development of bioactive surfaces to control cell behavior
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Transcript of development of bioactive surfaces to control cell behavior
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Faculté de génie
Département de génie chimique
DÉVELOPPEMENT DE SURFACES BIOACTIVES POUR CONTRÔLER LE COMPORTEMENT CELLULAIRE
UTILISATION DE PUCES DE POLYMÈRES ET DE BIOMOLÉCULES
DEVELOPMENT OF BIOACTIVE SURFACES TO CONTROL CELL BEHAVIOR
USE OF POLYMER AND BIOMOLECULE ARRAYS
Thèse de doctorat es sciences appliquées Spécialité génie chimique
Emmanuelle MONCHAUX
Sherbrooke (Québec), Canada Décembre 2007
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À mes parents, Yves et Marie
À mon frère et ma sœur, Géraud et Stéphanie
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Résumé
La compréhension et le contrôle des interactions entre les cellules et les surfaces des
matériaux sont essentiels pour la conception d’implants biocompatibles et fonctionnels. Ainsi,
des surfaces bioactives prévenant toute réaction biologique non-spécifique et fournissant des
signaux dirigeant le comportement cellulaire sont développées.
Les interactions entre les surfaces et les cellules endothéliales sont particulièrement
importantes pour des applications médicales telles que l’endothélisation des prothèses
vasculaires synthétiques et le développement de tissus artificiels vascularisés par génie
tissulaire. L’objectif de ce travail de recherche est donc de développer une surface bioactive
interagissant spécifiquement avec les cellules endothéliales par le développement et
l’utilisation de puces de polymères et de surfaces bioactives.
Des puces de carboxy-méthyl-dextran (CMD) sont développées afin de déterminer une
surface anti-adhérente résistant aux interactions non-spécifiques. Les surfaces de CMD
immobilisées selon différentes conditions sont caractérisées par spectroscopie des
photoélectrons-X (XPS), microscopie à force atomique (AFM), exposées à un mélange de
protéines puis à des fibroblastes afin d’identifier les conditions d’immobilisation influençant
leur structure et leur capacité à résister à l’adsorption non-spécifique de protéines et à
l’adhésion cellulaire. Une couche de CMD optimisée se révèle aussi résistante à l’adhésion
cellulaire que le polyéthylène glycol, polymère le plus utilisé pour prévenir toute interaction
non spécifique.
Des puces de surfaces bioactives ciblant spécifiquement les cellules endothéliales et
comprenant les séquences peptidiques REDV, SVVYGLR et/ou le facteur de croissance
VEGF sont fabriquées et exposées à des cellules endothéliales et des fibroblastes. Les
molécules immobilisées n’induisent pas d’adhésion sélective mais induisent la réorganisation
du cytosquelette et des adhésions focales spécifiquement chez les cellules endothéliales.
L’utilisation des puces de polymères a permis le développement d’une surface anti-
adhérente efficace pour la fabrication de surfaces bioactives et les puces de molécules
bioactives rendent possible l’étude de la réponse cellulaire face à des surfaces de composition
variée.
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Summary
Understanding and controlling interactions between cells and material surfaces are
essential to design biocompatible and functional medical implants. To this end, bioactive
surfaces preventing non-specific biological reactions and providing signals that guide cell
behavior are designed.
Interactions between surfaces and endothelial cells are particularly important for
biomedical applications such as endothelialization of vascular grafts and the development of
vascularized artificial tissues by tissue engineering. The aim of this research project is then to
develop a bioactive surface specifically interacting with endothelial cells by developing and
using arrays of polymers and bioactive surfaces.
Arrays of carboxy-methyl-dextran (CMD) are made to determine a low-fouling surface
preventing non-specific protein adsorption and cell adhesion. CMD layers are grafted with
various conditions and are characterized by X-ray photoelectron spectroscopy (XPS), atomic
force microscopy (AFM), exposed to proteins and then fibroblasts, to identify immobilization
parameters that influence layers structure and ability to resist non-specific interactions. An
optimized surface of CMD is obtained and is as resistant to cell adhesion as a layer of
poly(ethylene glycol), the most used low-fouling polymer.
Arrays of bioactive surfaces specifically directed toward endothelial cells and made
with peptide sequences REDV, SVVYGLR and/or vascular endothelial growth factor (VEGF)
are synthesized and exposed to endothelial cells and fibroblasts. Immobilized biomolecules do
not promote a selective endothelial adhesion but induce cytoskeleton and focal adhesion
reorganization specifically for endothelial cells.
Use of polymer arrays allowed the development of a low-fouling surface efficient for
the making of bioactive surfaces and biomolecule arrays allowed to study cell responses
toward surfaces of various molecular composition.
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Remerciements
Mes remerciements s'adressent tout d'abord à mon directeur de recherche, Patrick
Vermette, qui m'a permis de venir finir mes études à Sherbrooke. Merci pour m'avoir confié
ce sujet de recherche et soutenue financièrement.
Je remercie les autres membres de l'équipe, étudiants gradués, post-docs et assistants
de recherche, pour leurs conseils constructifs et la bonne ambiance de travail au laboratoire.
Je tiens à remercier tous les membres de mon jury de thèse qui me font l'honneur
d'évaluer mon travail.
J'adresse également mes remerciements au personnel technique du département de
génie chimique et du centre de caractérisation des matériaux de l'IMSI ainsi qu'au personnel
administratif du département de génie chimique pour leur bonne humeur et leurs coups de
main toujours bienvenus.
Je remercie tous les étudiants gradués et post-docs de génie chimique qui m'ont permis
de passer ces années à Sherbrooke dans une ambiance multinationale enrichissante et des plus
agréables.
Enfin, j'adresse mes remerciements les plus sincères à ma mère, mon frère et ma sœur
pour leur soutien durant ces années d'études, surtout dans les moments de doute. Merci…
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Table des matières
1. Mise en contexte
2. Effets des propriétés de surface des biomatériaux et de la bioactivation sur les cellules endothéliales
2.1. Résumé
2.2. Abstract
2.3. Introduction
2.4. The endothelial tissue and its matrix
2.4.1. Morphological and functional heterogeneity
2.4.2. Endothelial cells – matrix interactions
2.4.3 Formation of blood vessels
2.5. Interactions between endothelial cells and biomaterial surfaces
2.5.1. Surface properties and the mechanical environment
2.5.2. Pre-coatings made of matrix proteins
2.5.3. Grafting of peptides
2.5.4. Growth factors immobilization
2.6. Conclusions and perspectives
2.7. References
3. Développement de puces d’un dérivé du dextran pour identifier les propriétés physico-chimiques impliquées dans l’adsorption de protéines du sérum
3.1. Résumé
3.2. Abstract
3.3. Introduction
3.4. Experimental section
3.4.1. Materials
3.4.2 Fabrication of CMD arrays
3.4.3 Elemental composition of CMD spots by X-ray photoelectron spectroscopy (XPS)
3.4.4 CMD spot layer structure by AFM force measurements
3.4.5 CMD spot homogeneity and CMD fouling from serum measured by surface plasmon resonance (SPR) microscopy
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3.5. Results and discussion
3.5.1 Elemental composition of CMD spots by XPS
3.5.2 AFM colloidal probe interaction forces with HApp layers
3.5.3 AFM colloidal probe interaction forces with CMD spots grafted using high EDC+NHS/COOH ratios
3.5.4 AFM colloidal probe interaction forces with CMD spots grafted using low EDC+NHS/COOH ratios1
3.5.5 CMD spot homogeneity and CMD fouling from serum measured by SPR microscopy
3.6. Conclusions
3.7. Acknowledgements
3.8. References
4. Étude des mécanismes de résistance à l’adhésion cellulaire par utilisation de puces d’un dérivé du dextran
4.1. Résumé
4.2. Abstract
4.3. Introduction
4.4. Materials and methods
4.4.1 Carboxy-methyl-dextran (CMD) synthesis
4.4.2 Fabrication of arrays of CMD graft layers
4.4.3 Testing arrays of CMD graft layers towards cell responses
4.5. Results
4.5.1 Initial cell interaction with CMD graft layers: short-term assays
4.5.2 CMD resistance following 3 days of cell confluence
4.6. Discussion
4.7. Conclusions
4.8. Acknowledgements
4.9. References
5. Puces bioactives immobilisées sur des surfaces anti-adhérentes pour étudier l’adhésion spécifique des cellules endothéliales
5.1. Résumé
5.2. Abstract
5.3. Introduction
5.4. Materials and methods
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5.4.1. Carboxy-methyl-dextran (CMD) layers
5.4.2. Fabrication of arrays of bioactive molecules
5.4.3. Testing bioactive arrays towards cell responses
5.4.4. Cell adhesion assay
5.4.5. Cell labelling
5.5. Results
5.5.1. RGD peptide is required to initiate cell adhesion
5.5.2. REDV, SVVYGLR and VEGF specifically affect endothelial cells adhesion when co-immobilized with RGD
5.5.3. REDV, SVVYGLR and VEGF affect actin filaments organization and focal adhesion assembly in endothelial cells
5.6. Discussion
5.7. Conclusions
5.8. Acknowledgements
5.9. References
Conclusions et perspectives
Annexe A: Synthèse du carboxy-méthyl-dextran (CMD)
Annexe B : Imagerie par microscopie à force atomique (AFM) des puces de CMD hydrates
Annexe C : Marquages immunocytochimiques
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Liste des figures
Figure 3.1: Guiding system to localize individual CMD spots for XPS and AFM analyses. 48
Figure 3.2: a) High-resolution XPS C 1s spectra of a freshly deposited HApp layer and of CMD spots made under conditions no. 1 and no. 3. b) High-resolution XPS C 1s spectra of CMD layers produced under condition no. 1 on a fully covered surface and on an array. 51
Figure 3.3: Representative colloidal probe force measurements between a silica colloidal probe and HApp layers on a polymer array in two PBS solutions (pH 7.4). 53
Figure 3.4: Representative colloidal probe force measurements between a silica colloidal probe and spots of CMD grafted on HApp surface using high EDC+NHS/COOH ratios in two PBS solutions. (a) silica-CMD layers no. 1 and no. 6 and silica-CMD layers no. 3; (b) silica-CMD layers no. 8. 55
Figure 3.5: Representative colloidal probe force measurements between a silica colloidal probe and spots of CMD grafted on HApp surface using low EDC+NHS/COOH ratios in two PBS solutions. (a) silica-CMD layers no. 4 and no. 7; (b) silica-CMD layers no. 2 and no. 5. 57
Figure 3.6: Protein adsorption on spots of CMD graft layers evaluated by SPR microscopy. a) Image of CMD layers produced in conditions no. 2 and no. 5 obtained at plasmon angle θSPR (scale bars: 500µm). b) SPR angle shifts (°) resulting from FBS protein adsorption on CMD arrays. Condition no. 9 involves the use of 70kDa CMD, 25% carboxylation degree, high EDC+NHS/COOH ratio, and a 2mg/ml CMD solution concentration. 59
Figure 3.7: Schematic picture (not to scale) of immobilized CMD molecules. Final layer conformation can be controlled by electrolyte concentration and coupling agent ratio (EDC+NHS/COOH) during immobilization. CMD graft layers structures are hypothesized based on the present XPS and AFM results. 63
Figure 4.1: Optical microscope images of fibroblasts seeded on CMD arrays following 4h (a,b) and 12h (c,d) cell seeding on CMD spots made using condition no. 8 (a,c) and condition no. 3 (b,d). 73
Figure 4.2: Initial cell behavior on CMD arrays. Cells were fixed 4h following cell seeding. Cells on the HApp layer (a,b,c) and on a CMD spot (d,e,f: condition no.1 and g,h,i: condition no. 7). 74
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Figure 4.3: Initial cell behavior on CMD arrays. Cells were fixed 12h following cell seeding. Cells on the HApp layer (a-c), on cell-resistant CMD spots (d-f, condition no. 9), and on CMD layers made using condition no. 4 (g-i). 76
Figure 4.4: CMD spots surface coverage following 3-day exposition to confluent cells. Spot size was measured each day and compared to initial spot size. (*) Some spots were completely covered on Day 3, therefore the average initial size was not 1, resulting in larger standard deviations. 77
Figure 4.5: CMD spots exposed to confluent fibroblasts for a period of 3 days. CMD layers made using condition nos. 5 (a-c), 2 (d-f), and 9 (g-i). 78
Figure 4.6: Fibroblast actin cytoskeleton (a,d,g,j), focal adhesion formation (b,e,h,k) and human fibronectin deposition and reorganization (c,f,i,l) after 3 days of confluence on CMD spots. Cells found on HApp surfaces (a,b), at the edge of CMD spots made using condition nos. 1, 3, and 9 (c-f), on CMD spot made using condition no. 6 (g,h,i,k), and on CMD spot made using condition no. 5 (j,l). 79
Figure 4.7: Determination of an optimal cell-resistant CMD surface. CMD surfaces immobilized using optimized conditions and various carboxylation degrees (referred to as CMD-5, -25 and -50) and PEG surfaces made under cloud point conditions (PEG-CP) were exposed to confluent fibroblasts for a period of 3 days. A: spot size evolution. B: optical microscope images of spots over the 3 days. 81
Figure 5.1. Reaction scheme for the grafting of carboxy-methyl-dextran (CMD) to the HApp-modified surface and subsequent bioactive molecule (peptide or growth factor) immobilization. 98
Figure 5.2. Microscopy images of phalloidin-actin labeling of (A) human umbilical vein endothelial cells (HUVECs) and (B) human foreskin fibroblasts adhering on RGD spots, 6 h after cell seeding, and of (C) confluent HUVECs cultured on a RGD spot for 5 days with 10% serum. Scale bar is 250 µm. 99
Figure 5.3. (A) Human umbilical vein endothelial cells (HUVECs) and (B) human foreskin fibroblasts adhesion on bioactive spots 6 h after cell seeding. Significant difference compared with RGD alone (R25) at *p< 0.05 or ** p< 0.01. 100
Figure 5.4. Spreading levels measured for human umbilical vein endothelial cells (HUVECs) on bioactive spots 6 h after cell seeding. 100
Figure 5.5. Actin filaments labeled with phalloidin (A, C, E, G, I, K, N, O) and focal adhesions labeled with anti-vinculin (B, D, F, H, J, L, N, P) in human foreskin fibroblasts and human umbilical vein endothelial cells (HUVECs) on R25 (A, B, E, F), R25V (C, D, G, H), RE1 (I and J), RE1V (K and L), SV1 (M and N), and SV1V (O and P) 6 h after cell seeding. Scale bars are 25µm. 102
Figure A.1: Principe de la réaction de carboxy-méthylation du dextran. 114
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Figure A.2: Représentation schématique d’un monomère carboxylé de la chaîne de CMD. 115
Figure A.3: Spectre H1-RMN dans D2O du dextran non modifié. †: hydrogène anomérique. 116
Figure A.4: Spectre H1-RMN dans D2O du CMD obtenu avec 1M d’acide bromoacétique. †: hydrogène anomérique, *: hydrogènes du groupe carboxyméthyl. 116
Figure B.1: Image AFM de 1µm² de la surface HApp prise dans le PBS 150mM. 117
Figure B.2: Images AFM de 1µm² prises dans le PBS 150mM des “spots” de CMD de 70kDa immobilisés selon les conditions #5 (a) et #9 (b). 118
