JASON HABERMEHL
DEVELOPMENT AND VALIDATION OF A
COLLAGEN-BASED SCAFFOLD FOR VASCULAR
TISSUE ENGINEERING
Mémoire présenté
à la Faculté des études supérieures de l’Université Laval
dans le cadre du programme de maîtrise en génie de la métallurgie
pour l’obtention du grade de maître ès sciences (M.Sc.)
FACULTÉ DES SCIENCES ET DE GÉNIE
UNIVERSITÉ LAVAL
QUÉBEC
2005
© Jason Habermehl, 2005
Résumé
L’ingénierie tissulaire est une approche qui vise à combler le besoin de remplacements
d’organes. Ceci est d’autant plus vrai pour les artères, dont le besoin de remplacements
résulte en partie de la prévalence des maladies cardiovasculaires dans le monde
industrialisé. Pour ce faire, nous croyons qu’une des alternatives les plus prometteuse
implique l’utilisation d’une structure permettant l’échafaudage tridimensionnel de cellules
lors de la régénération. Le collagène possède plusieurs caractéristiques qui font en sorte
qu’il peut être un matériau convenable à la fabrication de cette structure.
Au cours de ce projet, un procédé d’extraction et de mise en solution de collagène type I à
partir de queues de rat a été choisi et validé. Ce collagène a été caractérisé par rapport à ses
propriétés moléculaires et ses performances mécaniques, biologiques et hématologiques.
Suivant cette validation, une méthode pour produire des gels en forme de disques et de
tubes ensemencés de cellules vasculaires a été élaborée. Ces structures ont été caractérisées
quant à la viabilité des cellules dans le gel, la capacité des gels à être manipulés ainsi qu‘à
leurs propriétés hématologiques, en milieu humide. Ce projet constitue une base pour des
recherches futures visant à améliorer la fonctionnalité et les propriétés structurales des
artères régénérées à base de collagène.
ii
Abstract
Tissue engineering provides insight into solving the organ shortage. This is especially the
case for small diameter vascular substitutes, with which a shortage is due in part to the
prevalence of cardiovascular disease in the industrialized world. For this, in our opinion,
one of the most appropriate approaches involves using a structure to guide the cells during
the regeneration phase. Collagen has many characteristics that make it suitable as a scaffold
material for vascular tissue engineering.
Two slightly different methods for extracting and processing collagen type I from rat tail
were compared with respect to the molecular structure of the collagen molecule, the
mechanical properties of thin films obtained from solvent evaporation and preliminary
cellular viability with fibroblasts seeded on these same collagen films. One of the above
methods was chosen and this collagen was then characterized with respect to cellular
viability with smooth muscle cells and endothelial cells and also with blood contact
assays. A method for producing three-dimensional gels seeded with vascular cells was
developed. Cell distribution and viability, preliminary compliance testing and blood
contact assays were performed on these gels. This project has provided the basis for further
studies in order to maximize cell functionality and the structural properties required for
implantation of collagen-gel-based vascular grafts.
Preface
My two favourite subjects at CEGEP were physics and biology. I chose to pursue an
undergraduate degree in physics engineering. I believed that this would provide me with
better career opportunities. Upon graduation, having realized that no immediate
employment was available for myself and most of my friends, I chose to return somewhat
to my previous interest in biology by entering the exciting and highly multidisciplinary
field of bioengineering by pursuing a masters degree. With no additional formation than a
basic biology CEGEP class, I had a lot to learn. I therefore needed the help of people from
many different fields, notably biology, without whom my project would not have been such
a success. I would like to thank the following people for their involvement in helping me
throughout my research.
Diego Mantovani, my project director for allowing me great freedom to pursue my interests
in choosing a masters project.
Gaetan Laroche, my project co-director, for guidance and invaluable input.
Francesca Boccafoschi, with whom I performed most of my work and who was my guiding
light and motivator throughout it all. Without her, I would have been lost in the dark.
Marie France Côté, who introduced me to rat tail collagen and for showing me all that I
know about extraction, processing and making gels.
Navneeta Rajan for taking over my project so that my stay at the LBB may be remembered
after my departure.
Joanna Skopinska, for her short stay among us and for helping me to learn more about
collagen.
Richard Janvier, for his patience in teaching the wonderful basics of cell culture to a
physics engineer with no prior experience.
My other colleagues at the laboratory for biomaterials and bioengineering, who although
may not all have contributed to my project per se, all helped to maintain an enjoyable work
iv
atmosphere and helped take my mind off work from time to time: Karine Vallières, for
enduring the proximity of such a close work space for so long and for all we have shared,
Pascale Chevalier for always enquiring about my extra-laboratory activities, MarieHaidopoulous for making the coffee breaks so much more enjoyable and for her great
ducks-on-ice imitation, Stéphane Turgeon, for initiating so many Ashton excursions and for
providing technical assistance, Jean Lagueux for his paternal figure status and assistance in
various fields, Christian Sara-Bournet for his out-of-the-blue but highly entertaining tidbits
of information and insights, Louis Gagné for making me look less ‘bio.’ in my eating
habits, Francois Lewis for making me realise my true orientations, Stéphanie Coulomb, and
all the others not specifically mentioned. Lastly, Jean-Claude who, although not actually
there in the physical sense, provided inspiration for myself and others with his deep and
mystical citations.
I would also like to mention: Simone Vesentini for welcoming me to Italy and helping to
make my stay there all the more enjoyable, Chiara for her short stay among us and for
showing me Québec’s Italian sister city: Bergamo, Sébastien Blanchet for his mechanical
abilities in making our custom-made cutter and Guy Bureau for technical assistance in
preparing the samples for mechanical testing.
And a final mention concerning the financial assistance that made my project possible:
NSERC (Natural Science and Engineering and Research Council, Canada)
FQRNT (Fonds Québécois de recherche sur la nature et les technologies, Québec)
NATO Science program
CFI (Canadian Foundation for Innovation, Canada)
MRI (Ministère des Relations Internationales, Québec)
Two articles make up the results section of the present thesis. I take this moment to indicate
my involvement in these works. The first part of Chapter 3 is an article published by
Macromolecular Bioscience (2005) of which I am the first author. My contribution
included much of the overall writing of the article and a good part of the experimentation. I
v
did not perform first hand the chemical analyses of collagen; this work was done by Joanna
Skopinska. I performed first hand the sample preparations, mechanical assays and gave a
helping hand for the cell tests. The second part of Chapter 3 is an article published by
Biomaterials (2005), of which I am the second author. I performed the sample preparations
and helped to some extent with the various cell analyses and blood contact assays. No
significant changes were performed on these first two articles before inclusion in this
thesis.
Objective of the project
The general theme in the laboratory involves researching various strategies for treating
arterial disease. The objective of this project, inexistent before my arrival, was to design
and develop a scaffold structure from collagen for a tissue-regenerated artery. Emphasis
was placed on working towards a completely biological arterial substitute which will
eventually be placed in a perfusion bioreactor and cultured therein.
Due to the groundbreaking nature of this project, the first requirement was to obtain and
validate a method with which to extract collagen. Once an adequate extraction and
processing method was mastered, the collagen solution thus obtained needed to be
characterized with respect to its molecular structure, mechanical properties and cell culture
feasibility to validate the method. Blood contact tests were also deemed necessary to
further validate the collagens adequateness for such an important blood contact application
as an arterial replacement.
Following these preliminary tests, a means to produce three-dimensional cylindrical
collagen tubes of adequate dimensions was required. In fact, collagen can form a gel under
the appropriate conditions. This gel can be seeded with various cell types and can be
formed into various shapes. A process of making tubes with this approach enabling cell
incorporation is paramount. These constructs were then tested with respect to the viability
of cells while embedded in collagen and with respect to their structural integrity.
vi
Structure of the thesis
The first section (Chapter 1) of this thesis introduces the concept of using collagen for
vascular tissue engineering. This section encompasses a revue of relevant literature and the
possibilities available for improvement of these vascular replacements. The second section
(Cahpter 2) details the methodology taken during the project. This section also presents
various unpublished results. The third section (Chapter 3), which constitutes the entirety of
the results, is a collection of two articles. The first article presents the validation of the
chosen collagen extracting and processing method. The second article presents further
preliminary validations of our collagen concerning cell compatibility and blood contact
properties. Finally, the conclusion presents a recapitulation of the project objectives, the
objectives met, the limitations and future works. Annexes A to C present the main
protocols.
Table of Contents
Chapter 1: Introduction .......................................................................................................1
Blood vessels ......................................................................................................................3
Basic blood vessel physiology........................................................................................3
Arteries............................................................................................................................6
Collagen ..........................................................................................................................7
Elastin .............................................................................................................................8
Requirements for arterial replacement................................................................................9
Tissue engineering ............................................................................................................11
Approaches to vascular tissue engineering...................................................................11
The scaffold ..................................................................................................................17
Collagen as a scaffold for vascular tissue engineering .....................................................20
Collagen properties .......................................................................................................20
Cell functionality and phenotype..................................................................................21
Modifications of collagen gel properties ......................................................................24
Vascularisation..............................................................................................................34
Conclusion ........................................................................................................................35
Chapter 2: Rationale, Methodology and unpublished results ........................................37
Collagen: source ...............................................................................................................39
Collagen: extraction and processing .................................................................................39
Collagen films: characterization .......................................................................................40
Collagen films: cell viability and hemocompatibility.......................................................42
Collagen gels (without cells): feasibility ..........................................................................42
Discs..............................................................................................................................43
Tubes.............................................................................................................................43
Collagen gels with cells ....................................................................................................46
Discs..............................................................................................................................46
Tubes.............................................................................................................................47
Chapter 3: Results ..............................................................................................................56
Preparation of a Ready-to-use, Stockable and Reconstituted collagen. ...........................56
Résumé..........................................................................................................................57
Abstract.........................................................................................................................58
Introduction...................................................................................................................59
Materials and Methods..................................................................................................61
Results...........................................................................................................................64
Discussion.....................................................................................................................72
Conclusion ....................................................................................................................74
Biological Performances of Collagen-based Scaffolds for Vascular Tissue Engineering75
Résumé..........................................................................................................................76
Abstract.........................................................................................................................77
Introduction...................................................................................................................78
Materials and Method ...................................................................................................80
Results...........................................................................................................................85
Discussion.....................................................................................................................91
List of Tables
Table 1-1 : Comparison of the major approaches using collagen for vascular tissue
engineering to date........................................................................................................16
Table 1-2 : The effects of various chemical signals on smooth muscle cell seeded collagen
gels or films. (a blank box indicates no relevant influence was found in the literature.)
......................................................................................................................................23
Table 2-1 : Disc-shaped collagen gel integrity depending on the relative amount of various
ingredients. Relative gel strength is given by 1 to 4, 4 being the strongest gel............43
Table 2-2 : Comparative table for the three main cell phases present in smooth muscle cells
after 1 week in collagen tubular gels (Scaffold) and in gelatin coated Petri dishes
(Control)........................................................................................................................51
Table 3-1 : Assignments of the spectral features underlying the Amide I infrared band. ....66
Table 3-2 : A1660/A1630 ratios measured from the curve fitted Amide I infrared feature of
collagen samples A and B before and after neutralization. ..........................................68
List of Figures
Figure 1.1 : The cardiovascular system ..................................................................................4
Figure 1.2 : The physiology of the typical artery and vein.....................................................5
Figure 1.3 : The structure of collagen. From the tropocollagen helix to the fiber..................8
Figure 1.4 : Tissue engineering. The basics..........................................................................17
Figure 2.1 : Collagen cell-seeded gel approach to vascular tissue regeneration ..................37
Figure 2.2 : Strategy of the research project.........................................................................38
Figure 2.3 : Custom-made cutter for thin film mechanical testing sample preparation. ......41
Figure 2.4 : Custom-made tube mandrel set-ups. A: Initial set-up for cylindrical dried-film
preparation, B: Modification of initial set-up for cell-seeded gels, C: Final set-up
allowing gas exchange for cell-seeded gels. Custom-made rotation device A: front
view, B: side view.........................................................................................................44
Figure 2.5 : Custom-made rotation device for cylindrical gels. A: isometric view, B: top
view, C: front view. ......................................................................................................45
Figure 2.6 : Gel contraction seeded with smooth muscle cells (SMC) and fibroblasts (FB).
......................................................................................................................................47
Figure 2.7 : Typical collagen sell-seeded tube obtained with the gas-exchange tube-mandrel
system ...........................................................................................................................49
Figure 2.8 : Masson’s trichrome on a longitudinal a) and transversal b) section of tubular
collagen gel scaffold seeded with smooth muscle cells after 1 week maturation.........50
Figure 2.9 : The various phases of the cell cycle..................................................................50
Figure 2.10 : Platelet adhesion on collagen gel seeded with smooth muscle cells without
(A-C) and with (B-D) endothelial cells layer on the surface. Images were acquired
with different magnifications: 200x (A-B) and 2000x (C-D).......................................52
Figure 3.1 : Curve fitting of the Amide I infrared peak of collagen samples A and B before
(a and b) and after neutralization (c and d). ..................................................................67
Figure 3.2 : Tensile properties of collagen films a) ultimate strength b) elastic modulus c)
maximum elongation. ...................................................................................................69
Figure 3.3 : Representative Stress-Strain curves of films made of acidic and neutral
collagen films................................................................................................................70
Figure 3.4 : Scanning electron microscopy (SEM) analyses of NIH 3T3 cells after 1 week
on neutral collagen A (a) and neutral collagen B (b)....................................................71
Figure 3.5 : MTT assay testing cytotoxicity of different collagens on NIH 3T3 after 24h
and 1 week. Control is considered as 100% viability. All results are significant with
respect to control and between pure collagen (A and B) and neutral collagen (A and B)
(p < 0.005).....................................................................................................................71
Figure 3.6 : Haemoglobin free test on collagen, Teflon and glass. * indicates that results of
collagen with respect to both reference materials are statistically significant..............86
Figure 3.7 : Fibrin formation. * indicates statistically significant results with respect to
blood. °indicates statistically significant results with respect to collagen....................87
Figure 3.8 : Thrombogenicity Index. * indicates significant results with respect to blood..88
Figure 3.9 : SEM analyses of platelets on collagen film. .....................................................89
Figure 3.10 : SEM analyses of endothelial cell (a) and smooth muscle cell (b) morphology
after 24 hours of contact with collagen films. ..............................................................89
x
Figure 3.11 : MTT results on endothelial cells (a) and smooth muscle cells (b). * indicates
significant results with respect to control. ....................................................................90
1
Chapter 1: Introduction
The human body is a complex assembly of organs and tissues. Like any living organism, it
and its components are not meant to last indefinitely. These organs can become damaged or
diseased affecting our quality of life. The arterial system is no exception. In fact,
cardiovascular diseases are the single most important cause of death in the world[1]. When
not treated, these diseases which affect the heart and blood vessels, end in heart attacks or
strokes. Cardiovascular diseases take many forms including coronary heart diseases such as
atherosclerosis and hypertension, heart muscle disease, arrhythmias, and valve
disorders[2]. Atherosclerosis for example is characterized by the accumulation of lipids,
cells and extra-cellular matrix molecules in the vessel wall[3]. Atherosclerosis of coronary
or peripheral vascular arteries is the largest cause of mortality in Canada and in all
developed countries[4]. Symptoms include angina, or chest pain, and vascular deficiencies.
Often, cases of atherosclerosis are not detected until complete occlusion results in heart
attack or stroke. In many cases, pharmaceutical treatments are not sufficient and more
drastic measures must be taken in order to return adequate blood flow through the arteries
and minimize the risk of occlusion and embolism.
For coronary arteries or peripheral arteries such as that below the knee, autologous graft
transplantation is the option of choice. The grafts usually consist of either mammary artery
or saphenous vein harvested from the patient. Unfortunately, all too often, autologous grafts
are not always available due to their limited number and their insufficient quality. Repeated
procedures and multiple bypasses are also problematic since all vessels appropriate for
replacement may have been employed during previous interventions. Furthermore,
saphenous vein in elderly patients is prone to thrombi, neointimal formation,
atherosclerosis or aneurysm[5]. Veins also lack vasomotor tone. Moreover, harvesting
veins and arteries leaves wounds that can break down and become infected. Another
alternative used in the past was fresh allografts. However, these allograft arteries or
veins are no longer used for coronary bypass surgery due to poor patency, rejection
2
complications, endothelial cell sloughing and reaction with leukocytes, and loss of
cellular reactivity[6].
For large diameter arteries, synthetic prostheses made of Teflon or Dacron have been a
relatively viable solution for the past few decades. These prostheses perform reasonably
well in high-flow and low resistance conditions[5]. But their use is limited to vascular
conduits larger than 5-6 mm in diameter due to the high risk of thrombus formation,
embolism and occlusion associated with their use in smaller vessels. Despite extensive
research on synthetic vascular prostheses, these have yet to prove suitable for long term
implantation. In fact, 65% of them must be explanted in the 10 years following
implantation[7]. Synthetic prostheses pose a high risk of foreign body reaction and
extensive use of anticoagulant/antithrombotic control is often required[5]. Synthetic
prostheses also lack a confluent endothelium. This single cell layer is what confers
haemostatic properties to natural arteries rendering them antithrombotic[8]. Therefore, the
importance of creating a suitable endothelium on the luminal surface of any small diameter
vessel substitute is paramount. Research into surface modifications with coatings of
proteins, polymer materials or cells is being pursued to increase the bioactivity of
prostheses in order to render then more suitable for endothelial cell seeding. These
approaches have been met with mixed success[5]. Also, these linings do not provide vital
vascular functions such as vascular responsiveness or other biological secretory functions
seen with normal blood vessels[8].
Recently, stents have been in clinical use when performing balloon angioplasty in order to
unclog blocked arteries using interventional surgery as opposed to classical surgery
methods. The stent remains in the artery upon removal of the balloon catheter thereby
offering structural support to the compressed layer of lipids and cells on the luminal surface
of the artery[9]. Although, some success has been observed with this method, many
complications remain, most notably, their limited use in extensively diseased blood vessels.
Furthermore, in stent restenosis is possible. In fact, in-stent restenosis (stenosis diameter
≥50%) occurs in 20-30 % of implanted stents[9].
3
Because of the problems associated with native, synthetic and modified grafts and
prostheses, much research has been done in order to create tissue-engineered vessels.
Tissue Engineering has been classified as an interdisciplinary field that applies the
principles of engineering and life sciences toward the development of biological substitutes
that restore, maintain, or improve tissue function[10]. The present chapter consists of a
review of the techniques available to generate an arterial replacement. Various approaches
are mentioned but one in particular, using reconstituted collagen as a scaffold, is discussed
in detail.
Blood vessels
Basic blood vessel physiology
Before going any further, it is important here to take a step back and get familiar with this
organ that needs replacement. The blood vessels are components of the cardiovascular
system. They are much more than an arrangement of pipes and tubes with blood flowing
through. There are many types of blood vessels in the body, each with a distinct link
between structure and function. The arteries transport blood from the heart to the other
organs and the peripheries of the body. The blood returns to the heart via veins. The arterial
and venous systems are linked through a capillary system. In addition to blood transport
functionality, blood vessels also possess biochemical functions making them organs in
themselves.
