A literature review by nikeborome

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                 Small-diameter vascular tissue engineering

                              A literature review



                                     R. van Lith

                               November 2004

                                     BMTE04.57




Part I of MSc-thesis

Supervisors:
Prof. Dr. L.H.E.H.Snoeckx
Dr. Ir. M.C.M. Rutten
Ir. M. Stekelenburg

Eindhoven University of Technology
Faculty of Biomedical Engineering



                                                              1
Table of contents
Introduction…..................................................................................................................... 4
Chapter 1: Nature of vascular disease................................................................................. 5
   1.1 Atherosclerosis.......................................................................................................... 5
   1.2 Endothelium dysfunction .......................................................................................... 6
   1.3 Inflammatory smooth muscle cell infiltration and proliferation............................... 6
   1.4 Role of shear stress ................................................................................................... 7
   2.1 Cell layers ................................................................................................................. 8
      2.1.1 Tunica adventitia................................................................................................ 8
      2.1.2 Tunica media...................................................................................................... 9
      2.1.3 Tunica intima ..................................................................................................... 9
   2.2 Biochemical properties ............................................................................................. 9
   2.3 Mechanical properties............................................................................................. 11
      2.3.1 Collagen ........................................................................................................... 12
      2.3.2 Elastin .............................................................................................................. 13
      2.3.3 Coronary artery specific properties.................................................................. 13
Chapter 3: Tissue engineered blood vessels ..................................................................... 15
   3.1 Requirements of tissue-engineered vessels............................................................. 15
      3.1.1 Mechanical Properties...................................................................................... 16
      3.1.2 Vasoactivity ..................................................................................................... 18
      3.1.3 Non-thrombogeneicity ..................................................................................... 18
   3.2 Approaches ............................................................................................................. 19
      3.2.1 A-cellular constructs ........................................................................................ 19
         Non-biodegradable synthetic constructs ............................................................... 19
         Biodegradable synthetic constructs....................................................................... 20
         Biological constructs............................................................................................. 21
         Concluding remarks .............................................................................................. 22
      3.2.2 Cellular constructs ........................................................................................... 22
         Non-biodegradable synthetic constructs ............................................................... 22
         Biodegradable synthetic constructs....................................................................... 23
         Biological constructs............................................................................................. 25
         Concluding remarks .............................................................................................. 26
      3.2.3 Other techniques .............................................................................................. 26
         Self-assembly method........................................................................................... 26
         Body cavities as bioreactors.................................................................................. 27
         Synthetic protein based polymers ......................................................................... 27
         Endovascular grafts............................................................................................... 28
      3.2.4 Concluding remarks ......................................................................................... 28
Chapter 4: Cell seeded synthetic grafts............................................................................. 29
   4.1 Criteria .................................................................................................................... 29
      4.1.1 Favourable surface chemistry .......................................................................... 29
      4.1.2 Sufficient mechanical compatibility ................................................................ 30
      4.1.3 Porous structure ............................................................................................... 31
      4.1.4 Commercial potential....................................................................................... 31
   4.2 Scaffold materials ................................................................................................... 31
      4.2.1 Non-biodegradable........................................................................................... 32


                                                                                                                                     2
        Dacron (PET) ........................................................................................................ 32
        Gore-Tex (PTFE) .................................................................................................. 32
        Polyurethane (PU)................................................................................................. 33
     4.2.2 Biodegradable .................................................................................................. 34
        PCL ....................................................................................................................... 34
        PLLA..................................................................................................................... 34
        P4HB..................................................................................................................... 34
        Dexon (PGA) ........................................................................................................ 35
        Hyaff-11................................................................................................................ 35
        Copolymers ........................................................................................................... 35
     4.2.3 Surface modifications ...................................................................................... 36
        Coatings ................................................................................................................ 36
        Preclotting ............................................................................................................. 37
        Chemical bondings................................................................................................ 37
        Surface treatment .................................................................................................. 37
     4.2.4 Concluding remarks ......................................................................................... 37
  4.3 Production methods ................................................................................................ 38
     4.3.1 Fiber bonding ................................................................................................... 38
     4.3.2 Solvent casting and particulate leaching.......................................................... 38
     4.3.3 Membrane lamination ...................................................................................... 39
     4.3.4 Melt molding.................................................................................................... 39
     4.3.5 Three-dimensional printing.............................................................................. 39
     4.3.6 Electrospinning ................................................................................................ 40
  4.4 Endothelial cell seeding .......................................................................................... 41
     4.4.1 Choice of cell source........................................................................................ 41
        Phenotypic appearance.......................................................................................... 42
        Proliferative capacity ............................................................................................ 42
        Antigenic variety................................................................................................... 44
        Function ................................................................................................................ 44
     4.4.2 Seeding of the scaffold..................................................................................... 45
        Gravitational seeding ............................................................................................ 45
        Hydrostatic seeding............................................................................................... 46
        Biological glue ...................................................................................................... 46
        Electrostatic seeding ............................................................................................. 46
     4.4.3 Characterization of cells .................................................................................. 47
        Phenotypical markers............................................................................................ 47
        Functional markers................................................................................................ 47
  4.5 Preconditioning and dynamical loading.................................................................. 49
     4.5.1 Mechanical protocols ....................................................................................... 49
     4.5.2 Results so far.................................................................................................... 52
     4.5.3 Concluding remarks ......................................................................................... 53
Chapter 5: Future directions.............................................................................................. 54
List of references............................................................................................................... 55
Uncited references ............................................................................................................ 66
Appendix A: List of abbreviations.................................................................................... 69
Appendix B: Leaders in the field ...................................................................................... 71



                                                                                                                                    3
Introduction
Deterioration of the cardiovascular system is still the most common cause of death in
the western world, and though recognized as a severe problem, it remains a growing
socio-economical problem. Atherosclerosis, the most common disease associated
with these cardiovascular problems, is a process that leads to narrowing of the
arteries by an expanding plaque within the intima of a vessel [61]. When it affects the
coronary arteries, the myocardium is weakened, eventually leading to a myocardial
infarct. The typical way to treat coronary disease is a coronary artery bypass graft
surgery (CABG), which usually involves the replacement of the diseased artery by a
substitute vein or artery (figure 1), which is harvested from the patient’s leg or, in
some cases, arm. The human saphenous vein (HSV) and internal mammary artery
(IMA) are the most popular conduits [19]. Clearly though, removal of a functional
vessel from its location in the body is suboptimal. Furthermore, in a fairly large
number of patients, other problems are present like disease of the replacement vessel
itself, usage in previous surgeries or even the need for multiple bypasses.

                                   Figure 1: Typical location for a coronary artery
                                   bypass graft.




A recently emerged area of research is the field of tissue engineering (TE), defined as
the application of knowledge from both engineering and life sciences to ultimately
create biological substitutes to restore, maintain or even improve tissue function.
Although this definition has been narrowed by some researchers, in this article I
would like to keep the original definition in mind.

The above mentioned issues concerning CABG have led to a so-called ‘holy grail’
amongst researchers in the field of tissue engineering, being the development of a
readily available small diameter tissue-engineered blood vessel (TEBV) that mimics
the native vessel in such a way that long-term patency can be achieved [61], [20]. This
article review aims at addressing recent developments in the field of small-diameter
blood vessel tissue engineering, questioning several proposed approaches that have
been subject of intensive research. A close look will be taken with regard to used
techniques, results and relevant conclusions. Furthermore, current opinions on the
requirements of a vascular graft for small-diameter vessels will be discussed. Finally,
a view on future developments and relevant areas of research with respect to small
caliber vascular grafts will be given.




                                                                                      4
Chapter 1: Nature of vascular disease
This chapter aims at clarifying some basic issues associated with common vascular
pathologies to which TEBV can provide a therapeutic answer. Also, it should give
insight in what features a TEBV should have to prevent problems like thrombosis,
aneurysms and stenosis as much as possible after implantation.

The most abundant pathology associated with the need for bypass surgery is
atherosclerosis, the primary cause of death in the western world. This disease is part
of the group of endotheliopathies (i.e. ailment of the endothelium) and will be
discussed in the next paragraph.

1.1 Atherosclerosis
Atherosclerosis is an inflammatory process leading to a thickening of the intima and
is accelerated in coronary artery bypass grafting (CABG). Besides being responsible
for the need for CABG in most cases, many implanted grafts suffer from occlusion
within 10 years after implantation, associated with atherosclerosis. This
atherosclerosis is again associated with neointimal thickening. At present it is widely
accepted that atherosclerosis can be considered as a response to vascular injury,
initiated by circulating factors and modulated by local anatomy and hemodynamics
[63], [84], [49] One of the substances responsible for vascular injury is oxidized low
density lipid (oxLDL). Infiltration into the vascular wall and subsequent oxidation is
influenced by many factors, including blood pressure, circulating concentration of
LDL, endothelial integrity (infiltration) and oxygen free radicals from endothelial cells
(ECs), smooth muscle cells (SMCs) or inflammatory cells (oxidation) [49], [84], [46].


Regions where loss (of part) of the endothelial barrier occurs are very prone to high
rates of LDL infiltration [46]. Most LDL infiltration occurs due to transcytosis though
[75]. Usually this LDL is filtered through the vessel wall, but a hydrodynamic
resistance to fluid flow is one factor that might delay this filtration. The
hydrodynamic resistance, a property of the extracellular matrix in an artery (ECM), is
increased during vasoconstriction [10]. Hence flow-mediated constriction mediated
by nitric oxide [65] and prostacyclin (PGI2) might explain the low LDL accumulation
rates in regions where the vessel wall is subjected to high shear.

Vessel wall cells are capable to induce oxidation of LDL when there is a lack of anti-
oxidants and/or when transition metal ions are present. When a sufficient number of
LDL molecules is oxidized, the death of ECs and SMCs can be provoked [59,44].
Furthermore, oxLDL can stimulate the expression of adhesion molecules on ECs and
production of monocyte chemotactic peptide-1 (MCP-1) from SMCs.




                                                                                       5
1.2 Endothelium dysfunction
Usually, endothelium favors vasodilatation when all factors and mediators are
correctly balanced. A dysfunctional endothelium however has a preference for
vasoconstriction as well as mitogenesis. Whenever endothelial cells are damaged the
result is an inability to synthesize the vasodilatory factors nitric oxide [65] and
prostacyclin (PGI2), resulting in a lack of anti-constrictive action. Dysfunctional
endothelium will ultimately start releasing endothelins, angiotensin II and
endothelial-derived constriction factor (EDCF), three vasoconstrictive mediators.
Together these events will lead to an intense vasoconstrictive cascade. Furthermore,
the potential of aggregating platelets to produce vasoconstrictors at the site of injury
promotes an even larger vasoconstriction.

The above stated is generally true for most vascular diseases. In atherosclerosis, a
decreased production of NO has been demonstrated as well as an increase in the
endothelin level. In the initial stage of atherosclerosis, endothelial dysfunction is one
of the first events that can be observed. This dysfunction is illustrated by a loss of
endothelium-dependent vasodilatation and an increased expression of adhesion
molecules. An impaired ability of relaxation has been ascribed to decreased NO levels
[117]. In bypass grafting, endothelial regrowth occurs rapidly, but impairment of
endothelium-dependent vasodilatation in the regrown endothelium suggests a defect
in receptor coupling to NO production [99]. The effects of NO are at least twofold:
firstly NO inhibits neointima formation and SMC proliferation and secondly it
inhibits leukocyte adhesion to ECs [90].

The second feature of endothelial dysfunction associated with atherosclerosis, i.e. the
increased expression of adhesion molecules, is illustrated by induction of vascular
cell adhesion molecule-1 (VCAM-1) and intracellular adhesion moleule-1 (ICAM-1),
responsible for binding of leukocytes. This endothelial activation is not only
associated with VCAM-1 and ICAM-1, but also with the production of other products
like von Willebrand factor (vWF) and plasminogen activator inhibitor-1 (PAI-1) and
release of them into the circulation.

1.3 Inflammatory smooth muscle cell infiltration and proliferation
Inflammatory insudation, found in models of vein-grafting [87], remains unclear.
Probably oxLDL promotes infiltration after adhesion of monocytes. The discovery of
platelet-derived growth factor (PDGF) has played a key role in the response to injury
hypothesis. PDGF is produced by aggregating platelets and acts as a potent mitogen
for SMCs. PDGF is a chemo-attractant for SMCs and thus stimulates SMC
hyperplasia by directing SMCs from the media to the intima. Also it promotes matrix
deposition. In vein grafting models, PDGF is considered the most important agent to
direct cells into the intima [22]

Ultimately, thrombotic occlusion is likely to occur in atherosclerotic coronary arteries


                                                                                       6
and vein grafts. This usually takes place by thrombus formation on a plaques surface
denuded of endothelium or in a plaque that has ruptured [116].

1.4 Role of shear stress
Although atherosclerosis is associated with systemic risk factors like high blood
pressure and high circulating LDL levels, the disease has a preference to develop at
specific locations in the vascular tree. Such arteries like the carotid artery bifurcation
and infrarenal, femoral and coronary arteries have certain regions of low shear stress
in common. Studies conducted in the last 3 decades have confirmed the low-shear
hypothesis of atherosclerosis first proposed by Caro and coworkers [10].

Direct measurements of shear stress values in the susceptible regions have revealed
values in the order of 2-10 dynes/cm2, while in the rest of the arterial tree shear
stress generally exceeds 15 dynes/cm2. It is now thought that physiological shear
stress values keep the phenotype of ECs in such a way that the production of
vasodilators is increased and that of adhesion molecules is decreased when compared
to subphysiological shear stress, for example. Shear stress of high enough values thus
keeps the endothelium in a atheroprotective state, being anti-proliferative, anti-
oxidant and anti-thrombotic [50,74].In this chapter, it has undoubtedly become clear
why a functional endothelial monolayer is essential when addressing the concept of
tissue-engineered blood vessels. If the pathophysiological processes accompanying
the vascular response to injury can be grasped, our total understanding of blood
vessel responses to a variety of stimuli can be elevated to a level that makes us capable
of producing true tissue equivalents.




                                                                                        7
Chapter 2: Blood vessel physiology
To be able to place all used techniques and proposed methods in the right perspective,
a basic understanding of the physiological characteristics of blood vessels is required.
Although complex differences exist in the cardiovascular system, regionally as well as
within organ systems, a common organization can be observed. In light of this review
however, the physiological characteristics of an artery will be the starting point.

2.1 Cell layers
Generally spoken, an artery consists of three distinct regions, all of them containing
specialized cells and extracellular matrix (ECM). Anatomically, from the outside of
the vessel to the inside, these regions are called the tunica adventitia, tunica media
and tunica intima. Furthermore, the tunica intima is lined with a specialized, single
layer of endothelial cells (see figure 2).

