and
exhibit more differentiated features [3]. Preconditioning of endothelia cells seeded on vascular grafts
with stepwise shear stress in vitro can be used, therefore, to improve cell retention and differentiation.
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It is also possible to generate vascular grafts with tightly adherent endothelia cell monolayer by
preconditioning cells under in vitro shear stress. To study the possible beneficial effects of shear stress on
EC attachment, polyurethane vascular grafts are first seeded with ECs with the aid of surface molecules
such as gelatin and RGD peptide coatings. The ECs are then cultured in vitro for 6 days with or without
continuous laminar shear stress: 1 to 2 dynes/cm2 for the first 3 days, followed by another 3 days of
approximately 25 dynes/cm2 shear. Shear stress is delivered in pulsatile flow (50-250 pulsation/min)
though it is calculate to represent a time-average value. A sudden short period of shearing (sustained 25
sec pulse) is applied to determine cell-substrate adhesion strength. Preconditioning under long-term shear
stress markedly enhances EC adhesion and retention on vascular grafts. As a consequent, these grafts,
with sustained endothelial monolayer, are less thrombogenic [12].
This discovery was followed by further in vivo studies in rats. Grafts that were seeded and
preconditioned under shear stress as follows: 1 dynes/cm2 for the 3 days, followed by 3 days of either 1 or
25 dynes/cm2 shear. Confluent monolayers of ECs are observed lining the lumen of the grafts. The grafts
are then implanted as aortic interposition grafts into syngeneic rats. After just 24 hours in vivo,
pretreatment with 25 dynes/cm2, but not with 0 or 1 dynes/cm2 shear stress, resulted in the retention of
fully confluent endothelial monolayers on graft surface. In shear stress conditioned grafts, immediate
graft thrombosis is inhibited. One week after implantation, however, macrophage infiltration is observed
underneath the luminal cell monolayer, and after 3 months, sub-endothelial intimal thickening is
observed. However, vessel thickening in grafts pretreated under 25 dynes/cm2 is less severe than grafts
pretreated under 1 dynes/cm2 shear, and both are significantly reduced compared to the control case (no
endothelial cells seeded) [4]. Therefore, shear stress pretreatment is able to enhance the retention of
endothelial monolayer, reduction of immediate thrombogenicity, and delayed neointimal thickness. This
procedure, possibly in combination with other methods that may enhance the endothelialization of
prosthetic vascular grafts, has significant potential in clinical applications to reduce the incidence rate of
graft thrombosis.
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4. Surface Materials
Perhaps the most significant disadvantage of ePTFE is still its chemical inertness, which makes
surface modifications difficult. Another practical choice for graft material is the new stress free poly
(carbonate-urea)urethane (CPU) graft with compliance similar to that of human artery. Seeding cell
density and incubation time as been optimized to allow complete endothelialization and high cellsubstrate strength that resists hydrodynamic stresses [15].
Human ECs seeded on compliant CPU and PTFE grafts are exposed to varying shear stress of up
to 14 dynes/cm2 using a pulsatile flow model for up to 6 hours. Seeding efficiency is significantly better
on CPU substrates, and CPU showed better EC attachment rate after 6 hr exposure to shear stress [5].
However, no in vivo studies have yet been done using this material.
Further tests are underway to evaluate the use of polyurethane vascular grafts, constructed from
poly(carbonate) polyurethane. These grafts are elastic, microporous tubes with self-healing properties. It
has been shown that the drug dipyridamole can be coupled to polyurethane surfaces. Dipyridamole is a
potent nontoxic inhibitor of platelet activation/aggregation, and also a strong inhibitor of vascular smooth
muscle cell proliferation. Surfaces modified with dipyridamole are found to have lower thrombogenicity
and adherence of blood platelets in vitro [1]. Seeding ECs on these modified surfaces resulted in
increased EC proliferation and spreading as compared to uncoated surfaces. Subsequent in vivo studies
were carried out in goats, as a bypass of the carotid artery (12 cm length), and in sheep, as an interposition
graft in the carotid artery (4 cm length) [2]. While dipyridamole coating was shown to have three main
beneficial effects in vitro: reduced thrombogenicity, reduced adherence of blood platelet, and
accommodation of a confluent monolayer of endothelial cells, the in vivo experiments were inconclusive.
There is low coated-graft patency in goat models (high incidence of occlusion), and significant
deterioration of the polyurethane material was observed in the sheep experiments, indicating lack of
structural stability of the grafts. Further studies are needed regarding the use of this material as
permanent vascular grafts.
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5. General Considerations
There is a significant loss of ECs from the surface of seeded prosthetic grafts after implantation.