In the majority of cardiovascular malfunctions, the problematic section of the
cardiovascular system is the artery. The size and function of arteries depends on their
location in the body (Figure 1.1)[11]. There are two main types of arteries: elastic and
muscular. The aorta and other larger vessels near the heart possess an important elastic
component giving them adequate properties to deal with high pressure and flow. Their
elastic component (elastin) can be as high as 40%. As the arteries are situated further and
further from the heart, this elastic component becomes secondary to the muscular
component. In the muscular arteries which range in diameter from 1mm to 1 cm, the elastic
component makes up roughly 10%.
4
Figure 1.1 : The cardiovascular system
All arteries have the same basic structure despite variations in the proportions of their
structural components. They are composed of three main layers, or tunicae, which are from
the inside out: the intima, the media, and the adventitia[12].
5
Figure 1.2 : The physiology of the typical artery and vein.
Intima
The intima is composed of a confluent monolayer of endothelial cells (EC) arranged
longitudinally, which form the endothelium. The endothelial cells are orientated parallel to
the direction of blood flow. This layer of cells assures the hemocompatibility and antithrombogenicity of the artery.
Media
The media is the major component of muscular arteries. It is composed of smooth muscle
cells (SMC) aligned concentrically along with collagen and elastin fibers and
proteoglycans. The media confers the majority of the mechanical properties to the artery
and is responsible for the peristalsis effect which assists with blood transport to the
peripheral organs. To this effect, smooth muscle cells contract and relax in response to biochemical and bio-mechanical stimuli.
6
Adventitia
The adventitia is mainly composed of collagen and fibroblast cells (FC) both arranged
longitudinally. The function of the adventitia is basically the same in all types of arteries. It
provides anchorage with connective tissues and maintains nutrient support and
vascularisation to the internal layers of the artery.
These different layers are separated by two elastic laminas which act as barriers. The
internal elastic lamina provides also a surface for endothelial cell attachment and supports
the endothelium in its role as a barrier to blood contact with the media.
Arteries
Functions
Evidently arteries have a main goal of transporting blood from the heart to the rest of the
body in order to assure adequate vascularisation of all organs. However, they do not act as
simple inanimate pipes to this effect. Arteries can vary their diameter in response to various
external stimuli, such as nerve impulses and hormonal signals, to regulate blood flow in
order to control the supply of blood to individual organs. This is especially true for
muscular arteries in which the cellular component confers the capacity to multiply by
twenty-five times the blood flow to skeletal muscles during physical exertion[13]. For their
part, the large arteries near the heart act as a pressure reservoir. During the systolic period,
the elastic fibers temporarily store mechanical energy and, in returning to their normal state
during the diastolic period, release this energy thereby maintaining continuous blood flow.
Composition
The extracellular matrix (ECM) in the vascular wall provides a structural framework for
the structural and functional properties of the vessel walls[14]. The ECM therefore affects
the elasticity, the resistance and the stretching capabilities of the blood vessel. The two
main structural protein constituents of the vascular ECM are collagen and elastin. The
7
collagens provide tensile stiffness and elastin the elastic properties. Another molecule, the
proteoglycans, contribute to compressibility; these combined with collagen and elastin, are
also responsible for the viscoelastic properties supplying the necessary elasticity to stretch
and recoil. These proteins also play essential roles in other functional requirements of
vessels such as hydration, ion filtration, growth factor bioavailability and cell-matrix
interactions. In fact, the absolute and relative quantities of collagen and elastin have a major
impact on the biomechanical properties of vessels. Collagen, with its very high tensile
strength, maintains the structural integrity of the vessel. These aforementioned
macromolecules are synthesized to some extent by the three vascular cell types
(endothelial, smooth muscle and fibroblast).
Collagen
Collagen is the main protein constituent of muscular arteries where it serves a major
structural role. It is in fact, the most abundant protein in the animal kingdom. It is present in
almost all tissues ranging from bones, cartilage, tendons, ligaments and all other soft tissues
(skin, muscles and all other organs). It is characterized by great tensile strength in
molecular and fiber form.
There are at least 20 types of collagen[15]. The human body is mainly composed of
collagens type I, II and III, however many other types are present. Collagens type I and III
are the major fibrillar collagens in blood vessels where they represent 60% and 30% of
vascular collagens respectively[14]. The basic unit of fibrillar collagens is the triple helix
formed by three intertwining amino-acid chains (Figure 1.3.3). Each chain is roughly 330
amino acids long and the overall molecule, called tropocollagen, is 300 nm long and has a
diameter of roughly 1.5 nm. The most abundant amino-acids are glycine (Gly), proline
(Pro) and hydroxyproline. These form a repeating pattern of Gly – X – Y where X is
usually proline and Y is usually hydroxyproline. The repetition of glycine in every third
position is the most essential factor determining collagens triple helical structure. Intra- and
inter-molecular hydrogen bonds are responsible for the stability of the triple helix. Such
bonds can be inter-chain hydrogen bonds coupled by NH groups of a glycyl residue with
the CO group of a residue in a neighboring chain. Bonds are also formed via water
8
molecules participating in the formation of additional hydrogen bonds with the help of
collagen hydroxyl groups[16]. The great strength of collagen fibers, however, originates
from the stable intermolecular covalent bonds between adjacent tropocollagen
molecules[17]. These fibers can range in diameter from 50-200 nm.
Figure 1.3 : The structure of collagen. From the tropocollagen helix to the fiber
Elastin
Elastin, an insoluble protein produced specifically by SMC in the media, confers elasticity
and supports transportation of metabolic substances in arterial walls[18]. Tropoelastin is a
protein of 750 to 800 residues long. This soluble precursor of elastin contains large
hydrophobic domains, dominated by aliphatic residues of proline, alanine, valine, leucine
and glycine, and smaller alanine-rich domains. Elastogenesis leads to the construction of
mature functional elastin within elastic fibers. Mature elastin is an insoluble polymer
constituted by several tropoelastin molecules cross-linked together[19]. Elastic fibres are
basically composed of amorphous elastin and insoluble microfibrils. During fiber
formation, the microfibrillar compound acts as a scaffold onto which elastin is
deposited[20]. Despite its hydrophobicity, elastin is highly hydrated in water causing it to
swell. Mature elastin is very stable and possesses an extremely low turnover rate.
9
Requirements for arterial replacement
The arteries are complex organs. In order to replace them, a viable long-term replacement
must possess many essential properties. According to Thomas[5], a replacement artery
must:
possess a confluent, adherent and quiescent endothelium to resist thrombosis
in vivo.
be infection-resistant.
be biocompatible (noninflammatory, non-toxic, noncarcinogenic,
nonimmunogenic) and biostable.
have appropriate mechanical properties. They must also possess good
suturability during implantation and be kink resistant.
possess appropriate vasoactive physiological properties enabling it to constrict
and relax in response to neural or chemical stimuli.
be manufactured relatively cheaply, in a relatively short time period, and in
sufficient quantities.
Biocompatibility is a general term used to describe the suitability of a material for exposure
to the body or bodily fluids. The specific meaning is dependent upon the particular
application or circumstances. In fact, there are no completely biocompatible materials. The
success of many medical devices and implants is limited by the interaction of the device
materials with the tissues that they contact. This interaction can be improved by various
means, but the inflammatory reaction of the body to foreign substances has yet to be
eliminated completely. In the case of vascular substitutes, the activation of the immune
system and the coagulation cascade can slow the healing process and the host’s ability to
integrate the graft into the natural circulatory system. Biocompatibility can be associated to
characteristics such as: non-inflammatory, non-toxic, non-carcinogenic and nonimmunogenic.
Mechanical attributes relate to burst strength, compliance and viscoelastic properties. That
is, they must possess sufficient strength to resist pre-implantation manipulations and resist
the stresses acted upon it once in the host. A typical burst strength considered satisfactory
in tissue engineering is 2000 mmHg. This burst strength is slightly higher than that of
10
human saphenous vein, a currently used graft, which is 1680 ± 307 mmHg[21]. However,
high burst strength is not sufficient. Vessel substitutes must also be compliant. Compliance
can be defined as the change in luminal volume in response to a change in pressure inside a
tube. Replacement arteries must dilate in a similar manner to natural arteries upon
application of luminal pressure. Furthermore, they must be viscoelastic in order to contract
to their initial diameter during low pressure periods, without permanent deformation
induced during high pressure periods induced by each heart beat.
Furthermore, according to Mitchell, a blood vessel replacement must have a highly
organized collagen matrix to impart tissue strength and must contain an elastin network to
provide compliance and recoil[22]. In order to meet these demanding characteristics,
tissue engineering has been deemed an interesting approach. This approach may prove to
provide the solution regarding the shortage of suitable small diameter arterial replacements.
11
Tissue engineering
Approaches to vascular tissue engineering
The following is a list of the main vascular tissue regeneration approaches. One approach
involves removing cellular components from native tissue and subsequently seeding these
acellular tissues with vascular cells. A second approach involves using the body’s natural
wound healing response to form a tissue-like structure around a synthetic mandrel inserted
in the body. Thirdly, cells can be used to produce ECM in vitro. These ECM/cell sheets can
form a tube which mimic the natural physiology of the artery. Lastly, various biological or
synthetic materials can be used as scaffolds to guide cell growth into the desired structure.
Acellular ECM tubes
Acellular tubes can be processed from allogenic and xenogenic arteries and veins[23,24],
human umbilical cord[25], or heparinized porcine intestinal submucosal layers[26].
Cellular components can be removed from these tissues without extensive degradation of
the extra-cellular matrix resulting in a naturally derived scaffold consisting of extra-cellular
matrix molecules. These tubes can become colonised by host cells post-implantation or
seeded and cultured prior to implantation. A scaffold processed in this manner from
intestinal submucosa consists mainly of type I collagen without other proteins, lipids or
nucleic acids[26]. This scaffold can be cross-linked with bovine collagen or synthetic
elastomer tubes[27] and has shown SMC and endothelial cell infiltration and remained
patent after 13 weeks of implantation in rabbits. A burst strength of 240 mmHg was
obtained which is considerably less than the human saphenous vein standard (~1600
mmHg). Another study demonstrated that partially devitalised collagen/elastin matrices
may be prepared, without denaturation of the protein, from blood vessels and that the
success of this method depends on the segment and type of blood vessel used[23].
12
In vivo cell self-assembly
Another approach involves stimulating cells in the peritoneal or pleural cavities to coat
inert tubing with bone-marrow-derived cells and mesothelium[28]. This approach, that uses
the body as a natural bioreactor, attempts to adapt the body’s natural wound healing
response to produce a hollow tube. The tubes inserted in the peritoneal cavity become
encapsulated with layers of myofibroblasts, derived from bone-marrow-derived cells, and
mesothelial cells, which possess antithrombotic properties similar to endothelial cells, to
resemble an inverted blood vessel wall[5]. Granulation tissue formation can take up to 23 weeks followed by its removal from the tubing mould. The myofibroblasts, which are
smooth muscle-like cells, have as much alpha-smooth muscle actin and desmin as a
native artery. Autogenic implantation in abdominal rat aortas demonstrated patency of at
least 4 months accompanied with a development of contractile responsiveness.
In vitro ECM production by cells
L’Heureux et al. used a completely different approach to constructing a tissue engineered
blood vessel (TEBV) by taking advantage of the natural ability of cells to produce their
own ECM[21]. Sheets of human SMC and sheets of FC were grown to post-confluence and
subsequently wrapped around a perforated synthetic tubular mandrel. The construct was
placed in a bioreactor providing luminal flow and cultured for extended time periods. EC
seeding was performed one week prior to implantation or in vitro testing. The overall
culture period of the constructs was 3 months. These constructs demonstrated adequate
mechanical strength (2000 mmHg), blood compatibility and suturability. Short term
implantation in a canine model demonstrated a 50% patency rate after 1 week implantation.
However, this approach lacks the appropriate medial structure of circumferentially aligned
collagen molecules and SMC with functional contractile abilities.
13
Tissue-engineering by in vitro seeding of a scaffold
Synthetic support
Although the disadvantages of using synthetic polymers for vascular tissue engineering
scaffolds were mentioned, they provide adequate mechanical results and therefore may be a
short-term option or at the very least another tool to study cell behaviour and in vitro tissue
formation. It is to them that we must compare any other regenerated tissues.
Niklason et al.[29,30], used a scaffold of degradable polyglycolic-acid (PGA) seeded with
SMC to construct a TEBV. PGA scaffolds were seeded with SMC and matured in
bioreactor at 165 pulses per min for 6-10 weeks. Endothelial coating was performed after 8
weeks. Implantation in pigs showed patency up to 4 weeks. It was demonstrated that
pulsations increased the production of collagen and improved patency as compared to static
culture. These pulsed constructs also demonstrated contraction in the presence of
appropriate chemical signals. Appropriate mechanical strength (2000 mmHg) as well as
suturability was achieved with this method. After 8 weeks of culture, PGA fragments
remained which resulted in a dedifferentiated SMC phenotype in the vicinity of these
fragments.
Hoerstrup et al.[31] used a similar approach by seeding fibroblasts and endothelial cells on
synthetic support of PGA scaffolds coated with poly-4-hydroxybutyrate (PHA). These
constructs were matured for up to 1 month in a bioreactor producing pulsed flow directly
through the lumen, thereby generating direct shear stress to the luminal surface as well as
periodical radial distension of the vessel wall. Maximal burst strength of 300 mmHg was
obtained. More recently, the same group[32] used human umbilical cord cells seeded on
PGA/PHA scaffolds to generate a pulmonary artery conduit. They demonstrated the
feasibility of using cells which exhibited myofibroblast-like characteristics.
With a similar approach, Jeong et al.[33] seeded vascular smooth muscle cells on tubular
rubber-like elastic degradable polymer poly-(lactide-co-caprolactone)(PCLC). These
constructs were matured in a pulsatile perfusion bioreactor incorporating shear stresses and
14
radial distension for up to 8 weeks. It is believed that the more rubber-like scaffold was
more beneficial than previous attempts in delivering mechanical signals to cells.
Biological support
The above in vitro seeded-scaffold attempts have based their tissue generation on synthetic
scaffolds which pose certain threats and disadvantages in vascular tissue engineering. The
following is a presentation of various attempts to replace synthetic scaffolds by biological
ones such as collagen and fibrin.
Weinberg et al.[34] were the first to attempt a completely biological TEBV using animal
collagen gels and cultured bovine endothelial cells and smooth muscle cells, and
fibroblasts. A three layered structure with endothelial cells seeded on inside of SMC seeded
collagen media with outer layer of adventitial fibroblasts was grown for 4 weeks. A Dacron
mesh sleeve was used for support between media and adventitia and the construct displayed
a burst strength of 300 mmHg after 3 weeks. Despite relatively weak mechanical results,
this study demonstrated the feasibility of this approach.
Hirai et al.[35,36] also used a collagen gel-based approach for autologous canine venous
substitutes. Smooth muscle cells were suspended in collagen and statically matured for 7
days. Following EC seeding, the construct was implanted into the posterior vena cava for
24 days after which 9 of 14 grafts remained patent. To prevent tearing, the grafts were
supported by a Dacron support. No burst strength measurements were performed; however
the effects of initial cell seeding density and initial collagen concentration on collagen
contraction were studied.
Ye et al.[37] demonstrated that similar two-dimensional cell seeded gels can be made with
from fibrin seeded with human myofibroblasts. Fibrin can be produced from blood
plasma. This same group then demonstrated its use for heart valves[38]. Although these
attempts were not aimed at blood vessel generation, they provided a basis for the following
studies. Grassl et al.[39] examined the feasibility of replacing collagen by fibrin for
vascular tissue replacements. They seeded fibrin gels with aortic rat SMC. SMC in these
15
constructs displayed a higher synthesis of collagen. However, the synthesized collagen is
not largely incorporated into the developing ECM. The same group pursued this further
with SMC seeded tubular gels of fibrin[40]. They found that fibrin media-equivalents
(ME) compacted more and were stronger and stiffer than collagen MEs.
Berglund et al.[41] demonstrated the feasibility of supporting collagen gels seeded with
neonatal fibroblasts with cross-linked collagen sheets. These constructs demonstrated burst
strengths of 650 mmHg due to the cross-linked supports but this strength deteriorated
during culture due to excessive degradation. The same group also used tubular collagen
gels seeded with rat aortic SMC and matured in a bioreactor with various biochemical
factors[42]. This group also studied the effect of a specific growth factor and mechanical
strain on the phenotype of SMC. They found that TGF-β strongly inhibited cell
proliferation and increased smooth muscle actin (SMA) expression, especially in the
presence of mechanical strain.
Summary
Collagen can be employed to fabricate arterial replacements in many different ways.
Cellular components of natural tissues such as blood vessels and intestinal tissue can be
removed to obtain a collagenous constructs. These constructs can be formed into tubular
structures and implanted in order to obtain in vivo cell in-growth. These constructs can also
be cell-seeded prior to implantation or implanted without cells with the goal of achieving
post-implantation cell migration. Another approach involves encapsulating tubes in vivo
with granulation tissue by using the body’s natural healing response. Thirdly, in vitro
cultured vascular cell sheets can be wrapped around a mandrel to form a tubular cell/ECM
structure. Lastly, a more traditional tissue engineering approach utilizes a scaffold structure
to provide a substrate onto which cells may adhere and proliferate into a tissue in vitro.
These last two tissue engineering approaches are presented below and are summarized in
Table 1-1.
16
Authors
Cells
Scaffold
No Scaffold
Synthetic Scaffold
L’Heureux et
al. , 1998 [21]
Niklason et al. ,
Hoerstrup et
Jeong et al., 2005
1999, 2003
al. , 2001 [31,32]
[33]
[29,30]
Human umbilical
Bovine aortic SMC
vein SMC and EC
and EC
and human skin FC
None
PGA
Biological Scaffold
Hirai et al., 1995, Grassl et al., 2002, Berglund et al.,
1996 [35,36]
2003 [39,40]
2003 [41]
Stegemenn et
al., 2003 [42]
Ovine carotid
artery
myofibroblasts and
EC
Rabbit SMC
Bovine aortic EC,
SMC and
adventitial FC
Canine jugular vein
SMC and EC
Aortic rat SMC
Neonatal human
dermal FC and
human coronary
EC
Rat aortic SMC
PGA coated with
P4HB
Rubber-like PCLC
Collagen gel
supported by
Dacron mesh
Collagen gel
supported by Dacron
mesh
Fibrin
Collagen gel
supported by crosslinked collagen film
Collagen gel
None
Implanted in canine
models
None
None
Pulsatile producing
radial stress
<100 mmHg after
3 weeks
Not measured
Not measured
650 mmHg
Not measured
Bioreactor
Non-pulsatile semiperfusion producing Pulsatile producing
radial stress
shear stresses and
radial distension
Pulsatile perfusion
Pulsatile perfusion in
in the lumen
the lumen producing
producing shear
shear stresses and
stresses and
pressure induced by the
pressure induced
flow
by the flow
Burst
strength
2000 mmHg after 3 2000 mmHg after 2
months
months
300 mmHg after 1
month
Major
Observations
Weinberg et
al. , 1986 [34]
Effects of dynamical Effect of pulsatile
mechanical stresses
flow
Not measured
Rubber-like scaffold
beneficial in delivering
mechanical signlas to
cells
Feasibility of
generating triple
layered structure
with collagen
Higher SMC seeding
density and lower
Feasibility of using
Fibrin stimulates
initial collagen
cross-linked
collagen production by
concentration
collagen to support
SMC
induced more rapid
gels
and prominent
shrikage.