2.1.1 Tunica adventitia

The tunica adventitia or shortly adventitia, the outermost layer of a blood vessel,
consists of mainly fibro-elastic connective tissue and ECM, supplying most of the
mechanical strength of the vessel and the structural integrity. Main components of
the adventitia are fibroblasts and elastin (hence fibro-elastic). Furthermore, the vasa
vasorum are observed in the adventitia, a network of small thin-walled blood vessels
needed to provide the larger arteries with the necessary nutrients and oxygen, since
simple diffusion of the latter from the lumen of the artery to the outside is limited.
Especially the media is dependent on the vasa vasorum. The ECM surrounding
vascular cells is a complex structure consisting mainly of collagen I and III, elastin
fibers, proteoglycans, hyaluronan and glycoproteins (e.g. laminin, fibronectin,
thrombospondin and tenascin). The external elastic lamina serves as a boundary layer
between the adventitia the smooth muscle cells of the medial layer.




 Figure 2: An overview of vascular anatomy.In the left picture,the structure of an artery (left) and avein (right) are drawn.
 On the right hand side is a cross-sectional view of an artery as seen with two different staining techniques.



                                                                                                                       8
2.1.2 Tunica media

The tunica media or media contains mainly smooth muscle cells and elastin fibers.
Especially the smooth muscle cells are highly abundant and make up the bulk of the
vessel wall thickness. These cell layers are usually highly organized in larger arteries
due to the function they have to fulfill in the movement of large volumes of
intravascular blood (i.e. compliance of a vessel). Smooth muscle cells are orientated
circumferentially, making it possible to strongly constrict. The elastin however
should not be underestimated for its role in compliance, since this material, as we
shall see later in this review, plays an important role in the visco-elastic behavior of a
vessel. For structural support, the media rests on an internal elastic lamina,
separating it from the innermost layer, the tunica intima.

2.1.3 Tunica intima

The tunica intima or intima is usually the thinnest structural layer present in a vessel
and is made of a single layer of endothelial cells (ECs) mounted on a basement
membrane. These endothelial (squamous) cells are oriented in a longitudinal way,
aligned with the direction of the main flow. Underneath this layer are a sub-
endothelial fibro-elastic connective tissue layer and an organized layer of internal
elastic lamina that provides flexibility and stability for the endothelial cells. Another
type of cells found in close proximity to endothelial cells is the class of pericytes.
These cells are multipotent stem cells that possess the ability to differentiate into
several different cell types. They therefore provide the endothelial cell layer with a
balanced cellular micro-environment. Other cells can also be present in the intimal
layer, such as lymphocytes and macrophages. The surface of the EC layer express
glycoproteins, together called the glycocalyx, who prohibit blood cells and plasma
proteins to migrate into the vessel wall under normal conditions, hence it is a
charged barrier [105].

The endothelial cell layer is nowadays considered as possibly the most important
agonist in the vessel wall. Instead of being a passive player, acting as merely a sieve-
like physical barrier against unwanted substances to penetrate into the vessel wall
from the bloodstream, it is now accepted that it has a critical role in regulating
thrombolysis as well as in coagulation, inflammatory and immunological processes.
The endothelial monolayer present in all vessels of the vasculature will be discussed
in greater detail in paragraph 2.2 [77,93].

2.2 Biochemical properties
The endothelium is involved in various physiological processes, such as hemostasis
maintenance, vasomotor tone regulation, inflammatory and immunological
regulation and the modulation of angiogenesis (see figure 3).




                                                                                        9
                                                             Figure 3: Anti-thrombogenic properties of
                                                             endothelial cells(Figure adopted from
                                                             Mitchell, 2003).




Regarding hemostasis, the endothelial layer already in a quiescent state releases
many agents that act against the formation of thrombi. Some of them include
heparan sulfate, prostacyclin (PGI2), nitric oxide [65] and ADPases. These all inhibit
platelet aggregation. Other factors that influence the coagulation process are tissue-
type plasminogen activator (t-PA) and plasminogen activator inhibitor-1 (PAI-1),
which have a positive effect on the digestion of platelet-fibrin clots. Besides
counteracting the thrombosis-cascade, the ECs also participate in regulating
thrombus-formation locally. The von Willebrand factor (vWF), a glycoprotein
responsible for carrying factor VIII in the plasma, mediates the adhesion of platelets
and their incorporation into evolving thrombi.

The second key process in which the endothelium is involved is the modulation of
vasoactivity [56]. Usually, endothelial layers favor dilatation of the vessel wall. Many
mediators are released that act as potent vasodilators, such as endothelium-derived
relaxation factor (EDRF), NO, PGI2 and endothelium-derived hyperpolarizing factor
(EDHF). Except NO, these substances are synthesized ‘de novo’ when an appropriate
trigger is induced (see figure 4).




                       Figure 4: Basic vasoactive mechanism in a vessel wall (Figure
                       adopted from Mitchell, 2003).




                                                                                                         10
Aside from vasodilatation, the endothelium also takes care of different situations
when rapid vasoconstriction is needed. Vasoconstrictive agents are mainly the
endothelins. These peptides are secreted by the endothelial cells primarily on their
abluminal side. When released, they cause an intense and prolonged constriction of
the vessel wall. A second constriction mediator is endothelium-derived constriction
factor (EDCF), which plays a minor role, together with angiotensin II, another
stimulating agent. It should also be mentioned that besides producing and releasing
factors themselves, endothelial cells promotes vasoactivity indirectly by binding
substances that are secreted elsewhere in the body, such as serotonin, acetylcholine,
histamine, insulin and of course norepinephrine. When bound, the action is
translated to the smooth muscle cells by the endothelium[56].

The third key process in which the endothelium plays a crucial role is the regulation
of inflammatory and immunological processes. A wide variety of cellular adhesion
molecules (CAMs) has been discovered, expressed on the glycocalyx surface of several
cells like leukocytes, platelets and vascular endothelial cells [40]. Whenever cellular
damage and/or infection occur, the CAMs get involved in the binding of leukocytes
and inflammatory mediators. Although an elaborate discussion of these CAMs is
beyond the scope of this review, it should be noted that they are vital for a correct
functioning of a vessel. Table 1 summarizes the known CAMs at present and their
function in the vessel [3,85].

 Table 1: Overview of known vascular CAMs.




The fourth and last key process in which endothelial cells in the vessel wall are
involved is angiogenesis or new vessel formation from pre-existing vessels. Though
the exact mechanisms still remain unclear, it is assumed that angiogenesis is
controlled by a balance of angiogenic and anti-angiogenic cytokines [15,65]. Vascular
endothelial growth factor (VEGF) and fibroblast growth factor (FGF) promote
angiogenesis [74]. Both promote mitosis of cells, and are produced by smooth muscle
cells and pericytes. Angiostatin (made from plasminogen) and endostatin inhibit
angiogenesis. Upon stimulation of angiogenic cytokines, both the endothelium and
SMCs proliferate and migrate in the direction of the angiogenic stimulus to form a
new intima and media.

2.3 Mechanical properties




                                                                                    11
The structural and mechanical properties of a blood vessel determine its strength and
reaction to hemodynamic forces. In the light of this review and in particular the part
of it coping with the tissue engineering [107] issues, these characteristics of vessels
will be covered first.

Blood vessels and in particular arteries bear three distinct mechanical characteristics,

 Figure 5: Typical stress-
 strain curve of native
 tissue and intended graf
 (Adopted from Shum-
 Ttm, 1999).




being elasticity, determined by the elastin content, tensile stiffness, determined by
the collagen content and compressibility which is taken care of by the
glucosaminoglycans (GAGs). The degree of contribution of elastin and collagen is not
only determined by the amounts of these proteins, but also by the orientation of the
fibers and the degree off crosslinking between individual fibers, meant to stabilize
the fibers. Collagen and elastin are part of the extracellular matrix (ECM) and may be
secreted by vascular smooth muscle cells in the medial layer of the artery wall and/or
fibroblasts in the adventitial layer. Together with the GAGs they make that an artery
has a visco-elastic nature, responding to stress as indicated by the following curve
(see figure 5) [101]. For clarification, the anticipated development of the curve for a
vessel conduit after implantation is depicted as well.

Arterial burst strength and compliance are both closely connected with the visco-
elastic behavior. The burst strength determines whether a vessel is capable of
withstanding the physiological pressures that exist in the cardiovascular system,
while the compliance of a vessel determines the way pulsatile wave energy is
dissipated and thus how the blood enters the downstream vasculature. Further in this
review we will focus on mechanical values that are relevant for tissue engineering of
vessels. First let us return to the two basic mediators of vascular mechanical strength
in a vessel.

2.3.1 Collagen

Mechanical strength of a blood vessel is mainly derived from collagen and elastin,
both produced by the SMCs in the media. The collagen fibers are the main providers
of the tensile stiffness of a vessel and therefore maintain the structural integrity,


                                                                                     12
largely determining the burst pressure. The synthesis of collagen is positively affected
by growth factors and cyclic strain. The fact that collagen is such a prevalent protein
in the vessel wall stimulated many research groups to use collagen substrates as
scaffolds for tissue engineered blood vessels (TEBV). However, collagen fibers only
provide mechanical strength when they are crosslinked in a highly organized way
(see figure 6). In coronary arteries, the collagen content varies between 40 % and
27%, depending on age (less when the person is older) and coronary artery (left or
right) [70].
                                                     Figure 6: Organization of collagen and
                                                     elastin in a vessel.




2.3.2 Elastin

Elastin, the other ECM protein secreted by the SMCs, is responsible for the elastic
properties of a blood vessel. It acts as a recoil protein that stretches and pulls the
vessel back to its original diameter, and thus contributes to a large extent to the
compliance of vessels. Vascular dilatation, which would be caused by the pressures of
blood flow, is prevented in vivo after implantation. Also, adequate transmission of
pulsatile wave energy is hindered when the compliance of the TEBV is too low, which
is often the case when synthetic grafts are used. The in vitro production of elastin
fibers still remains a challenge, although some synthesis of elastin after implantation
has been observed [64]. Consequently, the elasticity index remains below
requirement level. The elastin level in human coronary arteries varies between 1 to
12 %, again being lower at higher age. The ratio of collagen and elastin in the right
coronary artery decreases from 3.8 to 1.7 and in the left coronary artery from 5.4 to
2.6, similar reductions of about 50 % between the age of 20 and 60 [70].

2.3.3 Coronary artery specific properties

In the previous paragraphs, an attempt has been made to clarify the basic properties
of blood vessels in general. However, this review is specifically aimed at the research
done in the field of small-diameter vessel replacement. Therefore it is useful to briefly
describe the specific features of human coronary arteries and their hemodynamics
which are relevant to the TEBV construction process.




                                                                                              13
First of all the physical characteristics, important for the construction of the graft: a
typical coronary artery has a diameter that varies from 2.7-4.1 mm and a wall
thickness that varies from 0.56-1.25 mm, according to van Andel and coworkers [115].
Two important mechanical properties, essential for TEBVs to match as good as
possible, are the burst pressure, being about 2000 mmHg [79] and the vascular
compliance, varying between 0.010 mmHgand 0.052 mm2/mmHg at a mean
arterial pressure of 100 mmHg, with a mean compliance of 0.02 mm2/mmHg [121].

The hemodynamic characteristics of the coronary circulation are highly relevant to
the TE approach. Blood pressure varies between 80 and 120 mmHg with a mean of
100 mmHg by generalization, whereas the mean coronary blood flow rate is about
60-70 ml/min [121]. The pulsatile flow has the effect on the coronary arteries of
imposing a cyclic radial strain of 4-10 % per 100 mmHg (Greisler & Zilla, 1999),
giving a circumferential stress of 0.05-0.2 MPa and a shear stress on the
endothelium of 3.3-12.4 dynes/cm2 with a mean of 6.8 dynes/cm2 [17]. The latter is
interesting, since the shear stress values in the rest of the arterial circulation is much
higher, being generally spoken higher than 15 dynes/cm2. The shear stress in the
coronary circulation is closer to that in veins, which could explain the success of
saphenous vein grafts, even though the higher arterial blood pressure is detrimental
[48].




                                                                                       14
Chapter 3: Tissue engineered blood vessels


With the ongoing increase in percentage of the western population suffering from
cardiovascular disease, also the incidence number of coronary artery bypass grafting
is getting larger and larger. Therefore the development of an artificial construct that
mimics the native coronary artery after implantation seems increasingly interesting,
also from a commercial point of view. Although considerable success has been
achieved with such constructs for replacement of larger vessels, when applied to
small-diameter arteries with a radius of less than 3 mm, outcomes have been
disappointing. Hence a tissue engineered blood vessel (TEBV) that can replace an
autologous vein or artery without the at present still accompanying higher costs
(apart from still far from satisfactory results) is desirable. Though kind of macabre
the increasing number of diseased is likely to further reduce the costs of TEBVs.
Hence the future is bright for groups working on this domain! However, as noted the
results are still far from optimal and many aspects still have to be improved before a
coronary artery equivalent is produced that can be approved by the FDA. To
accomplish these goals several issues have to be addressed. In the next chapters some
of these issues and requirements will be discussed and elucidated.

3.1 Requirements of tissue-engineered vessels
Four basic TEBV requirements have to be met in order to mimic the functional
characteristics of a living blood vessel. TEBVs should be non-thrombogenic, non-
immunogenic, exhibit vasoactivity and possess mechanical properties matching those
of the native vessel [108,61,79]. To achieve this, Nerem et al. describes three essential
technologies to be developed [61].

1. Cell technology

This core technology concerns the source of cells to be used and the manipulation of
cell function. Obviously, the best choice would be autologous cells, although usually
allogenic cells are used in current research. For smooth muscle cells this may not be
a problem, as we will see in paragraph 4.3, but for ECs there should be a solution to
enhance immune acceptance. Control and manipulation of cell function, which is
achieved by either altering the cell’s environment or changing the cell’s genetic
features, are both an option. The first can be easily understood because SMCs and
ECs are all very much under influence of their environment, both biochemically and
mechanically. The second feature includes the change of a cell’s genetic ability for the
secretion of anti-thrombogenic substances, expression of elastin and immunological
acceptance.




                                                                                      15
2. Construct technology

In order to gain success in the field of TE, a neovessel should optimally resemble the
native vessel to be replaced. With regards to this matter it is not only important to be
able to engineer constructs that provide the right 3-dimensional structure and
mechanical properties, e.g. burst pressure, elasticity etc. Also, the matter of
reproducibility should be considered if in the future a methodology has to be
commercialized.

3. Integration into living system

For this, one should consider the choice of animal model, control of biological
responses that occur when implanted into a living system and immunogenic
acceptance. Is it possible to control the reaction mechanisms and in-vivo evolution of
the construct?

The accomplishment of these technologies should result in a functional TEBV that
fulfills the requirements for a blood vessel substitute (see table 2 [60]), being non-
thrombogeneicity, vasoactivity and proper mechanical properties.


                                                    Table 2: Requirements for a blood vessel
                                                    substitute (Adopted from Nerem, 2003).




3.1.1 Mechanical Properties

The mechanical properties problem is the first to be addressed, since a TEBV, being a
polymeric graft, a reconstructed neovessel or a combination of both, has to bear a
load immediately after implantation. The hemodynamic conditions inside the human
body require a blood vessel to have a certain strength (in terms of burst pressure) and
compliance [25]. In order too prevent rupture of the neovessel the burst pressure
should be at least 2000 mmHg. Although physiological pressures generally do not
exceed 250 mmHg, a large safety margin is desirable. For a correct energy dissipation
and prevention of thrombosis, the radial elasticity or compliance of the implanted
graft should match that of the native artery as close as possible [1,47].