Figure B.3: Images AFM de 1µm² prises dans le PBS 150mM des “spots” de CMD de 500kDa immobilisés selon les conditions #2 (a) et #4 (b). 118
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Liste des tableaux
TABLE 3.1: CMD immobilization conditions used to produce spots of CMD graft layers. 47
TABLE 3.2: Elemental composition of CMD arrays and fully covered control CMD surfaces on borosilicate glass determined by XPS analyses. 50
TABLE 4.1: CMD immobilization conditions used to produce arrays of CMD graft layers. 70
TABLE 4.2: Physico-chemical characterization of CMD layers and their resistance to protein adsorption and cell adhesion. 83
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Liste des abréviations
AFM: atomic force microscopy
Ang: angiopoietin
BSA: bovine serum albumin
CMD: carboxymethyl dextran
EC: vascular endothelial cell
ECM: extracellular matrix
FBS: foetal bovine serum
FEP: fluoroethylenepropylene (Teflon)
FGF: fibroblast growth factor
Fn: fibronectin
HApp: n-heptylamine plasma polymer
HFF: human foreskin fibroblast
HGF: hepatocyte growth factor
HS: heparan sulfate
HSPG: heparan sulfate proteoglycan
HUVEC: human umbilical vein endothelial cell
ICAM: intercellular adhesion molecule
Ln: laminin
MMP: matrix metalloproteinase
NMR: nuclear magnetic resonance
NO: nitric oxide
OEG: oligo(ethylene glycol)
PA: plasminogen activator
PBS: phosphate-buffered saline
PDGF: platelet-derived growth factor
PEG, PEO: poly(ethylene glycol), poly(ethylene oxide)
PETP: polyethyleneterephthalate (Dacron)
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PG: proteoglycan
PGI2: prostacyclin
PTFE: polytetrafluoroethylene
QCM: quartz crystal microbalance
SMC: smooth muscle cell
SPR: surface plasmon resonance
TCPS: tissue culture polystyrene
TGF: transforming growth factor
Tsp: thrombospondin
VCAM: vascular cell adhesion molecule
VEGF: vascular endothelial growth factor
VEGFR: VEGF receptor
Vn: vitronectin
XPS: X-ray photoelectron spectroscopy
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Chapitre 1
Mise en contexte
Comprendre, optimiser et contrôler les interactions entre les matériaux et leur
environnement biologique, pouvant contenir des protéines, des lipides et/ou des cellules, est
essentiel pour le développement de biomatériaux intégrés et fonctionnels une fois implantés
dans l’organisme.1,2 Dans le domaine biomédical, les interactions matériaux – biomolécules
sont responsables du succès ou de l’échec de l’implantation de prothèses, destinées à
remplacer ou améliorer une fonction vitale perdue ou défaillante. Ces interactions dictent aussi
le destin de biosenseurs, de membranes de dialyse, de systèmes de libération contrôlée de
principes actifs ou bien encore, ces interactions sont essentielles pour la conception de
structures polymériques permettant le développement contrôlé d’un tissu fonctionnel par génie
tissulaire.1,3
Surfaces bioactives
La plupart des biomatériaux implantés sont rapidement recouverts d’une couche de
protéines, via des interactions de type van der Waals, hydrophobiques, électrostatiques et des
liaisons hydrogènes. Des cellules, dont des monocytes, adhèrent à la couche de protéines pré-
adsorbées; l'activation des monocytes mène à une réaction inflammatoire prolongée et à la
formation d'une capsule fibreuse autour des implants, dont l'intégration et la fonction sont
alors altérées.4,5
Les surfaces des biomatériaux doivent donc être conçues pour prévenir toute
interaction non spécifique menant à une réponse chaotique et non désirée. Elles doivent aussi
pouvoir induire la réaction biologique désirée, interagissant de façon spécifique avec leur
environnement, pour une meilleure intégration des implants. Dans cette optique, il est possible
d’immobiliser une couche anti-adhérente sur les surfaces des biomatériaux pour empêcher
toute interaction non spécifique et ensuite, de greffer des molécules bioactives pré-
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déterminées, reconnues par les cellules environnantes et leur fournissant des signaux
spécifiques.
Les couches anti-adhérentes peuvent être utilisées seules lorsque aucune réaction n'est
désirée à la surface du biomatériau. Cela peut concerner les membranes de dialyse, pour
limiter leur obstruction, ou bien l'extrémité sensible des biosenseurs, afin d'augmenter le ratio
signal/bruit.5 Un des principaux objectifs de la modification de surfaces de biomatériaux par
immobilisation d'une couche anti-adhérente et le greffage subséquent de molécules bioactives
choisies est de favoriser l'intégration du biomatériau (c'est à dire une cicatrisation contrôlée) et
ainsi sa fonctionnalité. Des électrodes implantées dans le système nerveux induisent la
formation d'une cicatrice gliale qui gêne la mesure d l'activité neuronale (à long terme).6 La
surface de l'électrode peut être modifiée dans le but de limiter la prolifération des cellules
gliales et la formation de cette cicatrice tout en favorisant l'adhésion des neurones. De même,
l'implantation de prothèses vasculaires ou d'implants en contact avec le sang induit l'adhésion
des plaquettes à leur surface et la formation d'un caillot, et à plus long terme, d'autres
complications.7 Il est donc nécessaire de modifier la surface de ces implants de manière à
empêcher ces réactions et favoriser la formation d'une monocouche de cellules endothéliales
aux propriétés anti-thrombiques, c'est à dire dans leur état différencié. Enfin, les surfaces anti-
adhérentes sur lesquelles ont été greffées des biomolécules peuvent servir de modèles pour
l'étude de l'interaction entre une molécule et un type cellulaire et de ses conséquences sur le
comportement cellulaire.
Couches anti-adhérentes
Des couches anti-adhérentes composées de chaînes moléculaires sont immobilisées sur
les surfaces des biomatériaux pour prévenir toute interaction non spécifique menant à
l'adsorption de protéines et à l'adhésion cellulaire non spécifiques. Le caractère répulsif de ces
couches peut être expliqué par différents mécanismes.8 La résistance à l'adsorption protéique
peut être expliquée par un phénomène de répulsion stérique lorsqu'une protéine interagit avec
une couche dense et compressible.9,10 Elle peut également être due à la présence d'une couche
d'eau à l'interface “surface anti-adhérente/environnement”, prévenant toute interaction avec la
surface.11,12 La stabilité de cette couche d'eau dépend de l'orientation des fragments des
chaînes moléculaires à l'interface ainsi que de leur force d'interaction avec les molécules d'eau
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environnantes.12 Ainsi, la capacité des couches immobilisées à prévenir les interactions non
spécifiques dépend de leur structure. Ces couches, le plus souvent constituées de polymères
tels que le poly(éthylène glycol) (PEG) et les polysaccharides, sont généralement caractérisées
par une forte densité de chaînes, une faible densité de charge et présentent une interface
hydratée.12-15
Techniques de fabrication. Les surfaces de biomatériaux à modifier peuvent être
activées par oxydation, par silanisation (pour les surfaces contenant de la silice)16,17 ou bien
par déposition de molécules organiques par plasma (incorporation de groupements amines ou
carboxyles par exemple).15,18 Les groupements réactifs ainsi introduits à la surface permettent
l'immobilisation covalente des couches anti-adhérentes.
Les couches les plus étudiées et utilisées sont composées de PEG (aussi appelé oxyde
de polyéthylène, PEO), un polymère synthétique, hydrophile et neutre. L'efficacité des
couches de PEG dépend de la densité des chaînes à l'interface.9,15 Les chaînes de PEG portant
un groupement réactif tel qu'un aldéhyde ou une amine primaire à une extrémité, peuvent être
covalemment immobilisées sur une surface activée par l'action d'agents de couplage.19 Afin
d'augmenter la densité de chaînes l'interface, le PEG peut être immobilisé sous les conditions
de solvatation limite ou “cloud point”: sous ces conditions, les répulsions inter-chaînes sont
réduites durant la réaction d'immobilisation, résultant en une plus grande densité de chaînes à
la surface.15 Une forte densité de groupements oligo(éthylène glycol) (OEG) peut aussi être
obtenue par la formation de couches auto-assemblées (“self-assembled monolayers”, SAMs).13
Ces couches de structure et de chimie définies découlent de l'assemblage organisé de
molécules d'alkyle-silane ou d'alcane-thiol exposées respectivement à une surface de silice ou
à une surface métallique (or ou argent le plus souvent).20 Un bout de la chaîne moléculaire se
lie fortement à la surface alors que l'autre extrémité de la chaîne, qui peut être constituée d'une
variété de groupements fonctionnels, reste libre à la face externe de la couche. Les SAMs
terminées par des groupements OEG préviennent l'adsorption non spécifique de protéines.13
Les couches formées de polysaccharides reçoivent un intérêt grandissant puisque la
barrière protectrice présente à la surface des cellules, le glycocalyx, est composée de
polysaccharides. Ces polymères naturels, hautement hydratés, portent des groupements
hydroxyles qui peuvent être modifiés à différents degrés, soit par oxydation (formation de
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groupements aldéhydes),17,21 soit par addition de groupements réactifs (carboxyles, amines
primaires).18 Les chaînes de polysaccharides peuvent également être thiolatées et adsorbées
sur une surface d'or.14 Ainsi, des surfaces de biomatériaux recouvertes de dextran, alginate,
hyaluronane ou de leurs dérivés montrent un certain degré de résistance, selon leur structure à
la surface,11,18,21 et les couches de polysaccharides peuvent être aussi répulsives que le
PEG,22,23 tout en présentant plusieurs sites de liaison.
Immobilisation de biomolécules
Des biomolécules peuvent être greffées sur une couche anti-adhérente dans le but de
transmettre un signal promouvant l'adhésion cellulaire, la prolifération, la migration, la
différentiation ou bien la maintenance d'un phénotype différencié. Le signal est efficacement
transmis aux cellules si les biomolécules sont présentes dans une conformation stable,
reconnaissable par les récepteurs cellulaires, avec une densité et une orientation appropriées. Il
est aussi nécessaire de souligner l'importance de l'absence d'interactions non spécifiques pour
une transmission du signal optimale. Le but ultime de l'immobilisation de biomolécules est la
préservation de leur activité et de leur spécificité. Il peut s'agir de protéines, de peptides ou
séquences-signal, ou bien de fragments de protéine, immobilisés seuls ou en combinaison.
Modes d'immobilisation. La formation d'une liaison covalente entre la surface anti-
adhérente et les biomolécules est le mode de greffage le plus utilisé. La liaison se fait entre un
groupement fonctionnel exposé à la surface de la couche anti-adhérente et un groupement
fonctionnel présent dans un peptide (amine primaire, carboxyle ou thiol d'un acide aminé
terminal) ou bien exposé à l'extérieur d'une protéine (amine de la lysine, thiol de la cystéine),
via des agents de couplage.24,25 Si nécessaire, la liaison peut être faite via une molécule
d'espacement ou “spacer”, tel que le PEG, pour éloigner la molécule de la surface et
augmenter son exposition et sa liberté de conformation.26 La liaison covalente via l'utilisation
d'agents de couplage est une technique simple et polyvalente, cependant, ce mode
d'immobilisation est non spécifique et, dans le cas de protéines, offre peu de contrôle sur
l'orientation et l'activité de la molécule immobilisée.
Aussi, il existe des techniques de greffage par l'intermédiaire de complexes récepteur-
ligand reposant sur la reconnaissance spécifique des biomolécules, permettant l'attachement
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de molécules dans leur conformation native et avec une orientation connue. Ainsi, des
anticorps monoclonaux dirigés contre la biomolécule d'intérêt peuvent être immobilisés sur
une couche anti-adhérente;27 un système plus polyvalent repose sur l'utilisation de l'avidine
comme agent de liaison entre la surface et une biomolécule biotinylées.28 Enfin, des séquences
composées d'histidines peuvent être insérées dans des séquences protéiques (“histidine-tags”),
elles permettent l'attachement dans une orientation connue de ces molécules sur une couche
présentant des atomes métalliques chélatés.29
Enfin, l'incorporation de molécules bioactives dans une couche anti-adhérente lors de
la formation de celle-ci permet de contrôler leur densité à la surface.26,30 Ainsi, un hydrogel de
PEG bioactif peut résulter de la polymérisation de molécules de PEG acrylées et de molécules
de PEG liées à une séquence bioactive.26 Des molécules bioactives peuvent aussi être
incorporées dans une bicouche lipidique, composée de molécules amphiphiles (queue di-
alkyle lipidique et tête hydrophile contenant la séquence bioactive ou bien un groupement
OEG sur la couche extérieure).31,32 Ces structures, mimant les membranes cellulaires,
confèrent stabilité et mobilité aux biomolécules immobilisées.
Objectifs du travail
Les interactions spécifiques entre une surface et des cellules endothéliales sont
recherchées pour de nombreuses applications biomédicales. Par exemple, le recouvrement de
la surface luminale des prothèses vasculaires de petit diamètre par une couche de cellules
endothéliales stable et complète permettrait d’exposer au sang une surface non thrombogène
et ainsi de limiter leur occlusion.7,33 Également, l’interaction des cellules avec une surface
pourrait promouvoir l’infiltration de cellules endothéliales dans une structure polymérique
suivie de la formation de vaisseaux sanguins pour le développement d’un tissu fonctionnel par
génie tissulaire. Un tissu sain nécessite un apport constant en oxygène et en nutriments pour
croître et survivre.34,35 Enfin, inciter la formation d’un réseau de capillaires sanguins autour
d’un implant favoriserait son intégration dans l’organisme.4 Des surfaces favorisant
l’adhésion, la prolifération et/ou la migration plus ou moins contrôlées des cellules
endothéliales ont été développées mais peu permettent l’adhésion sélective et une action
6
spécifique sur le comportement de ces cellules dans des systèmes de co-culture en 2D voire en
3D, pour le développement de réseaux vasculaires fonctionnels.
L’objectif principal de ce travail est donc de développer et de caractériser une surface
bioactive modèle capable d’interagir spécifiquement avec des cellules endothéliales. Tout
d’abord, (i) le travail de recherche consiste à développer un revêtement de surface inerte,
résistant à l’adsorption non spécifique de protéines ainsi qu’à l’adhésion cellulaire. Pour ce
faire, des couches d’un polymère, immobilisées selon différentes conditions, sont caractérisées
et exposées à un mélange de protéines et à des cellules afin de déterminer les paramètres
influençant leur structure et leur résistance et également, de trouver une surface inerte
optimisée. Ensuite, (ii) des molécules bioactives spécifiques pour les cellules endothéliales
sont immobilisées sur ce revêtement et la réponse de différents types cellulaires exposés à ces
surfaces est étudiée.