Table 1-1 : Comparison of the major approaches using collagen for vascular tissue engineering to date.
TGF-β strongly
inhibited cell
proliferation and
increased SMA
expression,
especially in the
presence of
mechanical strain
The scaffold
As mentioned previously, one of main approaches to tissue engineering involves the use of
a scaffold to provide a matrix onto which the cells can organize and develop in the proper
environment prior to implantation (Figure 1.4). The scaffold provides an initial biochemical
substrate for the novel tissue until the cells can produce their own extra-cellular matrix[43].
This scaffold not only defines the three-dimensional space for the formation of new tissues
with appropriate structure, but serves also to provide tissues with appropriate functions[44].
Most cells are anchorage dependant and their growth is influenced by the substrate onto
which they are adhered. The scaffold surface properties such as chemistry and wettability
affect cell spreading and proliferation while scaffold structure affects cell spatial
arrangement and the transmission of biochemical and mechanical signals[45]. Also, various
signals provided by the scaffold may affect cellular gene expression. These signals may
include cell adhesion molecules, growth factors and mechanical signals.
Figure 1.4 : Tissue engineering. The basics.
Tissue engineering scaffolds must meet the following criteria[45]:
be biocompatible and meet the various nutritional and biological needs for the specific
cell populations.
be reproducible in three dimensional complex shapes.
18
be highly porous and structured to permit an adequate cell distribution for cell seeding
and permit diffusion of nutritional elements during cultivation.
be potentially tuneable with respect to their chemical, physical and mechanical
properties.
have controlled biodegradability.
In order to meet these requirements, the most common materials range from biological
scaffolds such as decellularized xenogenic matrices, small intestinal submucosa, fibrin
and collagen to synthetic materials such as Polyglycolic-acid (PGA), Polylactid-acid
(PLA), and the biologically derived polymer poly-4-hydroxybutyrate (P-4-HB).[44,43]
Different demands involved in various tissue engineering applications call for an educated
choice of scaffold material. Synthetic materials are attractive due to their controlled
manufacturability, mechanical properties and degradation times. For many tissues,
synthetic scaffolds appear to function quite well.
The cost involved in the conception and production of new materials precludes their
conception for very specific applications such as biomaterials. Often, research is conducted
with materials conceived for completely different applications. Inevitably, these materials
are not tailored to the specific requirements of biological applications. In vascular tissue
engineering, the use of a non-biodegradable synthetic material implies a complete loss of
vaso-activity and would also hinder the normal remodelling response of the vascular
system, thus becoming a physical barrier to long-term implantation[8]. Therefore, the
scaffold should be degraded or metabolized during the formation and organization of the
newly generated matrix at a rate in accordance with the rate at which newly synthesized
molecules are produced and incorporated in the regenerated tissue. This degradation should
not produce large debris which would cause occlusion problems down-stream nor
degradation products which would be toxic to the body. Although biodegradable synthetic
materials are available, biological materials are often favoured. One reason for this is that
vascular replacements that are composed in whole or in part of synthetic polymeric
materials remain at risk for bacterial colonization and subsequent infection, and are capable
of promoting a low-level, chronic inflammatory response that may lead to neointimal
hyperplasia[46]. These materials may fracture, induce immunological responses and are
difficult to anchor[47]. Also, the chemistry of the scaffold profoundly affects cellularity,
19
cellular gene expression and the overall tissue composition[48]. Natural ECM molecules
possess ligands to which cells adhere and are therefore more appropriate than synthetic
surfaces which do not possess natural binding sites. A major mechanism by which cells
bind to ECM is through integrins which exist as heterodimeric transmembrane
glycoproteins, consisting of an α and a β subunit. The binding and activation of integrins
promotes a signalling cascade within the cell that affects differentiation, activation, gene
expression, and proliferation[44]. The response of cells to external stimuli such as
mechanical stimulation has also been shown to be influenced by cell chemistry[49]. More
specifically for vascular tissue, compliance mismatch has been associated with graft
failure[50]. Hence, the compliance difference between synthetic materials and native
arteries may contribute to the high failure rate of synthetic prostheses.
20
Collagen as a scaffold for vascular tissue engineering
Collagen properties
Collagen is a prevalent protein in the vessel wall and, in-situ, plays a structural role as the
main load carrying element. It therefore constitutes a valid choice of scaffold material for
vascular tissue engineering. Furthermore, it is versatile and can be processed in a variety of
forms such as sheets, tubes, sponges, powder and fleece[51]. It can be solubilized by acidic
aqueous solution and can be engineered to exhibit customized properties. Collagen in the
form of thin sheets or gels has been shown to provide a suitable substrate for many
different
cell
types
such
as
renal[52],
hepatocytes[53],
epithelial[54],
smooth
muscle[34,55,56], endothelial[34] and fibroblast[57,58,59]. The following lists other
advantages and inconveniences of using collagen as a biomaterial[60].
Advantages:
9 available in large quantities and relatively easily purified
9 non-antigenic
9 adjustable biodegradability and bioresorbability by cross-linking
9 non-toxic and biocompatible
9 good resistance and tensile properties
9 compatible with synthetic polymers
Disadvantages:
8 high cost for collagen type I
8 high extracted collagen variability (cross-linking density, fiber size,
impurities, etc)
8 complex handling properties
8 variable enzymatic degradation
8 possible side effects and mineralization
21
As a scaffold for vascular tissue engineering, collagen is commonly used in the form of gel.
Collagen in acidic solution can be neutralized and mixed with culture medium. When
placed in an incubator at 37ºC for less than one hour, the solution forms a weak gel[57].
Various cell types can be suspended in this gel mould. Unfortunately, collagen in gel form
does not possess high mechanical strength, resulting in constructs that are too weak for
implantation. Although many research teams have supported collagen matrices with
synthetic materials, it has been mentioned previously why this is not a desirable approach.
Fortunately, this approach to vascular tissue generation is in its infancy and much can still
be improved without resorting to synthetic supports. Some of the main areas in which
improvements can be made with this approach include: cell phenotype and functionality,
collagen concentration and cell seeding density, collagen fiber orientation, cross-linking,
incorporating of other biological molecules, incorporating other biological support
structures, mechanical conditioning and vascularisation.
Cell functionality and phenotype
It is well known that cell growth and functionality is guided by the substrate. In mature
arteries, the collagen network plays a structural and signalling role in the physiological
response of SMC[56]. Scaffolds derived from natural biomaterials such as collagen have
intrinsic cell adhesion properties[48]. Arterial SMC have very different characteristics
when grown in culture than when found in normal vessels. The normal contractile
phenotype changes to a more proliferating, protein secreting mode when grown in
culture[61]. The morphology becomes less elongated, proliferation increases with each
passage in culture and the expression of smooth muscle actin (SMA) expression decreases
with each passage[62]. The factors that play a role in this phenotype change include the
form of the substrate, seeding density, the presence of ECM proteins and growth factors,
mechanical conditioning and the presence of endothelial cells.
The physical form of ECM molecules strongly influences gene expression of adherent
cells[48]. In collagen gels, SMC appear much more elongated, appear to be more in a
22
contractile state, and therefore grow more slowly than those grown on plastic or on twodimensional collagen sheets. SMA expression of SMC is also down regulated in collagen
gels. Cells seeded below confluence maintain a synthetic phenotype until confluence is
reached[63].
The ECM is known to affect cell function through both biochemical and mechanical
signalling pathways. Culture of SMC on different ECM molecules affects cell phenotype
by modulating morphology, proliferation and protein expression[62]. Collagen type IV,
elastin and laminin maintain cells in the contractile phenotype while fibronectin promotes
a change to a more synthetic phenotype[64,63].
Phenotype control can also be modulated by exogenous biochemical stimulation such as
growth factors. Growth factors are proteins which promote proliferation and migration of
cells via interactions with specific cell-membrane receptors. To this effect, ascorbic acid,
platelet-derived growth factor (PDGF), transforming growth factor β (TGF-β) and heparin
have been studied[65,62]. Table 1-2 shows an overall picture of the various growth factors
and their effects. SMC and FC reach confluence faster in the presence of ascorbic acid[65].
In gels, PDGF caused an increase in cell number and a decrease in SMA. These effects
were the opposite with cultures in two-dimensions. Although TGF-β causes an increase in
SMA and a decrease in proliferation on flat surfaces, this factor was found to have little
effect in three dimensions. Heparin slightly increases SMA in flat culture conditions but
has no effect in gels. It does, however, decrease the growth rate of cells in gels and on flat
surfaces. Therefore, the presence of a three-dimensional collagen matrix has an effect on
cell morphology, proliferation and SMA expression. This matrix also affects the cells
reactions to various biochemical signals. There is also a marked difference in SMA
between cells grown in disc-shaped gels as compared to tubular gels compacted around a
mandrel[42]. It is possible that the higher mechanical stresses in the latter case are sensed
by the cells which react accordingly. This may also explain the varied morphological
behaviour of SMC depending on their position in the disc-shaped gels and tubes[66]. In
discs, cells on the surface of gels are more flattened. Cells on the bottom of gels are more
organized into a cellular network than those in the middle which are more elongated. This
23
also applies to tubular gels in which the position of cells relative to the lumen may affect
phenotype.
Smooth
muscle
actin
Gel
Proliferation
compaction
↑
↓
↑
−
↑↑
−
PDGF (2D)
PDGF (3D)
Heparin (2D)
Heparin (3D)
TGF-β (2D)
TGF-β (3D)
↓
↑
↓
↓
−
N/A
↑
N/A
↓
N/A
↑
Elastin
synthesis
↑
↑
↑
↓
↓
↑
↑
Insulin
↑
Insulin-like growth factor-1
↑
Mechanical strain
References
Matrix
metalloproteinase
synthesis
↑
↑
↑
Ascorbate
Ascorbic Acid
Collagen
synthesis
[62]
[74]
[73]
↑
↑
[60,71,94]
[48]
↑
[42]
[65]
Table 1-2 : The effects of various chemical signals on smooth muscle cell seeded collagen
gels or films. (a blank box indicates no relevant influence was found in the literature.)
Cyclic distension induces SMC to a more pronounced contractile phenotype[67]. It has
been shown that SMC in the synthetic phenotype can revert to the contractile phenotype
after implantation[68]. Shear stress also affects EC functionality. These mechanical
environmental conditioning effects are discussed in more detail further on.
The engineering of a TEBV in this manner infers the co-culture of EC and SMC. This coculture adds increased complexity than simply referring to two simple cultures of each cell
type. EC and SMC can interact by two mechanisms: by body fluids and by direct
contact[61]. EC secrete both inhibitors and stimulants of SMC growth and SMA expression
such as TGF-β and PDGF. However, in co-culture, EC release more SMC growth
inhibitory factors such as a heparin-related glycosaminoglycan and transforming growth
factor β (TGF-β). Co-culture of these two cell types also results in completely different
24
endothelial properties. EC grown on SMC seeded collagen gels are more elongated than
those grown in culture dishes.
It is also possible with the collagen gel method to culture all three vascular cell types
together. It has even been demonstrated that it is unnecessary to separate SMC and FC prior
to gel formation. Vessel replacement constructs with a homogeneous mixture of SMC and
FC prior to implantation demonstrated a segregation of these two cell types after
implantation with SMC accumulating on the subendothelial layer and FC accumulating on
the outer layer[69].
Control over the phenotype of SMC is essential. Intimal hyperplasia caused by
excessive growth of SMC at the TEBV-artery interface is a major impediment to long-term
implantibility of TEBVs[18].
A major problem associated with TEBV without synthetic support is a lack of adequate
mechanical integrity, more specifically stiffness, strength and elasticity. Collagen tubular
constructs naturally stiffen during culture although this takes an extended period of time
and may not be sufficient. Therefore, many approaches exist to attempt to improve
mechanical properties of collagen vessel replacements. The main approaches to achieve this
are discussed below.
Modifications of collagen gel properties
Collagen concentration and cell-seeding density
When dealing with a collagenous tissue derived from a contracted gel, contraction
primarily determines the initial mechanical strength[35]. The two most basic factors that
affect collagen gel compaction are the collagen concentration in the initial solution and the
SMC seeding density. A higher SMC seeding density accompanied with relatively low
initial collagen concentration induced more rapid and prominent shrinkage[36]. Gel
compaction can also be affected by various biochemical and mechanical factors. Gel
compaction is increased with PDGF and TGF-β and decreased with heparin[62].
25
Mechanical stimulation in the form of cyclic distension caused an increase in
compaction[42]. Gel compaction is also affected by collagen type with collagen type I
promoting higher gel compaction than collagen type III[70].
Despite these findings, the overall effect of collagen concentration in initial solution is
somewhat limited. Greater strength can be achieved to some extent by increasing postcompaction collagen concentration. A higher cell-produced collagen concentration would
lead to increased mechanical strength. SMC are known to synthesize their own matrix
when grown on or in an adequate substrate and when in the appropriate phenotype[63]. In
this sense, an increase of endogenous collagen production by cells accompanied with an
appropriate rate of collagenase activity and enzymatic degradation of reconstituted collagen
would lead to improved mechanical properties. In vitro collagen synthesis can be
influenced by various biochemical factors such as: cyclic strains, growth factors, ascorbic
acid and amino acid supplementation[22]. Collagen synthesis is also increased by
mechanical stimulation[71,60].
Ascorbic acid is an essential cofactor for hydroxylation of proline and lysine residues in
collagens synthesized by human FC. It is also known to modulate the growth properties of
cells[37]. Ascorbic acid in low concentrations is essential for the production of collagen
and has been shown to increase type I collagen production by SMC and FC[72,48].
However, it also produces negative effects on elastin accumulation and production. These
effects are dependant on both the dose and the time of exposure. Combining ascorbate with
TGF-β and insulin increases collagen incorporation. TGF-β increases overall protein
synthesis while ascorbate increases the collagen fraction of protein synthesis[73]. Also,
TGF-β, plasmin and insulin were shown to increase collagen production by human dermal
fibroblasts[74].
The degradation of the matrix, involves enzymes known as matrix metalloproteinases
(MMPs) which include collagenases, gelatinases, metalloelastases, etc. Each degrades a
specific ECM component. These enzymes are produced by vascular cells and can be stored
in latent form until required[12]. Their activities are regulated by growth factors and
cytokines. These MMPs play a crucial role in vascular matrix remodeling and are essential
26
for cell migration, synthesis of new ECM components and regulation of growth factors.
MMP activity must be controlled and regulated to match that of collagen production in
order to maintain the structural integrity of TEBVs. MMP production, like all aspects
related cell function, may be modified by various biochemical and mechanical factors[42].
Fiber formation and alignment
Fibers of 5-50 µm can orientate and guide cells. This phenomenon was discovered by Paul
Weiss in 1945 and is called ‘contact guidance’. Collagen fibers have this ability to guide
and orientate cells. Collagen fibers also contribute to cell remodelling and are absorbable
over time as cells produce their own extracellular matrix[41]. The natural media is
composed of circumferentially aligned collagen fibers and SMC and this alignment is
important for both vasoactivity and structural integrity. Since cells can be orientated by
collagen fibers, by first of all guiding fibers to a circumferential orientation, the overall
media can aligned in this manner.
By using a cylindrical collagen gel seeded with smooth muscle cells, gel compaction can
occur around an inner mandrel and results in a natural circumferential orientation of both
collagen fibers and smooth muscle cells. This orientation can only occur if the gel is
allowed to contract longitudinally as well as circumferentially. This natural phenomenon
can be assisted by various means. Alignment can also be achieved using the effect of a
strong magnetic field during collagen fibrillogenesis[75]. Magnetic field fiber orientation
worked well with constructs without a central mandrel. However, mandrel compaction
proved to be more beneficial than magnetic pre-alignment on construct integrity. This
approach, however, may decrease the time to total contraction. More importantly, cyclic
distension via forced peristaltic flow through the lumen induces circumferential collagen
fibril/SMC orientation[76].
Cross-linking
Collagen can be cross-linked in order to increase its mechanical properties and molecular
stability. The most basic strategies introduce stable and covalent intermolecular cross-links
27
between collagen fibrils[77]. Cross-linking influences the strength, resorption rate, and
biocompatibility of biomaterials. Collagen molecules are endogenously cross-linked by
cells and can be artificially cross-linked in many different ways, although not all are
appropriate for vascular tissue engineering. There are two main conditions which must be
met to provide suitable cross-linking[77].
1.
Collagen fibrils must contain an amino-acid reactive group to be targeted by the
chemistry of the cross-linking agent and coupling mechanisms.
2.
The conditions such as temperature, pH and nature of solvent must not be
detrimental to the collagen molecule, fibrils or properties of the bioprosthesis.
There are basically three main ways to induce collagen cross-linking. These include
artificial methods, using biological compounds or inducing endogenous collagen crosslinking by cell.
Artificial Cross-linking
Chemical
Some chemical cross-linking agents include: formaldehyde, glutaraldehyde[78,79],
polyepoxy compounds[78,80], isocyantes[77], chromium and carbodiimide[81]. Although
the exact mechanisms may vary, these methods achieve cross-linking by forming covalent
bonds between adjacent collagen fibrils. Although effective at this, they are themselves
cytotoxic. Glutaraldehyde can autopolymerize and subsequently hydrolyze, releasing free
glutaraldehyde which is cytotoxic. Also, glutaraldehyde causes calcification and an
exaggerated increase of stiffness which is not beneficial for vascular tissue. Polyepoxy
compounds have been suggested as a replacement since they do not cause calcification nor
an exaggerated increase in stiffness[80]. In order to decrease cytoxicity, reagents such
carbodiimide (EDC) and N-hydroxysuccinimide (NHS) can be used. These reagents act
solely as catalysts for specific cross-linking reactions and are removed from the tissue
following the treatment. These above methods have been used with collagen materials with
which the goal is to perform subsequent in vitro or in vivo cell ingrowth. However, these
28
chemical methods, which are cytotoxic, may not be used on collagen gels suspended with
cells since the cells are present at the time of cross-linking.
Artificial cross-linking can also be achieved without the use of potentially harmful
chemical reagents.
Irradiation
Ultraviolet irradiation is efficient for the introduction of cross-links. Telopeptides on each
end of the collagen molecule are responsible for the photochemical reaction to ultraviolet
light. Collagen can be cross-linked in this manner in the form of films, fibers or while in
solution. In the latter, the viscosity of the solution increases with irradiation time and a gel
may be formed. Collagen is degraded by prolonged irradiation and the fibril formation
ability of collagen is easily deteriorated even after short periods of ultraviolet irradiation.
This irradiation also modifies the properties of collagen in collagen-cell interactions by
increasing cell growth. This effect can be deleterious for vascular tissue engineering.
Therefore, these approaches are mainly appropriate for collagen films and fibers. Recently,
a method involving visible-light photomediated cross-linking of collagen gels in the
presence of SMC was investigated[46]. Collagen was derivatized, through lysine and
hydroxylisine residues, with methacrylamide moieties. These moieties, in the presence of a
photo-initiator, were photochemically cross-linkable. This method showed some efficiency
in cross-linking while retaining collagen triple helical structure and high cell viability.