Compliance of a vessel is defined as the ratio of change in diameter over change in
blood pressure, expressed as the percentage diameter change per mmHg (%/mmHg
x 10-2). Writing this definition in mathematical form, this gives:


                                                                                               16
              ( D s − Dd )
C=                                                                                               (1)
           [( Ps − Pd ) ⋅ Ds ]

In this equation of compliance C, DS and DD depict the end-systolic and end-diastolic
diameter, respectively, whereas PS and PD describe the corresponding pressures.

A typical curve of the compliance of a human coronary artery in relation to arterial
pressure is depicted in figure 7. In this graph, the compliance is depicted as change
in cross-sectional surface per mmHg, another way to define this parameter. The
declining slope found in coronary arteries is lacking in grafts by definition [126].

                                                             Figure 7: Variation of compliance with
                                                             pressure in coronary arteries (Reproduced from
                                                             Zilla, 1999).




A normal radial elasticity of a human coronary artery results in a dilatational strain of
about 6% per 100 mmHg [126]. When this compliance property can not be achieved
when engineering a TEBV, the patency rate reduces greatly with increasing
compliance mismatch. Early trials with Dacron and PTFE grafts showed a
compliance mismatch much larger than with autografts like saphenous vein (see
figure 8).

                 120                                                    Figure 8: Relation between
                                          2                             compliance of native coronary
                                         R = 0.9407                     artery or coronary grafts and
                 100
                                                                        patency rate. Increasing values
                                                                        of compliance belong
   Patency (%)




                 80
                                                                        respectively to PTFE grafts,
                                                                        Dacron grafts, bovine
                 60
                                                                        heterografts, umbilical veins,
                                                                        saphenous veins and native
                 40                                                     arteries (graph data taken over
                                                                        from Zilla & Greisler, 1999).
                 20

                  0
                       0         0.02     0.04        0.06   0.08
                                  Compliance (%/mm Hg)


                                                                                                          17
It is clear that the patency rates of grafts will increase together with compliance in a
linear fashion. This graph indicates the importance of a sufficient compliance of the
graft to assure a good clinical outcome.
These essential mechanical properties can be achieved through a visco-elastic nature
of the newly constructed artery, which is mainly determined by the contents of
collagen and elastin in the TEBV, as mentioned earlier.
Finally an additional mechanical requirement should be mentioned. A good
suturability seems to be obvious, but nonetheless very important. The anastomosis
between the neovessel and the intact artery is often the location where intimal
hyperplasia occurs, also because of a compliance mismatch [45]. This can lead to
occlusion of the graft and therefore a shorter patency period. Furthermore, the suture
should hold under circumferential as well as longitudinal tensions, and should retain
axial and radial compliance and pulsatility.

3.1.2 Vasoactivity

The ability of a neovessel to adapt to changing hemodynamic forces and chemical
stimuli is directly correlated with the mechanical properties as mentioned above.
Grafts that are unable to dilate or contract in an appropriate way will consequently
lack the correct compliance upon changing hemodynamic situations. Therefore the
medial layer is indispensable in a TEBV. The endothelial cell layer is usually the layer
that transfers the stimuli to the medial layer since it provides a (not absolute) barrier
between the blood with vasoactive factors and the smooth muscle cells. Therefore this
layer seems indispensable as well. However, in vitro experiments have proven that
smooth muscle cells can behave in a vasoactive way similar to their behavior in ‘in
vivo’ conditions, even in the absence of an endothelium.

3.1.3 Non-thrombogeneicity

Possibly the most important requirement for successful substitution of a blood vessel
is a confluent, adherent endothelial layer to provide a non-thrombogenic lining.
Whenever blood comes into contact with another surface than the endothelium, there
is a risk of thrombosis [7]. Furthermore, loosely attached ECs could detach right after
implantation due to blood flow related shear stress. The EC layer also actively inhibits
thrombosis. This is achieved by thrombomodulin receptors, heparan sulfate,
proteoglycans and the secretion of NO, prostacyclin, protein S and t-PA, all of which
inhibit the clotting process (see fig. 3 in Chapter 2).

Aside these features, the endothelium also plays a role in blood pressure regulation,
angiogenesis and adhesion and transmigration of inflammatory cells. Hence, it can
be understood that the lack of a functional endothelium or an aberrant one can lead
to problems such as atherogenesis, bleeding disorders or graft rejection.

In conclusion, a TEBV for small-diameter arteries should behave vasoactively through
a non-thrombogenic luminal lining and exhibit correct mechanical properties,
matching the compliance of the native vessel as close as possible and possessing a


                                                                                      18
strength that allows a burst pressure of at least 2000 mmHg, thereby creating a
sufficiently large safety margin.

3.2 Approaches
The current constructs used for the development of TE neovessels can be classified in
two main categories: a-cellular and cellular constructs. Another division can be made
into synthetic or natural constructs. These classifications can be looked upon in a
chronological perspective as well, since basically investigators started with just a-
cellular constructs to replace blood vessels, whereas the need for a more adaptive
approach with incorporation of living cells was recognized only in a later stage. In the
next paragraph we will focus on the main disadvantages and advantages of the
methods, clinical results, if any, and future directions. Also, other recently developed
methods with great potential are briefly discussed.

3.2.1 A-cellular constructs

Non-biodegradable synthetic constructs

Synthetic polymer grafting has been very successful for the replacement of large-
diameter blood vessels. Such materials as Dacron (PET) and ePTFE have extensively
and successfully been used and were the standard biomaterials for synthetic grafts
over the past decades. However, for the replacement of small-diameter blood vessels,
they suffer from rapid occlusion, mainly because of the thrombogenic blood-
contacting surface. Basically, one could say that the healing process of polymeric
synthetic grafts is quite unfavorable, leading to problems like incomplete
endothelialization and intimal hyperplasia (IH), eventually resulting in the occlusion
of the graft.

Intimal hyperplasia is characterized by proliferation and migration of SMCs from the
medial wall to the intima, with subsequent synthesis of ECM proteins and other
matrix material [1,95]. The cause of IH is probably multifactorial, and factors involved
are a compliance mismatch between native artery, the site of anastomosis and the
graft, which leads to a disturbed flow pattern, turbulence and vessel wall injury.
When implanted as a replacement for small blood vessels, blood and tissue reactions
occur immediately, which leads to a cascade of events ultimately occluding the graft.
Dynamic protein adsorption/desorption, followed by platelet adhesion, inflammatory
cell infiltration, and EC and SMC migration are the first steps. These processes result
in the deposition of a compact fibrin layer, the so-called pseudo-intima, and of
foreign body giant cells, densely packed between the outer layer of the graft wall and
surrounding connective tissue. These initial effects keep evolving until final failure of
most grafts.

An alternative for ePTFE and Dacron grafts are the polyurethanes (PU), polymers
that gained popularity because they have a better biocompatibility than ePTFE and
Dacron [112].


                                                                                      19
Aside from this, PU, and especially segmented polyurethane (SPU) have far better
elastic properties than the traditional polymers. Some types of PUs have reached the
phase of clinical trial [112]. However, eventually the technique was abandoned
because of varying problems like occlusion or rapid biodegradation. Most recently, a
new PU named MyoLink has been produced, consisting of poly(carbonate)
polyurethane. This material has shown superior compliance characteristics as well as
adaptable mechanical properties. A final version of the MyoLink graft was compared
with native vein and artery grafts, as well as ePTFE and Dacron grafts (see figure 9).


                                                        Figure 9: Comparison of several grafts, indicating a
                                                        excellent compliance of the MyoLink graft (graph
                                                        reproduced from Tiwari, 2002).




This PU grafts show similar compliance properties as a native artery in the
physiological range of 30-100 mmHg [112]. Furthermore, this new PU better resists
biodegradation than other PUs. Other experiments with PU prostheses as small as 1.5
mm ID were performed by Zhang et al [125]. PU grafts were implanted in rats and
remained patent for minimal 8 weeks, indicating no initial occlusion problems.

Nonetheless, still no conclusion can be drawn as to whether PU grafts are truly
applicable for small blood vessel replacement. The significance of immunological
reactions to the polyurethane, still being a foreign material to the human body, has
yet to be proven irrelevant. Another major concern is the carcinogenic potential of the
degradation products, although recent results indicate only a non-significant
biodegradation [94].

Biodegradable synthetic constructs

In the case of non-degradable synthetic grafts tissue-material interactions continue,
and this eventually could lead to failure of the graft. Bioresorbable or biodegradable
polymers disappear over time, hence excluding foreign-body reactions. Therefore, it
should be theoretically possible to ‘in situ engineer’ a neoartery if enough load-
bearing capacity is present to resist initial dilatation, and if cells are attracted in such
way that desirable physiologic characteristics are already present before degradation
occurs.



                                                                                                  20
Two intensively studied polymeric materials are polyglycolic acid (PGA) and
polylactic acid (PLA). PLA is mainly used in its L-form (PLLA) because it has a high
mechanical strength. PGA is highly crystalline and hydrophilic. PLLA is more
hydrophobic and thus less prone to hydrolysis. In several clinical studies both
polymers have shown to have a good potential to eventually develop into a viable
alternative for diseased vessels. The main problem is the generation of a sufficiently
strong neo-tissue before total resorption of the polymer occurs, leading to aneurismal
dilatation. The combination of PGA or PLLA with other polymers was not successful
in solving this problem. Although a relatively compliant material could be introduced
into the body to induce circumferential orientation of the SMCs, tissue ingrowth
compromises the compliance in many cases.

Several methods have been developed to enhance the patency rates. Examples are the
linking of heparin to graft surfaces so that the thrombogenic activity is reduced. The
major problem is the duration of heparin activity due to premature release of the
compound or the presence of a physical barrier, created by adherent blood
components.

Other modifications are the coating of the luminal surface with carbon so that electro
negativity is improved and thus thrombus formation reduced. Another coating
material is fibrin glue. It is believed to improve endothelialization and other physical
and chemical variations. However, little clinical improvements have been observed.

Biological constructs

One of the main problems of a-cellular synthetic constructs is the danger of an
immunological response to the bulk material or the degraded debris. Furthermore, it
is impossible to mechanically remodel these constructs. The concept of natural
(biological) materials was therefore proposed as early as 1985 when Weinberg and
Bell proposed the concept of a collagen-based construct [120]. The first tested samples
had an unfavorable mechanical strength, so that several, yet unsuccessful attempts
have been made to reinforce these constructs, among others by improving the
alignment [41,28] or through glycation of the collagen fibers [24].

Another a-cellular approach was described by Badylak and coworkers [4]. Badylak
used the small intestinal submucosa (SIS), which is a cell-free layer of collagen
derived from the small intestine. When rolled it has a typical burst pressure of about
3500 mmHg, making it very strong. A SIS graft is about half as compliant as a canine
carotid artery, but much more compliant as a typical vein graft [82]. It is therefore a
potential candidate for application. Badylak’s group observed recruitment of a
neointima and endothelial monolayer within one month after implantation [83].
Other researchers like Huynh [31] and Kim [38] observed patency rates of 13 and 8
weeks in rabbits and rats, respectively.

Recently, investigators have been tempted to use constructs containing both collagen
and elastin fibers. A lacking elastin is often the reason for insufficient compliance.



                                                                                     21
One could therefore argue that decellularized arteries may offer more suitable grafts
than collagen grafts. Tamura et al. [106] reported a method for the decellularization
of an artery, resulting in a construct with aligned elastin and collagen fibers and a
burst pressure comparable to that of native vessels. Upon implantation very good
results were obtained. The conduits had normal layers of SMC and a lining of
endothelial cells covering the entire luminal surface. Even neovascularization was
observed in the adventitia. A coating of heparin was required to prevent direct
obstruction of the graft though.

Instead of using arteries for decellularization, the group of Schaner [91] tried human
saphenous veins. Advantages of veins include a thinner wall which is easy to
decellularize and thus also a thinner ECM for better migration of recipient cells and
nutrients. Besides, veins are easy to harvest from tissue donors. After
decellularization, Schaner found excellent burst pressure levels and suture retention
strength. Upon implantation in dogs an equally good patency time was observed.
Although this group refrained from using human veins and did not test the
constructs in a low-flow environment where thrombosis is more likely to occur, the
model seemed to be feasible and interesting for further research.

Concluding remarks

The use of a-cellular constructs thus far resulted in quite disappointing results in
terms of long-term patency, strength and immunogenic response. However, when
their mechanical and chemical characteristics are properly adapted they can be used
in practice. A-cellular constructs can be easily sterilized and time-consuming,
expensive and difficult tissue engineering is avoided. Upon implantation, a major
disadvantage is the incomprehensive recruitment of cells after implantation. This is
poorly understood so that it is also difficult to predict the development of a functional
neovessel. Another drawback is the lack of an endothelial lining when implanted,
leading to exaggerated embolization and immunogenic responses.

3.2.2 Cellular constructs

Non-biodegradable synthetic constructs

One of the first attempts to improve function and patency of polymeric grafts was to
seed endothelial cells onto the synthetic surface so that it became less thrombogenic.
This technique has been applied to Dacron and ePTFE grafts [16,92], although up till
now only autologous endothelial cells have been used because of immunological
reactions. Since the adhesion of EC to the polymer material was poor, several
attempts have been made to alter the surface properties of the polymer, such as the
use of a recombinant fibronectin-like adhesion factor [54] or of ammonia plasma
treatment, both on ePTFE grafts [102]. No satisfying improvements were observed.

In 2001, Seifalian and co-workers showed that a combined binding of heparin and
the RGD-sequence (amino acids sequence of arginine-glycin-aspartic acid) to the


                                                                                      22
surface of MyoLink improves cell attachment to over 75% [94]. Hence these
investigators succeeded in obtaining a vascular conduit with an optimal compliance
and sufficient EC-attachment and -retention properties. Although this CABG
substitute seems to be appropriate, it remains to be proven if in situ remodeling of
the graft is such that it continues to bear the correct properties, even with the
polymeric framework still inside. It is possible that a correct regulation of vasotone is
inhibited by the synthetic material, eventually leading to intimal hyperplasia.

Biodegradable synthetic constructs

A new approach was established using biodegradable polymeric scaffolds seeded with
cells. The general idea is that bioresorbable or biodegradable polymers are seeded
with (autologous or allogenic) cells in vitro. Further in vitro culture would stimulate
the development of a neovessel. When inserted in the body, the polymer continues to
degrade while the new tissue further develops into a functional vessel with an
optimal cell orientation, neovascularization and vessel wall organization. For example,
a cell-seeded polymer could be left to develop in vitro for one month, after which the
rest of the development will take place in vivo, leading to the formation of an
endothelium and a media containing collagen and elastin.

Although at present this technique seems very promising, a number of issues have to
be verified. For example, it has to be determined whether or not residual polymer
fragments have a negative effect on the microenvironment under in vivo
circumstances. Another issue is the way in which a cell-seeded scaffold has to be
preconditioned before placement. It is clear that an organized tissue with correct
mechanical and vasoactive properties can only be achieved when the appropriate
mechanical conditions are imposed. Still it remains unclear what these conditions
should ideally be. Several regimes have been investigated with various degrees of
success, as far as mechanical strength and vasoactivity are concerned. Sometimes the
preconditioning protocol seemed to be appropriate, but the in vivo hemodynamics
resulted in an overshoot of ECM production. Although it is clear that ECM deposition
is vital for the establishment of graft strength, the excessive matrix formation
indicates undesirable tissue remodeling. The balance is thus very fragile, not well
understood, and hence insufficiently controllable.