Originalité
L’originalité de ce projet de recherche repose sur la méthode expérimentale employée.
Des puces de polymères sont développées et utilisées pour la caractérisation et l’optimisation
des surfaces inertes tandis que des puces de molécules bioactives sont fabriquées pour l’étude
du comportement cellulaire face à des surfaces de composition moléculaire variée. Les puces
sont réalisées à l’aide d’un robot habituellement utilisé pour la fabrication de puces d’ADN ou
de protéines. La possibilité de déposer sur une même surface une multitude de solutions de
compositions différentes permet non seulement de caractériser simultanément les différentes
surfaces créées, mais aussi d’étudier et de comparer l’effet d’un plus grand nombre de
variables (facteurs influençant l’immobilisation du polymère, composition des surfaces
bioactives). Enfin, la polyvalence du robot permet de créer des surfaces avec une taille, un
espacement, et un nombre d’échantillons variés et déterminés selon les besoins de
l’expérience donnée.
7
Structure de la thèse
Ainsi, après avoir présenté dans un premier chapitre le contexte dans lequel s’inscrit ce
travail de recherche, une revue de littérature abordant les interactions entre les cellules
endothéliales et leur environnement in vivo et exposant les modifications de surfaces
expérimentées pour favoriser l’adhésion des cellules endothéliales est présentée dans le
deuxième chapitre. En effet, la connaissance et la compréhension des interactions entre les
cellules et leur environnement dans l’organisme d’une part et au laboratoire d’autre part sont
essentielles pour la conception de surfaces intelligentes permettant l’intégration d’un
biomatériau et guidant la réparation ou le développement d’un tissu fonctionnel. Ce chapitre
sera soumis à la revue Frontiers in Bioscience.
Dans le troisième chapitre, la fabrication de puces de polymères est détaillée, leurs
propriétés physico-chimiques sont caractérisées par spectroscopie des photoélectrons-X (XPS)
et par microscopie à force atomique (AFM). Les puces sont également exposées aux protéines
de sérum. L’analyse des résultats a permis de corréler les conditions d’immobilisation des
couches de polymère avec leur structure et leur capacité à repousser les protéines. Cette étude
a fait l’objet d’un article publié dans le journal Langmuir.
Dans le quatrième chapitre, les puces de polymères sont exposées à des cellules afin de
poursuivre la caractérisation des conditions plus ou moins résistantes. Le mécanisme
d’invasion des cellules selon le degré de résistance et la structure de la couche est étudié et un
revêtement inerte optimal est déterminé grâce à l’ensemble des résultats obtenus. Ce travail
est décrit dans un article sous presse pour le périodique Journal of Biomedical Material
Research - Part A.
Le cinquième chapitre est dédié à l’étude de puces bioactives synthétisées par
immobilisation de molécules spécifiques pour cellules endothéliales sur la couche inerte.
L’effet de la composition moléculaire des spots bioactifs sur l’adhésion et la morphologie des
cellules endothéliales et des fibroblastes est analysé. Ce travail a été publié dans le journal
Biomacromolecules.
Enfin, les conclusions et perspectives propres à ce travail seront données.
8
Références
(1) Ratner, B. D.; Hoffman, A. S.; Schoen, F. J.; Lemons, J. E. Biomaterials science, an
introduction to materials in medicine.; 2nd ed.; Elsevier: 2004.
(2) Lutolf, M. P.; Hubbell, J. A. Nat. Biotechnol. 2005, 23, 47-55.
(3) Langer, R.; Vacanti, J. P. Science 1993, 260, 920-926.
(4) Ratner, B. D. J. Control Release 2002, 78, 211-218.
(5) Wisniewski, N.; Reichert, M. Colloids Surf. B Biointerfaces 2000, 18, 197-219.
(6) Fawcett, J. W.; Asher, R. A. Brain Res Bull 1999, 49, 377-391.
(7) Merzkirch, C.; Davies, N.; Zilla, P. Anat Rec 2001, 263, 379-387.
(8) Morra, M. J. Biomater. Sci. Polym. Ed 2000, 11, 547-569.
(9) Jeon, S. I.; Lee, J. H.; Andrade, J. D.; DeGennes, P. G. Journal of Colloid and
Interface Science 1991, 142, 149-158.
(10) Halperin, A. Langmuir 1999, 15, 2525-2533.
(11) Morra, M.; Cassinelli, C. Langmuir 1999, 15, 4658-4663.
(12) Harder, P.; Grunze, M.; Dahint, R.; Whitesides, G. M.; Laibinis, P. E. J. Phys. Chem.
B 1998, 102, 426-436.
(13) Prime, K.; Whitesides, G. Science 1991, 252, 1164-1167.
(14) Frazier, R. A.; Matthijs, G.; Davies, M. C.; Roberts, C. J.; Schacht, E.; Tendler, S. J. Biomaterials 2000, 21, 957-966.
(15) Kingshott, P.; Thissen, H.; Griesser, H. J. Biomaterials 2002, 23, 2043-2056.
(16) Tasker Langmuir 1996, 12, 6436-6442.
(17) Martwiset, S.; Koh, A. E.; Chen, W. Langmuir 2006, 22, 8192-8196.
(18) McArthur, S. L.; McLean, K. M.; Kingshott, P.; St John, H. A. W.; Chatelier, R. C.; Griesser, H. J. Colloids and Surfaces B-Biointerfaces 2000, 17, 37-48.
(19) Vermette, P.; Meagher, L. Colloids and Surfaces B-Biointerfaces 2003, 28, 153-198.
(20) senaratne, W.; Andruzzi, L.; Ober, C. K. Biomacromolecules 2005, 6, 2427-2448.
(21) Massia, S. P.; Stark, J.; Letbetter, D. S. Biomaterials 2000, 21, 2253-2261.
9
(22) Osterberg, E.; Bergstrom, K.; Holmberg, K.; Schuman, T. P.; Riggs, J. A.; Burns, N. L.; Van Alstine, J. M.; Harris, J. M. J. Biomed. Mater. Res. 1995, 29, 741-747.
(23) Pallarola, D.; Domenianni, L.; Priano, G.; Battaglini, F. Electroanalysis 2007, 19, 690-697.
(24) Johnsson, B.; Lofas, S.; Lindquist, G. Anal. Biochem. 1991, 198, 268-277.
(25) Lahiri, J.; Isaacs, L.; Tien, J.; Whitesides, G. M. Anal Chem 1999, 71, 777-790.
(26) Hern, D. L.; Hubbell, J. A. J. Biomed. Mater. Res. 1998, 39, 266-276.
(27) Calonder, C.; Matthew, H. W.; Van Tassel, P. R. J. Biomed. Mater. Res. A 2005, 75, 316-323.
(28) Marie, R.; Beech, J. P.; Voros, J.; Tegenfeldt, J. O.; Hook, F. Langmuir 2006, 22, 10103-10108.
(29) Kato, K.; Sato, H.; Iwata, H. Langmuir 2005, 21, 7071-7075.
(30) Murugesan, G.; Ruegsegger, M. A.; Kligman, F.; Marchant, R. E.; Kottke-Marchant, K. Cell Commun. Adhes. 2002, 9, 59-73.
(31) Dillow, A. K.; Ochsenhirt, S. E.; McCarthy, J. B.; Fields, G. B.; Tirrell, M. Biomaterials 2001, 22, 1493-1505.
(32) Ochsenhirt, S. E.; Kokkoli, E.; McCarthy, J. B.; Tirrell, M. Biomaterials 2006, 27, 3863-3874.
(33) Dardik, A.; Liu, A.; Ballerman, B. J. J Vasc Surg 1999, 29, 157-167.
(34) Bouhadir, K. H.; Mooney, D. G. J. Drug Target. 2001, 9, 397-406.
(35) Martin, Y.; Vermette, P. Biomaterials 2005, 26, 7481-7503.
10
Chapitre 2
Effets des propriétés de surface des biomatériaux et
de la bioactivation sur les cellules endothéliales
2.1. Résumé
Les interactions entre les cellules endothéliales vasculaires (ECs) et les matériaux sont
centrales pour des applications biomédicales telles que l’endothélisation de prothèses
vasculaires ou la vascularisation de substituts tissulaires. Afin d’améliorer les résultats des
implants, les surfaces des biomatériaux sont conçues pour promouvoir l’adhésion des ECs et
diriger leur comportement. In vivo, les ECs recouvrent tous les vaisseaux sanguins ; leur
morphologie, leur fonction et la matrice associée sont localement adaptées et spécifiques au
micro-environnement. Pour induire l’adhésion et la croissance des ECs, des traitements
modifiant les propriétés physico-chimiques et mécaniques des surfaces des matériaux ont été
développés. Les matériaux peuvent aussi être recouverts de molécules bioactives telles que
des protéines de la matrice, des peptides et/ou des facteurs de croissance afin d’étudier et
contrôler le comportement des ECs. Le but de cette revue est de donner un aperçu des
connaissances actuelles au sujet des ECs et de leur environnement solide in vivo et de leurs
réponses face aux surfaces synthétiques in vitro.
11
Effects of biomaterials surface properties and
bioactivation on endothelial cells
Emmanuelle Monchaux, Patrick Vermette
2.2. Abstract
Interactions between vascular endothelial cells (ECs) and materials are central to
biomedical applications such as vascular graft endothelialization or vascularization of an
engineered tissue substitute. To improve implant success, biomaterial surfaces are designed to
promote EC adhesion and direct their response. In vivo, ECs line all blood vessels; their
morphology, function and associated matrix are locally adapted to and specific for the
microenvironment. To enhance EC adhesion and growth, surface treatments have been
developed that modify materials surface physico-chemical and mechanical properties.
Materials may also be coated with bioactive molecules such as proteins from the matrix,
peptides and/or growth factors to study and control EC behavior. The aim of this review is to
give an overview of current knowledge about EC and their solid environment in vivo and their
responses to synthetic surfaces in vitro.
12
2.3. Introduction
Adhesion of endothelial cells (ECs) on biomaterial surfaces and subsequent controlled
behavior are of increasing importance in the biomedical field with implication in the
endothelialization of vascular grafts or in the formation of a vascular network in engineered
tissue substitutes. Prosthetic vascular grafts used to replace small-diameter blood vessels
(diameter < 6mm) are characterized by a reduced patency and occlusion: thrombosis and
intimal hyperplasia are the main reasons for the high failure.1 The lack of a complete
endothelium covering blood contacting devices is a major contributing factor to both
phenomena. One approach to prevent thrombosis and to improve the hemocompatibility of
synthetic vascular grafts is to create a functional, quiescent monolayer of ECs on the graft
surface prior to implantation. Another solution is to develop implants that will enhance
endothelialization upon implantation.2 ECs in-growth and formation of a functional, mature
vascular network remain a challenge in tissue engineering research; this network is required
for the construction or regeneration of hybrid tissues.1,3 Similar to normal tissues, engineered
tissues need blood supply to grow and to remain viable. In addition, implant biocompatibility
could be improved by promoting a normal wound healing response including peri-implant
vascularization and reduced encapsulation.4
Biomaterial science and tissue engineering rely heavily on cell-material interactions:
surfaces may induce cell adhesion, determine cell fate and promote the regulated development
of functional structures.5 In vivo, cells are anchored to their extracellular matrix (ECM) and
cell-ECM interactions modulate cells survival and responses.6 Biomaterial surfaces should
thus be designed to mimic cells biological environment and the knowledge of the interactions
of cells with their natural environment in an organism is essential to develop implants that will
be integrated into the host organism.
The objectives of this review paper are to give an overview of ECs and their natural
environment in vivo and to present surface modifications and their effect on EC responses in
vitro. In the first part of this paper, normal endothelial tissue characteristics and functions are
presented, and interactions with its ECM in vivo are described with a particular interest in the
dynamic process of blood vessel formation through angiogenesis. The second part deals with
EC interactions with synthetic surfaces. A multitude of surface chemical modifications and
13
bioactive coatings have been developed to promote EC adhesion onto biomaterials. Substrate
properties and immobilization mode influence proteins binding to their receptors and
consequently, cell response. Combined immobilization of various signaling molecules and its
effects on cell responses are also discussed.
2.4. The endothelial tissue and its matrix
Vascular endothelial cells (ECs) are a specialized type of epithelial cells lining the
inner surface of blood vessels of the entire circulatory system, from the heart, arteries and
veins to smallest capillaries. They do not form a passive barrier between circulating blood and
surrounding tissues. In fact, the endothelium provides a non-thrombogenic surface,
communicates with the surrounding microenvironment and releases biochemical regulators.
ECs hence form a heterogeneous tissue, they exhibit a great diversity in morphology and
functions among the vascular tree depending on vessel type, tissue irrigated and activation
state. Interactions between ECs and the ECM more particularly are crucial for the
maintenance of ECs integrity and functions and for the controlled formation and regeneration
of blood vessels.
2.4.1. Morphological and functional heterogeneity
Morphological diversity
Walls of large vessels like arteries and veins consist of three layers: an inner intima
made up of a layer of ECs attached to their basement membrane, an intermediate media
mainly composed of smooth muscle cells (SMCs) and elastic fibers and an outer adventitia
made of collagenous ECM with fibroblasts, blood vessels and nerves. Arteries are muscular,
elastic blood vessels with thick walls that possess elastic laminae surrounding the intima and
the media, and that pulsate. Veins have thin walls, they do not pulsate but possess valves.7 In
vivo, vascular ECs experience fluid shear stress, the tangential component of hemodynamic
stresses. In large straight arteries of uniform geometry, the mean wall shear stress is between
10 and 20 dynes/cm² while in regions of non-uniform geometries (branches and arches)
transient shear stress can be as high as 50 dynes/cm² with pulsatile flow. ECs are thick and
14
aligned in the direction of blood flow in straight segments of arteries but not at branch points;
in veins, ECs are shorter and flat and are not aligned in the direction of blood flow.7
Arterioles and venules are intermediate vessels between capillaries and arteries and
capillaries and veins, respectively. Pre-capillary arterioles are completely surrounded by one
or two layers of SMCs and post-capillary venules are surrounded by pericytes embedded in
their basement membrane.8
Capillary microvessels represent the most abundant vessels in an organism and consist
of ECs surrounded by a basement membrane and occasional pericytes, allowing direct
physical contact between endothelial and tissue cells. ECs in capillaries are flattened,
elongated, they adapt to their microenvironment and acquire specialized characteristics to
accommodate local physiological requirements.7,9,10 Continuous endothelium consists of ECs
tightly connected to each other via tight junctions and surrounded by a complete basement
membrane. It is found in capillaries of the brain, heart, skin and lung, as well as in arteries and
veins. Further specialization of the continuous endothelium is observed in blood-brain, blood-
retina and blood-testis barriers with acquisition of complex tight junctions and highly
regulated transcellular transports. Fenestrated continuous capillaries are characterized by the
presence of small openings called fenestrae, and are found in capillaries with an increased
fluid exchange between blood and tissues: diaphragmed fenestrated capillaries are found in
endocrine and exocrine glands, gastric and intestinal mucosa, and renal tubules whereas non
diaphragmed fenestrated are present in renal glomerulus. Finally, discontinuous endothelium,
characterized by the presence of many large fenestrations with no diaphragm and gaps,
presents a poorly formed basement membrane (discontinuous or absent) and is found in more
restricted regions such as capillaries of the liver, spleen and bone marrow.