Dehydrothermal Treatment
Dehydrothermal treatment (DHT) is a physical method of cross-linking collagen fibers that
avoids potentially cytotoxic reaction products and provides moderate strength and
resorption rate[82]. Cross-linking is dependant on exhaustive removal of bound water from
collagen molecules which results in condensation reactions between the carboxyl and
amino groups on adjacent amino acid chains. This method can take at least three to five
days to complete and causes significant degradation to the collagen molecules. Evidently,
this method is not appropriate for collagen gels.
29
Glycation
Glycation is the nonenzymatic crosslinking of amine groups of collagen and other ECM
proteins brought about by reducing sugars, such as glucose and ribose. It results in
increased tissue stiffness and increased resistance to enzymatic degradation and is cell
tolerated. In order to increase glycation, it is simply necessary to increase the sugar
concentration in the medium, particularly that of ribose which is more effective than
glucose because it is about 17 times more available in open form[83].
Biological cross-linking compounds
Although not in current use in works related to vascular tissue engineering, there exist
various biological compounds capable of cross-linking collagen. Genipin, which is a
compound isolated from fruits of the gardenia plant (Gardenia jasminoides), cross-links
collagen in manner similar to gluraldehyde. The dialdehydes from Genipin react with εamino groups on lysine side chains of neighboring collagen molecules[77]. Another
compound, Nordihydroguaiaretic acid (NDGA), isolated from the creosote bush, produces
a different fixation mechanism. Cross-links are not formed between side chains of collagen
molecules but rather, NDGA polymerizes and is interpolated in the collagen network.
Fibers produced in this manner exhibit mechanical characteristics superior to those
obtained with all of the chemical reagents mentioned above. Furthermore, this method and
that using Genipin are less cytotoxic than the chemical reagents and create little foreign
body response or inflammatory reaction[77]. Enzymatic cross-linking methods, which can
be used for binding peptides and proteins, may also constitute a valid means for crosslinking collagen.[84]
Endogenous cross-linking
Cross-linking of collagen can be achieved by cells and their natural biochemical products.
In this sense, it is also possible to enhance the cells ability to cross-link. Endogenous crosslinking can be enhanced by ascorbic acid and ribose[35]. Lysly oxidase-mediated crosslinking is a significant contributor to stiffening. Enzymatic cross-linking of free amine
30
groups of lysine and hydroxylisine residues in collagen can be achieved by SMC-produced
lysyl oxidase[83]. Mechanical conditioning also influences the cells ability to cross-link
and organize collagen.
The overall objective of any cross-linking approach is to obtain adequate biomechanical
properties as well as a rate of degradation of collagen in tune with that of the repair process.
Incorporating other biological molecules
Other natural proteins have been assessed in their ability to enhance the mechanical
properties of collagen-gel based blood vessels.
Elastin
As mentioned previously, compliance mismatch has been associated with graft failure.
This would also be the case with regenerated vessels. Elastin contributes to compliance by
allowing stretch and recoil of the arterial wall with each pulse. Elastin fibers act in the
physiological range of deformation and provide the necessary resilience to recover from
deformations associated with pulsatile flow[85]. Consequently, an elastic network in a
TEBV would prevent vascular dilation in response to the continuous pressures exerted by
blood flow in vivo. Also, soluble elastin inhibits platelet aggregation induced by
collagen[22]. Therefore, incorporation of some form of elastin or stimulation of insoluble
elastin production in vitro would be greatly beneficial for a successful TEBV. Attempts to
do so are being pursued in one form or another.
Elastin is very difficult to process, due to purity and chemical contamination issues and a
strong tendency to calcify[20]. It is also very hydrophobic making it difficult to use in
matrix fabrication techniques such as collagen gels. Other forms of elastin may be used. It
was shown that incorporation of soluble α-elastin into collagen gels inhibits SMC
proliferation and migration limiting SMC hyperplasia without significant effects on
endothelial cell formation[18]. This soluble elastin protein also inhibits platelet aggregation
induced by collagen. Recombinant human elastin polypeptides may be expressed, produced
31
and purified. These polypeptides may prove to be a valuable component of a tissue
engineered vascular conduit[86]. A different approach combines the traditional cell-seeded
collagen gel approach with an auxiliary elastin scaffold isolated from porcine carotid
arteries[85]. This structure increased mechanical properties, especially creep resistance, as
compared to control gels. Collagen was also combined with elastin by lyophilizing various
mixtures of both molecules followed by cross-linking of the dried sponge[87]. An increase
of collagen ratio improved stiffness while increasing the elastin ratio improved elasticity.
Collagen appeared to act as glue, incorporating elastin into the scaffold and holding the
elastin fibers together. Pore sizes could be varied by changing the freezing rate prior to
lyophilization.
Due to the difficulties in processing elastin prior to incorporation in a TEBV, approaches
have been directed at enticing cells to increase their natural elastin production. SMC and
FC are known to produce elastin in vitro. ECM molecules likely influence elastogenesis.
Fibrillar and non-fibrillar glycoproteins and proteoglycans have been shown to increase
elastin synthesis[73]. Various biochemical factors also influence these cells to produce
more elastin. It was shown that TGF-β and insulin-like growth factor-1 increase
tropoelastin mRNA and protein synthesis by cultured cells. TGF-β also increases lysyl
oxidase activity which acts to crosslink tropoelastin into its mature fully functional
form[73]. However, ascorbic acid and ascorbate, which was shown to increase collagen
synthesis, inhibits elastin synthesis. It is important to note here that, elastin production by
SMC is maximal for neonatal cells and almost inexistent for adult cells[88]. This effect is
less severe with fibroblasts which retain their ability to produce elastin mRNA for longer
time periods. However, fibroblasts deposit little insoluble elastin.
Glycosaminoglycans
Most glycosaminoglycans (GAGs) are found in the form of proteoglycans on cell surfaces
and in the extracellular matrix[89]. The interactions between GAGs are essential for the
adhesion, migration, proliferation and differentiation of cells[90]. Incorporation of GAGs
into TEBV may therefore allow exploitation of their biocharacteristics and valorize
biomaterials like collagen. It was shown that various GAGs such as chondroitin sulfate
32
(CS), dermatan sulfate, heparan sulphate (HS) and heparin can be covalently cross-linked to
collagen using EDC-NHS[90]. This cross-linking method was mentioned above as suitable
for collagen cross-linking resulting in low cytotoxicity and adequate biocompatibility. This
study showed that CS in particular decreased the tensile strength of collagenous
matrices[91] forgoing it’s utilization to increase mechanical properties of gels. However,
other GAGs may prove to be more beneficial to this effect. The interest in incorporating
GAGs is not solely for mechanical integrity. For instance, heparin sulfate proteoglycans,
which are abundantly present in basement membranes, regulate cellular activities and
therefore may be used to achieve appropriate cellular adhesion and differentiation.
Furthermore, heparin, heparan sulphate and dermatan sulphate have been shown to exhibit
anticoagulant activity[92] and may reduce the need of systemic heparinization. Heparin
also improves endothelial cell proliferation and reduces proliferation of SMCs, thereby
helping to prevent the formation of intimal hyperplasia[93]. Various GAGs can also be
used as slow-release vehicles for growth factors and other biochemical molecules[94].
Mechanical conditioning
The influences of mechanical forces on cells in the vascular system have long been
recognized. Cyclic mechanical strain on two-dimensional substrates has been shown to
increase growth factor and matrix production and increase the expression of contractile
proteins[42]. These effects are also present in three-dimensions. It was demonstrated that
flow and cyclic stretch both have an effect on vascular biology[67]. Tension as a specific
form of mechanical stress also affects cell function and behaviour[76]. Also,
circumferential alignment of cells is extended throughout the media equivalent in
conditioned constructs. Unconditioned constructs demonstrated cell alignment mainly in
the immediate vicinity of the lumen.[95]. Dynamically stressed constructs have a higher
potential for reorganization and a higher orientation of SMC and collagen. Cell-mediated
remodelling of the tissue is thereby accelerated. The amplitude of strain also affects the
remodelling response with 15% distension having a larger effect than 5%[95]. Mechanical
stress enhances synthesis and secretion of proteins and mechanical stiffness, and has effects
on the level of gene expression. Tension as a specific form of mechanical stress affects the
33
cell function and behaviour[76]. Dynamic cyclic stretch improved mechanical integrity of
TEBV as compared to static flow and completely static culture[96]. Pulsed flow induces
increased production of collagen. The rate of pulsing also has an effect with 90 beats per
minute (bpm) inducing higher structural integrity than 165 bpm[97]. The presence of
MMP-1 was found to be greater in pulsed vessels and higher at 90 bpm than 165 bpm. It is
possible that pulsatile culture conditions induce a higher rate of collagenolysis.
Mechanical strain has been reported to stimulate extracellular matrix synthesis as well
as
secretion
of
enzymes
responsible
for
matrix
degradation
and
tissue
remodelling[71,95,60,97]. Mechanical stimulation also has an effect on the SMC reaction
to various biochemical factors and vice-versa[42]. Mechanical strain diminished the effect
of PDGF on both gel compaction and cell proliferation. The increase in cell proliferation
due to mechanical stress was suppressed by adding TGF-β.
Many teams have studied the effects of cyclic strain on a tubular tissue using an inner
plastic compliant tube as a mediator between the flow and the collagenous tissue. This
approach however, does not allow for shear stress which is an additional component of
mechanical conditioning. Shear stress, the tangential force acting in the direction of blood
flow on the surface of endothelial cells, also promotes the orientation and elongation of
these cells. It also has an effect on biochemical factors. In the regulation of vascular tone,
relaxation of SMCs is dependant on the integrity of the endothelium[98]. This is regulated
by the endothelium-derived relaxing factor (EDRF), whose release is regulated by shear
stress.
In order to simulate these mechanical factors that condition TEBV and improve their
properties, adequate bioreactors must be used. These bioreactors must provide complete
control over many different environmental factors such as temperature, culture medium,
chemical factors, and mechanical environmental forces such as shear stress due to flow,
cyclic radial distensions and perhaps even longitudinal tensile stress[99].
34
Vascularisation
A major impediment to vascularisation and nutrient supply and waste removal to and from
engineered tissues is the foreign body response which induces the formation of a fibrotic
capsule surrounding the implant. Molecules must diffuse through this barrier. By using
collagen as opposed to synthetic materials, this response is minimal, thereby favoring blood
supply to the tissue.
Furthermore, although tubular gels of collagen are relatively thin, the incorporation of a
microvascular network will insure adequate nutrient supply and excrement removal to and
from vascular cells throughout the entire thickness. A thickness of less than 1 mm is
required to block adequate diffusion of these elements without a microvascular
network[100].
Fortunately,
three-dimensional
collagen
scaffolds
can
support
vascularisation processes[60]. Cultured endothelial cells on or within ECM substrates such
as collagen, fibrin, fibronectin or laminin form capillary-like structures[101]. The ability
to form these tubes is dependant on collagen concentration and type[102]. Mechanical
factors such as cell adhesion and spreading also play a role. There also exist various proangiogenesis growth factors, such as acidic fibroblast growth factor (aFGF), basic
fibroblast growth factor (bFGF), vascular endothelial cell growth factor (VEGF),
angiopoietins and ephrins which may induce more-rapid and prominent capillary
formation[101,98]. VEGF has the most critical influence on vascular formation since it is
required to initiate the formation of immature vessels by vasculogenesis and angiogenic
sprouting[98].
35
Conclusion
The human body has developed over many centuries of evolution into a highly complex
system. We are no where near an understanding of the entire complexity nor, for that
matter, able predict the result of various factors on tissue formation and growth. It is
therefore all the more difficult to try and imitate evolution or the work of some higher
being and generate a tissue from scratch. Although some success has been achieved with
various tissues such as skin and cartilage, it is much more complicated when dealing with
three-dimensional tissues implicating multiple cell types. Generating a pseudophysiological artery is not a simple matter. A proper scaffold material must be found which
mimics the extra-cellular matrix onto which cells adhere and proliferate in vivo. One of the
most adequate materials to this effect is collagen. This collagen can be produced by cells or
reconstituted collagen can be extracted from various sources. Reconstituted collagen can be
processed into thin films or more appropriately cell-seeded gels which can be contracted by
the cells around a mandrel to form a tubular collagenous tissue.
Furthermore, the choice of cells is important. Each cell type affects the others. In order to
obtain adequate mechanical and functional tissue properties, all cell types must have their
proper functionality in order to influence the other cells in the correct manner. In order to
do this, the in vivo environment must be recreated as much as possible. To start with, the
scaffold material is essential to provide adequate cell adhesion, proliferation and
functionality. This is not sufficient however. The biochemical and mechanical environment
must also be recreated. Perhaps most importantly are the mechanical stimulations, such as
shear stress and cyclic distensions due to pulsed flow, which have great effects on both EC
and SMC. These forces may induce proper phenotype functionality in both cell types. By
doing so, EC and SMC can produce and transmit their proper biochemical signals.
Similarly, biochemical factors can alter tissue structure or function in such a way as to alter
the forces exerted on cells. Addition of other biochemical signals may be necessary in vitro
to replace those found in vivo but not produced by cells.
To date, mechanical integrity of collagenous tubes is not sufficient for implantation in the
arterial system. Burst strengths rarely exceed 2000 mmHg unless these constructs are
36
enhanced by a synthetic support. Nor do the tissue engineered arterial replacements
currently available possess adequate viscoelastic properties. Fatigue is also a factor as
implanted TEBV are exposed to conditions involving cyclical stresses for long time
periods. However, as discussed above, there are a multitude of factors at play that affect the
strength of these tissues such as collagen concentration and fiber alignment, cross-linking,
cell source and phenotype and biochemical and mechanical environmental factors.
Progress is continually being made in this relatively young field. As we progress, our
understanding of the requirements of cells increases which helps us to better improve the
next generation of arterial replacements in order to meet the challenges of tissue
engineering a vascular graft; that is providing a conduit that will have sufficient strength
not to burst with changes in blood pressure, a vessel wall that is elastic and can withstand
cyclic loading, matching compliance of the TEBV with the adjacent host vessel, and a
lining of the lumen that is antithrombotic[103].
This approach takes a great deal of time and would therefore not be designed for use in
emergency surgeries. However, with the increasingly efficient testing methods available
today, problems may be more and more predicted in advance leaving sufficient time to
culture a suitable TEBV with long-term capabilities.
The above chapter described the problem at hand concerning the lack of
suitable blood vessel replacements as well as some of the major approaches
to filling this lack. One approach in particular, involving using cell-seeded
collagen gels, was discussed in detail with a review of the possible means to
improve such tissue engineered blood vessels. The following chapter
presents this approach as it was undertaken in our laboratory including the
rationale, methodology and various unpublished result in order to work
towards the overall project objective which is to design and develop a
scaffold structure from collagen for a tissue-regenerated artery.
Chapter 2: Rationale, Methodology and unpublished
results
Collagen-based vascular models have been studied for many years. One variation of these
uses reconstituted collagen in the form of a hydrated gel which be used as a tissue
engineering scaffold material. This involves suspending cells in solubilized collagen,
followed by neutralization and incubation, thus initiating spontaneous reassembly of the
collagen into a gel. This process is known as fibrillogenesis. Subsequently to gel formation,
the cells reorganize the surrounding collagen matrix causing gel contraction, thereby
increasing the collagen density and mechanical integrity of the construct. This vascular
tissue engineering approach via a collagen gel/SMC suspension is very interesting and
worth exploiting. This approach has been presented and validated in Chapter 1. A schema
of the overall approach as performed in the author’s laboratory is depicted in Figure 2.1.
Figure 2.1 : Collagen cell-seeded gel approach to vascular tissue regeneration
A few major research institutes worldwide are currently using various modification of this
approach. Of all the groups working to this effect, the one from Georgia Tech[41] stands
out the most due to their use of a natural collagen reinforcement to add increased
38
mechanical integrity to their constructs. As mentioned in a previous section of this thesis,
many other factors can be modified with cell-seeded gel approach with respect to
reinforcements or with the actual cell suspension collagen gel. This approach is still in its
early stages and much research must still be performed. It is therefore an interesting and
worthwhile strategy with which this laboratory can go ahead. Furthermore, this project
aims at a long-term strategy which compliments our other, more immediate, strategies
dealing with finding suitable solutions for arterial inadequacies.
The work relative to this masters project and presented here in this thesis represents a first
step for the laboratory. It constitutes an exploratory work towards the conception,
realisation and characterisation of collagen scaffold structures for vascular tissue
regeneration. The initial goals were therefore on a more fundamental level. These included
choosing a collagen source and choosing and getting familiar with a collagen extraction
technique which should be validated for further studies. Subsequently, more tissue
engineering specific research was performed. Figure 2.2 depicts the manipulations
performed by the author for the present work.
Figure 2.2 : Strategy of the research project
39
Collagen: source
First of all, an adequate supply of reconstituted collagen type I was required. After
interactions with Marie-France Côté, a research assistant at the Saint-François d’Assise
Research Center, who was largely involved in research work concerning collagen with Dr.
C. Doillon, a researcher at the CHUL, a protocol for extracting collagen I from rat tail
tendon was chosen. This, combined with an adequate supply of rat tails from the Animal
facilities of the Hôpital Saint-Francois d’Assise (with the appropriate licences from the
Ethical Committee of the CHUQ) made this option an obvious choice. Furthermore, tendon
is more adequate than skin in preparing a homogeneous and strong fibril dispersion for
biomaterial fabrication making it more suitable for mechanically strong shaped products.
One drawback to tendon collagen is its higher platelet activation although the difference is
not extensive with respect to other collagens[78].
Collagen: extraction and processing
Many different processing methods were investigated in the past but the most basic and
appropriate one for tendon extraction is by solubilizing in acid. Basically, tendons are
removed from frozen rat tails using forceps and are soaked in Phosphate Buffer Saline
(PBS) for the duration of tendon extraction. The tendons are then rinsed in acetone and
isopropanol 70% for 5 minutes each. Tendons are then soaked in acetic acid 0.04N for 48
hours during which time they dissolve into a gel-like mass. The gel-like mass is mixed in a
blender and frozen, then lyophilized to form a dried collagen sponge. The sponge is stored
until needed at -80ºC. To obtain a working solution with a known concentration,
appropriate amounts of collagen sponge and acetic acid 0.04N are mixed in a blender. The
solution is then centrifuged to remove collagens other than type I and other undesirable
debris. To make this solution appropriate for thin film fabrication, degassing must be
performed. This is achieved with a vacuum in an Erlenmeyer flask. For eventual
applications with cells, collagen must be sterilized. This is achieved by dialysis with
chloroform 1% for one hour followed by continued dialysis in sterile acetic acid 0.04N for
6 days changing the solution every two days.
40
Collaborative scientific research was done with the research team of Dr. A. Sionkowska
from the Faculty of Chemistry of Nicolas Copernicus University in Torũn, Poland since
this research team has an extensive history of characterizing collagen films made from
collagen extracted from rat tail tendon[104,105,106]. Upon comparison of our respective
extraction and processing techniques, we observed significant differences in each method.
The processing method used by our collaborators in Torun does not include mixing in a
blender, freezing at -80ºC and lyophilization. The gel-like mass is immediately centrifuged
to obtain their final solution. The details of the two methods are presented in Annexe B.