 This approach however has not been without success. Recently promising results
were obtained by several groups. Niklason et al. [64,104] reported a burst pressure of
more than 2000 mmHg in constructs made of PGA meshes seeded with SMCs.
Under these conditions, the constructs were allowed to mature for 8 weeks after
which ECs were seeded onto the lumen. The authors applied a dynamic culturing
protocol by using pulsatile radial stresses (165 pulses per minute, maximal radial
strain 5%). In addition, the culture medium was supplemented with 20% fetal bovine
serum (FBS), ascorbic acid, proline, alanine, glycine and copper sulfate to ensure
proper nutrition and biochemical signaling. Pulsed constructs had very high burst
pressures, most likely because of high collagen content. Also, they exhibited some
vaso-activity in that they contracted when stimulated with serotonin. Compared to



                                                                                      23
native vessels however, the response was less prominent. There was also no elastin
present.

In vivo studies with this type of constructs have already started; the first results seem
to be promising. Even though it only concerns short-term results in animals and
many aspects still have to be elucidated, it still seems very promising. Attempts are
underway to induce elastin production by means of transfecting the SMCs with the
tropo-elastin gene.

At MIT in Boston, the group of Shum-Tim [101] used a PGA scaffold supplemented
with an outer layer made of polyhydroxyoctanoate (PHO). They cultured mixed cell
populations of FBs, ECs and SMCs and seeded such populations onto the luminal
surface of the grafts. Seven days later the grafts were implanted in lambs. Evaluation
of the constructs after a 3-5 months period revealed a structure of the implanted graft
similar to that of native arteries. An endothelial lining was present, as well as that the
cells in the TEBV generated collagen, elastic fibers and the von Willebrand factor,
indicating a differentiated EC-type. Furthermore in TEBV’s implanted for almost 6
months the stress-strain profile started to approach that of native tissue. Although
optimization of cell attachment and degradation time need to be further established,
the ability of a cell-seeded graft to remodel into a physiological vessel has been proven.

Hoerstrup et al. [29] designed a tubular scaffold made of PGA coated with the novel
material poly-4-hydroxy-butyrate (P4HB). The investigators seeded tubular constructs
of 5 mm diameter with myofibroblasts. After 4 days of incubation, ECs were seeded
onto the luminal surface. During the first week, an increasing flow and pressure was
applied to minimize the ‘wash-out’ effect of flow after seeding of the graft. An
increasing burst pressure, collagen content and DNA content was observed when the
constructs were pulsed compared to static controls. The feasibility of using this type
of construct was demonstrated although exact culture conditions remain to be
determined.

The group of Opitz [67] made a P4HB scaffold using a mold the size of an ovine
aorta. The polymer solution was poured in the mold, after which the
dichloromethane was allowed to dissolve. Before cell loading, the scaffolds were
precoated with collagen I, Matrigel, and gelatin and then placed in a roller mixer.
This precoating provided a better cell attachment. vSMC loading was done by
suspending 107 cells in 80 ml of medium and allowing them to attach overnight at 37
degrees in the roller mixer. A coating of ECs completed the in vitro culture. Pulsatile
cultivation of the construct resulted in a complete colonization with vSMCs, with
expression of SMC specific markers like α-SMA, calponin and caldesmon, although
the first two of these markers were less prominent as compared to native ovine
arteries. Also collagen and elastin contents amounted to a small percentage of that in
native arteries. In our opinion this undesirable effect could be prevented by varying
the culture conditions. One of the most important aspects of the work of Opitz is the
demonstration of redifferentiation of vSMCs, indicating a reversible phenotypic
modulation of SMCs even when cultured in vitro. Since beneficial features of SMCs



                                                                                      24
are otherwise lost because of the in vitro culturing, this could be of great value.

In summary, the latest progress in the field of biodegradable polymers as scaffolds
for TEBVs has been very encouraging. Some questions remain to be answered. For
example, what culturing conditions are best to promote the development of a
compliance-matching graft? Furthermore, how could the issue of lack of elastin
content be tackled? The same can be said about the long maturation time needed for
optimizing a construct in vitro. Once these problems solved the capability of
remodeling and vaso-activity makes this approach very appealing.

Biological constructs

Biological grafts on which cells are seeded could offer some advantages over the
earlier described constructs. First off all the problem of remodeling associated with
synthetic non-biodegradable grafts could be circumvented. Secondly the imminent
danger of residual polymer fragments would be circumvented. As mentioned earlier,
Weinberg & Bell developed a collagen-based constructs. More recently they produced
an adventitia of fibroblasts and collagen, a media from SMCs and collagen and an
intima from ECs. Nevertheless, due to inherent weakness of collagen gel additional
Dacron support sleeves were necessary to withstand physiological pressures. This
addition however limited both remodeling and vaso-activity. Therefore, investigators
focused their efforts on the improvement of collagen-based constructs without using
synthetic support materials. Some investigators aimed at improving strength with
magnetic prealignment and mandrel contraction; however no sufficient strength
could be achieved. Recently Girton et al. demonstrated that glycation of the collagen
matrix through culturing in high glucose medium significantly improved the
construct strength [24]. According to their measurements, a construct of 1 mm
thickness and 3 mm diameter has a burst pressure of 225 mmHg. This strength is
still far from the one observed in native arteries and insufficient for TEBVs to serve
as arterial replacements.

Berglund and his coworkers [6] proposed the use of a so-called construct sleeve-
hybrid (CSH). To this end, an a-cellular collagen support sleeve was fabricated by
dehydrating a tubular collagen suspension. A collagen suspension with human
fibroblasts was then cast upon the support sleeve. As such the support sleeve
provides the initial strength required for a vessel conduit, while during cell
maturation the sleeve itself is remodeled to be incorporated in the construct as a
whole. Berglund observed a positive effect of the a-cellular support sleeve on the
strength of the CSH. The same was true for crosslinking of the collagen fibers
through glutaraldehyde treatment. Both in supported and crosslinked constructs the
burst pressures were as high as 700 mmHg. Although this is much higher than in
earlier constructs, it still remains below the burst strength of native arteries, which
lies in the order of 2000 mmHg. Aside from this the crosslinked constructs seemed
to be more prone to abrupt failure than non-crosslinked constructs.




                                                                                      25
Berglund’s experiments clearly show the feasibility of tailoring a support sleeve made
of collagen such that compliance and mechanical strength are optimal, whereas the
biodegradation process of co-cultured layers of a-cellular and cellular collagen is
comparable to that in synthetic polymer-based constructs, eventually leading to a
completely cellularized vessel conduit.

In summary, at present collagen-based biological constructs in combination with an
a-cellular collagen support sleeve seem to have the best potential for use as vascular
graft, although they still lack the desired strength. The consequence of the absence of
any synthetic material is a reduced strength.

Concluding remarks

By reviewing the concept of cell-seeded scaffolds, whether synthetic or biological, it
becomes clear that biodegradable materials are superior over others both from an
immunological and remodeling point of view. Although potentially ideal many
questions remain to be answered. Correct culturing conditions are essential for the
in-vitro development into a vessel-like conduit. These conditions have a profound
effect on cell orientation, strength, compliance and visco-elastic behavior of the
construct, on the formation of ECM and the expression of other proteins.
Furthermore, the degradation process should be characterized in more detail in order
to predict the level of the eventually desired strength. Hence the choice of scaffold
material, whether coated, modified or otherwise enhanced, is crucial. All factors have
to be adequately tuned in order to obtain a TEBV suitable for commercial use.

Even with all these drawbacks, the polymeric biodegradable scaffold seeded with
living cells offers a great potential and has the benefit over biological constructs in
terms of higher initial strength and better controllability.

3.2.3 Other techniques

Self-assembly method

In 1998 L’Heureux et al. proposed an exciting new methodology, now known as the
self-assembly method [42]. In short the procedure involves a number of consecutive
steps, i.e. ①the construction of an inner membrane consisting of a dehydrated
fibroblast layer, ② the wrapping of another layer around the inner membrane,
providing the media, and ③ a second layer of fibroblasts to make up the adventitial
layer and also intended to provide the necessary strength. Finally, the three layers
were wrapped around an inert cylindrical mandrel made of PTFE.

Each layer was created by culturing the required cell source (FBs, SMCs etc.) in a
flask for a certain amount of time. When a sheet was formed, it was rolled around the
mandrel. New layers were added as described above. In some cases the intraluminal
side was covered with endothelial cells. After the formation of a cohesive construct,
L’Heureux and colleagues tested the mechanical strength and/or integrity. Very


                                                                                    26
encouraging results were obtained as a supraphysiologocal burst pressure of over
2000 mmHg was reached. Furthermore, constructs were organized strikingly similar
to native human arteries. Observed was a dense collagenous matrix, ECM proteins
such as fibronectin and laminin, and even elastin fibers produced in the adventitial
layer. Especially the presence of elastin could be of utmost importance in favoring
this new methodology above traditional ones. Furthermore, desmin was produced by
the SMCs, which is a feature normally lost in cultured SMCs. Some questions
however still remain to be answered, such as about the right conditions for obtaining
optimal compliance and visco-elastic properties. The major drawback of this
technique is the extremely long in-vitro maturation period of at least three months,
which could hamper the progress towards clinical applications.

Body cavities as bioreactors

Recently, Julie Campbell’s group [11,14,8] published the first results of a new
alternative. They investigated the possibility of body cavities such as the peritoneal
cavity to behave as a ‘native bioreactor’ for TEBV maturation. The underlying
assumption is that foreign body material invokes an inflammatory reaction.
Haematopoetic cells which float in the peritoneal cavity attach to the foreign body,
differentiate into myofibroblasts and form a capsule through the production of ECM.
In the peritoneal cavity of dogs polymer tubes of up to 25 cm long were implanted for
a period of three weeks. The tubes, around which a thick homogeneous capsule had
formed was tested for mechanical properties and characterized biochemically. Burst
pressure exceeded 2500 mmHg, while α-SMA and collagen as well as small amounts
of elastin, were found. Campbell reimplanted these conduits as replacements for
femoral arteries in the same dogs, which was associated with remarkably good
patency rates. Harvested TEBVs not only exhibited α-SMA expression as before, but
also synthesized myosin and smoothelin. Vasa vasorum were localized in the
adventitia. In a minor group thrombosis was found. Even so, the possibility of
growing tubes suitable for vascular replacement inside the receivers own body with
the use of minimal invasive surgery has been proven. More research is needed to
understand all processes taking place in these cavities, as well as to evaluate this
procedure in human cavities.

Synthetic protein based polymers

Recently, a new class of polymers based on synthetic proteins, has been described.
For instance, recombinant DNA technology allows the production of a polymer based
on a core sequence of elastin. Furthermore, it has been shown that the degradation
rate of such materials can theoretically be controlled by an incorporated chemical
clock, achieved through the addition of certain chemical components. Also the
attachment characteristics are better controlled by hooking special RGD sequences to
the polymer [124].




                                                                                   27
Endovascular grafts

Similar to the replacement of a diseased or occluded vessel segment, it is theoretically
feasible to insert a’stent-like’ graft inside a diseased vessel [124]. After introduction,
gradual deterioration of the native vessel segment and optimization of the
endovascular graft might occur. However, some aspects remain unclear. For instance,
these grafts need to be particularly thin-walled for fitting into a delivery sheet or
catheter. Also, porosity and tissue reactions might be very different from
conventional graft placing.

3.2.4 Concluding remarks

Although ePTFE and Dacron are still the material of choice for large-diameter grafts,
their application in small-diameter grafts is still problematic. Recent developments
have been such that in the next decade a living vascular graft with controllable,
predictable and desirable characteristics is likely to be constructed. This can be
achieved by culturing blood vessel cells on scaffolds, whether biological or synthetic.
The in vitro development of scaffolds and suitable grafts has proven to be most
encouraging. The correct hemodynamic and biomechanical preconditioning in
combination with possible incorporation of bioactive agents and/or genetic
engineering can lead to an optimal prepared neovessel at the time it needs to be
implanted. At present the development of correct preconditioning protocols is the
main objective of investigation in most laboratories. Even though recent results
indicate a possible role for biological scaffolds, at this moment synthetic scaffolds are
more likely to result in a fast clinically applicable tool to replace autologous native
veins in bypass surgery.




                                                                                      28
Chapter 4: Cell seeded synthetic grafts


The technique of cell-seeded synthetic grafts seems the most promising and
especially in recent years promising results have been achieved by several research
groups. Synthetic materials are preferable due to a higher controllability, e.g. a highly
predictable lot-to-lot uniformity. Synthetic polymers can be further engineered to give
a wider range of properties. These two advantages over natural materials is the reason
why in this review focus will be laid on this methodology. Several aspects that are of
major influence in possible success in TEBV for bypass surgery will be discussed.
The scaffold material at first seems to have the biggest influence, since this defines
the attachibility of cells to the scaffold surface, degradation time in the body (if any)
and development of the new to be formed vessel. Therefore it is important to define
the graft criteria that deserve the most attention before looking at the various
biomaterials that could be used.

4.1 Criteria
A biodegradable graft to be used for the seeding of cells which are subsequently
preconditioned should meet the following criteria:

1. Favourable surface chemistry

2. Sufficient mechanical compatibility

3. Porous structure

4. Commercial potential



4.1.1 Favourable surface chemistry

This seems to be the most obvious criterion and has been addressed extensively in
the previous chapters. In short, the polymer surface should promote endothelial cell
attachment and facilitate growth and proliferation of desirable cell types as SMCs and
fibroblasts. The polymer and its degradation products should not be harmful in any
way, i.e. they should not elicit any form of immunological reaction or inflammation,
leading to such unwanted effects as intimal hyperplasia (an important cause of graft
failure). The polymer is preferably inert and anti-coagulant, making the problem of
incomplete endothelial monolayer formation less problematic.




                                                                                      29
4.1.2 Sufficient mechanical compatibility

Although the range of acceptable mechanical properties of the scaffold material does
not seem to impose large restraints on its design, the importance should not be
underestimated. At first, when cells are seeded onto the scaffold surface and
conditioning is started, the scaffold itself needs to be sufficiently strong to withstand
the imposed conditioning regime. Furthermore, upon graft implantation into the
body, mechanical support is vital. In that case the graft is directly forced to react
correctly to a pulsatile pressure. This means usually the ability to withstand
physiological pressure of 80-120 mmHg and a pulsatile circumferential stretch.
Under physiological conditions the radial elasticity of coronary arteries should
counteract a radial strain of about 6%. Also, since the compliance of a vessel is
essential for successful graft applications, the visco-elastic behavior should be
matched as good as possible. Another aspect that is part of the mechanical
compatibility criterion are the degradation kinetics of the biomaterial. Hypothetically,
the combination of the polymer’s mechanical properties with those of the newly
formed cell/vessel material remains constant over time if the scaffold material
degrades in an ideal way (see figure 10).