Functional diversity
Transport function. Capillaries form the main site of exchange of nutrients between
blood and tissues. They use several specific transport mechanisms to meet the metabolic needs
of the surrounding tissue cells. Fluids and small solutes move passively across the barrier via
the paracellular pathway, regulated by intercellular tight junctions, whereas macromolecules
use transcellular transports, controlled by the presence of specific membrane receptors or
vesicular carriers such as caveolae and vesiculo-vacuolar organelles.9,11 Spatial heterogeneity
15
of permeability depends on differences in junctional properties and presence or absence of
fenestrae and gaps.
Vasomotricity. Transport may be regulated by blood perfusion which is locally
controlled by vasomotricity of pre-capillary arterioles. Generally speaking, ECs finely control
blood flow in response to metabolic demand (oxygen tension and glucose concentration, for
instance), cytokines and shear stress, by acting on SMCs in vessel walls of arterioles and large
vessels. This regulation is short and local, through the production and catabolism of
vasoactive molecules by ECs, either vasodilators such as nitric oxide (NO), prostacyclin
(PGI2) or vasoconstrictors such as endothelins.11
Host defense and inflammation. The endothelium actively participates in the
inflammatory response following tissue infection or irritation, mainly at post-capillary venule
sites where cell-cell junctions are looser. Activated ECs (i) secrete vasoactive molecules to
locally increase permeability, (ii) express receptors for immune cells adhesion such as
vascular cell adhesion molecule (VCAM) and intercellular adhesion molecule (ICAM) and
(iii) secrete cytokines for the recruitment of leucocytes and induction of angiogenesis. EC
activation allows adhesion and transmigration of leucocytes to inflammation sites and
neovascularization of the injured tissue.9
Vascular hemostasis. The endothelium lining arteries, veins and all blood vessels
provides a non-thrombogenic, anti-coagulant surface by the secretion and/or surface
expression of several regulatory factors that maintain blood in a fluid state. An intact EC
monolayer is covered by a layer of glycocalyx containing anti-coagulant heparan sulfate
proteoglycans (HSPGs) and anti-thrombic thrombomodulin. ECs also secrete vasodilators that
prevent platelet adhesion. When a vascular lesion occurs, platelets adhere to exposed vessel
walls, ECs and surrounding cells secrete pro-coagulant molecules leading to the formation of
a fibrin clot and finally EC produce fibrinolytic effectors to limit clot formation. The nature of
factors involved in vascular hemostasis depend on location in the vascular tree.11
16
2.4.2. Endothelial cells – matrix interactions
Cell-ECM interactions
ECs interactions with their underlying ECM are essential for maintenance of cell
integrity and functional activity, and for the formation of functional mature blood vessels. The
ECM provides mechanical support and biochemical cues for cell adhesion, migration,
proliferation and differentiation via interactions with cell membrane receptors and through
growth factor sequestering. Matrix proteins and more particularly adhesive ones such as
fibronectin (Fn), laminins (Ln) and vitronectin (Vn) possess many binding domains capable of
interacting with other ECM proteins as well as with cell surface receptors. Defined amino acid
sequences present within ECM molecules specifically bind cell surface receptors to trigger
various intracellular pathways. Cell-ECM adhesions are mediated primarily by integrin
receptors, heterodimeric transmembrane proteins composed of α and β subunits that connect
the ECM molecules to the cell cytoskeleton. When bound to their specific ligand, integrins
cluster, form focal adhesion structures, mediate cell anchorage to the underlying matrix, and
can also initiate signaling cascades transduced to the nucleus.12 These events may affect many
aspects of the cell responses such as proliferation, differentiation, migration and survival. A
single cell binding motif can be found within several proteins, such as the most investigated
Arg-Gly-Asp (RGD) sequence present in Fn, Vn and Ln among others. A protein is able to
bind several receptors through various sequences which exposition depends on protein self-
assembly into fibers or a network, its association with other ECM molecules or its proteolytic
degradation. Moreover, cell membrane receptors frequently associate with other receptors
such as integrins or growth factor receptors, allowing integration of diverse signaling
pathways. Hence, signals transduced to cell nucleus depend on the set of membrane receptors
expressed by cells, as well as on the composition, the structure and the spatial organization of
the underlying ECM which are characteristic of a tissue at a given time.6
Vascular basement membrane
Basement membranes are specialized types of ECM, highly cross-linked and organized
in a sheet-like structure that separate the epithelium from the connective tissue. They function
as selective filters, maintain mature tissue function and define spatial organization during
17
tissue development and reconstruction following tissue injury, by regulation of cell growth,
differentiation, and migration.13 The upper layers, called basal lamina, are secreted by
epithelial cells and consist of a network of Ln and a network of collagen type IV
interconnected via nidogen/entactin. The lower layer of the basement membrane is secreted by
cells from the underlying connective tissue and contains fibrils of collagen type I and type III
and Fn.13,14
The Ln network assembly is necessary for the basal lamina formation and plays an
essential role in cell adhesion and signaling.14 The laminins are a family of heterotrimeric
molecules composed of α-, β- and γ-chains. Ln α-chains possess many receptor binding sites
and they are expressed in a tissue-specific and developmentally regulated manner, conferring
heterogeneity among basement membranes.13,15 ECs express only 2 Ln α-chains: Ln α4 which
is a component of Ln-8 (α4β1γ1) and Ln-9 (α4β2γ1) and α5 which is a component of Ln–10
(α5β1γ1), Ln-11 (α5β2γ1) and Ln-15 (α5β2γ3). Ln α4 is the predominant α chain found in
vascular basement membranes and is expressed by all types of ECs, both during development
in the embryo and in the adult, while Ln α5 is detectable in basement membranes of quiescent
mature vessels, primarily in capillaries and some venules after birth and is not associated with
a fenestrated endothelium.16,17 Ln-10 is believed to be involved in vessel maturation and
stability.16,18 ECs bind to the Ln network via integrin receptors including integrins α3β1 and
α6β1 for both Ln α4 and α5 chains and αvβ3 and αvβ5 for Ln α5 chain via exposed RGD
sites. Ln α5 also binds the α-dystroglycan and the Lutheran blood group transmembrane
glycoproteins.16,19
Collagen type IV network provides the scaffold mechanical resistance.13 In addition,
network forming collagen type VIII, closely associated with human vascular basement
membranes and therefore used as a marker for blood vessels, maintains vascular basement
membranes in an open porous structure.16 Ln and collagen IV networks are principally linked
by the nidogen/entactin-2 isoform.16
Proteoglycans (PGs), proteins with glycosaminoglycan side chains, associated with
basement membranes play a structural role in maintaining tissue architecture via interactions
with matrix proteins, help in selective filtration, sequester soluble growth factors via their
heparan sulfate (HS) side chains and thus help in regulating cell differentiation.13,14 In addition
to perlecan, agrin and collagen type XVIII, heparan sulfate proteoglycans (HSPGs) associated
18
with basement membranes, leprecan and collagen type XV chondroitin sulfate PGs are
detected in vascular basement membranes.14-16
Thrombospondins (Tsp) belong to a group of proteins called matricellular, which
interact with cell receptors and matrix components and are defined as regulators of cell
function; they are expressed at high levels during development and in response to injury.20
Tsp-1 and -2 are strong inhibitors of angiogenesis, they are both involved in pathological
conditions and Tsp-2 is associated with developing vessels.16,20 von Willebrand factor, a
thrombogenic molecule released by EC that favors platelet adhesion, is primarily found in
veins basement membranes.11 Finally, as ECs in capillaries may be in close contact to
surrounding tissue cells, they interact with composite basement membranes containing other
isoforms including Ln α-chains produced by surrounding cells.16
2.4.3 Formation of blood vessels
Context and initiation
Blood vessels in the embryo develop through vasculogenesis i.e., through in situ
differentiation of mesodermal precursor cells, called angioblasts, into ECs that assemble into a
primary capillary plexus. The primitive network is then grown, remodeled and stabilized by
the process of angiogenesis into a complex, mature, and functional network.21 In the adult,
ECs interact with a laminin-rich ECM that maintains mature vessel in a stable quiescent state.
During regulated physiological processes such as endometrium vascularization or wound
repair, ECs undergo rapid proliferation to form new vessels following matrix remodeling via
sprouting angiogenesis. Activated ECs degrade the underlying basement membrane, migrate
and proliferate in the perivascular matrix, form tubular structures that become mature and
functional.22 Under these conditions, angiogenesis is transitory and highly regulated, spatially
and temporally. However, many diseases such as arthritis, diabetes and tumor growth are
driven by a persistent unregulated angiogenesis. Thus, control of the angiogenic process is
essential and relies especially on regulated cell-matrix interactions. Gene knock-out
experiments in mice and in vitro/in vivo experiments allowed to define a model mechanism of
sprouting angiogenesis.
19
Local hypoxia, hypoglycemia, shear stress or inflammation induce the release of pro-
angiogenic vascular endothelial growth factor (VEGF) and fibroblast growth factor (FGF)-2
that in turn attract ECs.18,22-24 Activated ECs locally increase vascular permeability thereby
allowing extravasation of plasma proteins that lay down a provisional matrix rich in fibrin and
fibronectin. They also secrete proteinases such as plasminogen activators (PA) and matrix
metalloproteinases (MMP) that degrade ECM proteins and liberate growth factors sequestered
within the ECM. Angiopoietin (Ang)-2, antagonist of the Tie2 receptor, destabilizes existing
vessels probably by loosening ECs adhesion with local basement membrane and
periendothelial cells that surround and support blood vessels.21
Proliferation and migration
Local ECM remodeling not only creates a free path for ECs to migrate towards the
angiogenic stimulus, but also induces the participation of a new set of cell-matrix interactions,
eliciting new signals. The release of matrix-bound growth factors such as FGF-2 and heparin-
binding forms of VEGF induces angiogenic signals promoting EC proliferation and migration
and modulates cell integrin expression.23-25 Matrix protein cleavage or conformational changes
following remodeling expose cryptic sites within matrix proteins that alter their function
promoting EC proliferation and migration.25 Collagen type I and type IV usually interact with
α1β1 and α2β1 integrins whereas cleaved molecules expose cryptic sites interacting with
αvβ3.15 Proteolytic degradation also induce production of soluble fragments such as
endostatin, from collagen type XVIII, that can exert anti-angiogenic effects by inhibiting EC
proliferation and migration and thus permit control of the angiogenic process.25 Finally, ECs
previously exposed to a laminin-rich stabilizing ECM, then interact with a new set of ECM
molecules from the fibronectin-rich provisional matrix, enhancing proliferation. Formation of
new blood vessels in embryos and in adult organisms relies upon different endothelial
integrins and ECM ligands: in the embryo, a successful vascular development depends on Fn
and its major receptor α5β1, but not on αvβ3 and αvβ5, which are up-regulated in adult and
pathological angiogenesis.18,24,26 Non-proliferative cells located at migrating tip are exposed to
the interstitial matrix rich in collagen type I and type III fibers.24 Concurrently, ECs
reorganize into a cord-like structures, which acquire a lumen and form interconnected tubes.17
20
Stabilization and maturation
Vessel mature or regress depending on their use in the network, which is modeled by
blood flow generated forces. Hemodynamic forces induce modifications of cell-cell and cell-
ECM adhesions and are believed to up-regulate growth factors such as platelet-derived growth
factor (PDGF).8,22,27 ECs recruit PDGF receptor-α-expressing periendothelial cells and
contact between ECs and mural cells triggers the activation of transforming growth factor
(TGF)-β that inhibits ECs proliferation and migration, induces SMCs differentiation,
stimulates basement membrane production and deposition and alters integrin
profiles.8,18,22,23,27 Ang-1, a Tie2 ligand, stabilizes EC-EC interactions and adhesions of mural
cells with ECs.21,23 Mural cells adhesion and deposition of a complete stable basement
membrane provide stability against rupture or regression in absence of VEGF, except for a
fenestrated endothelium.23 EC interaction with periendothelial cells is essential for deposition
of a complete stable basement membrane.
Further vessel specialization is realized by EC interaction with repelling cues for
arterio-venous determination and guided vessel branching, heterotypic interaction with
periendothelial cells such as astrocytes involved in the blood-brain barrier formation,
interaction with organ-specific growth factors such as endocrine gland (EG)-VEGF for
endocrine gland capillaries specialization, and other non determined processes.7-10
2.5. Interactions between endothelial cells and biomaterial surfaces
Endothelial cell adhesion on biomaterial surfaces is required to provide a non-
thrombogenic surface to vascular prosthesis, to subsequently induce formation of blood
vessels around implanted biomedical devices improving their biocompatibility or to promote
vascularization of growing (hybrid) tissues for regenerative medicine. Cell-material
interactions must then incite EC adhesion, but also allow cells to maintain their differentiated
functional state and, in some cases, guide spouting i.e., regulated cell proliferation and
migration.
The ability of a material to support cell adhesion depends on its surface properties.
Treatments to modify a surface chemistry or topography have been used to induce protein
adsorption and subsequent cell adhesion. Surface attachment of bioactive molecules such as
21
ECM proteins, adhesive peptide sequences and/or growth factors have been used to promote
EC adhesion onto materials and to control cell processes such as proliferation, migration,
differentiation and survival.