Variables that may influence cell retention following exposure to flow include cell density at time of
seeding and post-seeding incubation time. In general, delayed exposure to shear stress, extended postseeding incubation time, is shown to improve cell retention, and maximal retention is usually achieved at
confluency, regardless of initial seeding density [9].
6. Tissue Engineering – More Novel Approaches
The demand for healthy, functional, vessel replacement has prompted research beyond polymerbased vascular grafts. Collagen-based vascular grafts are now been studied for their potentials. Collagen
gels can be fabricated into tubular structure to provide the basic conduit on which smooth muscle cells
can be seeded. Endothelial cells are then used to line the inner lumen [22]. This method demonstrates the
used of a naturally derived material to provide mechanical properties. A vascular graft formed without
scaffold support has also been developed. Sheets of smooth muscle cells and fibroblasts can be grown in
culture and rolled into tubular shape. The inner surface is then seeded with endothelial cells, forming a
three-layered cellular vascular architecture that can also withstand physiological pressure loadings.
However, in vivo studies with these grafts so far show only 50% patency at 1 week.
Smooth muscle cells can also be cultured onto porous and degradable poly-glycolid tubular
structure and cultured in bioreactors that can impart pulsatile radial distensions that mimic in vivo
conditions [11]. The bioresorbable scaffold is eventually replaced by smooth muscle cells, and
endothelial cells are then seeded on the inside surface to yield a model of the native arteries with
comparable mechanical strength.
To ensure that blood vessels have the required mechanical properties immediately upon
implantation, non-degradable scaffold has been also be used. The scaffolds are selected to optimize cell
attachment and ECM deposition while maintaining the elastic properties for in vivo functions.
Polyurethane scaffolds can be constructed with the appropriate porosity, pore size, and thickness to
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provide biocompatibility, tensile stiffness and elasticity similar to native vessels. Smooth muscle cells are
seeded on the scaffold and cultured under fluid flow followed by endothelial cell seeding on top [14].
Following culture under fluid flow, the endothelial cells generated a confluent monolayer. In vivo studies
in dogs revealed graft patency and endothelial lining retained for up to 4 weeks. The use of a
biocompatible scaffold and a dynamic culturing environment that provides mechanical stresses are
important in the growth and development of such tissue constructs.
7. Conclusions
The need for a healthy, physiologically compatible blood vessel replacement graft in the
application of bypass surgery is significant and represents a large commercial opportunity and substantial
clinical impact. The more practical approach currently available involves the implantation of polymerbased tubular vascular grafts, which are working fairly successfully for large-diameter vessels, but fail in
small-diameter transplants due to thrombosis and hyperplasia. Attempts to overcome this barrier focus
primarily on the seeding and formation of an endothelial cell monolayer on the inside surface. The
presence of a vasoactive endothelium enhances endothelial cell adhesion, spreading, and proliferation
while reducing platelet recruitment and vessel thrombogenicity. To accomplish this, polymer-based
grafts are modified with bio-molecular as well as chemical coatings that would aid in promoting the
appropriate cell-substrate interactions. The strength of EC attachment on graft surfaces must also be
tested to withstand under physiological flow conditions and fluid dynamics forces.
Advantages to seeding vascular grafts with endothelial cells include the immediate presence of a
vasoactive endothelium that can prevent immediate clotting and also functions as mediator in vasodilation
and smooth muscle cell growth inhibition. This procedure also poses implications in gene therapy
applications. Given the ability to plant ECs directly on grafts, it is possible to seed genetically modified
ECs on grafts that would come in direct contact with blood flow. The relatively simple structure of an
endothelial lining on the lumen of a tubular vascular graft may also ease the process of in situ, endothelial
cell-specific gene modifications.
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Finally, novel tissue engineering approaches used to regenerate blood vessels on a cellular basis
focuses mainly on 3-dimensional co-culturing of cell types native to blood vessel in an attempt to
reconstruct the biological structure of the vessel wall. These measures also hold great promises and
together, the construction of the optimal biocompatible vascular graft will eventually be made feasible
and practical for clinical use.
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1. Aldenhoff YB, Blezer R, Lindhout T, and Koole LH, Photo-immobilization of dipyridamole
(Persantin) at the surface of polyurethane biomaterials: reduction of in-vitro thrombogenicity,
Biomaterials, 18(2):167-72 (1997).
2. Aldenhoff YBJ, van der Veen FH, Woorst J, Habets J, Poole-Warren LA, Koole LH, Performance of
a polyurethane vascular prosthesis carrying a dipyridamole (Persantin) coating on its lumenal surface,
J Biomed Mater Res, 54(2):224-33 (2000).