                                                          Figure 10: Ideal combination of
                                                          tissue formation and polymer
                                                          degradation of cell-seeded polymeric
                                                          constructs after implantation (figure
                                                          modified from Berglund, 2003).




Slightly modifying this criterion means that the degradation rate has to be such that
tissue integration is optimized and that the strength of the graft remains above a
minimal threshold. This modification is especially necessary when part of the
degradation and neovessel formation has to take place in vivo, which is most likely
the case in future applications.




                                                                                              30
4.1.3 Porous structure

An underestimated feature of scaffold material is its porous structure. Since the ‘60s
it has been realized that a porous structure exhibits a crucial effect on the neovessel
formation. However, still no conclusive reports have been published on the used
structure in terms of optimal porosity, pore size and spatial distribution. It is widely
accepted that a porous inner surface is necessary for the proper anchoring of a
neointima, while a porous outer surface attracts tissue to infiltrate. Recently it has
been demonstrated that a minimum pore size of 20 µm is required for cells in
general to migrate into a scaffold, whereas a pore size of over 60 µm is less favorable
[88]. Hence tissue ingrowth has to be encouraged by the biomaterial through a
sufficiently large pore size. A pore size that is too big bears a higher risk for
inflammation (if exposed to the blood of course, what is generally not intended when
using a TEBV) and is less prone to cellular adhesion. While cellular ingrowth is
promoted by pore sizes between 20 and 60 µm, neovascularization, a desirable event
after graft implantation, is favored by pore sizes around 5 µm [88]. A combination of
these pore sizes does not seem to be feasible, although a gradually increasing or
decreasing pore size from the inside to the outside of a polymer graft is desirable. It
might be concluded that also in terms of porous structure an optimum exists which
has to be determined for every novel scaffold material.

4.1.4 Commercial potential

This fourth and final criterion may seem inappropriate to assess at this stage, since
the primary goal of the whole TEBV approach is to create an improvement in
coronary artery bypass grafting. However, even if a methodology which meets all the
requirements can be developed, its application will depend solely on the industrial
interest. This industrial interest is usually guided by the trivial question: how
profitable can it be? Since acceptable interventions exist, a graft not only needs to lead
to substantial improvement when compared to conventional grafts; it should also be
relatively cheap. A successful outcome is therefore based on the cost-benefit analysis
which – in turn - depends on several factors like processability, off-the-shelf
availability and sterilizability.

4.2 Scaffold materials
The choice of scaffold material is essential when aiming at a cell-seeded TEBV
construct, since the material’s properties determine many critical aspects like
mechanical properties, stability, cell attachment, migration and proliferation and
inflammation. There are two options for polymeric scaffold materials, biodegradable
and non-biodegradable.




                                                                                       31
4.2.1 Non-biodegradable

Dacron (PET)

Polyethyleneterephtalate (PET), better known as Dacron, has been one of the
traditional polymers used for constructing vascular grafts. It is a strong material with
a tensile strength of about 175 MPa and a tensile modulus of about 14 GPa. It is
available in two forms, woven or knitted, the latter having a much higher porosity and
greater distensibility. Impregnation of Dacron knitted grafts with albumin, collagen
or gelatin to seal the pores for prevention against blood leakage is possible [51]. Due to
its relative stability the material was considered to be a good candidate for vascular
reconstruction. The highly crystalline and hydrophobic nature of Dacron both prevent
hydrolysis of a graft, leading to a potential of residing inside the human body for
decades. In large diameter grafting, Dacron has been used extensively and with
considerable result. In small-diameter grafts however, Dacron grafts suffer from fast
occlusion because of high reactivity with blood and vascular tissue, which in turn
leads to inflammation and neo-intimal proliferation. Also the endothelial retention
rate when seeded with ECs was poor, as observed by Turner [114]. Furthermore,
knitted Dacron grafts are prone to dilatation when implanted in an arterial
environment [66]. Nowadays, Dacron is generally considered to be inappropriate for
small-diameter grafting, although surface modifications and possibly coatings could
improve endothelial retention rates and decrease inflammation responses.




   Figure 11: Structural formula of poly ethylene terephtalate (PET).




Gore-Tex (PTFE)

Gore-Tex, or PTFE, is by far the most commonly used material in implants, but in an
extended form (ePTFE). This is mainly due to its excellent biostability and
biocompatibility in vivo. Usually, biological deterioration is absent when ePTFE is
used for grafting. When used in tubular grafts the electronegative surface of ePTFE
minimizes the reaction with blood components. Again, just like in Dacron, the highly
crystalline and hydrophobic nature yields a stable product by preventing hydrolysis.
Tubular grafts made from ePTFE are produced by an extrusion, drawing and
sintering process [43] and consist of fibrils and nodules, controllable to different pore
sizes. Typical material properties are a tensile strength of about 14 MPa and a tensile
modulus of 0.5 GPa.

The controllable pore size is a major benefit in grafting, since this pore size
substantially affects the process of endothelialization. However, still this


                                                                                       32
endothelialization remains a problem. Even the highest levels of endothelialization
are far from sufficient, with levels of only 14 % retention (close to that of Dacron
grafts [114] and the thrombogenecity of the graft prevents high patency rates in small-
diameter grafts.



   Figure 12: Structural formula of poly tetrafluoroethylene (PTFE).


Although used massively in large- and medium-diameter grafts, for small-diameter
surgery, ePTFE grafts seem inappropriate. Still many researchers believe to be able to
modify the surface of ePTFE in such a way that endothelialization becomes easier an
thrombogenecity decreases [5], [111], [124].

Polyurethane (PU)

Polyurethanes comprise a large group of polymers with very diverse characteristics.
Besides the urethane group present in PUs, other groups are usually present. Usually
PUs are copolymers made of three different regions to differentiate mechanical
properties. These regions consist of a hard domain derived from a diisocyanate, a
chain extender and a soft domain, usually polyol. The soft part is responsible for
flexibility, whereas the hard region imparts strength. Good elastic and compliant
properties and acceptable biocompatibility, together with relatively easy possibilities
for modification makes PU an extremely attractive material for vascular grafts.

PUs have a wide range of mechanical properties due to the presence of soft, tough
elastomers and strong, rigid polymers. When PUs are considered as permanent
implants, possible in vivo deterioration is present because of hydrolytical instability
and oxidative degradation of the soft segment. Potential carcinogenic effect of its
degradation products come into play here, implying that care has to be taken before
using PUs in clinical trials. Relatively hard polymers will not suffer from this
problem much though. Recently, oxidatively and hydrolytically stable polycarbonate-
based PUs have been proposed for vascular grafts. In a non-woven PU graft it was
demonstrated that endothelialization, early stabilization of neo-intimal proliferation
and a neo-intima thickness was better than in ePTFE grafts [36]. At this very moment
this type of graft is undergoing clinical trials.

Yet it remains undecided whether PUs of any kind are favorable when compared to
Dacron and ePTFE although their mechanical properties are generally superior to
that of Dacron and ePTFE. Optimal porosity can also be achieved and controlled in
certain PUs [125]. The largest problem remains the potential instability and
carcinogenicity of the PU debris.




                                                                                    33
4.2.2 Biodegradable

Of course many biodegradable polymers have been tested for their potential in
vascular scaffolding, but most of them have been abandoned in an early stage of the
development process. Some however continue to be at interest of researchers.

PCL

Poly ε-caprolactone (PCL) is a semi-crystalline polymer, synthesized following ring-
opening polymerization of the monoester. Since it is tissue compatible, PCL has
achieved a tremendous success as a biodegradable suture. Recently it also attracted
the attention of investigators for use as vascular scaffold, since it exhibits a better
elasticity than PGA or PLLA. The elastomeric properties required for certain tissues
are not present though. Recently Serrano and colleagues [97] demonstrated good
adhesion, growth, viability, morphology and mitochondrial activity of adhered
endothelial cells, indicating potential for vascular tissue engineering. A disadvantage
of PCL is the degradation time, which is in the order of two years.




 Figure 13: Structural formula of poly-ε-caprolactone (PCL).


PLLA

Polylactic acid (PLLA) is another potential scaffold material. The L-homopolymer is a
semi crystalline, natural occurring hydrolytic product, and has a high tensile strength
and low elongation characteristics. This makes PLLA more suitable for load-bearing
applications such as in the cardiovascular system. PLLA is also slightly hydrophobic,
so that it is more soluble than PGA in organic solvents and thus easier to process.
Also the hydrolytic degradation rate is low due this hydrophobic nature. The typical
degradation period of PLLA is in the order of two years.

P4HB

A new biomaterial is poly-4-hydroxybutyrate (P4HB) [103]. P4HB is not only a
naturally occurring polymer, but also easily synthesized in vitro. It is a substance
belonging to the group of polyhydroxy-alkanoates and is characterized by a rough,
porous surface, ideal for cellular attachment. Its pore sizes can be varied from 80-
400 nm, a range well suited for different TE applications. The degradation rate in
vivo is in the order of months, making it exceptionally attractive for vascular TE. As
described only recently, the potentials of P4HB are largely unknown. When used as a
coating P4HB yielded good results [29].




                                                                                    34
   Figure 14: Structural formula of polyhydroxy butyrate (PHB).
Dexon (PGA)

Dexon or poly-glycolic acid (PGA) has been used for decades as a suture material. It is
highly crystalline in nature, and is insoluble in most organic solvents. Furthermore,
the porosity of PGA scaffolds can be up as high as 90%, so that cellular ingrowth can
easily be promoted. After implantation PGA scaffolds tend to loose their mechanical
strength over a period of only 2-4 weeks and are completely absorbed after 4-6
months, due to their hydrophilic nature. On PGA grafts implanted in rabbits a
confluent layer of ECs and myofibroblasts (MFs) was formed within 4 weeks. Within
one year 10 % of the implanted grafts showed aneurysmal dilatation [26].

 The implantation of PGA is associated with a considerable inflammatory response, a
problem which should be considered when using PGA in small-diameter
reconstructed vessels.




 Figure 15: Structural formula of polyflycolic acid


Hyaff-11

A novel material is Hyaff-11, an esterified form of hyaluronan. Hyaff-11 is a highly
hydrophobic polymer which can easily be shaped into a 3-dimensional scaffold by
weaving or spinning. Hyaff-11 does not provoke inflammation upon degradation in
contrast to other materials like PGA and Dacron. Furthermore, upon hydrolysis the
polymer becomes more and more hydrophilic, eventually turning into a gel similar to
native hyaluronan. In about 2 months Hyaff-11 is completely degraded, an excellent
time-frame for tissue engineering applications. Although still in an early stage of
investigation, Turner and co-workers [114] have demonstrated that Hyaff-11 supports
attachment, proliferation and migration of ECs through the scaffold. The latter
occurred when the fabric was not compressed, resulting in pore sizes of about 50-
200 µm. When these scaffolds are compressed cells cannot penetrate, so that it is
more suitable for the formation of an endothelial monolayer. Also, the degradation
time of compressed scaffolds is longer than that of the not compressed scaffolds
because of the production of more ECM, which holds the fibers together. EC
retention and proliferation rates were excellent with over 1.5 times the initial seeding
concentration after 10 days, making it an attractive material to investigate more
extensively in the future.

Copolymers

As some homopolymers do lack some required features for application in vascular
grafting, the problem can be overcome through the combination of different types of
homopolymers.



                                                                                     35
A copolymer of glycolide and PCL, named PGCL was first used by Lee and coworkers
[43]. PGCL has a degradation time appropriate for vascular grafting, namely in the
order of a few months. Lee showed that PGCL is very elastic in nature, making it
interesting to achieve a desirable compliance in tubular constructs.

Another copolymer, made out of a combination of PLLA and PCL (ratio 50:50) has
been investigated [57,123,69,35] and even applied clinically for the reconstruction of
a part of the pulmonary artery. At 8 weeks after implantation an excellent patency
was observed [27].

A great variety of other copolymers have been investigated of which combinations of
PLLA and PU and PGA and PLLA were most popular, but still no ultimate blend has
been found.

4.2.3 Surface modifications

In spite of the wide range of biopolymers that has been used or experimented with to
improve biocompatibility, mechanical structure and processability, cellular adhesion
rate is still poor. Cells seeded onto the lumen of polymer grafts, especially endothelial
cells, generally adhere poorly to the surfaces of all above-mentioned polymers.
Furthermore, upon implantation, the endothelial cells are exposed to a pulsatile flow,
which causes endothelial cell losses of up to 70 %. Cellular retention rates need to be
as high as possible, since every lost EC will increase the risk of thrombogenic effects.
Methods used so far for improving this retention include the preconditioning of the
seeded scaffold with an increasing (pulsatile) flow and electrostatic charging.
However, the best results have been made with surface modifications of the graft to
improve EC adhesion and retention. These surface modifications can consist of a
variety of methods, like the coating of a graft before seeding with a substance that
promotes EC-graft adhesion forces, preclotting the graft or treatment of its surface
[122].

Coatings

A truly vast variety of coating substances have been used by TE groups, with the
number continuously increasing. However, a short summary of the most commonly
used chemical coatings includes mainly well-known natural materials. Popular for
example is the coating of polymer grafts with collagen and fibronectin, but also
laminin, gelatin and other ECM components. In the case of collagen, the material can
provide a fibrous matrix and cell-attachment promoting matrix proteins. Laminin, an
ECM glycoprotein that regulates a variety of tissue processes, has been inconsistently
found to enhance adhesion of ECs. Another large natural-occurring molecule, gelatin,
also yielded various different outcomes. The application of not only a part of the ECM
like collagen, but the complete ECM has been applied by Kidd and co-workers [89],
who found a much higher patency in small-diameter ePTFE grafts precoated with
ECM.




                                                                                      36
Fibronectin was the most successful peptide up till now. This glycoprotein is
synthesized by many cells, including ECs, and it is thought to facilitate attachment
and expansion up to confluency of ECs . Furthermore, the effect of other coatings like
collagen and gelatin will be enhanced when combined with fibronectin.

Preclotting

It is thought that cellular attachment can be improved by preclotting the graft. To
preclot the vessel, one uses the patient’s own plasma, blood, serum or a fibrin glue
[39]. The effects of these substances have been compared with coatings like
fibronectin and laminin, with varying results. The use of fibrin glue as a kind of
embedding substance has yielded very promising results when used on ePTFE and
implanted in the infrainguinal position, giving evidence for usage in small-diameter
grafts.

Chemical bondings

Chemical bonding of such surface molecules as peptides provides a more or less of-
the-shelf availability of vascular grafts. This method yields promising results. The
RGD-sequence for example is a recognition site used by ECs to attach to the ECM.
RGD used in chemical bonding has provided very good results, comparable to those
of fibronectin. Also, RGD enhances DNA activity of ECs, and thus promotes
endothelial function [30]. Heparin has also been investigated and shown to enhance
cellular adhesion [39].

Surface treatment

Besides modifying a graft by adding a substance, it is also possible to physically alter
graft surfaces by treating the top layer only. Instead of altering the entire graft
material, the surface can be modified by using irradiation or plasma treatment.
Bacakova and coworkers found promising results after fluorine ion radiation of
polystyrene [81]. Another possibility is to change surface properties by treating it with
certain gases like argon or even oxygen [76].