2.5.1. Surface properties and the mechanical environment
Physico-chemical properties
Cell attachment to synthetic surfaces relies heavily on the presence of adsorbed
proteins from the media containing serum, from blood plasma or from cellular secretion of
matrix proteins. The physico-chemical properties of the surface (hydrophilicity, chemical
composition, charge) determine the composition of the adsorbed protein layer as well as the
amount and conformation of adsorbed proteins. Protein adsorption results from a combination
of interactions with material surface including hydrophobic interactions, electrostatic forces,
hydrogen bonding and van der Waals forces.28
It is generally observed that ECs adhere and spread on moderately hydrophilic surfaces
such as tissue culture polystyrene (TCPS) and glass, while EC adhesion is reduced or even
absent on hydrophobic surfaces such as polytetrafluoroethylene (PTFE),
polyethyleneterephthalate (PETP, Dacron) and fluoroethylenepropylene (FEP, Teflon).29-31
Differences in surface hydrophilicity result in quantitative and qualitative variations in the
composition of the adsorbed protein layer. Proteins preferentially bind to different surfaces
dependent on their nature, and adsorption onto a surface induces protein conformation
changes, more or less important, depending on the surface properties.32 Hydrophobic surfaces
exert strong attraction with hydrophobic parts of the protein (inside); proteins strongly and
usually irreversibly adhere to these surfaces and may undergo denaturation i.e., disruption of
native conformational state, altering exposition of cell binding sites naturally exposed on the
outside of the protein. When materials are exposed to blood plasma or serum, albumin
strongly binds to hydrophobic surfaces while conformationally active adhesive proteins such
as Fn and Vn preferentially adsorb on hydrophilic surfaces.33,34 As adsorption onto
hydrophilic surfaces is reversible, proteins can be displaced.29
Many polymeric biomaterials used for clinical applications (vascular grafts) are
hydrophobic (PTFE, PETP, FEP, polyurethane). A solution to promote cell adhesion is to
increase surface hydrophilicity by chemical treatments (UV exposition, alkaline hydrolysis) or
22
plasma treatment (ammonia or oxygen plasma treatment), all resulting in the addition of polar
groups and charges.35-37 TCPS and Primaria, routinely used for cell culture, have been made
hydrophilic and cell adhesive by plasma treatment. Since EC adhesion to hydrophobic
polymeric surfaces has been enhanced by plasma modification using either nitrogen- or
oxygen-containing monomers: modified surfaces are more hydrophilic, adsorb more adhesive
proteins such as Fn and support cell adhesion, spreading and growth.38-41 Moreover, EC
adhering on plasma treated surfaces show improved anti-coagulant and fibrinolytic activities
and better resist to shear stress induced detachment.38,40
Surface exposition of various chemical groups and charges affect initial cell adhesion.
In presence of serum, initial cell attachment on nitrogen rich surfaces such as Primaria is a
result of adsorption of Vn and Fn, while cell adhesion on oxygen-rich surfaces such as TCPS
is mediated by adsorbed Vn only, as Fn adsorption on these surfaces is sub-optimal for cell
adhesion.39,40,42,43 Oxygen-rich surfaces present negatively charged groups while nitrogen rich
surfaces rather expose positively charged groups (at physiological pH). Fn is an acidic protein
overall negatively charged that more abundantly adsorbs on positively charged surfaces.
Introduction of electrical charges on a surface can enhance protein adsorption via electro-
attractive forces. Addition of a polyelectrolyte film on a poorly adhesive surface, either as a
monolayer or a multilayer film resulting from alternate adsorption of polycations and
polyanions, enhances cell attachment.44,45 In case of weak polyelectrolyte gels, EC adhesion
increases with charge density while cells highly adhere and proliferate on strong
polyeletrolytes, they also expose higher amount of anti-platelet HSPGs and are more resistant
to shear stress than confluent ECs attached on glass or TCPS.44,46,47
The mechanical environment
Cells sense and respond to underlying substrate stiffness, to topography and to fluid
flow. External mechanical forces can be sensed by cytoskeleton linked receptors or the
cytoskeleton structure itself. The mechanical coupling of the cytoskeleton with cell-ECM and
cell-cell adhesions allow transduction of mechanical signals throughout the cell.48 Upon
adhesion, cells form adhesive contacts with surfaces and pull on the substrate. Increased
surface stiffness is associated with increased cell contractility.49 On stiff surfaces, cells have
an increased focal adhesion formation and a more organized cytoskeleton with formation of
23
actin stress fibers, resulting in cell spreading and a higher adhesion strength than on soft
substrates.45,49,50 Substrate stiffness and cell contractility appear to regulate EC proliferation
and differentiation: ECs form tube-like structures on soft malleable substrates while they
proliferate on rigid surfaces.45,51,52
ECs are also sensitive to surface topography.53 When attached to a surface presenting
micro- or even nano-features, they emit filopodia and lamellipodia, sensory organs of cells,
and undergo morphological changes.54 On grooved surfaces, ECs elongate, they align parallel
to the channel direction; cell orientation persists until near confluence is reached and increases
with channel depth.55-57 Cell actin stress fibers and focal adhesions align parallel to channel
direction and focal contacts are preferentially located at feature edges.56 Their localization
correlates with preferential Fn fibrils formation at edges of grooves, pillars or wells.58
However, enhancement of cell elongation and orientation associated with actin stress fibers
alignment has also been observed on wave features, in absence of sharp edges.59 Grooved and
waved topographies induce polarization of cell movement: cell orientation and directed
movement in response to substrate topography may be referred to as “contact guidance”.53
Random topography or increased surface roughness have been shown to enhance EC adhesion
but also Fn and Vn adsorption.60,61 Hence, surface roughness and topography may affect cell
morphology directly by directing focal adhesions and stress fibers assembly, but they may
also influence cell behavior indirectly by altering surface protein adsorption.
The shear stress applied to the luminal surface of cells can be sensed by the cell
membrane and the associated receptors and this stress can be transmitted throughout the cell
to cell-matrix and cell-cell adhesions. As a result, shear stress induces EC alignment and an
increase of stress fibers and remodels cell-matrix adhesions to increase adhesion strength. In
confluent ECs, shear stress increases the size of focal adhesions and activates integrins such as
αvβ3 and α5β1.48 In non confluent ECs, shear stress promotes cell spreading, then
lamellipodial extension in the flow direction, cell-cell adhesion dissociation and enhances
directed cell migration. The existing focal adhesions under the main cell body increase in size
while new transient focal adhesions assemble under the lamellipodia and align with the flow
direction.48,62 Shear stress preconditioning may be used to orient ECs or to increase cell
adhesion strength to the substrate.63
24
2.5.2. Pre-coatings made of matrix proteins
Protein coatings
Surface treatment may enhance cell adhesion by modifying protein adsorption.
However, the exact composition of the adsorbed protein layer is not known and cannot be
controlled or reproduced. Moreover, surface modification may not be sufficient for the
formation of a confluent endothelial monolayer. Hence, individual components of the ECM
such as Fn, collagen type I and IV have been used as surface coatings to facilitate cell
attachment onto biomaterials and to promote their biocompatibility.29,34 The most used and
studied coating consist in an adsorbed or immobilized layer of Fn which strongly enhance EC
adhesion and spreading, the formation of focal adhesions, and the organization of actin
filaments into stress fibers, via interaction with its main receptors, integrins α5β1 and αvβ3.
ECs anchored to Fn proliferate, migrate, resist shear stress induced detachment and keep their
differentiated phenotype.29,64 Vn and fibrinogen, both αvβ3 ligands, enhance EC adhesion,
spreading and motility.25,65,66 Collagen type I or I/III and collagen type IV promote EC
adhesion and retention and induce EC migration to a higher extent than Fn via α1β1 and
α2β1.65,67-70 Laminins also support EC attachment and spreading but to a lesser extent than
Fn, collagen and Vn.65 ECs adhere more strongly and spread on Ln-10 rather than Ln-1 or Ln-
8, with no formation of focal or even fibrillar adhesions, indicative of a motile phenotype.19,71
Ln-10 also better supports EC migration.71
Osteopontin and tenascin, two matricellular proteins mainly detected during
development and in response to tissue injury i.e., associated to remodeling, promote EC
adhesion and migration via αvβ3 integrin in particular.20,25,72 On the contrary, Tsp and
SPARC are anti-adhesive proteins, they induce an intermediate state of adhesion associated
with a disruption of focal adhesions formation; they inhibit migration and VEGF-induced
proliferation.20
The nature of the coated protein and the various expressed integrins may differentially
regulate intracellular signaling. Cell attachment to a surface and subsequent behavior also
depend on protein density, conformation and spatial organization which, in turn, depend on its
interaction with the underlying substrate.
25
Protein-surface interaction and cell adhesion
Collagen type I molecules self-assemble into fibrillar structures or into a network only
when adsorbed on hydrophobic surfaces.73,74 ECs adhering to collagen fibrils have a reduced
spreading and a weak cytoskeleton organization corresponding to a motile phenotype, while
ECs on a collagen type I network spread and present actin stress fibers, indicative of strong
adhesion.74 Furthermore, collagen type IV network formation on a hydrophobic surface
enhances EC adhesion and retention.68 Collagen spatial rearrangement may expose or hinder
cell binding sites, thus changing amino acid sequence, density and spatial distribution of
exposed motives. In the same way, Fn adsorbed on a hydrophilic surface support higher EC
adhesion than when adsorbed on a hydrophobic surface and enhances cell retention.64,75 The
protein exposes a higher density of cell binding sites when exposed to a hydrophilic substrate,
consequently, cells may form more bonds with the underlying surface resulting in a higher
adhesion strength.75-77
Cell adhesion studies on hydrophilic model surfaces exposing neutral, positively
charged or negatively charged groups and precoated with Fn showed that chemistry alter Fn
conformation and strongly influences subsequent integrin binding. Depending on surface
chemistry, Fn binds with different affinity levels α5β1 and/or αvβ3 integrins. Accordingly,
focal adhesions assembly, composition and then elicited signals differ depending on surface
chemistry and Fn conformation.78-80 Furthermore, Fn adsorbed onto hydrophilic negatively
charged surfaces is weakly bound and can be displaced. The protein is then rearranged into
fibrils by ECs via interaction of its N-terminal domain with cellular α5β1.81 Fn fibrils
formation is associated with formation of fibrillar adhesions rich in α5β1 integrin and an
increased assembly of focal adhesions. On the contrary, Fn covalently attached to a surface or
adsorbed to mild hydrophobic or positively charged surfaces strongly interacts with the
substrate and cannot be rearranged into fibrils by cells. EC attached on strongly bound Fn
present fewer focal adhesions than when anchored to weakly bound Fn.36,82 Similarly,
fibroblasts adhering on a strongly bound Fn layer cannot rearrange Fn into fibrils, they present
focal contacts rich in α5β1 and show an increased adhesion strength associated with an
increased Fn-α5β1 bonds density.83,84 On the other hand, fibroblasts exposed to weakly bound
Fn form fibrillar adhesions associated with Fn fibrils and focal adhesions rich in αvβ3 integrin
26
and they present a higher motility.84,85 Hence, integrin binding to Fn result in different cell
behavior depending on substrate pliability. Similarly, when exposed to FGF-2 for a few days,
ECs attached on weakly bound Fn rearrange their Fn matrix into cords and form capillary-like
structures, while ECs anchored to Fn strongly interacting with the underlying substrate
proliferate and form a confluent monolayer resistant to shear stress detachment.36,64 Then, EC
formation of capillary-like structures on a surface seems to be related to cell anchorage
strength to surface that in turn depends on the matrix rigidity as well as on its biochemical
composition.
Strategies to immobilize and orient proteins in order to favor ligand-integrin binding
and then cell adhesion strength have been developed. Most current protein immobilization
procedures involving covalent chemical coupling result in a random distribution of protein
orientations on the surface, leaving a significant fraction of binding sites inaccessible to cells,
and may be responsible for loss of protein biological activity. Since, an intermediate layer of
antibodies or peptides directed towards a region away from the Fn cell binding site have been
used to orient Fn molecules.86,87 An alternative approach is to take advantage of natural
interactions between ECM molecules to both bind and orient a biomolecule naturally, thereby
enhancing its biological activity. Osteopontin naturally binds collagen type I and adsorption of
osteopontin on a collagen I intermediate layer enhances EC adhesion.88 Likewise, Fn naturally
interacts with HSPGs in the matrix and Fn interaction with heparin, a natural analog of
HSPGs, increases exposition of its cell binding and VEGF binding sites, enhancing VEGF-
induced EC proliferation and migration.76,89
2.5.3. Grafting of peptides
Sequence and receptor selectivity
Since the discovery of amino acid sequences within ECM proteins specifically
recognized by cell receptors, many researchers have immobilized cell recognition peptides
directly onto material surfaces in order to control elicited signals and subsequent cell
behavior. In addition to providing binding specificity, peptides present the advantage of being
conformationally stable, easy to synthesize and to modify. They usually are firmly linked to
surfaces either directly or via a spacer arm to enhance their steric availability and
conformational freedom, thus promoting their binding. Control of the desired signal and
27
subsequent cell responses are better achieved if there is minimal non-specific protein
adsorption. For that purpose, peptides can be immobilized on a low-fouling layer that resist
(or at least limit) non-specific protein adsorption. These layers can be made of ethylene glycol
or poly(ethylene glycol) (PEG, often also referred as to poly(ethylene oxide)) layers,90,91
polysaccharides layers,92-94 and phospholipid bilayers.95-97
RGD is the most effective and most often used peptide sequence to promote adhesion
on synthetic surfaces.98 The RGD sequence was first identified in the 10th Fn type III repeat
(FnIII-10), then in other matrix proteins such as Vn, Ln and collagen and many integrins bind
molecules in a RGD-dependent manner such as integrins α5β1 and αvβ3, which are
predominant cell adhesion mediators. Amino acids flanking the RGD sequence and
conformation of the amino loop are mainly responsible for their different integrin affinity and
selectivity. Cyclic RGD peptides or kinked sequences such as GRGDY and GRGDVY
selectively bind the αvβ3 integrin while linear RGD sequences such as GRGDSP
preferentially mediate cell adhesion via α5β1.96-98 Cell integrin expression pattern determines
relative adhesion strength on a peptide surface. ECs, SMCs and fibroblasts equally adhere and
spread on linear RGD sequences while cyclic RGD peptides preferentially mediate ECs and
SMCs adhesion via αvβ3. Concerning integrin selectivity, some other non-RGD signal
recognition sequences are used in integrin binding. REDV peptide, derived from Fn
alternatively spliced CS5 domain, selectively addresses integrin α4β1 expressed in a small
number of cell lines. REDV specifically mediates adhesion of ECs but not SMCs nor
fibroblasts via α4β1.99-101 The SVVYGLR sequence, exposed after proteolytic cleavage of
osteopontin, mediates EC adhesion and migration via integrin α9β1.102 Integrin subunits α4
and α9 are both responsible for inhibition of cell spreading and enhanced cell migration.103,104
YIGSR, from Ln β1-chain, selectively binds the 67kDa laminin-binding receptor and
promotes cell adhesion with no formation of focal adhesions or actin stress fibers.99,105
Finally, sequences from Fn C-terminal heparin-binding domain II such as WQPPRARI and
SPPRRARVT bind cell surface proteoglycans syndecan-1 and syndecan-4, components of
focal adhesions. Indeed, these immobilized peptides support cell adhesion with formation of
focal adhesions and actin stress fibers.106
28
Immobilization and cell responses
Adhesive signals. Ligand density and distribution on the surface influence cell
behavior. EC adhesion, spreading, focal adhesion formation and proliferation increase with
RGD peptide or Fn surface density while EC migration is maximal for an intermediate
concentration of RGD or Fn.91,93-95,98,107-109 Cell migration necessitates a turn-over of focal
adhesions at the front edge. At low ligand density, cells cannot form new focal adhesions at
the front efficiently, while at high ligand densities cells can not break focal adhesions. RGD
clustering, obtained by immobilization on a branched molecule, or increased peptide mobility
in the layer both enhance integrin binding and cell response with minimum RGD densities,
probably by favoring integrin clustering.96,110 EC polarization and directed migration is
achieved on either RGD or Fn gradients or collagen or Fn stripes; migration speed increases
by increasing gradient slope and decreasing stripes width, respectively.98,108,111,112 Finally, ECs
immobilized on RGD surfaces preserve their differentiated phenotype and form a non-
thrombogenic surface.99,113
Combination with synergic peptides and protein fragments. As peptides only represent
minimal binding sequences, they possess only a fraction of the activity of the entire protein
and cells adhering on peptide surfaces usually show a decreased response that is probably
related to non-optimized ligand conformation and/or absence of synergy sites reducing affinity
for receptors. A solution is to co-immobilize the minimal RGD adhesive peptide with its
synergy site PHSRN for increased adhesion. The PHSRN sequence found within the FnIII-9
domain of Fn does not support cell adhesion on its own but acts synergistically with RGD to
increase α5β1 binding, cell adhesion and spreading.90,96,97 A peptide containing both the RGD
cell binding and the PHSRN synergic sequences connected by a linker of appropriate length
promotes higher EC adhesion and spreading than mixed RGD and PHSRN.114 Equally,
combination of the RGD peptide with syndecan-binding sequences from the Fn C-terminal
heparin-binding domain II enhances EC proliferation and migration.106,115,116 EC response is
even increased if cells adhere on peptides distributed as clusters on the surface.115 Another
alternative is to immobilize a small Fn fragment containing the cell binding and synergy
sites.117,118 Likewise, the REDV sequence from the CS5 domain of Fn does not support EC
adhesion when incorporated in an artificial protein.119 However, incorporation of the entire
29
CS5 region restores EC adhesion, probably by allowing the REDV binding sequence to adopt
a recognizable conformation.120
Another important objective in EC-material interaction studies is to be able to control
EC proliferation and migration, which both require cell adhesion. To enhance cell migration,
the RGD adhesive sequence can be co-immobilized with a pro-migratory peptide or fragment.