Qcoll refers to the Quebec extraction and processing method presented above, whereas
Tcoll refers to the Torũn method.
Collagen films: characterization
The final properties of collagenous products depend on the processing method[78] and
therefore, the variations in these two methods were deemed too important to ignore. It was
initially unknown what kind of changes may occur to the collagen molecule and/or to the
properties of the final products we wish to obtain from this collagen. The additional steps in
the Qcoll method may alter collagen properties. This method also has an increased
processing time before the stock solution can be obtained. However, we feel that this
processing method is preferable to the alternative method. The foundation of this
conclusion is explained in detail in an article published in 2005 by Macromolecular
Bioscience entitled: ‘‘Preparation of a ready-to-use, stockable and neutralized reconstituted
collagen. Effects of the preparation method on properties of collagen films’’. In this article,
which constitutes the first part of Chapter 3 of this present work, we compare Infra-red
spectra, tensile testing properties and fibroblast viability on acidic and neutral collagen
films. Generally, results indicate that no significant differences exist between collagen
produced by either of these two methods.
Validation of the processing method using thin films is appropriate considering their
relative ease of reproducible fabrication, their ease of characterization and their planned use
as a reinforcement for the final three-dimensional collagen gel construct. Making these thin
films in a reproducible manner for homogenous mechanical assays initially proved to be
41
troublesome. Before degassing, it is impossible to obtain flat films without bubbles.
Degassing solved this problem but the substrate on which thin films are dried is also
important. Many plastic types were tried to allow proper and homogeneous film fabrication
and which permitted removal of the film without damage. Of all the substrates tested, noncell-treated Petri dishes proved to be the option of choice. Not only did they provide nonadhesive properties allowing film removal but they allowed using large enough volumes of
collagen solution to make thicker films.
Once films were fabricated, samples for traction tests needed to be cut. A proper sample
shape was determined from the ASTM standard concerning the Standard Test Method for
Tensile Properties of Thin Plastic Sheeting[107]. This standard was judged to be the most
appropriate for the situation. This rectangular sample shape (50 mm x 5 mm) was also used
to some extent in literature[108,109]. To cut the samples in a reproducible manner while
limiting defects which may cause erroneous results, a custom-made punch (Figure 2.3) was
fabricated by Sebastien Blanchette of the ‘Départment de génie des mines, de la métallurgie
et des matériaux’ of Laval University.
Figure 2.3 : Custom-made cutter for thin film mechanical testing sample preparation.
Parallel to results presented in the first part of Chapter 3, an electrophoresis gel was used to
compare the collagen solutions obtained with the Qcoll and Tcoll methods to that of type I
rat tail collagen purchased from Sigma. This test demonstrated no differences in the
molecular size of components obtained with either extraction method or that obtained from
Sigma (results not shown).
42
Collagen films: cell viability and hemocompatibility
Once validation of the extraction method was performed, it was important to determine the
validity of neutral collagen films as substrates for the main cell types which will be used in
the project. Therefore, the behavior of smooth muscle cells and endothelial cells seeded on
neutral collagen films was investigated.
Reconstituted collagen from rat tail tendon is the major non-cellular component of our
tissue engineered arteries. In the course of maturation and eventual implantation, it will
almost surely come into contact with blood. A confluent endothelium is a major goal of the
overall project, however, complete endothelial cell coverage can not be guaranteed,
especially under flow conditions. Furthermore, the rate at which reconstituted collagen is
degraded and replaced by cell synthesized collagen is yet unknown. It is valid to believe
that reconstituted collagen will still be present in the vessel substitute upon implantation. It
is therefore important to validate the use of our collagen in contact with blood. Collagen in
the body is a natural coagulant and collagen type I from rat tail tendon has been shown to
have an increased platelet activation capacity[81]. Because of this, it is all the more
important to valid its use as a scaffold material for vascular tissue engineering. This
validation was performed and has been published in 2005 by Biomaterials in an article
entitled: ‘‘Biological Performances of Collagen-based Scaffolds for Vascular Tissue
Engineering’’. This article constitutes the second part of Chapter 3 of this present work. In
this article, smooth muscle cells and endothelial cells are grown on neutral collagen films
and analysed by MTT and scanning electron microscopy (SEM) for morphology and
proliferation. Also, various blood contact assays were performed. Results indicate that
neutral collagen films provide an adequate substrate for vascular cells and do not enhance
clot formation or platelet adhesion.
Collagen gels (without cells): feasibility
Having had no previous experience working with cells, I started out the project
experimenting with collagen and its gelling properties without cells. This was done in the
form of discs and in the form of cylinders. Collagen gels without cells do not contract.
They are also much weaker than their cell-seeded counterparts.
43
Discs
Disc-shaped gels without cells were made in 12-well plates for preliminary testing. The
protocol obtained from Marie-France Côté for collagen gels (Annexe C) was varied in
order to optimize the strength of the gels. No appropriate method was available to measure
the resistance of these collagen gels. This was not a major objective of this work and
therefore, non-quantifiable observations were used to characterize gel integrity in the goal
of becoming familiar with collagen and its gelling properties. The strength is given by
numbers 1 through 4, with 4 being indicative of a stronger gel. These observations were
done by manipulating the gels and determining the overall integrity and ability to be
manipulated. The results are presented in Table 2-1.
Gel #
Collagen 3 mg/ml
NaOH
PBS
1
50
25
0
2
75
25
0
3
50
25
25
Percentage of final volume (%)
4
5
6
7
8
50
50
50
50
50
25
25
25
0.5
0.5
12.5
0
0
12.5
25
NaCO3
0
0
0
4
0
0
12.5
2
Relative gel strength
12.5
3
0
1
12.5
1
0
1
9
50
0.5
25
10
50
12.5
12.5
11
50
12.5
12.5
0
1
0
2
12.5
3
Table 2-1 : Disc-shaped collagen gel integrity depending on the relative amount of various
ingredients. Relative gel strength is given by 1 to 4, 4 being the strongest gel.
Tubes
Cylindrical gels were also made without cells in a custom-made tube-mandrel system. The
tubes consisted of plastic round-bottomed tubes with a rubber stopper part-way down. The
tubes were pierced near the bottom to allow air entry which is necessary to allow removal
of the gel from the tube. A mandrel with a second rubber stopper was inserted in the tube
during gel formation (Figure 2.4A). Since these gels do not contract, once adequate
gelation took place, they were removed from the tubes while still on the mandrel and let to
dry horizontally while in rotation on the rotation device (Figure 2.5). The gels dried in this
manner and shrunk to form dried tubes around the mandrel. Immediately after gel
formation, all gels were 5 cm long. The average length of the gels after removal from the
tube was 4.0 ± 0.3 cm. The manipulations required to remove the gel and mandrel from the
tube caused some medium to seep out of the gel, thereby causing shrinkage. The average
length of the dried cylindrical films was 2.2 ± 1.0 cm. This reduction in length can be
44
avoided by making the gels adhere to rubber stoppers at each end of the mandrel.
Reproducibility was difficult to achieve. A means of obtaining longer coatings on the
mandrels is to perform multiple layers. In this way, it is possible to completely cover the 5
cm long mandrel.
Figure 2.4 : Custom-made tube mandrel set-ups. A: Initial set-up for cylindrical dried-film
preparation, B: Modification of initial set-up for cell-seeded gels, C: Final set-up allowing
gas exchange for cell-seeded gels. Custom-made rotation device A: front view, B: side
view.
45
Figure 2.5 : Custom-made rotation device for cylindrical gels. A: isometric view, B: top
view, C: front view.
The goal of making tubes in this manner was that it may provide a way to incorporate a
collagenous support structure on the lumen of the cell-seeded gels produced subsequently.
Another option is to make flat thin films and roll them onto the mandrel. This is a feasible
option but has multiple problems concerning the rolling process and maintaining sterility.
On the other hand, the cylindrical films can be dried under the hood to maintain sterility.
Also, by forming the film directly on the mandrel with which we will be subsequently
forming the gel, it is possible to avoid certain problems with rolling the films. Furthermore,
making films with culture medium, in order to induce fibrillogenesis, proved to be
troublesome. The films were brittle and were highly inhomogeneous and therefore would
prove quite difficult to roll without breaking. The same is true with gels dried around a
mandrel, but since the method does not require removing the cylindrical dried film, this is
not a major problem. Once the collagen-SMC hydrated gel is poured into the tube
46
containing the mandrel, the film will become rehydrated and no longer brittle. Inducing
fibrillogenesis may be beneficial from a mechanical standpoint and a cellular reaction
standpoint. A thin layer of neutral collagen may also prove beneficial to prevent excessive
mandrel contraction. This layer can be obtained by simply dipping the mandrel
successively in a neutral collagen solution.
Collagen gels with cells
Following validation of the collagen processing method using thin films and initial threedimensional gels without cells, we performed preliminary assays of collagen gels seeded
prior to gelation with smooth muscle cells. Smooth muscle cells were extracted from pig
aorta using the protocol in Annexe B. Collagen gels were made using a protocol provided
by Marie-France Côté (Annexe C). Briefly, reconstituted collagen is mixed with
appropriate amounts of concentrated medium, serums and a suspension of smooth muscle
cells in cell culture medium. Adding salt and increasing the pH to a physiological level,
causes collagen to form a gel after incubation. The time for incubation depends on the
volume and composition of the gel. An incubation time of 1 hour was sufficient for all our
studies.
Discs
Gels in the form of discs were made with the above protocol seeded with porcine aortic
SMC, 3T3 FC and porcine aortic EC. First of all, the extent of gel contraction with smooth
muscle cells and fibroblasts was determined by measuring the diameter of gels in a 6-well
plate (Figure 2.6).
47
Diameter (cm)
3
SMC
FB
SMC + FB
2.5
2
1.5
1
0.5
0
0h
24h
48h
72h
1 week
Time
Figure 2.6 : Gel contraction seeded with smooth muscle cells (SMC) and fibroblasts (FB).
Tubes
Fabrication and cell viability testing
It has been suggested that cell distribution and attachment in three-dimensions are mainly
determined by gravity[47]. In literature, other groups working with cylindrical gels remove
the gels from the tubular mould after 24 hours and let the gel mature in a Petri dish. This
does not appear to be an ideal method to obtain a homogeneous cell distribution. It was
therefore deemed necessary to fabricate a tube-mandrel system permitting rotation of the
gel for extended periods of maturation. The custom-made rotation device presented above
(Figure 2.5) was used once again. However, the tube-mandrel system required
modifications.
Many experiments were performed to find an optimal way of obtaining gels in this manner.
Mainly, pre-gel solutions were poured in a plastic test tube with an inner mandrel. The test
tube was initially closed at both ends with rubber stoppers which also kept the inner
mandrel as much as possible in the center of the test tube (Figure 2.4B). Histology and cell
cycle analyses were performed on these gels to determine cell viability. Cells grown in
cell-treated Petri dishes were used as a control. The initial set-up did not allow for adequate
48
gas exchange limiting the time with which gels could mature. It was initially thought that
by changing the medium every day, cells would not suffer substantially from this lack of
gas exchange. Preliminary cell-cycle analyses using a Fluorescence Activated Cell Sorter
(FACS) demonstrated that the opposite was true. Cells were not viable after three days.
Histology using Hemotoxylin-eosin staining confirmed these results showing rounded cells
(results not shown). Information concerning the basics behind cell cycle analyses will be
provided below.
It was therefore necessary to modify our initial set-up to allow for continual gas exchange.
This is not evident when dealing with a tube completely filled with medium in horizontal
rotation. An initial approach consisted of using CO2 independent culture medium in the
tubes. This medium, which is designed to require less gas exchange than ‘normal’ culture
medium did not provide an adequate solution as cell viability, as determined from cellcycle analysis, for more than one day was not improved compared to the initial set-up
(results not shown).
Finally, the following set-up was found and validated. A 15 ml test tube was used as a
mould. The bottom was cut out and replaced by a rubber stopper and mandrel. In order to
close the tube and allow gas exchange, a cap from a 25 cm2 tissue-culture flask was
employed (Figure 2.4C). To our knowledge, caps with a membrane for gas exchange are
not available for test tubes and this was therefore the most appropriate solution. The caps,
even those from the same company, were not entirely compatible with the tubes. Parafilm
was extensively used to assure an adequate seal and minimize the risk of contamination. A
representative collagen tube obtained in this manner is shown in Figure 2.7.
49
Figure 2.7 : Typical collagen sell-seeded tube obtained with the gas-exchange tube-mandrel
system
In order to demonstrate that this set-up allowing gas-exchange is adequate for extended cell
growth, more analyses were performed.
Masson’s trichrome staining was used to provide histological observations. Briefly, three
different dyes were used in order to differentiate between cells and extracellular matrix:
Celestine blue solution was used for cellular staining (dark blue-black), while acid fuchsin
solution and methyl blue solution were used for respectively staining elastin (pink) and
collagen (blue). Figure 2.8A shows the Masson’s trichrome staining used for both smooth
muscle cells and extracellular matrix detection (collagen and elastin). Observation at low
magnification showed several areas with a darker coloration revealing a higher collagen
concentration. These zones are mainly located in the proximity of cells, thereby indicating
synthesis of neo-collagen by smooth muscle cells. Figure 2.8B shows smooth muscle cells
spread in the scaffold, forming a cellular network. Elastin was not detected in any of the
samples.
50
Figure 2.8 : Masson’s trichrome on a longitudinal a) and transversal b) section of tubular
collagen gel scaffold seeded with smooth muscle cells after 1 week maturation.
Another method of determining the viability of cells in a three-dimensional structure such
as this is with cell cycle analyses. Briefly, collagen gels are digested with the appropriate
enzymes (collagenase type I) and the remaining cells are labeled and analysed with a BD
FACS flow cytometer (Becton Dickinson). This analysis method provides a means to
compare the proportion of cells that are in each phase of the cell cycle (Figure 2.9). The G1
and G2 phases are characterized by protein and RNA synthesis, the S phase involves DNA
synthesis and the M phase is when the cell divides. G0 is a ‘resting’ stage when the cell is
not actively dividing. The timing of the cell cycle and the relative lengths of the various
stages depend on the specific type of cell and the growth conditions. Moreover, cell cycle
analyses give information on the presence of damaged cells.
Figure 2.9 : The various phases of the cell cycle
51
The cell cycle analyses in Table 2-2 show no significant differences between the cells
grown in the tubular moulds and the control cells grown in two-dimensional culture. The
G0G1 and S phases have statistically similar percentages indicating a very similar cell cycle
state. Only the percentage of cells in the phase G2M is significantly higher for cells grown
in the tubular mould. Furthermore, results show few damaged cells in both cases (data not
shown).
Table 2-2 : Comparative table for the three main cell phases present in smooth muscle cells
after 1 week in collagen tubular gels (Scaffold) and in gelatin coated Petri dishes (Control).
In order to verify if the endothelial cells still demonstrate their natural anti-thrombotic
properties once seeded on the surface of the scaffold, platelet adhesion testing was
performed on gels with and without endothelial seeding. Endothelial cells reached a
confluent layer after one week, thereby confirming the scaffolds as an adequate surface for
adhesion and proliferation. Figure 2.10 shows the platelet adhesion on scaffolds before and
after endothelial cell seeding. In absence of an endothelium, a high number of platelets
adhered to the collagen surface and appeared highly activated and totally spread. On the
other hand, scaffolds after endothelial cell seeding completely inhibited platelet adhesion
and activation. In fact, single adhered not-spread platelets were observed on the surface.
This phenomenon was limited almost exclusively to intercellular gaps. The count of
adhered platelets confirmed a highly significant difference between the scaffolds before and
after endothelial cell seeding. A mean of 64±4 platelets adhered and spread on the nonendothelialized scaffold, whereas more than 5 fold lesser cells were counted on the surface
in the presence of an endothelial cell layer. The low standard deviation indicates a very
high reproducibility among all the samples tested.
52
Figure 2.10 : Platelet adhesion on collagen gel seeded with smooth muscle cells without
(A-C) and with (B-D) endothelial cells layer on the surface. Images were acquired with
different magnifications: 200x (A-B) and 2000x (C-D).
Mechanical properties
The ultimate goal of the gel/cell tubular structure is implantation and maturation in a
bioreactor. This maturation period allows the cells to contract the collagen gel and produce
ECM molecules in the ultimate goal that they will remodel the gel into a more structurally
and functionally adequate construct for vascular tissue replacement. For the time being,
statically grown gels will be tested for structural strength with the custom-made
compliance testing apparatus in the laboratory (Figure 2.11). The apparatus measures the
diameter of a cylindrical tube as a function of the pressure applied by a perfusion medium.
Therefore, information similar to a stress-strain curve can be obtained for these samples as
well as an indication of the burst strength of the tubes. This will provide information as to
the maximal pressure supported by the tube before implantation in the bioreactor.
53
Unfortunately, the actual compliance of these tubes was undetermined at this time.
Although the tubes were able to be manipulated and inserted in the compliance apparatus,
it was impossible to obtain a pressure read-out. The permeability of the gels was too high.
The water inserted in the lumen seeped through the walls thereby making it impossible with
the available set-up to obtain a detectable luminal pressure increase. Instead, the dilation of
these tubes was determined by simply pumping water through the tubes and measuring the
change in diameter with a laser micrometer. The initial diameter was roughly 8 mm and the
tube expanded to a diameter of 11 mm. At this time, a small tear in the gel was observed.
Figure 2.12 shows the gel before and after dilatation.
Figure 2.11 : Compliance apparatus
Figure 2.12 : Collagen tubular gels in the compliance apparatus. A: before dilatation, B:
after dilatation
54
Although not attempted in this work, fatigue tests can and will eventually be performed by
measuring the compliance at various time periods following cyclical stretching periods.
The parameters of these tests will be similar to those which will be applied with the
bioreactor. The frequency of oscillations must not exceed 200 beats/min and the pressure
will most likely be in the range of 0-100 mmHg since initial tubes are relatively weak. The
compliance apparatus does not maintain sterility or adequate conditions for cell growth.
Therefore, these tests will be performed for relatively short time periods (a maximum of a
couple hours).
Eventually, the collagen/cell tubes will be tested with the same apparatus after maturation
in the bioreactor (Figure 2.13). The bioreactor and its capacities have been extensively
developed by Katia Bilodeau in her MSc thesis entitled: Conception et validation d’un
bioréacteur spécifique à la regeneration du tissue artériel sous contraintes mécaniques
(2003)[110]. The goal being to obtain properties of the collagen/gel tubes similar to those
of natural blood vessels, ideally those of small diameter arteries. This aspect has been
covered in Chapter 1.
Figure 2.13 : Tissue engineered blood vessel (TEBV) bioreactor module.
The above chapter provides a detailed description of the rationale and
methodology undertaken during the masters presented in this thesis.
Various unpublished results are also presented here. The following chapter
55
presents more detailed published results pertaining to the validation of the
processing method for collagen extraction described in the above chapter.
56
Results
Preparation of a Ready-to-use, Stockable and Reconstituted
collagen.
J. Habermehla, J. Skopinskab, F. Boccafoschia,c, A. Sionkowskab, H. Kaczmarekb, G.
Larochea, and D. Mantovania
a
Laboratory for Biomaterials and Bioengineering, Laval University and University Hospital
Research Center, Québec City, G1K 7P4, Canada
b
Faculty of Chemistry, Nicolas Copernicus University, Torũn, 87-100, Poland
c
Human Anatomy Laboratory, University of Eastern Piedmont “A. Avogadro”, Novara,
28100, Italy
Macromolecular Bioscience 2005; 5:821-828.