4.2.4 Concluding remarks

Several materials such as P4HB or copolymers of PCL, PLLA, PGA and/or PUs seem
to be appropriate, but at present no ideal substance can be identified. Besides the
convential polymers, other, new materials have been discovered and found to be
potentially suitable in arterial reconstruction [34]. Furthermore, surface modifications
can enhance graft performance through alternation of their boundary structure and
character. The search for an ideal scaffold material will probably continue for many
years, although it is quite probable that several materials are suitable for usage,
depending on the intended application.




                                                                                      37
4.3 Production methods
Good scaffolds should allow and promote cells to attach, proliferate and in some
cases migrate. Success not only depends on a proper selection of the scaffold material,
but also on the processing technique to createa 3-dimensional construct. The
processing technique should certainly not decrease the biocompatibility of the
material. Ideally, the technique should allow to control the porous structure and to
yield reproducible scaffolds. A variety of methodologies have been developed to fit the
intended applications.

4.3.1 Fiber bonding

To promote structural stability of a scaffold two ways of fiber bonding have been
developed. These techniques have mainly been applied to bonding of PGA fibers. The
first method involves casting a polymer solution (e.g. PLLA) over a PGA mesh and
dissolving the solvent afterwards. Subsequently, the PGA fibers embedded in the
PLLA matrix is heated for a given period of time, resulting in bonding of the PGA
fibers. The PLLA matrix is required to ascertain a fiber-like structure of the PGA by
confinement and to prevent the collapse of the PGA mesh [55]. Problems involved in
this technique are choice of solvent, immiscibility of the polymers and lack of control
of porosity.

The second technique is a coating technique. By spraying a PGA mesh with a
polymer solution (e.g. PLLA), a layer of PLLA is deposited on the PGA, thereby
interconnecting PGA fibers [58]. This methodology has successfully been applied to
create hollow tubes and is therefore interesting when vascular constructs are
concerned.

4.3.2 Solvent casting and particulate leaching

A new technique to overcome some of the drawbacks of the fiber bonding techniques
was developed by Mikos in 1993 [55]. With this method, porosity, crystallinity and the
surface:volume ratio is controllable.

Briefly, sieved slat particles are dispersed in a polymer solution and cast into a glass
container, after which the solvent is left to evaporate. The salt is then removed from
the resulting polymer/salt composite by either leaching out the salt by placing it in
water or by first heating the composite and subsequent annealing by cooling at a slow
pace and then leaching out the salt.

The porosity of these membranes can be controlled by choosing the appropriate size
of the salt particles and the amount of salt used. The crystallinity can be controlled by
using the second method to remove the salt from the composite. One major
disadvantage of this methodology however, is the fact that only thin wafers of
membranes can be produced. By polyethyleneglycol incorporation the material can be
made more flexible [119], so that it can be rolled up to form a tubular construct which


                                                                                      38
still bears the same porous features. The mechanical characteristics of the material
however, are inherently changed.

Another concern is the seeding of cells onto this material. The material may be
highly hydrophobic, leading to bad attachment of cells. This can be circumvented by
pretreating the material by soaking in PVA [58] or by prewetting the material through
soaking it ethanol for one hour.

4.3.3 Membrane lamination

This relatively simple technique is used for the creation of three-dimensional
scaffolds from thin membranes which are made by using the particulate leaching
technique. Membranes are basically glued on top of each other by coating a small
quantity of chloroform (or any solvent appropriate) on the boundary layer. The top
part of each layer will be solved shortly and will solidify again, hereby connecting the
individual membranes. By cutting the membranes, the layer-by-layer fabrication
results in any 3-dimensional shape desired. However, this technique is only possible
when the porous structure of the original membrane is preserved.

4.3.4 Melt molding

This technique offers many advantages over the membrane lamination method. A
mixture of the desired polymer with (sieved) gelatin or salt particles is poured into a
Teflon mold in the desired shape (e.g. tubular). By heating the mold with polymer
composite above glass temperature, a porous structure is formed. After removal of
the salt or gelatin by placing the construct in water, a polymer structure the shape of
the mold is left [109]. Likewise to the particulate leaching method, porosity of the
structure can be controlled by varying the particle size and amount.

4.3.5 Three-dimensional printing

Three-dimensional printing aims at creating a shape similar to the tissue that needs
to be replaced. The three-dimensional printing (3DP) technique belongs to the group
of the solid free-form techniques [72]. In short, in 3DP a thin layer of polymer powder
is spread over a piston. Then, an inkjet printer (or other suitable printing device)
prints a liquid binder (‘glue’) like chloroform onto the powder layer. The piston is
then lowered and a new powder layer is dropped onto the former one and the process
repeats itself. Powder spraying and ‘glue’ printing can be well controlled so that 3-
dimensional constructs can be fabricated very precisely. Again, incorporation of salt
in the powder again is possible to achieve porous constructs.




                                                                                     39
4.3.6 Electrospinning

Electrospinning (ELSP) is a rather novel technique to produce fibers of diameter
ranging from micrometers to as small as 100 nm under a high-voltage electrostatic
field operated between the metal tip of a syringe and a metallic collector (see figure
16). A polymer solution is deposited randomly by a projected jet from the needle onto
a metal collector. The metal capillary is charged to induce the electrostatic field [73].
The physical properties of the fibers can easily be controlled by changing the tip
diameter, the distance between tip and collector, concentration of the polymer
solution and the flow rate of the solution into the needle tip. Translation, rotation of
the needle and/or collector can be used to control the shape of the constructs.

In this way the construction of tubular constructs of any kind of polymer (when a
suitable solvent is available) is possible. Recently, even the electrospinning of vascular
polymers like collagen has been investigated [53]. Novel methods, such as alternating
solutions or combining multiple needles with different polymer solutions at the same
time [37], provide means to use mixed constructs and layered tubular constructs.




 Figure 16: schematic representation of the electrospinning process. The target substrate can be any form, e.g.
 a sheet, tubular mandrel or even more complex shapes like valvular.



In conclusion, at present various techniques are available for the fabrication of
polymer scaffolds, although no ideal method can be identified. Each technique has
advantages and drawbacks, leading to different choices of method based on
intentions. Otherwise, for the fabrication of porous tubular constructs many
techniques can be abandoned. In fact, only 3-DP, melt molding and electrospinning
seem to be appropriate. Electrospinning and 3DP seem to be the most promising
techniques in incorporating superior controllability of the constructs’ characteristics.
In terms of handling and costs of methods, electrospinning seems to offer more
potential. The printing technique might turn out to yield better results in terms of
reproducible constructs.


                                                                                                                  40
4.4 Endothelial cell seeding
From the former chapters it becomes clear that grafts which are not lined with ECs or
which are ineffective in attracting ECs are prone to rapid failure after implantation.
Hence it is very likely that more and more research groups will shift their attention to
the EC-seeded graft, whether synthetic or biological. If this is the case, one of the first
aspects to consider is the source of the ECs to be used.

4.4.1 Choice of cell source

The source of cells to be seeded onto or into synthetic grafts is not yet unequivocally
established [62]. In table 4 is summarized which cell sources have been used the last
8 years in experiments aimed at vascular grafting.

It can be appreciated that a wide variety of animal cell sources are in use for vascular
grafting Most researchers however are inclined to use human-derived cells, which
appears a logical choice since ultimately the grafts should be implanted in humans.
Also, it is now known that for example the luminal recruitment of ECs occurs must
easier and leads to a faster monolayer formation when ECs of animal origin are
involved, emphasizing the essence of using human-derived cells. Especially HUVECs
are popular, probably because of the relative ease to obtain and culture them. Other
investigators prefer stem cells, which have the potential to differentiate into any kind
of cell type. Therefore, they offer a very attractive option. Some ethical aspects are
concerned with their use, not to mention the lack of knowledge of the correct
protocols for triggering differentiation towards a certain cell type.

Healthy ECs maintain a delicate balance in the vasculature by regulating growth
promotion and inhibition, vasoactivity, coagulation and cell adherence. Although
these mechanisms are controlled by the endothelium in general, one has to keep in
mind that all ECs do this in the same way. It is acknowledged that ECs are
heterogeneous and might vary within the species [33] , but also within the individual
subject, in phenotypical appearance, function and response to influences like growth
factors and mechanical stimulation. (see figure 16)




                                                                                        41
                                                   Figure 17: Aspects of endothelial cells that might
                                                   vary among species and harvesting location.




The differences that might occur between endothelial cells from the same species can
be divided in several subcategories and were thoroughly reviewed by Thorin and
Shreeve in 1998 and Aird in 2003 and will be shortly summarized hereafter [110,2].

Phenotypic appearance

Within one species the shape of ECs can vary from cobblestone to round, spindle-
shaped or polymorphic, depending on location in the vasculature. The size of
endothelial cells can differ significantly from vessel to vessel, as found in pig aorta
and coronary artery cells.

Proliferative capacity

In culture the proliferative capacity of ECs may vary, depending on the in situ
location of the harvested cells and the age of the donor. The proliferative capacity is
essential for several vital mechanisms involved in graft patency like
neovascularization of the TEBV. For example the proliferative capacity of ECs from
venules and capillaries was shown to be indispensable for vascularization of
transplanted limbs, while the endothelium of larger arteries did not participate in this
process.




                                                                                            42
 Table 4: Overview of used cell source in recent vascular graft related research.




First author     Year    Graft material              Cell source           Graft form   Abbreviation list
Kleinert         1996    ePTFE                       CMVEC                 tube
Imbert           1997    PET                         HUVEC                 Patch
Kawamoto         1997    SPU                         BAEC                  tube         Animal
Dunkern          1998    PTFE                        HUVEC                 Tube         OVEC          Ovine vascular endothelial cells
Niklason         1999    PGA                         BAEC                  Tube         MES           Mouse embryonic stem cells
Steinhoff        2000    Decellularized artery       HVEC                  Tube         CEPC          Canine endothelial progenitor cells
Kaibara          2000    SPU                         BAEC                  Tube         PJVEC         Porcine jugular vein endothelial cells
Salacinski       2000    CPU                         HUVEC                 Tube         RAEC          Rat aortic endothelial cells
Teebken          2000    decellularized artery       HVEC                  Tube         CJVEC         Canine jugular vein endothelial cells
Fernandez        2001    PET                         HUVEC                 Tube         CVEC          Canine saphenous vein endothelial cells
Hoerstrup        2001    PGA+P4HB                    OVEC                  Tube         OCEC          Ovine carotid endothelial cells
Meinhart         2001    ePTFE                       HCVEC                 Tube         BAEC          Bovine aortic endothelial cells
Salacinski       2001    PU                          HMVEC                 Tube         CMVEC         Canine microvascular endothelial cells
Arts             2002    ePTFE                       CMVEC                 Tube
He               2002    (S)PU                       CJVEC                 Tube         Human
Hodde            2002    SIS                         HMVEC                 Tube         HUVEC         Human umbilical vein endothelial cells
Krijgsman        2002    CPU                         HUVEC                 Patch        HCVEC         Human cephalic vein endothelial cells
Kumar            2002    PTFE+PET                    HUVEC                 Patch        HUCEC         Human umbilical cord endothelial cells
Pu               2002    PET+PTFE                    HUVEC                 Patch        HMVEC         Human microvascular endothelial cells
Tiwari           2002    PU                          HUVEC                 Tube         HVEC          Human saphenous vein endothelial cells
Xiong            2002    PGA+PLA+PHB                 BAEC                  Patch        HJVEC         Human jugular vein endothelial cells
Conklin          2003    decellularized artery       HMVEC+HEPC            Tube         HUAEC         Human umbilical artery endothelial cells
Deutsch          2003    PET                         HUVEC                 Tube         HAEC          Human aortic endothelial cells
He               2003    SPU                         CEPC                  Tube         HEPC          Human endothelial progenitor cells
He               2003    SPU                         CEPC                  Tube
Hsu              2003    PU                          HUVEC+HUAEC           Tube
Liu              2003    PGA                         HUCEC                 Tube
Miller           2003    PLGA                        RAEC                  Patch
Pan              2003    PGA+collagen                Rabbit aortic EC      Tube
Shen             2003    PGA                         MES                   Tube
Shirota          2003    SPU                         HEPC                  Tube
Tiwari           2003    PU                          HUVEC                 Tube
Iwai             2004    PLGA                        CVEC                  Patch
Koike            2004    collagen                    HUVEC                 Tube
Mall             2004    ePTFE                       PJVEC                 Tube
Opitz            2004    P4HB                        OCEC                  Tube
Turner           2004    Hyaff-11                    HVEC                  Patch
Zhu              2004    PU                          HUVEC                 Tube




                                                                                                                               43
Antigenic variety

Since the endothelium actively contributes to inflammatory and immunological
processes, it is clear that ECs used in grafts should have close to ideal properties to
prevent occlusion of the artery. A heterogeneous distribution of vascular adhesion
molecules in endothelium from different regions of the human cardiovascular
system shows that care has to be taken when choosing the correct graft cells [71].

Function

A critical aspect in determining the right EC-source is the endothelial function.
Within species, researchers found differences in the basal release of vasoactive
factors, flow-dependent responses, agonist-induced endothelium-mediated
vasoactivity and receptor-mediated responses. Also on pathological conditions such as
hypertension the response of ECs differs, depending on their location.

The above aspects may be correlated in many ways. For example, the functional
heterogeneity might be correlated with morphological and antigen expression
heterogeneity. It is thus possible that all ECs, when placed and cultured in exactly
identical environments, will ultimately transform into cells with identical appearance
and functionality. Although this would solve many issues associated with the TE
business, regretfully no group has made such observations so far. Possibly only
endothelial progenitor cells can be stimulated to differentiate along several pathway
and ECs loose most of the ability to change their functionality because of their
differentiated state.

All of these factors are to some extent responsible for the patency ability of a TEBV
and all lead to an obvious conclusion. At present when seeding a graft with ECs to
generate a TEBV for vessel replacement, it seems almost inevitable to use cells that by
their very nature express identical properties as the proper vessel cells. This directly
implies that for the use in a coronary bypass TEBV, ECs from a human coronary
artery should be used (HCAECs). Of course further research will elucidate the true
necessity of the latter, since it induces quite some restraints, especially in terms of
isolation and cultivation of this type of cells. For these reasons HCAECs have been
scarcely used in experiments involving EC seeding, and not at all with respect to
TEBVs. Hopefully it will be proven that ECs from other locations can provide similar
patency opportunities as native ones.

With respect to the source of the smooth muscle cell some aspects have to be
mentioned. Mostly actions of the ECs depend on the SMCs or other cells they
communicate with. However, no true evidence exists that SMCs themselves show
signs of heterogeneity. The lack of evidence so far does not imply non-existence and
intensive research might result in new insights in this matter. In fact, it seems more
likely that also for SMCs some heterogeneity exists, albeit only between different


                                                                                     44
species or maybe on just a negligible level within the same species. Therefore it is
recommended to use cells harvested from a human source at least. For investigation
of endothelium ‘in vitro’, closely modelling the ‘in vivo’ status of the majority of the
blood vessels, ECs derived from a not too specialized vasculature is superior. Yet, no
established superior models exist for EC research ‘in vitro’ and as the interest in
blood vessels is still growing, we can expect a flow of new EC lines with increasingly
beneficial properties in the near future.