Combination of RGD and YIGSR peptides leads to an increased EC migration.90 In addition,
Fn N-terminal heparin-binding domain I interacts with α5β1, helping with the formation of Fn
fibrils and promotes EC migration when immobilized.81 Hence, co-immobilized Fn cell-
binding domain and heparin-binding I fragment enhance EC migration.2
2.5.4. Growth factors immobilization
Cross-talk between integrins and growth factor receptors
Cell adhesion receptors such as integrins and growth factor receptors share a number
of important signaling molecules, and the collaborative or mutual activation of integrins and
growth factor receptors through their association results in signaling synergism and reciprocal
potentiation.121,122 Once integrin and growth factor receptor are both bound to their specific
ligands, they may cluster and act cooperatively. This results in the enhancement of growth
factor-dependent responses (cell proliferation, motility or survival) when cells are attached to
the appropriate matrix protein i.e., via the receptor-associated integrin. This type of
collaboration has been observed for the VEGF receptor (VEGFR)-2, which physically
associates with the αvβ3 integrin, inducing increased VEGFR-2 activation and EC
proliferation when anchored to Vn, an αvβ3 ligand.123 Collaborative activation of integrins
and growth factor receptors also results in increased expression and activation of integrins
associated to collaborative receptors. Soluble VEGF enhances αvβ3 and α5β1 activation and
subsequent EC adhesion and migration on Vn/osteopontin and Fn, respectively. Likewise,
soluble FGF-2 promotes EC adhesion and migration on Vn via αvβ3.72,124 In addition to
collaborative activation, integrin engagement by a specific matrix protein can trigger ligand-
independent activation of growth factor receptors and, on the other hand, growth factors can
induce unligated integrins to propagate signals.121,122 Synergism of integrin and growth factor
receptors signaling maybe relevant in dynamic, multistep processes such as tissue
30
development or regeneration, where cells undergo controlled migration and proliferation, and
may differentiate. Hence, co-immobilization of integrin adhesive ligands and growth factors
may provide a way to promote and achieve endothelialization of a biomaterial.
Surface-bound growth factors
Immobilized VEGF, a vascular specific growth factor, binds integrins αvβ3, α3β1 and
α9β1 which mediate EC adhesion and migration.125,126 In combination with Fn, covalently
immobilized VEGF enhances EC proliferation70 and immobilization of a Fn fragment
containing both α5β1 and VEGF binding domains promotes VEGF adsorption, increased EC
migration and proliferation.89 Immobilized FGF-2 and hepatocyte growth factor (HGF) also
promote strong and prolonged EC proliferation.127-129 However, surface-bound angiopoietins,
regulators of angiogenesis, mediate cell adhesion, in part via integrins, but only Ang-1 could
enhance EC spreading, focal adhesion formation and migration.130 A question remains
whereas to covalently/firmly link growth factors to the surface or not. Soluble growth factors
bind receptors on the apical side of the cell, they may form collaborative associations with
integrins and they are internalized, which immediately ceases the stimulated cell growth. On
the other hand, adsorbed and covalently bound growth factors are exposed to the basal side,
they can directly bind integrins and, owing to their increased local concentration, immobilized
growth factors may induce formation of integrin-growth factor receptor complexes that do not
form in the presence of soluble growth factors,126,128 thus eliciting different intracellular
signals. Moreover, firmly bound growth factors are probably not internalized, which explains
their sustained activity. In vivo, blood vessel formation is a multistep process both locally and
temporally controlled; interactions of cells with extracellular molecules, such as matrix
proteins and growth factors, are usually transient. Non-permanent immobilization of growth
factors may be required to achieve formation of a stable endothelium lining or mature and
functional blood vessels. It may be preferable to bind growth factors via natural interactions
such as VEGF-Fn,89 or to use recombinant growth factors with increased affinity for an ECM
protein such as collagen type I,129 or else, to chelate growth factors to a surface via surface-
bound metal ions.131
31
2.6. Conclusions and perspectives
Endothelial cells lining (ECs) the luminal surface of blood vessels share common
characteristics, provide a non-thrombogenic surface and actively respond to stimuli from their
microenvironment. However, they show great morphological and functional heterogeneity
among the vascular tree depending on the vessel type and vascular bed. This diversity results
from both local interactions with their microenvironment and epigenetic variations, mostly
between macro- and micro-vascular ECs.9,10 It should be mentioned that consequently, in
vitro, ECs originating from micro- or macro-vessels respond differently to a same stimulus.132
In addition, in vivo, ECs constantly interact with matrix proteins which either maintain the
endothelium integrity and functionality, or regulate, step by step, the formation of new blood
vessels, by modulating response to soluble signals, during development or in physiological
processes such as endometrium vascularization or wound healing.
Engineered surfaces have been designed to mimic in vivo EC environment in a
simplified way to promote EC attachment, survival and phenotype preservation (non-
thrombogenicity, angiogenic sprouting ability) for medical applications. In the past decades,
EC have been exposed to various surfaces, pre-coated or not with adhesive matrix proteins or
cell binding peptides, and results have shown that:
1. Material surface chemistry and the immobilization mode affect protein conformation,
orientation and anchorage strength that in turn influence receptor affinity and
selectivity as well as cell contractility and behavior;
2. Peptide sequence and conformation enable targeting a specific receptor;
3. Cell-surface interactions can only be controlled if the ligand(s) is(are) immobilized on
a low-fouling surface that prevents non-specific protein adsorption and thus “noise”
signaling;
4. As adhesive peptides represent only the minimal binding sequence isolated from native
proteins, co-immobilization with their synergic sequence or immobilization of a
complete protein fragment enhance cell adhesion;
5. Integrin receptors collaboratively associate with other receptors, either other integrins
or growth factor receptors, to coordinate and potentiate their signals; it should be
32
advantageous to combine an adhesive signal such as RGD or Fn with a pro-migratory
peptide/fragment or a growth factor to enhance EC migration and proliferation;
6. Directed cell migration can be achieved by physical or chemical contact guidance,
immobilization of a gradient of adhesive molecules or cell exposition to fluid flow;
7. Formation of capillary-like structures on a surface depends on the cell-substrate
adhesion strength, which relies on biochemical and mechanical signals.
Finally, the successful design of a biomaterial surface requires understanding and
control of substrate physico-chemical properties, of protein-surface interactions, and of cell
signaling in vivo and in vitro.
Endothelial cells retention on a surface correlate with strong binding to the surface and
can be increased by shear stress pre-conditioning.63 Cell engineering by molecular biology
techniques may be a way to artificially enhance EC adhesion to surfaces. For instance,
modified ECs that co-express an adhesive matrix protein and VEGF better resist shear stress
detachment.133 Likewise, biotinylated-ECs incubated with streptavidin and exposed to
surfaces presenting Fn and biotinylated-albumin, thus presenting a dual ligand system, show
an increased retention when exposed to fluid flow.134 In addition to designing surfaces,
molecular biology now enables researchers to engineer cells.
On the other hand, vascular in-growth, desired for implant integration or tissue
engineering matrix vascularization, requires cell proliferation and migration. These cell
processes should be tightly regulated, both spatially and temporally, to achieve the formation
of a functional and stable endothelium. Local control can be obtained by surface patterning
with matrix and growth factor cues, while temporal control may be accomplished by binding
molecules to the substrate via linkers cleaved in response to cell secretion of proteases, for
instance.5 Moreover, to acquire tissue specific vascular specialization or blood vessels
stability, it would be interesting to co-culture ECs with tissue cells or mural cells such as
pericytes or SMCs on patterned surfaces.8 Printing cell type specific domains would allow to
simultaneous study and control cell-surface and heterotypic cell-cell interactions.135,136
33
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41
Chapitre 3
Développement de puces d’un dérivé du dextran
pour identifier les propriétés physico-chimiques
impliquées dans l’adsorption de
protéines du sérum
3.1. Résumé
Afin de contrôler l’adsorption de protéines sur une surface, des revêtements
polymériques anti-adhérents tels que le poly(éthylène glycol) (PEG ou PEO) et les
polysaccharides sont utilisés. Leur capacité à résister à l’adsorption de protéines est liée à leur
structure et donc à leur mode d’immobilisation. Une technologie de puces de polymères a été
développée pour étudier les diverses structures de couches de carboxy-méthyl-dextran (CMD)
immobilisées selon différentes conditions. Les puces de CMD ont été analysées par
spectroscopie des photoélectrons-X (XPS) et par mesure de force réalisée par microscopie à
force atomique (AFM). L’adsorption de protéines de sérum a été étudiée directement sur les
puces de CMD par résonance des plasmons de surface (SPR). La caractérisation physico-
chimique a révélé que la densité des points de liaison influence le recouvrement de la surface
et la quantité de molécules adsorbées, et que la concentration de sel contrôle la structure du
polymère chargé, formant des couches écrasées ou étendues. Les expériences d’adsorption de
protéines ont indiqué que les couches de CMD résistantes sont denses et présentent des
chaînes flexibles et étendues. Cette étude souligne l’utilité des puces de polymères pour
étudier la diversité structurelle de couches minces et pour lier leur propriétés physico-
chimiques à leur capacité à résister à l’adsorption non-spécifique de protéines.
42
Development of dextran-derivative arrays to identify
physico-chemical properties involved in
biofouling from serum
Emmanuelle Monchaux, Patrick Vermette
3.2. Abstract
To control protein adsorption on surfaces, low-fouling polymer coatings like
poly(ethylene oxide) (PEG or PEO) and polysaccharides are used. Their ability to resist
protein adsorption is related to the layer structure, hence the immobilization mode. A polymer
array technology was developed to study structural diversity of carboxy-methyl-dextran
(CMD) layers, whose immobilization conditions were varied. CMD arrays were analyzed by
X-ray photoelectron spectroscopy (XPS) and by atomic force microscopy (AFM) colloidal
probe force measurements. Serum protein adsorption was studied directly on the CMD arrays
using surface plasmon resonance (SPR) microscopy. Physico-chemical characterization
revealed that pinning density regulates surface coverage and the amount of adsorbed
molecules, and that salt concentration influences the surface structure of the charged polymer,
forming extended or short layers. Protein adsorption experiments from serum showed that
repulsive CMD layers are dense, with extended flexible chains. The present study underlines
the usefulness of polymer arrays to study structural diversity of thin graft layers and to relate
their physico-chemical properties to their resistance towards non-specific protein adsorption.
Keywords: polymer arrays; polymer chips; dextran derivatives; carboxy-methyl-dextran;
surface characterization by AFM, SPR, and XPS; surface properties affecting protein
adsorption.
43
3.3. Introduction
Understanding and controlling the interactions between naturally-derived or synthetic
materials and biological fluids or cells is of increasing importance in a variety of fields. For
example, the interactions between biomolecules and solid surfaces are central to numerous
analytical systems, including immunodiagnostics, DNA and protein micro-arrays, biosensors
and microfluidic “lab-on-a-chip” devices. Performance of biomedical devices including
cardiovascular replacements (e.g., catheters, valves, stents), contact lenses and intra-ocular
lenses, shunts, membranes and drug delivery devices also depend on the control over
biomolecule-surface interactions. Non-specific adsorption can often lead to failure of a
medical device or to that of a biosensor exposed to complex biological environments, altering
their functions or inducing their rejection.1,2
Low-fouling polymers have been immobilized on solid surfaces to reduce these non-
specific interactions with the aim to enhance signal-to-noise ratios when a specific biological
response is investigated or intended. The most popular of these coatings is made of
poly(ethylene glycol) or poly(ethylene oxide) (PEG/PEO), a synthetic polymer which protein-
rejecting ability appears to result from its neutrality, high hydrophilicity and mobility.3,4
Another category of low-fouling layers is composed of polysaccharides like hyaluronic acid,
alginic acid and dextran.5-11 They are highly hydrated and flexible natural polymers. For these
properties, some polysaccharides show low level of non-specific protein adsorption and can be
as good as PEG layers.6 Moreover, the high concentration of hydroxyl groups along their
chain allows incorporation of active groups without significantly affecting their
hydrophilicity.
Mechanisms of polymer layer ability to limit protein adsorption can be explained in
two ways.12 Firstly, low-fouling properties of thin graft layers can be attributed to steric
repulsion effects,13 which in turn are related to layer thickness and surface coverage. Proteins
interacting with such a surface cause chain compression and desolvation. The resulting
entropic loss prevents proteins to adsorb on the surface.13,14 However, this theory does not
take into account chemical interactions between solvent water molecules and polymer chains.