57
Résumé
Le collagène est largement utilisé pour des applications biomédicales. Plusieurs méthodes
existent pour extraire le collagène des queues de rat. Malheureusement, la méthode
traditionnelle ne permet pas de l’entreposer facilement à long terme ni de contrôler
adéquatement sa concentration en solution. Du collagène a été extrait, à partir de queues de
rat et mis en solution acide par deux méthodes différentes. Des échantillons en forme de
films minces ont été fabriqués par évaporation de solvant. Pour améliorer la viabilité
cellulaire, des solutions de collagène ont été neutralisées avec du NaOH jusqu’à un pH
physiologique. Des caractérisations par spectroscopie infrarouge, de traction et de la
morphologie et la cytotoxicité avec des cellules fibroblastes ont été élaborées pour valider
les variations entre les deux procédés. Ces essais ont démontré que lors de la synthèse de
collagène en solution, la lyophilization et le mélange n’ont pas eu d’effet sur les propriétés
finales du collagène. Cette méthode d’extraction est donc adéquate et permet de mieux
conserver le collagène et d’exercer un meilleur contrôle sur sa concentration en solution. La
neutralisation préalable à la production des films a conduit à une diminution significative
des propriétés mécaniques mais a amélioré de beaucoup leurs performances biologiques.
58
Abstract
Collagen is a widely used material in biomedical applications. Although processes that
prepare collagen and collagen-based materials that show suitable properties after extraction
exist, a ready-to-use, easily stockable, with tailored collagen concentration has not yet been
developed. Using rat tail tendons, acid soluble collagen solutions were prepared by two
different methods. To improve cell viability of pure collagen films, solutions with
physiological pH were also prepared by mixing with NaOH. Specimens in the form of thin
sheets were then fabricated by solvent evaporation. Next, infrared spectroscopy, tensile
testing techniques as well as human fibroblast cell morphology and cytotoxicity were used
to validate the significant variations in the processes. The results demonstrated that, during
the synthesis of collagen stock solution, lyophilization and mechanical blending had little
effect on the final properties and therefore offers a method for obtaining solutions with a
more homogeneous and modifiable collagen concentration and longer storage time.
Neutralizing the stock solution with NaOH prior to solvent evaporation provided films that
had lower mechanical properties but significantly improved biological performance.
59
Introduction
Reconstituted collagen has been investigated for a variety of applications including drug
delivery, burn or wound cover dressings[111] or as a substrate for tissue
engineering[34,41]. Collagen is a significant material for biomedical applications. It has
low antigenicity[112], low inflammatory and cytotoxic responses[112,52,113], good
haemostatic properties[113], and it promotes cell growth[114,113]. Other notable
properties that are useful for tissue engineering applications include its high tensile strength
and its controllable biodegradability[79]. Collagen is the most abundant protein in animals,
and it is known to provide the majority of the structural and mechanical support to tissues
and organs[106]. Over 20 types of collagen exist[15], and their characteristic conformation
is the triple helix, consisting of three polypeptide chains coiled around each other. These
chains are composed of a repeating basic triplet: Gly-X-Y, where X is typically proline and
Y is hydroxyproline. Intra- and inter-molecular hydrogen bonds are responsible for the
stability of the triple helix. Such bonds can be inter-chain hydrogen bonds coupled by the
NH groups of a glycyl residue with the CO group of a residue in the neighbouring chain.
Bonds are also formed via water molecules participating in the formation of additional
hydrogen bonds with the help of collagen hydroxyl groups[16]. The great strength of
collagen fibers, however, originates from the stable intermolecular covalent bonds between
adjacent tropocollagen molecules[17].
Collagen is versatile and can be processed in a variety of forms such as sheets, tubes,
sponges, powder and fleece[51]. It can be solubilized into acidic aqueous solution and can
also be engineered to exhibit customized properties[60]. Collagen in the form of thin sheets
or gels has been shown to provide a suitable substrate for many different cell types such as
renal[52], hepatocytes[53], epithelial[54], smooth muscle [34,55,56], endothelial[34] and
fibroblast[57,58,59]. For several applications, collagen is generally processed using
aqueous-based liquid solutions,
and the properties of collagen-based materials are
influenced by the source of collagen and the methods of preparation[106]. First of all, the
extraction of collagen from its natural source is a multi-step process, and is generally
considered time-consuming. Secondly, collagen in acid-soluble solution may have a limited
60
shelf-life, while long-term storage and further manipulations could increase the risk of
contamination of the stocked solution. Furthermore, simple dissolution of tendons in acidic
acid does not favour control of collagen concentration. Therefore, a production process of
ready-to-use collagen involving an intermediate step that favours long-term stockability
and collagen concentration controllability is certainly in demand.
In this study, we investigated the effect of lyophilization and of mechanical blending of
collagenous acid-soluble solutions as compared to simply dissolving rat tail tendons in acid
solution. In fact, lyophilization provides a means to conserve extracted collagen for a much
longer period of time while limiting the risk of contamination in comparison to liquid
solutions. Moreover, it provides a means by which modifiable and reproducible collagen
solutions, with respect to collagen concentration, can be easily obtained. Although
variations of these two protocols already exist, such a comparison has never been
investigated. However, this is required in order to monitor that the variations in the
extraction methods do not significantly affect the final collagen solution, nor the properties
of the thin films produced from such solutions. In fact, acidic solution has been used to
fabricate solvent evaporated thin films, which have been extensively characterised
[104,106]. In this study, although their low pH was not ideal for cell culture, acidic films
were used as substrates for assessing fibroblast cell viability. Cell compatibility was
improved by neutralizing collagen solutions. Moreover, neutralization is hypothesized to
increase the mechanical properties of collagen specimens by increasing the stability of the
collagen molecule [115,116]. Therefore, the objective of this work was to compare
chemical and mechanical properties as well as cell compatibility of collagen films, which
are prepared by two methods differing in that one method involves lyophilization and
mechanical blending.
61
Materials and Methods
Collagen solutions
Two different solutions of collagen (A and B) were prepared and compared in this study.
Collagen type I was extracted from rat tail for both solutions. This method was previously
described by Erhmann et al.[117] Briefly, it involved extracting tendons from young albino
rats and their dissolution in acetic acid. For collagen solution A, tendons were rinsed in
deionised water and dissolved in 0.04 M acetic acid, following a method previously
developed by our group.[105] After 72 hours at 4ºC, the solution was centrifuged, degassed
in vacuum and sterilized. The final pH of this solution was 3.5. Aqueous solutions were
analyzed by the Biuretic Protein Method to determine the collagen concentration. Collagen
solution B, a modified method of the above involving lyophilization for improved longterm storage and reproducibility of subsequent collagen solutions, involved dissolving
tendons in 0.02 M acetic acid for 72 hours at 4ºC followed by blending on ice to form a
homogeneous viscous solution. This solution was frozen at -20ºC and lyophilized. The
lyophilized sponge was mixed in a blender with 0.02 M acetic acid at a dry weight to
solution ratio of 4mg/ml. The resulting solution was centrifuged and degassed in vacuum
followed by sterilization. The B solution had a final pH of 3.7.
Sterilization
Both collagen solutions were placed in dialysis bags (Spectra/Por 1, MWCO : 6 – 8,000)
and soaked in acetic acid 0.02 M for one hour followed by one hour in chloroform 1% in
water. Dialysis was continued in sterile 0.02 M acetic acid for 4 to 5 days. This solution
was changed daily.
Neutral solutions
Neutral solutions were prepared by adjusting collagen solutions to a pH of 6.8 by adding
NaOH (1% v/v) on ice and were then immediately made into films with the method
described below.
62
Preparation of collagen thin films
Samples were prepared in the form of thin films by solvent evaporation from collagen
solutions poured into non-treated Petri dishes (Fisher Scientific) at room temperature. The
volume of solvent was adjusted to obtain films with the same mass per area of collagen.
The resulting films were peeled off and used in the following chemical and mechanical
studies. Films made similarly on glass cover slides (Bellco Glass inc.) were used for cell
culture assays.
Infrared spectroscopy
The infrared spectra of films were collected on a Nicolet, Magna 550 Fourier transform
spectrophotometer (Thermo Nicolet, Madison, WI, USA) equipped with a deuterated
triglycine sulfate (DTGS) detector and a germanium-coated potassium bromide beam
splitter. One hundred scans were routinely acquired with a spectral resolution of 4 cm-1.
The ATR mode was employed using a Split Pea attachment (Harrick Scientific Corp.,
Ossining, NY, USA) equipped with a silicon hemispherical 3 mm-diameter internal
reflection element (IRE). To obtain more detailed information about the secondary structure
of macromolecules, we carried out the decomposition of the overlapping components under
the amide I counter bands. The curve fitting was performed with GRAMS (Salem, NH,
USA), using a linear baseline and Gaussian peak-fit components. The position of
components was defined by second derivative analysis in compliance with literature data
and the bandwidth at half maximum was fixed between 25 and 30 cm-1 which is a typical
width range for peaks related to the vibration of carbonyl groups.
Mechanical characterization
Films were cut into strips 50 x 5 mm with a custom made cutter. Thickness was measured
with a micrometer (Mitutoyo, Japan). Films were 23 μm thick (±15%) for acidic collagen
films and were 35 μm thick (±15%) for neutral films. Tensile tests were performed with an
Instron 5848 microtester apparatus (Canton, MA, USA) at a cross-head speed of 5 mm/min.
The ends of the strips were placed between pneumatic grips at a pressure of 345 kPa (50
63
psi). A gauge length of 25 mm was used. Samples were conditioned at room temperature
and a relative humidity of 20% for 4 days prior to testing. Samples were weakly stressed to
0.15 N one minute before testing to ensure the removal of crimps and folds. Crosshead
displacement was measured with an optical extensometer. Data was recorded with Merlin
operating software and analysed with Excel (Microsoft, WA, USA). Results are expressed
as mean ± standard deviation.
Cell culture
NIH 3T3 fibroblasts were cultivated in DMEM medium supplemented with 10% foetal
bovine serum, glutamine (2 mM), penicillin (100 U/ml), streptomycin (100 μg/ml),
fungizone (0.25 μg/ml) (all from Sigma Aldrich, Milwaukee, USA). Cells were routinely
maintained at 37ºC in 5% CO2 – humidified atmosphere and harvested by trypsinization
when a monolayer was reached. Cells were seeded onto specimens with 3.5x103 cells in 50
μl medium. Medium was added after 1 hour of adhesion. Cells seeded onto cover slides
appropriately treated for cell culture were used as a control.
Scanning Electron Microscopy (SEM)
Cell morphology after 24h and 1 week was investigated by SEM. Medium was removed,
samples were washed twice in 0.15 M cacodylate buffer and fixed for 30 minutes at 4ºC
with Karnowsky solution (2% paraformaldehyde and 2.5% glutaraldehyde in 0.15 M
cacodylate buffer, pH 7.4). Following fixation, samples were treated for 30 minutes with a
2% osmium tetroxide in 0.15M cacodylate buffer solution. Samples were then dehydrated
with graded ethanol (from 50% to 100%), soaked for 30 minutes in hexamethyldisilazane,
dried, mounted on appropriate stubs with colloidal silver and sputter-coated with goldpalladium. Images were collected using a JSM-35CF (JEOL) scanning electron
microscope.
64
Cytotoxicity
After 24h and 1 week, cell viability was assessed using the MTT colorimetric assay
(Sigma Chemical Co., St. Louis, MO). Briefly, MTT, [3-(4,5 dimethylthiazol-2-yl)-2,5diphenyl-2H-tetrazolium bromide], is reduced to purple formazan by mitochondrial
dehydrogenase in cells indicating normal cell metabolism. After each time point, 100 μl of
MTT (1 mg/ml in PBS) was added to the medium and samples were incubated for another
4h. After this final incubation, medium was removed and purple formazan crystals were
solubilized with 100 μl DMSO at room temperature. The optical density (OD) at a
wavelength of 550 nm was determined using an ELISA reader (BioRad mod. 450). All tests
were performed in triplicate and repeated twice. The cytotoxicity rates were calculated
from the OD readings using controls as 100%. Results are expressed as mean ± standard
deviation. Statistical significance was determined by the Student’s t-test (p<0.005).
Results
Infrared spectroscopy
The infrared spectra of proteins exhibit several features characteristic of the molecular
organization of these molecules. Generally speaking, amino acids are linked together
through peptide bonds giving rise to several infrared active vibration modes such as amide
A and B (near 3330 and 3080 cm-1, respectively) and amide I, II, and III (near 1650, 1550
and 1250 cm-1, respectively, see Table 3-1).[118] Several studies have demonstrated that
the amide I spectral feature can be used to determine the secondary structure of proteins.
This fairly broad spectral component, which arises mainly from stretching vibrations of
C=O groups in the peptide groups, is composed of several underlying features - each of
them characteristic of a particular protein secondary structure. For example, the antiparallel
β-sheet structure of proteins is observed at 1624 cm-1, while the random coil conformation
gives rise to an infrared absorption at 1654 cm-1.[119] However, the infrared peak
assignment of the features underlying the amide I component in the spectrum of collagen is
somewhat different due to the particular triple helical structure of this protein. Nevertheless,
65
several studies have been performed in order to establish the relationship between the
collagen secondary structure and the band shape of the amide I infrared feature.
66
Peak Frequency (cm-1)
Ref.
A
B
A/Neutral
B/Neutral
Helices of aggregated
121
1694
1696
1693
1693
collagen-like peptides
Denatured collagen
120
1682
1683
1682
1684
Carbonyl groups from
123
1672
1674
1673
1675
collagen lateral chains
Collagen helices
17
1657
1657
1657
1657
Denatured collagen
120
1652
1651
1650
1651
H2O bending mode
123
1643
1643
1641
1642
vibration
Denatured collagen
122, 123
1628
1628
1630
1630
Denatured collagen
120
1612
1612
NA
NA
Table 3-1 : Assignments of the spectral features underlying the Amide I infrared band.
Table 3-1 displays the peak assignments, obtained from literature, for each of the spectral
features underlying the amide I peak of collagen samples. These data were thereafter used
to curvefit the amide I peak of the infrared spectra of collagen samples recorded in the
present work. As can be seen in Table 3-1, the infrared component observed near 1660 cm-1
is assigned to the triple helical hydrogen bonded conformation of collagen
chains,[120,121,122] while those observed near 1680, 1650, 1628, and 1612 cm-1 are most
likely due to denatured collagen.[123] In addition, the spectral component located near
1695 cm-1 is associated with helices of aggregated collagen-like peptides also found in
gelatin.[124,123] Finally, the feature observed near 1640 cm-1 can be assigned to the H2O
bending mode vibration. The structural properties of collagen samples were of particular
interest in the current study. To study this, the ratio of the integrated intensity of the 1660
cm-1 infrared spectral component over that of the peak at 1630 cm-1 (A1660/A1630) can be
used as a sensitive probe to monitor the degree of order within the various protein films.
Figure 3.1 shows the curve fitted amide I spectral features of the various collagen samples
investigated. As can be seen in this figure, eight components of about 25-30 cm-1 of width
at half maximum were required to fit the amide I peak in the infrared spectra of both acidic
collagen samples while seven were required for the neutralized samples. This is a clear
67
indication that the preparation methods for the acidic collagen samples lead to a partial
denaturisation of the protein. Indeed, the infrared amide I peak of either native collagen or
gelatin (denatured collagen) can be fitted using only four components with geometrical
characteristics similar to those utilized in the current study.
Figure 3.1 : Curve fitting of the Amide I infrared peak of collagen samples A and B before
(a and b) and after neutralization (c and d).
The peak assignments aforementioned along with comparison of Figures 3.1a and 3.1b
indicate that both samples A and B present an almost identical conformation in an acidic
environment as the curve fitting is very similar. In addition, the most important underlying
feature of the overall amide I peak of both acidic collagen samples is clearly the peak
assigned to denatured collagen, therefore demonstrating that the acidic environment drove
the collagen conformation to a less ordered structure.
68
Neutralizing both samples A and B with NaOH clearly modified the secondary structure of
collagen as shown in Figures 3.1c and 3.1d as an increase of the 1660 cm-1 band was
observed at the expense of the 1630 cm-1 band. It is therefore obvious that neutralization of
collagen leads to a more ordered structure. Moreover, this ordering effect seems to be
almost identical while neutralizing either A or B collagen samples. This reordering effect
could probably be due to a partial reformation of the collagen triple helix in the neutral
environment.
These visual observations were confirmed in a more quantitative basis in Table 3-2 where
the A1660/A1630 ratios are reported for all collagen samples. As shown in this table,
neutralization of samples A and B leads to an increase of the A1660/A1630 ratio from 0.80 to
1.61 and 0.73 to 1.60 respectively, therefore indicating a more ordered protein molecular
structure.
Area of Bands
A
B
A/Neutral
B/Neutral
A1660
2.18
2.40
1.71
2.07
A1630
2.98
3.00
1.07
1.29
0.73
0.79
1.59
1.60
A1660/ A1630
Table 3-2 : A1660/A1630 ratios measured from the curve fitted Amide I infrared feature of
collagen samples A and B before and after neutralization.
Mechanical properties
Tensile strength and elastic modulus (Figs. 3.2a, b) were not significantly different for
acidic collagen films but were affected significantly by neutralization. Neutral films were
weaker and less rigid. Maximum elongation (Fig. 3.2c) was basically the same for all
samples tested.
69
Figure 3.2 : Tensile properties of collagen films a) ultimate strength b) elastic modulus c)
maximum elongation.
Figure 3.3 shows a typical stress-strain curve of the collagen films. Acidic collagen exhibits
a typical reverse exponential shape. Neutral films exhibited a distinct two region curve as
previously observed for hydrated collagen gels.[125] The stress-strain curves presented an
exponential behaviour in the low stress region followed by a linear region until failure.
70
100
A
Stress (MPa)
80
B
60
A/Neutral
A/NaOH
40
B/NaOH
B/Neutral
20
0
0
2
4
6
8
10
Strain (%)
Figure 3.3 : Representative Stress-Strain curves of films made of acidic and neutral
collagen films.
Qualitative response of fibroblast NIH 3T3 cells to collagen films
To asses the biological performance of collagen films, the response of fibroblast NIH 3T3
cells grown on acidic (pH = 3.5) and neutral (pH = 6.8) films was investigated. After 24h,
cells seeded on both types of acidic collagen showed a rounded morphology indicating an
inadequate environment for cell growth (data not shown). No significant morphological
changes, with respect to controls, were observed with cells in contact with neutral collagen
films for all time periods studied. After 24h, cells were well spread and showed numerous
filopodia, indicating appropriate cell adhesion to the surfaces (data not shown). A
subconfluent monolayer was present after 1 week of culture. (Fig. 3.4a, b)
71
Figure 3.4 : Scanning electron microscopy (SEM) analyses of NIH 3T3 cells after 1 week
on neutral collagen A (a) and neutral collagen B (b).
Cytotoxicity
MTT results are presented in Figure 3.5. Results obtained after 24h showed a significant
decrease in viability of cells seeded on acidic collagen, as was expected, due to low pH.