As said, a novel solution is the use of stem cells [98,100,12,52]. These cells possess
the ability to differentiate into several cell types, including SMCs and ECs, if properly
stimulated by mechanical forces and/or chemical influences. At present it is still
unclear how stem cells should be stimulated to trigger correct development.
Furthermore, the use of stem cells is still surrounded by ethical issues, making it
sub-optimal and impractical. In future though stem cell technology will be further
developed, and in turn lead to better control and harvesting opportunities.

4.4.2 Seeding of the scaffold

The endothelial cell seeding process deserves some attention. Clearly it can be
concluded that an endothelial cell lining is an absolute prerequisite for a successful
TEBV. However, the capability of a graft to acquire a confluent monolayer with a
good viability factor is based on the cell seeding process. EC concentration, cell
source and method of seeding all have a substantial impact on the resulting final
layer; hence they should be selected carefully. Finally, in vitro mechanical
conditioning of TEBVs seems to enhance several properties of TEBVs that are
required for successful implantation.

In general, in order to engineer a neovessel, viable cells, whether SMCs, ECs,
fibroblasts or others, have to be seeded onto or into a three-dimensional structure and
further cultivated in a suitable bioreactor. Basic requirements for the seeding of cells
include a high yield to maximize cell utilization, a spatial uniform distribution of
attached cells and a high kinetic rate to minimize the time in suspension. Some
general guidelines have been developed for the seeding of different cell types onto
scaffold material. However, this is a somewhat artificial standard, since cell
attachment rate, survival rate and cultivation characteristics all depend on the source
of the cell line, type of scaffold material as well as on the cultivation conditions and
the seeding technique.

In this chapter the currently applied techniques for the seeding of cells onto
polymeric scaffolds will be briefly summarized [73].

Gravitational seeding

The most basic technique used for seeding cells on the luminal surface of a vascular
graft involves the use of plain gravitational forces. In general it means that cells
suspended in seeding medium are delivered to a horizontally positioned graft, after


                                                                                      45
which the graft is rotated periodically or continuously over a prolonged seeding
period. Over the years many variations to this concept were introduced including the
use of a special coating of the graft surface to promote cell attachment and the
application of slowly evolving hemodynamic forces. Recently it has been shown that
seeding of a polymeric scaffold by suspending cells in an already flowing medium
might promote cellular adhesion more than static seeding. This is a more logical
paradigm when the physiological circulation characteristics in the human body in
which floating ECs adhere quite easily is kept in mind. Possibly the small shear
forces induced by a small flow triggers certain adhesion molecules, but this is yet
undetermined.

Hydrostatic seeding

Hydrostatic seeding is based on a pressure difference, either by applying an internal
pressure or by an external vacuum across the graft wall to force (endothelial) cells
onto the luminal surface of the graft. Although simple, some disadvantages are
associated to this technique. First of all, during the seeding process the graft pores
have to be kept open, while they have to be sealed off by a preclotting process prior to
implantation. Secondly, after the preclotting procedure the surface of the graft may
be inherently rough and prothrombotic, leading to possible thrombosis upon
implantation. Finally, the cells are merely trapped at the surface by the seeding
process instead of actually attached to it, making them prone to detachment when
exposed to physiological shear forces.

Biological glue

A biological coating of the luminal surface to promote cell adhesion has been
recognized to be beneficial. Adhesive ECM proteins like fibronectin or collagen can
be seen as biological glue proteins because they help in attaching cells to the inner
wall of the graft. The problem associated with this technique is the thrombogeneicity
of non-endothelialized spots on the inner wall.

Electrostatic seeding

Electrostatic seeding is a novel method that might prove to be more efficient when it
comes to EC attachment. This methodology is based upon the assumption that
negatively charged polymers exhibit a repellent force on negatively charged ECs. In
practice a temporary positive field is applied to the surface of a normally negatively
charged polymer. The first results show that this technique is superior over
gravitational and hydrostatic seeding in terms of time needed for cell attachment and
number of attached cell, as well as of cellular retention. Upon implantation of such
grafts possible incomplete endothelialization does not necessarily provoke
thrombogenic spots. Possible positive effects on other adhesion stimulating factors
like cytoskeleton rearrangement have to be investigated before the true significance
of this technique can be confirmed.




                                                                                     46
4.4.3 Characterization of cells

Seeding ECs on the surface of an intended TEBV requires the characterization of the
endothelial cell layer in order to assess the success of a specific method. ECs have
specific markers, many of which can be used to characterize them in a fast and
relatively simple way. It would go too far to discuss every cell-specific marker in detail.
Therefore, only the most commonly used will be pointed out, as well as some basic
ideas regarding the characterization process.

Generally spoken, ECs can be characterized by phenotype and/or function. If one
chooses human ECs for scaffold seeding the requirements should include a cell
lifespan according to the duration of the experiment, a stable expression of at least
some markers like Weibel-Palade(WP) bodies, von Willebrand factor (vWF), acLDL
uptake, cobblestone morphology, ACE activity and PECAM-I expression. Last but not
least as few as possible indications for tumorogenic conversion should be apparent.

Best-known markers in these categories are the following:

Phenotypical markers

Markers related to the EC phenotype are defined as molecules expressed by the EC as
well as physical characteristics present in or on the cell. Of all the physical features
only a few are needed to distinguish ECs from other cells. One of them are WP
bodies which are present in all ECs, regardless their origin. In granules inside the
WP bodies, P-selectin and the von Willebrand factor (vWF) are colocalized. WP
bodies store vWF, which binds and stabilizes factor VIII and is a cofactor for platelet
binding in damaged vessel walls. Thrombomodulin (TM), an antithrombotic factor
also known as CD141, and VE-cadherin, an adhesive protein also known as CD144,
are two other, exclusively endothelial markers. Further EC phenotypic features
include the production of prostacyclin (PGI2) and the expression of platelet cellular
adhesion molecule-I (PECAM-I), also known as CD31. Although frequently used the
latter two markers are not unique for ECs.

Functional markers

Functional markers of endothelial cells are defined as cellular processes that can be
monitored to ascertain a true endothelial nature of the observed cells. Out of the
many processes in endothelial cells, three processes are particularly useful in
distinguishing ECs. These are the binding of Ulex Europaeus Agluttin-I (UEA-I), the
activity of the Angiotensin Converting Enzyme (ACE), and the uptake of acetylated
LDL (acLDL). The latter two are correlated and not restricted to ECs though. See table
5 for a more complete overview of endothelial markers [23].

In graft research the most frequently used cell-specific markers are vWF/factor VIII,
PECAM-I (CD31), VE-cadherin and UEA-I binding [80]. Important with respect to the
coagulation process is the expression of vWF and of thrombomodulin (TM).


                                                                                       47
       Table 5: Overview of endothelial cell markers.


                                                  Species2                      Cell Type3
Factor VIII–related antigen                       h/m                           ECs (irregularly expressed by capillaries and tumor vessels), platelets, megakaryocytes
CD31/PECAM-1                                      h/m                           ECs, platelets, megakaryocytes, B and T lymphocyte subsets, monocytes, neutrophils
Angiotensin-converting enzyme                     h/m                           ECs, epithelial cells, monocyte-macrophages, T lymphocytes
Type I scavenger receptor (acetylated-LDL uptake) h/m                           ECs, macrophages, SMCs, pericytes, fibroblasts
Ulex europaeus I agglutinin binding/O(H) blood-   h                             ECs, erythrocytes
type antigen
Bandeirea simplicifolia lectin binding            m                             ECs
Griffonia simplicifolia agglutinin binding        m                             ECs
Weibal-Palade bodies                              h/m                           ECs
Vascular endothelial cadherin                     h/m                           ECs, trophoblasts, PLN sinus macrophages
CD34                                              h/m                           ECs, hemopoietic precursors
CD102/ICAM-2                                      h/m                           ECs, lymphocytes, monocytes, platelets
CD36                                              h                             Microvascular ECs, monocyte-macrophages, erythroid cells, platelets, megakaryocytes
CD73/VAP-2                                        h                             ECs, T and B lymphocytes, tonsillar epithelium
S-ENDO 1/MUC18                                    h/m                           ECs, SMCs, dendritic cells, leukocytes, melanoma cells, carcinoma cells
Thrombomodulin                                    h/m                           ECs, SMCs
HEMCAM                                            m                             Microvascular ECs, hemopoietic progenitors
Sca-1                                             m                             ECs, hemopoietic precursors
AAMP                                              h                             ECs, cytotrophoblasts, mononuclear inflammatory cells, melanoma cells,
                                                                                adenocarcinoma cells
h indicates human; m, murine; AAMP, angio-associated migratory cell protein; SMC, smooth muscle cell, and PLN, peripheral lymph nodes.

1
 Endothelial markers are those most commonly used and/or relatively restricted to the endothelial lineage. Most of them are not expressed by all kinds of vessels or in all tissues; markers
ubiquitously expressed are marked as EC.



                                                                                                                                                                               48
For example, using immunohistochemical techniques, Shen et al. [98] investigated
the expression of vWF and CD34, while L’Heureux [42] studied the expression of
vWF and the uptake of acLDL. Immunofluorescent staining can be done by labeling
cells with anti-vWF, anti-CD31, anti-CD34, anti-VE-cadherin or anti-TM. The latter
three are used for assessing the monolayer integrity. Flow cytometry can be used for
analyzing thrombomodulin.

To be complete, it should be mentioned that the intense communication of ECs with
SMCs makes it desirable to distinguish both cell types as clearly as possible.
Therefore it is important that also SMCs can be characterized. Most common
markers for SMCs include calponin, myosin heavy chain, vimentin, desmin and α-
smooth muscle actin (α-SMA). Especially α-SMA is a powerful indicator, though also
desmin and calponin are used frequently for SMC identification.

4.5 Preconditioning and dynamical loading
It is well known that cells respond to their environment through adaptation and
remodeling, which can be triggered by mechanical and chemical stimuli. Mechanical
stimuli are necessary for a correct development of a tissue towards a mechanically
stable, well-defined structure. Initially, cell-seeded scaffolds lack these properties.
Usually a scaffold can be defined quite well, but the seeded cells are randomly
oriented and positioned, leading to lack of mechanical strength sufficient for bearing
hemodynamic loads in vivo. This has led to the acknowledgement of a mechanical
loading protocol to be imposed on the TEBV prior to implantation [89]. This
mechanical protocol can vary from simple quasi-static flow via static pressure to cyclic
radial straining. Even pulsatile flow can be envisaged to mimic physiological
conditions.

4.5.1 Mechanical protocols

Although it is appreciated that a dynamical environment has a large influence on
cellular behavior, investigators have struggled to find the right mechanical loading
conditions. For instance, it seems to be logical to apply a physiological pulsatile flow
to coronary artery grafs. However, when such a load is imposed to a construct with
randomly oriented cells, cell detachment and thus deterioration of the endothelial
monolayer occurs. Maybe the different cell types in a graft require varying loading
regimes until that specific cell type reaches the correct status. For example, when
SMCs are seeded with a layer ECs on top, a pulsatile flow might be optimal for the
ECs, but inadequate for the underlying disoriented SMCs. It is this type of
consideration that has driven researchers to perform experiments with different
loading conditions. An overview of recent publications in this domain is presented in
table 6. It becomes clear that results of different loading regimes include almost
exclusively effects on cell retention percentages and proliferation rates. To the best of
our knowledge only six publications report the effect of a dynamical load on the
mechanical properties of the construct.


                                                                                      49
     Table 6: Overview of applied conditioning protocols with respect to vascular grafting.

Author               Cell source       Mechanical load                 Loading parameters             Graft             Diameter (mm)   Outcome
Ott (1995)           BAEC              shear stress                    day 1-3: 1 -> 2 dynes/cm2      PU                      1.5       Better retention when preconditioned
                                                                       day 4-6: 25 dynes/cm2
Braddon (2002)       HAEC              shear stress                    10 or 30 dynes/cm2             PGLA                   sheet      Increasing expression of ICAM-I
                     HDF                                               6, 24, 48 hr                                                     Higher expression with 30 dynes/cm2
Carnagey (2003)      HUVEC             shear stress                    a) 300 ml.min 2 hr             PTFE                     4        Gradual flow build-up increases cell retention 20-40 %
                                                                       b) increasing to 300 ml/min    PhotoFix OCA             4        Better cell retention on PhotoFIx grafts
                                                                       in 8 hr (6 dynes/cm2)
                                                                       c) phys. 68 ml/min for 24 hr
Imberti (2002)       HASMC             preconditioning cyclic strain   10 % radial strain             collagen+HASMC           3        Shear+strain both reduce EC proliferation
                                                                       1 Hz                                                             Pre-straining reduces EC proliferation
                                                                       3 days                                                           Effect is greater when flow parallel to collagen fibers
                     HAEC              shear stress                    10 dynes/cm2                   collagen +HASMC          3
                                                                       2 days
Tiwari (2001)        HUVEC             Pulsatile flow                  120/60 mmHg                    PU                      6         EC retention >65 %
                                                                       229 ml/min
                                                                       7.5 dynes/cm3
                                                                       6 hour

Salacinski (2000)    HUVEC             pulsatile flow                  209 ml/min                     CPU                      5        70 % of Ecs stayed attached
                                                                       7.5 dynes/cm2
                                                                       1 Hz
                                                                       120/60 mmHg
                                                                       6 hr
Fernandez (2001)     HUVEC             pulsatile flow                  340 ml/min                     PET                     6         87 % retention
                                                                       130/60 mmHg
                                                                       0.27 N/m2
                                                                       3 hr
Hoerstrup (2001)     OVEC              Increasing pulsatile flow       125 ml/min -> 750 ml/min       PGA+P4HB                 5        Increasing burst strength to 326 mmHg
                     OVMF                                              30 mmHg -> 55 mmHg
                                                                       4, 7, 14,21,28 days
Dunkern (1999)       HUVEC             increasing pulsatile flow       3 ml/min -> 20 ml/min          ePTFE                    4        Better attachment of Ecs
                                                                       1 dynes/cm2 -> 6.6 dynes/cm2
                                                                       24 hr
Wang (2001)          HAEC              cyclic strain                   elongation 10 %                Si                     sheet      Reorientation perpendicular




                                                                                                                                                                               50
    Table 5 continued.


                                                              uniaxial strain 10 %
                                                              biaxial strain 10%
Vouyouka (2002)     RASMC         static pressure             130 mmHg                       coculture           dish   Pressure decreases SMC proliferation
                    RAEC                                      1, 3, 5 days                                              Ecs promote SMC proliferation
                                                                                                                        Pressure inhibits positive EC effect
Jeong (2004)        Rabbit ASMC   pulsatile flow              130 ml/min                     PLCL                 4     Better SM attachment
                                                              25 mmHg                                                   Upregulation SMA
                                                              1 Hz                                                      Radial alignment
                                                              5 % radial distention
                                                              1, 3, 5, 8 weeks
Rashid (2004)       HUCSMC        pulsatile flow              preconditioning sub-arterial   PU                   4     Increased cell retention when preconditioned
                                                              1 week
Benbrahim (1996)    HVEC          pulsatile flow              125/75 mmHg                    Si                   6     Alignment of HVEC and BASMC in flow direction
                    HUVEC                                     120 ml/min
                    BAEC                                      1 Hz
                    BASMC                                     0.05 %/ mmHg
                                                              4, 24 hr
Cummings (2003)     RASMC         cyclic strain               10 % radial strain             collagen-fibrin +    5     Increased gel compaction
                                                              1 Hz                           RASMC                      Augmentation of mechanical properties
                                                              4 days
Seliktar (2000)     RASMC         cyclic strain               10 % radial strain             collagen             3     Increased contraction
                                                              1 Hz                                                      Increased strength
                                                              4, 8 days
Opitz (2004)        OCEC          Increasing Pulsatile flow   15/10 mmHg -> 60/40 mmHg       P4HB                 15    Mechanical strength after 2 weeks close
                    OCSMC                                     Final flow 3 l/min                                        To native aorta
                                                              1 Hz
Niklason (2001)     BASMC         Pulsatile flow              165 bpm                        PGA                 3.1    Compliance as well as burst strength closer to native
                    BAEC                                      270/-30 mmHg                                              artery after pulsatile flow
                                                              5 % radial distension
Niklason (1999)     BASMC         Pulsatile flow              165 bpm                        PGA                 3.1    After8 weeks similar morphology to native
                    BAEC                                      5 % radial strain
Seliktar (2003)     HASMC         Cyclic strain               10 % radial strain             collagen            3.55   Matrix remodeling positively influenced
                    RASMC                                     4, 8 days
                    HDF
Solan (2003)        PSMC          Pulsatile flow              90 or 165 bpm                  PGA                  3     Increased collagen production when pulsed
                                                              145/-5 mmHg or 150/-30 mHg
                                                              1.44 % strain
                                                              7 weeks




                                                                                                                                                              51
It can also be appreciated from the table that the examined regimes include the effects of
shear stress, cyclic radial strain or pulsatile flow only, and a gradually increasing pulsatile
flow or preconditioning of the SMC layer by cyclic strain before seeding the ECs and
applying shear stress subsequently. One group examined the effect of static pressure on
the proliferation behavior of SMCs and ECs [118].