Accordingly, a second explanation associates protein repellence with interactions between
interfacial water molecules and hydrophilic polymer moieties, through hydrogen bonding, for
44
instance.15 Conformation freedom of polymer chains may not be necessary for thin graft
layers to be low-fouling, but molecule orientation and strength of interaction with surrounding
water molecules and hence interfacial water layer stability are also believed to be involved in
the ability of a layer to reject proteins.12,15
Then, polymer graft layers ability to prevent non-specific interactions lies in layer
structure (density, thickness, conformation, interfacial properties, for instance), underlying the
importance to understand and control layer architecture. A polymer can adsorb to a surface in
trains or loops and tails,16 and the resulting conformation on the solid surfaces can be
controlled by varying the immobilization conditions, such as polymer conformation in
solution, surface affinity, and pinning density.
The aim of this study was to investigate the effect of immobilization conditions on the
physicochemical properties of carboxy-methyl-dextran (CMD) graft layers and on the bio-
fouling level. This study investigates the relationship between CMD graft layer structures and
the level of biofouling from serum. The polysaccharide studied here was an anionic derivative
of dextran i.e., CMD, with various carboxyl density on the chain. CMD is a highly flexible
polysaccharide, due to its α(1-6) linkages, present in an extended coil conformation in pure
water due to intra-molecular charge repulsion and present in a random coil form in electrolyte
solutions owing to the salt screening effect. High-molecular weight dextrans present very few
long branches, whose content increases with molecular weight.17 Hence, CMD conformation
depends on its molecular weight, its charge density and on solution salt concentration.
In this study, CMDs were immobilized on plasma-modified surfaces bearing amine
groups, via their carboxylic groups, using carbodiimide chemistry. To simultaneously study
parameters affecting CMD graft layer properties, polymer arrays were developed. These
polymer arrays can also be referred to as polymer chips and are based on DNA and protein
micro-array technology. To develop optimal protein repellent CMD graft layers, it is
important to identify the immobilization parameters that can lead to the best low-fouling CMD
layers and how such repulsion can be optimized. Therefore, polymer arrays were used to
allow the high-throughput study of a large numbers of CMD coatings on a same solid
substrate offering diversity in term of CMD layer physico-chemistry and limiting batch-to-
batch variations. This method was used to relate immobilization conditions with the resulting
CMD layer physico-chemical properties and with the level of protein adsorption onto these
45
different polymer coatings. X-ray photoelectron spectroscopy (XPS) was used to screen the
effect of CMD immobilization conditions on the surface chemical composition directly on the
CMD arrays. These XPS results were compared to those of surfaces fully covered (i.e., by
using no arrays) by CMD graft layers made using the same immobilization conditions. Atomic
force microscopy (AFM) colloidal probe force measurements were used to study the impact of
the immobilization conditions on the force profiles obtained between a silica sphere and the
CMD graft layers. Finally, surface plasmon resonance (SPR) microscopy was used to
investigate the homogeneity of the CMD spots across the arrays and also to investigate the
relative protein adsorption onto these different CMD spots directly on the arrays.
3.4. Experimental section
3.4.1. Materials
Borosilicate glass substrates (Assistent, Sondheim, Germany) were washed in
Sparkleen solution (Fisher Scientific Canada, Ottawa, ON, Canada), thoroughly rinsed with
Milli-Q water (Millipore Canada, Nepean, ON, Canada) with a resistivity of not less than 18.2
MΩ.cm and dried with 0.2-µm filtered air. For SPR analysis, LaSFN9 substrates were
purchased from ResTec (Framersheim, Germany). Before metal vaporization, the SPR
substrates were washed in Sparkleen solution and rinsed in Milli-Q water, in HPLC-grade
ethanol, and finally in Milli-Q water before being air-dried with 0.2-µm filtered air.
Vaporization of a 5-nm chromium layer and of a 48-nm gold layer was done by electron beam
evaporation in vacuum. Before utilization, substrates were cleaned under UV, immersed in
HPLC-grade ethanol, rinsed in Milli-Q water and dried with 0.2-µm filtered air. Metallized
samples could be reused for few experiments, then, an additional cleaning step was done in
piranha solution (3:1 ratio H2SO4 / H2O2) to remove any organic materials.
Carboxy-methyl-dextrans (CMDs) with different ratios of carboxyl groups to sugar
residues were prepared as follows. Five (5) g of dextran (70 and 500 kDa MW, cat. no. 17-
0280 and no. 17-0320, Amersham Biosciences, Uppsala Sweden) were dissolved in 25 ml of a
1M NaOH solution, then bromoacetic acid (cat. no. 3016, Lancaster Synthesis Inc., Pelham,
NH, USA) was added to yield a final concentration of either 0.125, 0.5 or 1M. The solution
46
was stirred overnight and dialyzed against Milli-Q water, 1M HCl, and finally Milli-Q water
for 24 hours each. Solutions were then lyophilized and stored at -20°C. The carboxylation
degree was determined by 1H NMR: carboxylation degrees obtained were 5%, 25%, and 50%
for 70-kDa dextran treated with 0.125M, 0.5M and 1M bromoacetic acid, respectively, and
5% and 40% of carboxylated residues for 500-kDa dextran treated with 0.125M and 1M
bromoacetic acid, respectively.
Phosphate buffered saline at pH 7.4 (PBS) used in these experiments was prepared in
Milli-Q water using NaCl (0.137M), KCl (0.003M), Na2HPO4 (0.008M), and KH2PO4
(0.002M). This 150mM solution was diluted in Milli-Q-water to make a 10mM solution.
3.4.2 Fabrication of CMD arrays
For polymer chips, the precisely delimited region of low-fouling polymers were made
using a non-contact dispensing robot (BioChip Arrayer, Perkin Elmer Life and Analytical
Sciences, Boston, MA) with arraying capabilities such as those used to manufacture DNA and
protein chips. Each piezo-tip is able to deliver approximately 350 pL per pulse at a maximum
rate of 500 pulses per second. The number of spots depends on the surface area of the solid
substrate and the surface area that is needed for the application.
To investigate structure diversity of carbohydrate macromolecules, CMD preparations
of different compositions (MW, %COOH, solution concentration) and immobilization
conditions (NaCl/monomer constituting the dextran and EDC+NHS/COOH ratios) were
printed on solid surfaces previously coated with an amine-bearing surface. As many
parameters were tested and as CMD arrays were characterized by various techniques, to
reduce analysis time, eight CMD conditions were randomly chosen by Design-Expert 6.0
software (StatEase Inc, Minneapolis, MN, USA) based on a quarter 2k factorial design model.
Conditions studied are presented in Table 3.1.
Clean borosilicate and SPR substrates were first surface-modified by plasma
polymerization of n-heptylamine (99.5% purity, catalog no. 126802, Sigma-Aldrich, Oakville,
Canada) in a custom-built plasma reactor.18 Briefly, the reactor chamber is made of a glass
cylinder (28-cm inside diameter, 50-cm height). Two disk-shaped horizontal copper electrodes
(14-cm diameter, 5-mm thickness) are maintained in the chamber by vertical copper rods (top
rod: 7-mm diameter with an adjustable vertical position; bottom rod: 2.5-cm diameter, 17-cm
47
length). Samples were placed onto the bottom electrode. Conditions of operation used were an
initial monomer pressure of 0.040 Torr, an ignition power of 80 W, an excitation frequency of
50 Hz, a deposition time of 45 sec, and a distance between the electrodes of 10cm. The
resulting n-heptylamine plasma polymer (HApp) layer exposes functional amine groups that
were used to graft CMD layers using carbodiimide chemistry.
TABLE 3.1: CMD immobilization conditions used to produce spots of CMD graft layers.
Conditions no. 1 no. 2 no. 3 no. 4 no. 5 no. 6 no. 7 no. 8
MW (kD) 70 500 70 500 70 500 70 500
%COOH 5 5 50 40 5 5 50 40
C (mg/mL) 1 1 1 1 3 3 3 3
NaCl:monomer 10 1 1 10 10 1 1 10
EDC:COOH 10 0.1 10 0.1 0.1 10 0.1 10
10M NaOH a + + + +
µI b (mM) 60 6 70 150 185 20 80 365
a + indicates when 10M NaOH was added to enhance the dissolution of CMD. b µI stands for solution final ionic strength.
NaCl was added to CMD solutions at varying ratios (Table 3.1); 10M NaOH was
added to enhance CMD dissolution in some systems (as indicated in Table 3.1). CMD
solutions were filtered (0.45-µm pore size filters) to eliminate CMD aggregates. Then,
solutions were activated with (1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC, catalog
no. E1769, Sigma-Aldrich) and N-hydroxysuccinimide (NHS, catalog no. H7377, Sigma-
Aldrich) for 10 minutes. EDC/NHS ratio was fixed to 1 for all CMD conditions. Solutions
were then dispensed with the BioChip Arrayer on HApp-coated surfaces. The piezo-tips were
set to dispense 50, 1000 or 1500 drops of solution per spot for AFM, XPS and SPR analyses,
respectively, as determined by the areas of analysis required by these techniques. As the CMD
grafting reaction is a wet-based chemistry, the reaction was allowed to proceed overnight in a
humidity saturated air chamber, preventing any evaporation or condensation. CMD arrays
were rinsed with 1M NaCl solution (2x12 hours), then with Milli-Q water (2x12 hours), dried
and stored under argon atmosphere until required. Because it is extremely difficult (almost
impossible) to see where the CMD spots have been deposited on the solid substrates, for XPS
48
and AFM analyses CMD arrays were made with exactly the same design as a pattern-food
colorant array (Fig. 3.1). Briefly, a transparent borosilicate substrate was glued to a previously
made food colorant array with epoxy glue, the assembly was coated by HApp and CMD spots
were deposited just above colorant spots.
Figure 3.1: Guiding system to localize individual CMD spots for XPS and AFM analyses.
In parallel, to compare the surface composition of the CMD spots with fully CMD-
covered surfaces, some substrates freshly covered by HApp layers were immersed in activated
CMD solutions for 2 hours, under agitation, to produce surfaces uniformly covered (i.e., no
arrays) by CMD graft layers. These surfaces were rinsed and preserved like polymer arrays.
3.4.3 Elemental composition of CMD spots by X-ray photoelectron spectroscopy (XPS)
XPS analysis of arrays was performed using an AXIS HS spectrometer (Kratos
Analytical Ltd., Manchester, GB) equipped with a monochromatic Al Kα source at a power of
120 W. The pressure in the main vacuum chamber during analysis was typically 2x10-8 torr.
The elemental composition of the analyzed surface areas was obtained from survey spectra
collected at a pass energy of 120 eV. High-resolution C 1s spectra were collected at 40 eV.
Atomic concentrations of each element were calculated using CasaXPS (Casa Software Ltd)
by determining the relevant integral peak areas, and applying the sensitivity factors supplied
by the instrument manufacturer; a Shirley background was used. To compare the high-
resolution C 1s peak positions, the spectra were shifted to ensure that the leading edges of the
fitted aliphatic CHx component were coincident. All spectral intensities were normalized to a
maximal intensity corresponding to the full height of the fitted aliphatic CHx (285.0 eV)
component peak. Survey scans and C 1s high-resolution spectra were collected on each spot
(at least three per condition).
Transparent substrate
Corner-guide
CMD spots
Coloured solution spots
Glue
49
3.4.4 CMD spot layer structure by AFM force measurements
Atomic Force Microscopy experiments were performed using a Digital Nanoscope IIIa
Bioscope (Veeco Instruments, Santa Barbara, CA, USA). The interaction forces between a
silica particle and immobilized hydrated CMD layers forming the arrays were measured using
the colloidal probe method developed by Ducker et al.19 Silica colloidal spheres (mean
diameter of 4.12 µm, Bangs Laboratories, Inc. Fishers, IN, USA) were attached to AFM Si3N4
cantilevers with a spring constant of 0.12 N/m via an epoxy adhesive (EP30MED, Master
Bond, Hackensack, NJ, USA). The spring constant of the modified cantilevers was determined
to be 0.1399 N/m using the resonance method proposed by Cleveland et al.20 Cantilever
deflection versus travel distance curves were converted to plots of force-versus-separation
distance using a home-made software. The linear region of the deflection curve was used to
define zero separation distance.
Force measurements between silica colloidal probes and the CMD graft layers on the
arrays were collected at room temperature in 10mM or 150mM PBS, allowing for
equilibration for at least 1 hour before measurement. Three different positions were studied
per spot, for each buffer, two spots per condition were analyzed, and at least ten force curves
were obtained for each position.
3.4.5 CMD spot homogeneity and CMD fouling from serum measured by surface
plasmon resonance (SPR) microscopy
The SPR instrument from Resonant Probes GmbH (Goslar, Germany) uses a
monochromatic laser light source at a wavelength of 633 nm. The beam was directed through
a glass prism to the gold-coated glass substrate optically coupled to the prism using immersion
oil (catalog no. 1812X, Cargille Laboratories Inc., NJ, USA). The reflected laser light
intensity (including the SPR signal) was collected by a two-dimensional CCD camera coupled
to a zoom lens as a function of the angle of incidence. The resulting data was processed by a
computer to obtain experimental plasmon angles. For each polymer spot, angular scan curves
were recorded from at least three sampling areas located centrally in each spot and at least two
areas from surrounding HApp layers.
For serum adsorption assays, SPR data were recorded in air before and after serum
adsorption. Following a first SPR analysis of the CMD arrays prior to serum incubation, CMD
50
arrays were rinsed in 150mM PBS and incubated in a 50% (v/v) fetal bovine serum (FBS,
catalog no. F1051, Sigma-Aldrich) solution made in 150mM PBS for 1 hour at room
temperature under agitation. Serum is a biological fluid composed of wide range of proteins
and molecules and was chosen to test general protein resistance of CMD layers. Substrates
were then rinsed with 150mM PBS, Milli-Q water, and dried with 0.2-µm filtered air before
the second SPR analysis. The shift in plasmon angle, θSPR, was measured following serum
adsorption. The experiment was repeated twice.
3.5. Results and discussion
3.5.1 Elemental composition of CMD spots by XPS
A high-precision robot designed to produce DNA micro-arrays was successfully
utilized to spot CMD on HApp-coated borosilicate surfaces. CMD arrays were rinsed, dried
and analyzed by XPS. Survey and high-resolution C 1s spectra were taken on CMD spots and
on HApp background surfaces between the CMD spots.
TABLE 3.2: Elemental composition of CMD arrays and fully covered control CMD surfaces on borosilicate glass determined by XPS analyses.