After 1 week, both acidic collagens showed complete cytotoxicity. Cells seeded on neutral
films were deemed viable. Even though neutral collagen films revealed a lower percentage
compared to control, morphologic analyses showed good cell behaviour. Results are
statistically significant.
Cell viability (%)
100
90
24 h
80
1 week
70
60
50
40
30
20
10
0
A
B
A/Neutral
B/Neutral
Figure 3.5 : MTT assay testing cytotoxicity of different collagens on NIH 3T3 after 24h
and 1 week. Control is considered as 100% viability. All results are significant with respect
to control and between pure collagen (A and B) and neutral collagen (A and B) (p < 0.005).
72
Discussion
Infrared spectroscopy is one of the most powerful methods for studying the protein
structure and the conformational mobility of polypeptide chains. Of particular interest is the
analysis of the band shape of the Amide I spectral feature which is particularly sensitive to
the secondary structure of proteins[120,104].
In connection with the objectives of this study, FTIR spectroscopy enabled characterization
of the secondary structure of collagen samples from rat tail tendons obtained using two
different methods. Despite sensible differences in the preparation procedures, it turned out
that both A and B samples exhibit a very similar structure with an important content of
denatured collagen. Not surprisingly, neutralization of these samples promoted a
conformational change of the protein structure for both collagen samples demonstrated by a
clear increase of the helical content.
In these circumstances, one could hypothesize that the collagen conformational change
observed upon neutralization should be accompanied by a modification of mechanical
properties, in agreement with previous findings showing that collagen type I fibrils are
more stable and possess higher tensile properties at higher pH[126]. Unexpectedly, the
present results suggest that the partial disordered to helical transition is accompanied with a
decrease of collagen film mechanical properties. The different form of the collagen samples
investigated previously[126], and those evaluated in the present study, can be taken into
account to explain this opposite behaviour. In addition, one should emphasize that the
relative intensity of the infrared peak assigned to the water molecule bending mode
vibration near 1640 cm-1 is about 1.5 times more important (Figure 3.1) in acidic collagen
samples than in neutralized ones. Therefore, it seems that more water is bound within the
acidic collagen films. This bound water has been shown to increase the stiffness of the
collagen molecule by facilitating hydrogen bond formation between polar telopeptides in
adjacent collagen molecules[127]. Although neutral films were significantly less resistant
with respect to ultimate strength, their decreased elastic modulus may be beneficial for
vascular tissue engineering as this property, and the inherent compliance, is generally
much lower for natural arteries.
73
As expected, cells seeded on acidic collagen showed a rounded morphology typical of
apoptotic cells. In fact, pH is a critical factor for cell growth and in some cases could be
associated with apoptosis induction[128]. Even if further analyses are needed to confirm a
programmed cell death reaction induced by acid surface contact, there is no doubt that
acidic collagen cytotoxicity is caused by the low pH. In fact, in the case of acidic collagen,
the increased cell mortality, due to prolonged contact with an inadequate surface, is evident
after 1 week. Collagen neutralization decreased toxicity. In fact, cells show spindle-shape
morphology similar to control. A lower viability percentage is due to slower cell
proliferation when in contact with neutralized collagen, and not to an eventual apoptosis
induction as in the case of acidic collagen.
Other experiments have been conducted with similar neutralized collagen films
supplemented with an adequate amount of cell culture medium (data not shown). These
conditions allowed the same proliferation rate as the control. Unfortunately, although this
method has been shown to work extremely well for gels, dried films were too
inhomogeneous and brittle to consider for chemical and mechanical testing. Consequently,
this solution has been discarded for the time being.
Our group has previously shown that B collagen films display good biological results with
other cell types as well as good results when in contact with blood[129]. In fact, smooth
muscle cells and endothelial cells displayed good cell spreading and proliferation on neutral
collagen films. Furthermore, collagen films presented low thrombogenicity and displayed
low platelet adhesion during static assays.
Conclusion
Thus, we can conclude that lyophilization and mechanical blending during the synthesis of
collagen stock solution has no significant effects on final collagen properties and therefore
offers a method for obtaining solutions with homogeneous and modifiable collagen
concentrations and with more favourable storage capabilities. Neutralizing stock solution
prior to solvent evaporation provided films that had lower mechanical properties but
significantly improved biological performances. These findings are very encouraging for
future applications of this collagen for biomedical applications, most notably for vascular
tissue engineering scaffolds.
The above section presented results validating the collagen extraction and
processing method. The following section presents more validation results
pertaining to the biological and blood contact performances of the collagen
validated in the above section.
75
Biological Performances of Collagen-based Scaffolds for
Vascular Tissue Engineering
F. Boccafoschia,b, J. Habermehla, S. Vesentinic, D. Mantovania
a
Laboratory for Biomaterials and Bioengineering, Laval University, Québec City, G1K
7P4, Canada and Research Center St. François d’Assise Hospital, Québec City, G1L 3L5,
Canada
b
Human Anatomy Laboratory, Department of Medical Sciences, Research Center for
Biocompatibility, University of Eastern Piedmont “A. Avogadro”, Novara, Italy
c
Department of Bioengineering, Politecnico di Milano, Milan, Italy
Biomaterials 2005; 26 : 7410-7417.
76
Résumé
L’utilisation du collagène dans le domaine biomédical est largement répandue. Il représente
un matériau d’échafaudage alternatif pour l’ingénierie de tissus vasculaires. Ce travail
présente des études d’hémocompatibilité et de viabilité cellulaire sur des films de collagène
reconstitué. Tout d’abord, des essais de ‘haemoglobin-free’, de thromboélastographe et
d’adhésion plaquettaire ont été effectués pour déterminer la performance du collagène en
contact avec le sang. Par la suite, des échantillons ensemencés avec des cellules
musculaires lisses et des cellules endothéliales ont été observés par microscopie
électronique à balayage et avec l’essai MTT. Ces essais démontrent que le collagène de
type I soluble en milieu acide et neutralisé n’augmente pas la coagulation, n’altère pas les
propriétés viscoélastiques du sang et n’active que faiblement l’adhésion et l’agrégation des
plaquettes. De plus, ces films sont des substrats adéquats pour supporter l’adhésion et la
prolifération des cellules musculaires lisses et endothéliales.
77
Abstract
Collagen is widely used for biomedical applications and it could represent a valid
alternative scaffold material for vascular tissue engineering. In this work, reconstituted
collagen films were prepared from neutralized acid-soluble solutions for subsequent
haemocompatibility and cell viability performance assays. First, haemoglobin-free,
thrombelastography and platelet adhesion tests were performed in order to investigate the
blood contact performance. Secondly, specimens were seeded with endothelial cells and
smooth muscle cells, and cell viability tests were carried out by MTT and SEM. Results
show that neutralized acid-soluble type I collagen films do not enhance blood coagulation,
do not alter normal viscoelastic properties of blood and slightly activate platelet adhesion
and aggregation. Cell culture shows that the samples are adequate substrates to support the
adhesion and proliferation of endothelial and smooth muscle cells.
78
Introduction
According to the World Health Report (2003), 16.7 million people around the globe die of
cardiovascular diseases each year, nearly one-third of total global deaths[1]. Cardiovascular
diseases are often caused by atherosclerosis, which could lead to evolutive arterial diseases
like thrombosis and aneurysms[3]. When the substitution of a diseased vessel is necessary,
artificial prostheses, made of natural or synthetic material, are therefore required. Different
materials are currently used as synthetic vascular substitutes; for instance, Teflon is used
for prostheses of medium diameter (around 6mm) vessels[130]. However, no biomaterial
has yet shown satisfactory performances when in contact with blood for long time periods.
This leads to clinical complications, such as aneurysms, thrombosis or restenosis[131,132].
Moreover, no appropriate biomaterial has yet been developed for small diameter vessel
replacement.
Tissue engineering has already shown a strong potential for regenerating and repairing
chronic wounds[133], burns[134] and, at the experimental level, cartilage defects[135] and
represents a promising field in vascular reconstruction. The overall goal of vascular tissue
engineering is to obtain the same mechanical and biological properties as a native vessel.
Different approaches have been taken using collagen as a main scaffold constituent. In
1986, Weinberg and Bell[34] performed their groundbreaking study using a scaffold based
on collagen for blood vessel reconstruction. They used an in vitro model generated using
multiple layers of collagen integrated with a Dacron mesh to provide the necessary tensile
strength. Smooth muscle cells were cultured in the graft and endothelial cells were used to
line the inner lumen. Various polymeric materials have also been used for mechanical
support in addition to collagen such as PGA and PLGA. Moreover, cells have been grown
under dynamic conditions in particular bioreactors, to closely mimic in vivo
conditions[136]. Research has also been done using decellularized vein as a potential
scaffold for vascular tissue engineering. This process preserves the extracellular matrix,
the basement membrane structure and sufficient strength for vascular grafting[137]. It is
clear that collagen plays a crucial role in all these studies, and that it could represent a
suitable candidate to scaffold tissue regeneration.
79
Finally, collagen is one of the main proteins forming the vascular extracellular matrix.
It is primarily produced by the smooth muscle cells of the media, and the fibroblasts of the
adventitia. Functionally, collagen fibers impose constraints on the elongation of large
vessels under pressure, limiting the distension of the vessel. It also provides attachment to
smooth muscle cells, transmitting force around the circumference of the vessel[12].
Collagen has low antigenicity, low inflammatory and cytotoxic responses[113] and is also
biodegradable[79]. However, there exists a lack of knowledge, about the biological
performance of collagen-based scaffolds when in contact with blood and cells. Therefore,
the aim of this work was to study neutralized collagens impact when in contact with blood
and cells.
80
Materials and Method
Sample Preparation
Specimens consist of flat sheets of collagen type I, which was extracted from rat tail
collagen. Briefly, rat tail tendons were dissolved in acetic acid for 48 hours, then the gellike mass was mixed in a blender, frozen and lyophilized. The lyophilized sponge was
mixed again with 0.02 N acetic acid at a dry weight to solution ratio of 4 mg/ml. The
resulting solution was centrifuged, degassed in vacuum and sterilized. Sterilization was
performed in dialysis bags (Spectra/Por 1, MWCO:6-8,000, Spectrum Laboratories Inc.,
California, USA) and soaked in acetic acid 0.02 N for one hour followed by one hour in
chloroform 1% in water. Dialysis was continued in sterile acetic acid 0.02 N for 4 to 5
days. The resulting sterile solution was neutralized to a pH of 6.8 with NaOH 1%
immediately prior to use. Samples were prepared in the form of thin films by solvent
evaporation from neutral collagen solutions poured onto glass coverslips.
Blood Performances
Reference materials
Teflon which is a commonly employed material for vascular prostheses[138], was used as a
reference material for clotting time measurement and thrombelastography. Glass, although
not used for surgical purposes but due to its negative reaction when in contact with
blood[92], was considered as a negative control.
Blood
Native whole blood was collected from 11 healthy male donors having taken no medication
for at least 10 days prior to donation. 40 ml of blood, collected in tubes containing sodium
citrate, was used for thrombelastograph assays and platelet extraction. 10 ml of blood
without anticoagulants was used for clotting time assessment.
81
Clotting Time
The haemoglobin free method for clotting time measurement is described elsewhere[139].
Briefly, 0.1 ml of untreated blood was immediately dropped onto the specimens. After 10,
20, 30, 40 and 50 minutes, 5 ml of distilled water was added to each specimen and
incubated for 5 minutes. Red blood cells not entrapped in a thrombus were haemolyzed and
free haemoglobin molecules in the water were colorimetrically measured by monitoring the
absorbance at 570 nm using a spectrophotometer ELISA reader (BioRad mod.450,
Mississauga, Ontario, Canada). Absorbance values were converted into a percentage of the
maximum amount of haemoglobin free present in the water. Statistical significance was
determined by the Student’s t-test.
Thrombelastograph (TEG)
400 µl of blood was exposed to the surface of collagen and controls for 30 minutes at 37°C.
Then, 300 µl of this blood was placed in TEG stainless steel cups and 60 μl of 1.29%
isotonic CaCl2 solution were added for recalcification. TEG traces were obtained with a
Thrombelastograph D (Hellige GMBH, Germany) and analyses were conducted following
manufacturer’s instructions. Briefly, the cylindrical cup containing blood, oscillates
through an angle of 4°45’ and is heated to a temperature of 37°C. Each rotation cycle of the
cup lasts 10 seconds. A stainless-steel cylindrical piston is suspended from a torsion wire
and immersed in the cup. The torque of the cup is transmitted to the piston through the
fibrin fibers which gradually form between the piston and the wall of the cup; the rotation
of the piston becomes increasingly stronger, as the clot becomes more solid.
Thrombelastography measures the elastic properties of blood clots as they form. The
strength of the clot is graphically represented over time as a characteristic cigar shaped
figure. There are three main parameters of the TEG trace which are used in this study: r, K
and MA. The time for the initial fibrin formation is given by r. The time from the
beginning of clot formation until 20 mm of amplitude (k) is characteristic of the dynamics
of clot formation. The maximum amplitude (MA) represents the strength of the clot which
is dependent on the number and function of platelets and their interaction with fibrin[140].
82
Combining k and MA leads to the thrombogenicity index. Statistical significance was
determined by the Student’s t-test.
Platelet Adhesion
In vitro testing was performed to investigate the morphology, aggregation and
pseudopodium of the adherent platelets. After centrifuging, platelet-rich plasma was
obtained and a suspension of 3x105 platelets was placed on the samples and incubated for
30 minutes at 37°C. Samples were then prepared for SEM analysis as described later.
Images are representative of all results obtained.
Biological performances
Cell Culture
Endothelial cells and smooth muscle cells were isolated from porcine aorta as described
elsewhere[141,142]. Briefly, after removing the external adventitia, the aorta was incubated
for 25 minutes at 37°C with collagenase type IA 130 U/ml (Gibco, Invitrogen Corporation,
Burlington, Ontario, Canada). Endothelial cells were isolated and seeded on Petri dishes
coated with 0.2% porcine gelatine. DMEM growth medium was supplemented with 10%
porcine serum, glutamine (2 mM), penicillin (100 U/ml), streptomycin (100 μg/ml),
fungizone (0.25 μg/ml), and heparin (0.09 g/L). (all from Sigma Aldrich, Milwaukee,
USA). As already reported in literature, a fluorescent die labelling acetylated-low density
lipoprotein (Dil-Ac-LDL, Biomedical Technologies Inc., Stoughton, Massachusetts, USA)
was used to confirm the sole presence of endothelial cells[143].
Following the first digestion, the aorta was incubated overnight with collagenase type II
270 U/ml (Worthington Biochemical Corporation, Lakewood, New Jersey, USA). Smooth
muscle cells were subsequently extracted and seeded in tissue-culture flasks coated with
porcine gelatine 0.2%. DMEM growth medium was supplemented with 15% porcine
serum, 10% foetal bovine serum, glutamine (2 mM), penicillin (100 U/ml), streptomycin
(100 μg/ml), and fungizone (0.25 μg/ml) (all from Sigma Aldrich, Milwaukee, USA).
83
Positive staining using an anti α-smooth muscle actin-specific antibody was used to
confirm the smooth muscle phenotype of the extracted cells (Boehringer Mannheim,
Mannheim, Germany)[144].
Cells were maintained at 37ºC in 5% CO2-humified atmosphere and harvested by
trypsinization when a monolayer was reached. All experiments were performed using cells
between the 4th and 8th passage and 2x105 cells were seeded on each sample. Cells used as
controls were seeded on Petri dishes or cover slides previously treated with porcine gelatine
0.2% (Sigma Aldrich, Milwaukee, USA).
Scanning Electron Microscopy (SEM)
Cell morphology after adhesion and growth on collagen was investigated by SEM. Medium
was removed, samples were washed twice in 0.15 M cacodylate buffer and fixed for 30
minutes at 4ºC with Karnowsky solution (2% paraformaldehyde and 2.5% glutaraldehyde
in 0.15 M cacodylate buffer, pH 7.2-7.4). Following fixation, samples were treated for 30
minutes with 1% osmium tetroxide in 0.15 M cacodylate buffer solution. Samples were
then dehydrated with graded ethanol (from 50% to 100%), soaked for 30 minutes in
hexamethyldisilazane, dried and sputter-coated with gold-palladium. Images were collected
using a scanning electron microscope (JSM-35CF, JEOL, Tokyo, Japan).
Cytotoxicity Assay
After three growth periods (24h, 48h and 72h) cell viability was assessed using the MTT
colorimetric assay (Sigma Chemical Co., St. Louis, MO). Briefly, {3-(4,5 dimethylthiazol2-yl)-2,5-diphenyl-2H-tetrazolium bromide}(MTT) is reduced to purple formazan by
mitochondrial dehydrogenase in cells indicating normal metabolism. After each time point,
MTT (1 mg/ml in PBS) was added to the medium (1:10) and samples were further
incubated for 4 hours. Medium was removed and purple formazan crystals were solubilized
with 200 μl acidic isopropanol (0.04 M HCl in isopropanol) at room temperature. The
optical density at a wavelength of 570 nm was determined using an ELISA reader (BioRad
mod.450, Mississauga, Ontario, Canada). All tests were performed in triplicate. Results are
84
expressed as mean ± standard deviation. Statistical significance was determined by the
Student’s t-test.
85
Results
Clotting time
During clot formation, the quantity of haemoglobin entrapped by fibrin increases with
time as the clot forms. Therefore, the amount of free haemoglobin available for haemolysis
decreases. This can be used as a significant indicator for evaluating the potential of a
substrate to not induce immediate clotting. In fact, the longer the clotting time is, better the
compatibility between the substrate and blood is. The blood clotting profiles for collagen,
Teflon and glass are presented in Figure 3.6. As collagen consistently failed to initiate
fibrin clotting after 10 minutes, the maximum absorbance value was taken as 100% for
comparison purposes. Clot formation was found to be complete at approximately 25%.
Results show significant differences between collagen and Teflon after 20 minutes of
contact. In fact, haemoglobin free in water remains higher for collagen until 50 minutes
after contact, at which time the clot is completely formed. Teflon forms a stable clot 20
minutes after contact. Glass, as expected, forms a stable clot within the first 10 minutes
after contact.
86
Figure 3.6 : Haemoglobin free test on collagen, Teflon and glass. * indicates that results of
collagen with respect to both reference materials are statistically significant.
Fibrin formation
The parameter “r” obtained from the thrombelastograph profile indicates the time, in
minutes, from the beginning of the test to initial fibrin formation. Fibrin formation times
are shown in Figure 3.7. In this respect, blood in contact with collagen behaves identically
to that of blood having had no contact with control surfaces. Teflon and glass initiate
coagulation significantly faster than collagen.
87
Figure 3.7 : Fibrin formation. * indicates statistically significant results with respect to
blood. °indicates statistically significant results with respect to collagen.
Thrombogenicity Index
The thrombogenicity index informs about the dynamics of blood coagulation as well as the
strength of the final clot. This index is significantly lower for blood after contact with
collagen, thus showing that platelets and fibrin weakly interact in blood after contact with
collagen. The index for Teflon displays a similar trend but it is highly affected by interdonor variability, and therefore the variance is important. Glass shows strong and stable
clot formation, which indicates activation of the coagulation cascade.