4.5.2 Results so far

Although considerable efforts have been made to improve the construction of TEBVs,
an optimal conditioning protocol is still not at hand. In establishing a construct, the
cells first need to form a confluent mono- or multilayer, depending on the cells.
When seeding conditions are optimized, the retention of the cells after loading is
essential. After a static incubation period necessary for the attachment of cells to the
scaffold, shear stress, cyclic strain or a combination of both is applied. A pulsatile
flow conditioning close to physiological circumstances showed a decrease in cellular
retention to 65, 70 or 87 % (depending on the study) of the initial attachment density
[113,86,21]. Even better results were obtained by a gradually increase in flow rate.

Four independent studies show comparable positive results on different
combinations of graft material and cell type/source. Ott [68] used PU grafts seeded
with BAECs. During the first three days these investigators demonstrated a better EC
retention when the shear stress was confined to only 1 or 2 dynes/cm2 before
switching to a shear stress of 25 dynes/cm2. Dunkern and coworkers showed a better
retention of HUVECs on ePTFE grafts when a pulsatile flow was gradually increased
from 3 to 20 ml/min with an increasing shear stress of 1 to 6.6 dynes/cm2 [18].
Carnagey and co-workers experienced a 20-40 % increase of cell retention when
HUVECs were seeded onto PTFE or OCA grafts [9]. This group imposed an
increasing flow that started with 68 ml/min which is the average physiological flow
rate in coronary arteries, and gradually increased it to 300 ml/min. Shear stress rose
to 6 dynes/cm2, similar to in vivo values. Interesting to mention is that in another
experiment with a constant flow rate of 68 ml/min an even better retention was
found, although the alignment of the cells was less good [9].

Similar results have been obtained with SMCs. Rashid and colleagues used PU grafts
and HUCSMCs and observed a better cell retention when the construct was
preconditioned with a subarterial pulsatile flow during one week [78].

After the optimization of conditions for cell retention, a next step is to modulate the
loading conditions to enhance alignment of cells, appropriate production of cellular
components and to optimize mechanical strength of the constructs. So far, only six
studies have addressed the effects of dynamical loading on mechanical strength of
tubular vascular grafts. Cummings and colleagues found indications of the beneficial
effects of cyclic radial strain on rodent aortic SMCs (RASMCs) which were seeded
into a collagen-fibrin scaffold. A 10% radial strain during 4 days increased the
ultimate tensile stress, linear modulus and toughness with 30, 20 and 35%,



                                                                                            52
respectively [13]. In a collagen gel scaffold in which RASMCs were embedded the
same group found that the construct gained 200% in ultimate tensile stress after 8
days of conditioning with 10% radial strain. At the same time the modulus gained
100% [96].

Hoerstrup and his group used a PGA graft coated with P4HB and on which ovine
myofibroblasts (OCMF) were seeded [29]. After 4 days of static incubation ovine ECs
(OCEC) were seeded into the lumen. Over a period of 28 subsequent days the flow
rate was expanded from 125 to 750 ml/min with a concomitant pressure increase of
30 to 55 mmHg. Burst strength increased from a starting value of 180 to 320 mmHg.
In contrast, in static cultures burst strength decreased to 50 mmHg.

4.5.3 Concluding remarks

At present, it seems too early to define ideal mechanical loading regimes for TEBVs.
However, from the recently obtained results we can conclude that a gradually
increasing pulsatile flow seems to be the best choice for a proper development of a
graft into a strong construct as well as for a cell retention percentage warranting a
proper cell lining.

Some reservations have to be made regarding this conclusion. First, for SMCs an
alternative approach could be more effective. Cyclic radial strain on seeded SMCs
already proved to be beneficial for mechanical properties. Therefore it might be an
option to precondition already seeded SMCs with a straining protocol before seeding
ECs on top of them. This concept has been applied by Imberti [32]. Unfortunately
only EC proliferation was investigated, which was reduced by the pre-straining. The
influence on the mechanical strength has not yet been evaluated and remains a goal
for future researchers.

Secondly, besides the sequence of loading events also the amount of load has to be
determined. Applying physiological forces seems to be logical, but might not to be so
in practice. Even if such a protocol works, whether with a gradually increasing load or
not, applying higher loads decreases the maturation time of the graft. Since this
aspect is crucial for commercial success, it should not be overlooked. At this moment,
we propose to precondition SMCs with a cyclic straining protocol, followed by EC
seeding and the application of a gradually increasing pulsatile flow. This flow might
be even increased to supraphysiological loads, so that after a maturation time the flow
could be gradually decreased again towards more physiological values to let the tissue
adjust correctly prior to implantation again.




                                                                                   53
Chapter 5: Future directions


Over the past decades, novel strategies for creating better constructs for coronary
artery bypass surgery have emerged one after another. Slowly the conviction arose
that only a living TEBV enables successful future application. However, despite
numerous creative attempts and promising prototypes, no adequate product has yet
been established. Some questions still remain unsolved. For example, to what extent
is it actually necessary to have a vasoactive construct to implant? Of course, for exactly
mimicking a real vessel it is a requirement. However, to achieve a sufficient patency,
possibly it is not a necessity. Deutsch et al. already reported an excellent 9 year
patency of EC-seeded ePTFE grafts [16]. Another question is whether the mechanical
properties of the grafts should be similar to that of native vessels. At present it seems
that the compliance of a graft is a crucial feature. The strength however must not
exactly match that of native arteries as long as the graft can withstand physiological
forces.

Another aspect of TEBVs is the source of cells that are used on cell-seeded constructs.
Especially for ECs, it appears that autologous cells are a necessity, based on
immunogenic responses upon implantation and varying cellular responses to certain
stimuli amongst different species. Even then, the origin or autologous cells deserves
further attention. Due to varying cell behavior on different locations, perhaps only
cells from few locations are suitable. Autologous ECs from coronary arteries are
difficult to obtain though. Thus, further research is needed regarding the appropriate
cell choice and possible manipulation of cell behavior.

Finally, the choice of methodology remains subject of many discussions. It seems
there are two tracks at the moment, being cell-seeded synthetic grafting and
biological grafting, whether seeded or unseeded. These two tracks are investigated
separately, and no definite winner (if any) can be appointed at this moment. However,
in our opinion the most recent developments point in favor of cell-seeded synthetic
grafting, due to a significant improvement in preclinical results. A further
improvement on cellular retention after seeding, tuning of ideal polymer degradation
in vivo is required, as well as optimization of conditioning regimes prior to
implantation.

Even with all the ongoing efforts and positive prospects with current methodologies
clinical TEBV application is still a long way to go,. However, I would like to state that
investigators should be permanently encouraged to find novel, creative solutions
and/or unorthodox, sophisticated methodologies to solve existing problems and
bottlenecks. The ‘holy grail’ shall be found!




                                                                                       54
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Appendix A: List of abbreviations


Abbreviation   Explanation
3DP            3-dimensional printing
ACE            Angiotensin converting enzyme
acLDL          Acetylated low density lipid
BAEC           Bovine aortic endothelial cells
CABG           Coronary artery bypass grafting
CAM            Cellular adhesion molecule
CMVEC          Canine microvascular endothelial cells
CPU            Poly(carbonate) urethane
CSH            Collagen sleeve hybrid
EC             Endothelial cell
ECM            Extracellular matrix
EDCF           Endothelial-derived constriction factor
EDHF           Endothelial-derived hyperpolarization factor
EDRF           Endothelial-derived relaxation factor
EPC            Endothelial progenitor cells
ePTFE          Expanded PTFE
FB             Fibroblast
FBS            Fetal bovine serum
FGF            Fibroblast growth factor
GAG            Glucosaminoglycan
HAEC           Human aortic endothelial cells
HCAEC          Human coronary artery endothelial cell
HJVEC          Human jugular vein endothelial cells
HMVEC          Human microvascular endothelial cell
HSV            Human saphenous vein
HUAEC          Human umbilical artery endothelial cell
HUVEC          Human umbilical vein endothelial cell
HVEC           Human saphenous vein endothelial cell
ICAM-1         Intercellular adhesion molecule
IH             Intimal hyperplasia
IMA            Internal mammary artery
LDL            Low density lipid
MCP-1          Monocyte chemoattractant protein
MFB            Myofibroblast
NO             Nitric oxide



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OVEC      Ovine vascular endothelial cells
P4HB      Poly-4-hydroxy butyrate
PAI-1     Plasminogen activator inhibitor
PCL       Poly- -caprolactone
PDGF      Platelet-derived growth factor
PECAM-1   Platelet endothelial cell adhesion molecule
PET       Poly ethyleneterephtalate
PGA       Poly glycolic acid
PGI2      Prostacyclin
PHO       Poly hydroxyoctanoate
PJVEC     Porcine jugular vein endothelial cells
PLA       Poly lactic acid
PLCL      Poly(l-lactic-co-caprolactone)
PLGA      Poly(l-lactic-co-glycolic acid)
PLLA      Poly-l-lactic acid
PTFE      Poly tetrafluoroethylene
PU        Polyurethane
RAEC      Rat aortic endothelial cells
RGD       Arginine-glycine-aspartic acid
SIS       Small intestine submucosa
SMC       Smooth muscle cell
SPU       Segmented polyurethane
TE        Tissue engineering
TEBV      Tissue-engineered blood vessel
TM        Thrombomodulin
t-PA      Tissue-type plasminogen activator
UEA-1     Ulex europeaus agglutinin
VCAM-1    Vascular cellular adhesion molecule
VEGF      Vascular endothelial growth factor
VSMC      Vascular smooth muscle cell
vWF       Von Willebrand factor
WP        Weibel-Palade
α-SMA     Alpha-smooth muscle actin




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Appendix B: Leaders in the field

Institution                                         City         Country       Department                                   Researchers
University of Queensland                            Brisbane     Australia     Centre for Research in Vascular Biology      Julie H. Campbell
                                                                                                                            Gordon R. Campbell
Laval University                                    Quebec       Canada        Laboratoire d’Organogenese Experimentale     Francois A. Auger
                                                                                                                            Lucie Germain
                                                                                                                            M. Kodama
Shanghai 2nd Medical University                     Shanghai     China         Shanghai Key Laboratory of Tissue            Y. L. Cao
                                                                               Engineering                                  G. Shen
Friedrich-Schiller University                       Jena         Germany       Department of Cario, Thoracic and Vascular   Franka Opitz
                                                                               Surgery                                      U. A. Stock
Hannover Medical School                             Hannover     Germany       Leibniz Research Laboratories for            Omke E. Teebken
                                                                               Biotechnology & Artificial Organs
Kyushu University                                   Fukuoka      Japan         Division of Biomedical Engineering           T. Matsuda
                                                                                                                            S. Kidoaki
                                                                                                                            H. He
Tokyo Women’s Medical University                    Tokyo        Japan         Cardiovascular Surgery                       T. Shin’oka
                                                                                                                            G. Matsumura
National Cardiovascular Center Research Institute   Osaka        Japan         Department of Bioengineering                 K. Takamizawa
Kyoto University                                    Kyoto        Japan         Graduate School of Medicine                  M. Komeda
                                                                                                                            N. Tamura
Osaka University                                    Osaka        Japan         Division of Cardiovascular Surgery           Hikaru Matsuda
                                                                                                                            S. Iwai
Maastricht University                               Maastricht   Netherland    Center for Biomaterials Research             Leo H. Koole
                                                                 s
National University of Singapore                    Singapore    Singapore     Nanobioengineering Laboratories              S. Ramakrishna
Hanyang University                                  Seoul        South-        School of Chemical Engineering               Young M. Lee
                                                                 Korea                                                      Sung I. Jeong
Korea Institute of Science and Technology           Seoul        South-        Bionaterials Research Centre                 Young H. Kim
                                                                 Korea                                                      S. H. Kim
University of Zurich                                Zurich       Switzerland   Laboratory for Cardiovascular Tissue         Simon P. Hoerstrup
                                                                               Engineering and Cell Transplantation




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University College London               London         UK    Vascular Haemodynamics Laboratory              Alexander Seifalian
                                                                                                            G. Hamilton
                                                                                                            H. Salacinski
Manchester University                   Manchester     UK    UK Centre for Tissue Engineering               Ann E. Canfield
                                                                                                            N. J. Turner
Georgia Institute of Technology         Atlanta        USA   Institute for Bioengineering and Biosciences   Robert M. Nerem
                                                                                                            A. Sambanis
                                                                                                            D. Seliktar
Massachusetts Institute of Technology   Cambridge      USA   Department of Chemical Engineering             Robert S. Langer
                                                                                                            Joseph P. Vacanti
Duke University                         Durham         USA   Centre for Biomolecular and Tissue             Laura E. Niklason
                                                             Engineering                                    Shannon L. Mitchell
Harvard Medical School                  Boston         USA   Edwin L. Steele Laboratory                     K. Rakesh
Thomas Jefferson University             Philadelphia   USA   Orthopedic Research                            Paul J. DiMuzio
                                                                                                            Patrick J. Schaner
Rice University                         Houston        USA   Department of Bioengineering                   J. L. West




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