3. Ballermann BJ and Ott MJ, Adhesion and differentiation of endothelial cells by exposure to chronic
shear stress: a vascular graft model, Blood Purif, 13(3-4):125-34 (1995).
4. Dardik A, Liu A, and Ballermann MD, Chronic in vitro shear stress stimulates endothelial cell
retention on prosthetic vascular grafts and reduces subsequent in vivo neointimal thickness, J Vasc
Surg, 29(1):157-167 (1999).
5. Giudiceandrea A, Seifalian AM, Krijgsman B, and Hamilton G, Effect of prolonge pulsatile shear
stress in vitro on endothelial cell seeded PTFE and compliant polyurethane vascular grafts, Eur J
Vasc Endovasc Surg, 15(2):147-54 (1998).
6. Herring M., Dilley R., Jerslid R., Seeding arterial prosthesis with vascular endothelium, The nature of
the libing, Ann Surg, 1979, 190:84-90.
7. Hubbell JA, Massia SP, Desai NP, Drumheller PD, Endothelial cell-selective materials for tissue
engineering in the vascular graft via a new receptor, Biotechnology, 9:568-572 (1991).
8. Massia SP and Hubbell JA, Human endothelial cell interactions with surface-coupled adhesion
peptides on a nonadhesive glass substrate and two polymeric biomaterials, J Biomed Mater Res,
25:223-242 (1991).
9. Miyata T, Conte MS, Trodel LA, Mason D, Whittemore AD, and Birinyi LK, Delayed exposure to
pulsatile shear stress improves retention of human saphenous vein endothelial cells on seeded ePTFE
grafts, J Sur Res, 50:485-493 (1991).
10. Nerem RM, Alexander RW, Chappell DC, Medford RM, Varner SE, and Taylor WR, The study of
the influence of flow on vascular endothelial biology, Am J Med Sci, 316(3):169-175 (1998).
11. Nicklason LE, Gao J, Abbott WM, Hirschi KK, Houser S, Marini R, and Langer R, Functional
arteries grown in vitro, Science, 284: 489-493 (1999).
12. Ott MJ, and Ballermann BJ, Shear stress-conditioned, endothelial cell-seeded vascular grafts:
improved cell adherence in response to in vitro shear stress, Surgery, 117(3):334-339 (1995).
13. Ramires PA, Mirenghi L, Romano AR, Palumbo F, and Nicolardi G, Plasma-treated PET surfaces
improve the biocompatibility of human endothelial cells, J Biomed Mater Res, 51(3):535-9 (2000).
14. Ratcliffe A, Tissue engineering of vascular grafts, Matrix Bio, 19: 353-357 (2000).
15. Salacinski HJ, Tai NR, Punshon G, Giudiceandrea A, Hamilton G, Seifalian AM, Optimal
endothelialization of a new compliant poly(carbonate-urea) urethane vascular graft with effect of
physiological shear stress, Eur J Vasc Endovasc Surg, 20(4):342-52 (2000).
14
16. Sank A, Rostami K, Weaver F, Ertl D, Yellin A, Nimni M, and Tuan T, New evidence and new hope
concerning endothelial seeding of vascular grafts, Am J Surg, 164:199-204 (1992).
17. Schneider A, Chandra M, Lazarovici G, Vlodavsky I, Merin G, Uretzky G, Borman JB, and Schalb
H, Naturally produced extracellular matrix is an excellent sustrate for canine endothelial cell
proliferation and resistance to shear stress on PTFE vascular graft, Thromb Haemost, 78:1392-8
(1997).
18. Thomson GJL, Vohra RK, Carr MH, and Walker MG, Adult human endothelial cell seeding using
expanded polytetrafluoroethylene vascular grafts: a comparison of four substrate, Surgery, 109(1):2027 (1990).
19. Tseng DY and Edelman ER, Effects of amide and amine plasma-treated ePTFE vascular grafts on
endothelial cell lining in an artificial circulatory system, J Biomed Mater Res, 42(2):188-98 (1998).
20. Voorhees AB, Jaretzki A, Blakemore AH. The use of tubes constructed from Vinyon N cloth in
bridging arterial defects. Ann Surg 1952, 135:332.
21. Walluscheck KP, Steinhoff G, Kelm S, Haverich A, Improved endothelial cell attachment on ePTFE
vascular grafts pretreated with synthetic RGD-containing peptides, Eur J Vasc Endovasc Surg,
12(3):321-30 (1996).
22. Weinberg CB and Bell E, A blood vessel model constructed from collage and cultured vascular cells,
Science, 231:397-400 (1986).
23. Ziegler T and Nerem RM, Tissue engineering a blood vessel: regulation of vascular biology by
mechanical stresses, J Cell Biochem, 56:204-209 (1994).