88
Figure 3.8 : Thrombogenicity Index. * indicates significant results with respect to blood.
Platelet adhesion
Although SEM does not allow quantification of adhered platelets, it is possible to visually
compare surfaces after platelet adhesion and gain insight into platelet activation and
aggregation. A SEM image of a representative zone of the highest platelet density for
collagen is shown in Figure 3.9a. It is important to note that platelets adhere in a nonhomogeneous way on collagen substrates, and several parts of the specimens were
completely free of platelets (data not shown). Figures 3.9b and c show platelet
morphologies after adhesion to collagen. In Figure 3.9b cells express a rounded
morphology with no pseudopodium. Small aggregates were also observed on the surface,
and they are shown in Figure 3.9c. It is interesting to note that large aggregates as well as
long pseudopodia are entirely absent, showing low activation.
89
Figure 3.9 : SEM analyses of platelets on collagen film.
Cell Morphology
As observed by SEM, endothelial cells (Fig. 3.10a) and smooth muscle cells (Fig. 3.10b)
were well attached and spread after 24 hours. Both cell types exhibited solid and numerous
focal contacts to the collagen substrate. The cobblestone morphology, typical of endothelial
cells, as well as the spindle morphology of smooth muscle cells, are well defined. After 72
hours of proliferation on collagen, a complete monolayer was reached for both cell types
(data not shown).
Figure 3.10 : SEM analyses of endothelial cell (a) and smooth muscle cell (b) morphology
after 24 hours of contact with collagen films.
90
Cytotoxicity assay
MTT results obtained with endothelial cells and smooth muscle cells seeded on collagen
film are presented in figure 3.11a and b respectively. Endothelial cells are proliferating, as
demonstrated by the increased absorbance with time. However, cell growth is significantly
slower on collagen than on controls. Smooth muscle cells, on the other hand exhibit
significantly lower proliferation than on the control after 72 hours, once a complete
monolayer has been formed.
Figure 3.11 : MTT results on endothelial cells (a) and smooth muscle cells (b). * indicates
significant results with respect to control.
91
Discussion
It is well known that, in native blood vessels, the endothelium possesses natural
anticoagulant and antithrombotic properties[98]. One of the major downfalls of prosthetic
vascular grafts is the inability to obtain a confluent endothelium on the luminal surface. To
improve antithrombotic performances of vascular replacements, tissue-engineering attempts
to mimic as close as possible the behavior of a native vessel. A successful tissue-engineered
arterial replacement, apart from adequate mechanical properties, requires two fundamental
biological qualities: thromboresistance and adequate cell presence and functionality. In this
study we investigated blood-material and cell-material interactions in order to evaluate the
potential of acid-soluble type I collagen from rat tail tendon to be considered as a natural
scaffold for vascular regeneration. Teflon, due to its current satisfactory use for medium
diameter vascular prostheses, and due to its demonstrated low capacity for platelet
activation compared to other polymeric prosthetic materials, was considered as a positive
control[145].
Thrombelastography is the only test that evaluates the entire dynamics of clot formation.
By measuring the viscoelastic properties of blood that characterize the strength of the clot
over time, it gives information about several steps in the coagulation process. An important
indicator of coagulant activity is the parameter r, which indicates the time required for
fibrin formation. Blood having had no contact with control materials has a variability
between 5 and 8 minutes for the initiation of fibrin formation. This variability does not
allow to conclude that significant differences of fibrin formation time exist when
compared with collagen and Teflon. However, it is notable that collagen has an r value
significantly higher than Teflon, with a small variability indicating constant positive
performances. Glass, in terms of fibrin formation, enhances clot formation in a shorter
time when compared to all other materials. The thrombogenicity index takes into account
the dynamics of clot formation as well as the strength of the interactions between platelets
and fibrin. A low index indicates a low thrombogenic potential. This is the case for blood
after contact with collagen which demonstrates a weak interaction between platelets and
92
fibrin, meaning that platelets are not activated to enhance the coagulation cascade. As with
fibrin formation time, the low standard deviation indicates that all assays with collagen
have the same trend. Clotting time, measured by the haemoglobin free test, confirms that
collagen is less efficient than Teflon in enhancing clot formation. Moreover, morphological
analyses show a low percentage of platelet adhesion and the absence of aggregates, proving
that good performances obtained concerning blood reactions are due to the modest
activation of platelets on collagen.
Collagen in the body is a natural coagulant and plays an important role in haemostasis and
thrombogenesis[146]. However, in this study, neutral collagen films have demonstrated
good performances in not enhancing clot formation. This discrepancy might be due to the
structure of the collagen substrate and the absence of von Willebrand factor. In fact, it has
been shown that different collagen preparations vary in their ability to aggregate platelets
and support their adhesion. Fibrillar collagen is much more potent in inducing platelet
activation than the same collagen type in acid soluble form[147]. It was concluded that the
quaternary structure of collagen is important in inducing platelet activation[146]. Although,
structural aspects of fibrils mediate platelet adhesion and aggregation, biochemical factors
may also play a dominating role. Recently, von Willebrand factor, which constitutes a
major difference between acid-soluble and fibrillar collagen, acts synergistically with
collagen fibrils in the activation of the coagulation cascade[148]. This protein, not present
in acid-soluble collagen, has been purposed as the main mediator for the initial capture of
platelets to the site of vascular injury[149]. An enzymatic test (Asserachrom vWF;
Boehringer-Mannheim, Mannheim, Germany) was performed (data not shown) to insure
that our acid-soluble collagen was absolutely free of von Willebrand factor. In fact, it is
widely known that the first step in platelet-collagen interactions is the binding of platelet
glycoprotein Ib (GPIb) to von Willebrand factor. Only subsequently will platelets adhere to
collagen with specific cell surface receptors[150]. It has also been reported that an
enhanced level of von Willebrand factor increases platelet adhesion and aggregation on
collagen surfaces[151]. This may explain why collagen, which naturally has exceptional
haemostatic properties, in this particular context, does not enhance the activation of the
coagulation cascade.
93
The use of collagen as a substrate for endothelial and smooth muscle cells is not an
innovative strategy[67,55]. However, we considered this test in the aim to prove that acidsoluble collagen, extracted from rat tails followed by lyophilization, constitutes an
acceptable substrate for cell growth. We previously performed preliminary tests
demonstrating the adequate biological performance of the substrate for fibroblasts (data not
shown). This study continues in this direction using two primary cell lines, which will
potentially be used for vascular regeneration. Morphological analyses show good cell
spreading and, even if cell growth is slightly slower than on the control, cells are
proliferating on the surface. A possible explanation for this difference in the dynamics of
cell growth could be due to the different stimuli induced by collagen on cellular
metabolism, which may increase extracellular matrix production at the expense of
proliferation. Moreover, the good performances shown by this collagen solution as a
substrate for various cell cultures allows us to speculate as to it’s suitability for different
tissue engineering applications, involving for example myocytes or hepatocytes and
eventually growth in a three-dimensional environment.
The results obtained, indicate that collagen films present low thrombogenicity and that they
support cell spreading. Therefore, this could be promising and this type of collagen can
represent an adequate substrate to successfully scaffold vascular tissue during regeneration
or engineering.
The above chapter presented results validating collagen as a suitable
substrate for vascular cells and blood contact applications. The following
chapter concludes the entire thesis and presents the future challenges and
various suggested approaches for meeting a few of the many challenges.
Conclusion
The overall objective of this project was to design and develop a completely biological
scaffold structure from collagen for a vascular graft. The resulting scaffold is seeded with
cells in order to more efficiently generate many characteristics of natural arteries. The
challenges were numerous. The above-mentioned generated tissue must meet several
requirements. Most notably, the mechanical strength must be sufficient to withstand
suturing and resist bursting when subjected to blood flow. It must also possess a confluent
endothelium to resist thrombosis. To mimic the functionality of a native vessel, it must also
be vasoactive. Some requirements more specific to the scaffold include being
biocompatible, having a controlled biodegradability, and being able to support viable cell
growth and proper cell differentiation. Many factors influence collagen and the cells
eventually seeded inside which may improve various properties to meet these requirements.
Most notably, dynamic conditioning via a perfusion bioreactor is one of the most influential
methods to improve mechanical strength and proper cell differentiation.
Before these factors may be studied, many preliminary operations were performed. Firstly,
a feasibility study was carried out in order to assure that collagen may provide an adequate
and feasible scaffold structure for vascular cells in the tissue engineering of a vascular
graft. Secondly, a source of collagen type I as well as processing methods for obtaining
acid-soluble collagen solution were obtained. Two processing methods were validated and
compared with respect to collagen molecular conformation, mechanical properties and
preliminary cell viability with fibroblast cells. Both processing methods were found to
provide collagen solutions and films with almost identical properties.
One of the above methods was chosen for further studies due to its appropriateness for
long-term storage and greater controllability of collagen concentration in solution. Then,
neutral collagen films were tested for cell viability with smooth muscle cells and
endothelial cells. These films were also tested by some blood contact assays. Neutral
collagens proved to provide an adequate substrate for these cell types. These films did not
induce coagulation and did not induce platelet adherence and activation.
95
Following these preliminary assays, collagen gels were made and tested. The cell-seeded
disc-shaped gels were tested initially for gel contraction with smooth muscle cells and
fibroblasts. Also, platelet adhesion and aggregation was studied on discs of smooth muscle
cells with and without an endothelial layer. Then, cylindrical gels were made and tested for
cell viability. An objective of the approach taken was to continually rotate horizontally the
cylindrical gels for the initial maturation period during which time the cell-mediated cell
contraction occurs around the mandrel. This was a major challenge as a means needed to be
found which allows this rotation but does not deprive the cells of nutrients and especially
gas exchange. Finally, an appropriate cylindrical mould-mandrel system was elaborated
which demonstrated good cell viability for up to 2 weeks. Cell-cycle analysis demonstrated
positive results. These results were supported by histology which showed good cell
distribution and spreading. After 1 week of maturation, these tubular gels were able to be
manipulated and placed in a compliance apparatus which measured the dilation of the tube
due to an increase in luminal volume.
Collagen-based vascular tissue engineering is a promising avenue for answering the need
for a suitable arterial replacement. However, there are many limitations that may or may
not be overcome as our understanding increases. Starting form the beginning, adequate
sources of cells and collagen must be found. Autologous biopsies can be used to obtain a
small amount of cells which can be multiplied in vitro. This process is very time consuming
and will not be appropriate for most applications. However, any non-autologous cell-source
implies the need to engineer the immune acceptance of the cells. There is also the question
of using stem cells or not. The source of collagen is somewhat more limited. Autologous
collagen cannot be obtained in sufficient quantities. The most frequently used collagen for
biomedical applications is bovine dermal collagen. From skin, Atelocollagen, which is
produced by removing the telopeptides from both ends of the collagen molecule by pepsin,
can be obtained and possess the lowest antigenicity of all the collagens[78]. However,
removal of the telopeptides makes it more difficult to obtain a mechanically solid
biomaterial. This is an important factor as one of the most important limiting factors with
this approach is a lack of the mechanical strength necessary for implantation as an arterial
substitute. Apart from a lack of mechanical strength, another issue is the biological reaction
96
to the graft upon implantation. The biochemical features of a blood vessel graft must be
controlled so that it may be positively identified as an active organ and integrated
structurally and functionally with in the vascular system[10]. One of the most important
issues is the time required to manufacture these grafts. This takes weeks to months to make
and mature. For autologous grafts, this excludes emergency surgeries. For this approach to
be successful, potential pathologies must be detected well in advance or a system must be
in place to produce and stock autologous grafts for undetermined future availability. A
more feasible approach would be to use non-autologous grafts which may be produced in
greater quantities and stored. This requires intensive reliability and up-scaling, to meet the
tremendous demand, in the manufacturing process. This also requires a means to reliably
freeze and store the grafts until they are needed.
Clinical implantation as a potential bypass material of these grafts is still more than a
decade away. In the mean time, there are multiple avenues which may be taken to attempt
to improve the mechanical and biological properties as well as to reduce the processing
time of these grafts. With each attempt comes an increased understanding which further
facilitates the next approach. In this way, the advances in tissue engineering are a valuable
biological tool with which improved comprehension can be obtained about the various
factors that influence cells and their surrounding tissues. The repercussions of these works
can be felt not only in the field of tissue engineering but in the biological field as whole.
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First Edition Paper 2004; In Press:
Annexe A: Protocols for rat tail collagen extraction and
solution fabrication
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1
Qcoll
Wash rat tails in distilled water
Extraction of tendons from rat tail
Place tendons in PBS during extraction
Rinse in Acetone for 5 min.
Rinse in isopropanol 70% for 5 min.
Solubilize tendons in 0.02N acetic
acid.(48 hrs) at 4ºC
Blend gel and freeze at -20ºC
Lyophilize1
Blend lyophilized collagen 4mg/ml
with 0.02N acetic acid
Centrifuge at 15000 rpm for 45 min at
4ºC. Conserve supernatant.
Degas 10-12 min/100ml
Dialysis in acetic acid for 1 day2
Sterilize collagen by dialysis in
Chloroform 1% for 1 hour
Dialysis in sterile acetic acid for 3 days
changing solution every day
Remove from dialysis bags and
conserve at 4ºC3
•
•
•
Tcoll
Wash rat tails in distilled water
Extraction of tendons from rat tail
Place tendons in PBS during extraction
•
Solubilize tendons in 0.4M acetic
acid.(48 hrs) at 4ºC
•
Centrifuge at 15000 rpm for 45 min at
4ºC. Conserve supernatant.
Degas 10-12 min/100ml
Dialysis in acetic acid for 1 day2
Sterilize collagen by dialysis in
Chloroform 1% for 1 hour
Dialysis in sterile acetic acid for 3 days
changing solution every day
Remove from dialysis bags and
conserve at 4ºC3
•
•
•
•
•
The collagen in this state can be stored at -70°C.
protocol for rinsing the dialysis bags:
1- Cut an length of bag (about 12 inches)
2- Roll the bag and rince in distilled water. Change water and repeat.
3- Tie a knot in one end of the bag.
4- Rinse the inside of bag with distilled water three times.
5- Pour collagen into the bag.
6- Close upper end of bag with the proper clamp.
7- Immerse bag in acetic acid 0.02N.
3
The collagen in this state can be kept for up to 1 year at 4°
2
108
Annexe B: Protocol for Smooth muscle cell and
endothelial cell extraction from pig aorta
(Source: Chiara Arrigoni)
1.
Prepare the following solutions:
MEM 2X Ab
2% fungizone
2% pen-strep
fill with DMEM
D-PBS 2X Ab
2% fungizone
2% pen-strep
fill with D-PBS
D-PBS 5X Ab
5% fungizone
5% pen-strep
fill with D-PBS
MEDIUM pSMC
MEDIUM PAEC
DMEM (+ 4 mM glutammina, 1
DMEM (+ 4 mM
month)
glutammina, 1 month)
10% FBS
15% PS
10% PS
1% pen-strep
1% pen-strep
1% fungizone
1% fungizone
5mM HEPES acido
5mM HEPES acido
2.
Remove the aorta from the animal. Make sure to leave the vessel as complete as
possible to avoid contamination.
3.
Soak the vessel in DMEM 2X at 4ºC on ice during transport.
4.
Prepare:
5.
Place the vessel in large Petri dish with DPBS 2X Ab.
6.
Remove all fat and collateral vessels with sterile tweezers and scissors. Maintain
vessel hydrated with DPBS. Keep cleaning until the vessel is smooth and white. Avoid
touching the inside of the vessel.
7.
Place in DPBS 5X Ab and put in fridge for 30 min.
8.
Remove the adventitia by separating it from the media from one end and pulling
with tweezers. It should be able to be pulled off ‘like a sock’.
9.
If the adventitia is adequately removed, the vessel should have lost some
consistency, but remains integral and is white.
10.
Place vessel in DPBS 5X Ab and leave for 15 min at 4ºC.
11.
Prepare the collagenase solution 1 A. Mix 130 U/ml in DPBS 2X Ab at 37ºC.
12.
Cut the vessel into pieces 2 cm long.
13.
Prepare an adequate number of p55 Petri dishes with 2.5ml of collagenase solution.
The bottom must be barely covered in order to make contact solely with the endothelium.
14.
Place the endothelial surface of the vessel pieces down on the surface of the Petri’s.
Take care to remove any air bubbles between the two surfaces. ATT: here it is suggested
to start the pieces at 5 min intervals to leave sufficient time to perform subsequent
operations.
15.
Place at 37ºC for 25 min lifting and moving the piece once to favor the action of the
collagenase.
16.
Prepare a sufficient number of 50 ml syringes with 18G needles (one for each piece)
with 40 ml DMEM 2X Ab maintained at 4ºC.
109
17.
Prepare a sufficient number of Corning tubes with 3 ml of FBS each.
18.
Remove the pieces, one by one, from their dishes and hold them over sterile tubes
with tweezers. With a syringe of 40 ml DMEM 2X Ab, spray the endothelial surface of
the vessel, rinsing the endothelial cells into the test tube. Use as much pressure as
possible.
19.
Put the vessel pieces into a p55 Petri dish with DMEM 2X Ab.
20.
Centrifuge the endothelial solution for 10 min at 1000rpm at 22ºC (acceleration 9,
braking 5).
21.
In the meantime, chop the remaining vessel (which should contain solely the media)
into small homogeneous pieces.
22.
Put the media pieces into 2 T50 flasks, adding to each 25 ml of collagenase type II
at a concentration of 0.05% in DMEM, and incubate overnight.
23.
Remove the corning tubes from the centrifuge and suspend the cells in 1.5 ml
complete medium in plates with wells of 9.6cm2.
The Next morning: Remove the soluble part from the flasks and centrifuge
110
Annexe C: Protocol for Collagen Gel Fabrication
(Source: Marie-France Côté)
The collagen used for these gels has been extracted from rat tail tendon by the method
described elsewhere. It is kept in acid soluble liquid form at 4ºC for up to 1 year prior to
usage. The collagen concentration is roughly 3.5 mg/ml and the Ph is roughly 7. The
protocol makes 5 ml.
Solution:
-1 ml DMEM 5X without NaHCO3 [20%]
-0.5 ml FBS (Fetal Bovine Serum) [10%] 1
-0.5 ml NaHCO3 0.26M 2 [10%]
-0.02 ml NaOH 1M [0.4%]
-0.1 ml H2O [2%]
Preparation of the gel:
a- Put the solution (2.120 ml) in a centrifuge tube.
b- Add 2 ml rat tail collagen at 3.5mg/ml [40%]
c- Mix gently
d- Add 0.88 ml of cell suspension [17.6%]
e- Mix well
f- Let gel for 1 hour 3 at 37ºC.
gAfter the gel has solidified, place a small amount of medium on top of the gel to avoid
excessive drying.
Note: Be careful not to make too much solution at a time. The gel solidifies rapidly.
For varying volumes of needed gel, the same volume proportions (in [ ]) were used in
experiments involving cells.
1
FBS only is used when fibroblasts are the suspended cells. For Smooth Muscle Cells (SMC) 5% Fetal
Bovine Serum (FBS) is used along with 5% Porcine serum (PS) for a total of 10% serum.
2
NaHCO3 is used here as a buffer
3
Variation from original protocol: Marie-France only incubates for 10 minutes. For the geometry of my
samples, this is insufficient.