Atomic concentration % Atomic ratio
C1s O1s N1s O/C N/C
Fresh HApp 86.7 4.7 8.5 0.05 0.10
Array HApp 81.8 10.5 7.75 0.13 0.10
no. 1 78.4 14.7 7.0 0.19 0.09
no. 2 77.9 15.8 6.4 0.20 0.08
no. 3 76.7 14.4 8.9 0.19 0.12
no. 4 82.0 10.6 7.5 0.13 0.09
no. 5 82.4 10.9 6.8 0.13 0.08
no. 6 75.3 18.4 6.3 0.24 0.08
no. 7 81.2 10.5 8.3 0.13 0.10
no. 8 78.3 13.8 8.0 0.18 0.10
no. 1 surface 79.2 15.4 5.4 0.19 0.07
no. 7 surface 83.1 10.2 6.7 0.12 0.08
51
XPS analysis of freshly deposited HApp layers on borosilicate substrates indicated a
polymer rich in hydrocarbon- and nitrogen-containing species (Table 3.2). The elemental
composition obtained on CMD spots revealed a marked increase in oxygen content relative to
that observed on freshly deposited HApp layers, confirming the successful attachment of the
polysaccharide to the HApp surface. Figure 3.2a shows representative examples of high-
resolution C 1s spectra indicating the introduction of oxygen containing carbon species C-O
(286,5 eV), C=O (287,9 eV) and O-C=O (289,2 eV), also confirming the grafting of CMD for
all the immobilization conditions tested in this study.
Figure 3.2: a) High-resolution XPS C 1s spectra of a freshly deposited HApp layer and of CMD spots made under conditions no. 1 and no. 3. b) High-resolution XPS C 1s spectra of CMD layers produced under condition no. 1 on a fully covered surface and on an array.
XPS analyses revealed that HApp layer on the array has a higher oxygen concentration
than that of freshly deposited HApp layer bearing no CMD spot. This is due to spontaneous
oxidation of the HApp layers via residual radicals while exposed to air21 between the plasma
deposition, the making of the CMD arrays, and XPS analyses. Hence, CMD solutions were
immediately deposited on HApp layers after removal from the plasma reactor, because post-
plasma surface oxidation is a rapid process.
Finally, it was verified whether or not the surface compositions of the CMD spots were
similar to those of CMD graft layers immobilized on fully covered surfaces and produced
using the same immobilization conditions than those of CMD spots. Fully covered control
282 284 286 288 290 292
BE (eV)
Rela
tive inte
nsity
a
HApp
HApp + CMD #1
HApp + CMD #3
282 284 286 288 290 292
BE (eV)
Rela
tive inte
nsity
b
HApp + CMD #1 surface
HApp + CMD #1 array
52
CMD surfaces made according to the immobilization no. 1 and no. 7 were analyzed by XPS.
XPS analyses revealed higher oxygen concentration on the uniform CMD surfaces made using
condition no. 1 (i.e., using no array) than on the corresponding CMD spots of the arrays. But,
the O/C ratios were similar when comparing spots and fully covered surfaces for the two
conditions tested. High-resolution C 1s spectra of CMD layers immobilized using condition
no. 1 on arrays and on fully covered surfaces are presented in Figure 3.2b and show similar
profiles, with a lower O-C=O signal (at 289,2 eV) for the CMD spots. The difference may lie
in the absence of agitation during the immobilization reaction on arrays that could result in a
slightly less efficient immobilization.
Following the development of high-throughput technologies that include nucleic acid-
and protein-based micro-arrays, researchers have immobilized carbohydrates in an array
fashion for specific oligosaccharide-protein22 and polysaccharide-antibody interactions.23 The
latter study showed that dextrans preserve their immunological properties, and hence chain
conformation when adsorbed on solid substrates to create arrays. Here, O/C ratios obtained
suggest that polysaccharides can be immobilized on arrays and may be used as acceptable
models of the corresponding uniform surfaces.
Of particular interest is the effect of the EDC+NHS/COOH ratios on the O/C ratio as
measured by XPS analyses of the CMD spots. Comparison of the O/C ratios of CMD layers
immobilized using high EDC+NHS/COOH ratios (conditions no. 1, no. 3, no. 6, and no. 8) to
those of CMD layers grafted using small EDC+NHS/COOH ratios (conditions no. 4, no. 5,
and no. 7) indicates a lower grafting density for the latter, as less carboxyl groups were
activated in CMD solutions and available to be involved in covalent linkage. Accordingly,
CMD molecules immobilized on a plasma surface via a polyamine spacer layer were bound in
higher amount than on HApp, owing to a higher grafting density.7,24 CMD coatings produced
under conditions no. 2 and no. 6 presented higher oxygen amount than their counterparts.
These conditions involved the use of very low ionic strength (Table 3.1) and therefore the
quasi-absence of a screening effect, possibly resulting in an electrostatic attraction between
the negatively charged CMD molecules and the positively charged freshly deposited HApp
layers exposing partially protonated amine groups. The electrostatic attraction should lead to
an improved diffusion towards the surface and a reorientation of the polymer chains so that
maximum carboxyl groups are in contact with the surface.16 It can be hypothesized that it
53
resulted in an improved immobilization efficiency with maximum carboxyl groups involved in
covalent linkages. If electrolyte concentration is high enough for charge screening, there is no
electro-sorption during the immobilization and, we believe, carboxyl groups will be randomly
distributed in the layer, reducing immobilization efficiency.
However, XPS results should be carefully interpreted because CMD layers are
probably collapsed when analyzed in the XPS chamber. Therefore, it is very difficult to gather
structural information on the CMD spots based on XPS analyses.
3.5.2 AFM colloidal probe interaction forces with HApp layers
Figure 3.3: Representative colloidal probe force measurements between a silica colloidal probe and HApp layers on a polymer array in two PBS solutions (pH 7.4).
Figure 3.3 shows AFM surface force measurements performed between a silica
colloidal probe and HApp layers on the array, at two different electrolyte concentrations. At
10mM PBS, the force-versus-separation curves show a jump-to-contact probably due to van
der Waals forces acting between the silica colloidal probe and the HApp layers. A strong
attraction was seen on retraction. The oxidized plasma surface bears some amide groups,21
which are very weak bases, probably not protonated at pH 7.4. Hence, the surface-oxidized
HApp layer is almost neutral, explaining why there is no electrostatic attraction between the
negatively charged silica sphere and the oxidized HApp surface on approach. However, at
150mM, interaction forces experienced between the silica colloidal probe and the HApp
surfaces were repulsive over a small range, with a jump-to-contact before the hard-wall
Separation distance (nm)
0 5 10 15 20
F/R
(m
N/m
)
-0,04
-0,02
0,00
0,02
0,04
0,06
0,08
0,10
HApp-10mM
HApp-150mM
54
repulsion. Ionic species present in the buffer may have adsorbed on the surface during the
equilibration step before measurements.25
3.5.3 AFM colloidal probe interaction forces with CMD spots grafted using high
EDC+NHS/COOH ratios
Surface force measurements recorded during the approach between silica colloidal
probes and spots of CMD grafted on HApp layers using high EDC+NHS/COOH ratios are
shown in Figure 3.4. An exponential repulsion between the colloidal probe and the CMD layer
was observed at all electrolyte concentrations for conditions no. 1, no. 3, and no. 6 (Fig. 3.4a).
No adhesion was observed on retraction of the probe from the surface. The exponential
repulsion exhibited a marked dependence on electrolyte concentration, reducing in range and
distance with increasing ionic strength. However, the decay lengths of the force profiles
presented do not match the theoretically calculated value for purely electrostatic interactions
between surfaces immersed in salt solutions of the concentration used here, with higher
discrepancy for CMD layers made with low carboxylated CMDs. For instance, the interaction
forces in 10mM PBS displayed a decay length of approximately 13.2 nm for conditions no. 1
and no. 6 compared to a value of 2.5 nm for purely electrostatic forces.26 Thus, the
interactions measured for the CMD surfaces were not purely electrostatic in nature. While
repulsion has an electrostatic component, it can be concluded that the interaction force profiles
presented here are largely a result of compression of the covalently attached CMD layers by
the silica sphere. Polysaccharides are highly hydrated polymers with extended flexible chains,
thus compressible. Moreover, CMD molecules are weak polyelectrolytes bearing negative
charges repelling each other, intra- and inter-molecularly, which causes chain extension in
solution. When electrolyte concentration increases, the diffuse ionic atmosphere screens
charges, and the polymer chains collapse in a more random coil conformation.16,27 This
explains why chain extension and layer compressibility depend on ionic strength.
CMD layers made of highly carboxylated CMD produced under condition no. 3 appear
to have smaller chain extension than those made of low-carboxylated CMDs (i.e., conditions
no. 1 and no. 6) at all electrolyte concentrations tested here. Carboxyl density seems to
regulate the number of pinning points on the polymer chain, thus the size of loops and tails
protruding in solution. A possible explanation is that CMD spots made using conditions no. 1
55
and no. 6 have long, flexible loops and tails, while those made using condition no. 3 have
shorter ones, hence a shorter extension in solution, due to higher pinning density.
Figure 3.4: Representative colloidal probe force measurements between a silica colloidal probe and spots of CMD grafted on HApp surface using high EDC+NHS/COOH ratios in two PBS solutions. (a) silica-CMD layers no. 1 and no. 6 and silica-CMD layers no. 3; (b) silica-CMD layers no. 8.
Force-versus-separation curves of CMD spots made using condition no. 8 differ from
the ones previously presented: no exponential repulsion was observed between the colloidal
probe and the surface, but rather a small jump-to-contact possibly resulting from van der
Waals attraction was noted,26 followed by a relatively small compression dependent on
electrolyte concentration (Fig. 3.4b). Also, an adhesion was observed on retraction in both
buffers. This behavior suggests an absence of electrostatic repulsion between the silica sphere
and the CMD layer, despite a high carboxyl group density, and a reduced extension of
polysaccharide chains in solution. In Table 3.1, it can be found that the immobilization
condition no.8 also differs from other conditions in that the ionic strength during the grafting
reaction was really high. Reduction of intra- and inter-molecular electrostatic repulsion may
have lead to the adsorption of random coiled shrinked chains, packed close to the surface. The
presence of a high-density of carboxyl groups and the use of a high EDC+NHS/COOH ratio
induced the formation of many covalent bonds between the surface and CMD chains, forming
Separation distance (nm)
0 20 40 60
F/R
(m
N/m
)
0,0
0,2
0,4
0,6
0,8
1,0
1,2
1,4
1,6CMD layers #1,6 -10mM
CMD layers #1,6 -150mM
CMD layers #3 -10mM
CMD layers #3 -150mM
Separation distance (nm)
0 5 10 15 20
F/R
(m
N/m
)-0,05
0,00
0,05
0,10
0,15
0,20CMD layers #8 -10mM
CMD layers #8 -150mM
a b
56
many contacts with the surface (trains) and few short extensions in solution (loops and tails).
High density of train segments came with a high surface area occupied per polymer chain i.e.,
fewer molecules were immobilized per unit area, resulting in a relatively low O/C ratio
obtained by XPS. Moreover, as most carboxyl groups may have been engaged in covalent
linkage with the surface, exposed charged groups and electrostatic interactions with the silica
sphere were reduced. The layer was densely packed close to the surface, but had loops which
extension and conformation freedom were reduced even at 10mM, resulting in a high polymer
free energy. To lower the polymer free energy, there may have been displacement of Na+ ions
towards the layer or a reduced dissociation of carboxyl groups to decrease electrostatic
repulsion between polymer segments, and the need for chain extension. Both carboxyl groups
engagement in trains and charges neutralization in loops would prevent electrostatic
interactions with the silica sphere.
Hence, CMD graft layers that exerted repulsion towards silica spheres seemed able to
do so via electrostatic interactions through charge exposition, and via steric-entropic repulsion
owing to some chain extension and compressibility in solution.13
3.5.4 AFM colloidal probe interaction forces with CMD spots grafted using low
EDC+NHS/COOH ratios
Force measurements recorded between a silica colloidal probe and spots of CMD made
using condition no. 7 showed a small attractive jump, resulting from van der Waals attraction
in both electrolyte concentrations (Fig. 3.5a). Following the jump between the two surfaces,
hard wall repulsion was observed with a minimal compressibility that was dependent on
electrolyte concentration, being larger for the lower electrolyte concentration. Strong
attraction was observed on retraction. The lack of polymer extension into solution, shown by
surface force measurements along with the XPS data, suggests a minimal amount of surface-
immobilized CMD on the surface. Only few quite extended molecules seem randomly
attached on the surface. It can be hypothesized that the erratic curves collected during those
interaction force measurements are indicative of an irregular distribution of loops and tails
sizes.
On the contrary, interaction forces between CMD layers made using condition no. 4
with silica colloidal probes showed an exponential repulsion between the probe and the spots,
57
with small discontinuities indicative of attractive jumps. Adhesion on retraction was observed.
Decay length measurements suggest that the exponential repulsion is partly electrostatic in
nature, but also due to the presence of a compressible layer, despite immobilization of few
molecules on the surface (measured decay lengths of 5.2 nm compared to a theoretical value
of 2.5 nm). Immobilization condition no. 4 involved a high electrolyte concentration. As
previously mentioned for conditions no. 8, in such conditions, polysaccharide molecules
adsorb on the surface as shrinked molecules close to the surface, whereas for CMD layers
produced using condition no. 7, molecules adsorbed on the surface in a quite extended
conformation. As EDC+NHS/COOH ratio was low, only few carboxyl groups in contact with
the surface were engaged in covalent linkage, probably leading to the formation of no trains
but of short loops and tails. The resulting layers would seem not as diffuse as those made with
conditions #7. They were more regular.
Figure 3.5: Representative colloidal probe force measurements between a silica colloidal probe and spots of CMD grafted on HApp surface using low EDC+NHS/COOH ratios in two PBS solutions. (a) silica-CMD layers no. 4 and no. 7; (b) silica-CMD layers no. 2 and no. 5.
Figure 3.5b shows interaction forces experienced between a silica colloidal probe and
layers made of low-carboxylated CMD molecules (conditions no. 2 and no. 5) immobilized
with a small EDC+NHS/COOH ratio. As for layers made of low-carboxylated CMD
Separation distance (nm)
0 5 10 15 20
F/R
(m
N/m
)
0,0
0,1
0,2
0,3CMD layers #7 -10mM
CMD layers #7 -150mM
CMD layers #4 -10mM
CMD layers #4 -150mM
Separation distance (nm)
0 10 20 30 40
F/R
(m
N/m
)
0,0
0,1
0,2
0,3
0,4
0,5
0,6 CMD layers #2 -10mM
CMD layers #2 -150mM
CMD layers #5 -10mM
CMD layers #5 -150mM
a b
58
molecules (conditions no. 1 and no. 6), they exhibit exponential and long-range repulsion, but
almost not dependent on electrolyte concentration. Also, chains do not collapse at high
electrolyte concentration. They behave like neutral polysaccharides, which extension and
repulsion are not due to electrostatic interactions but rather to chains hydration. The low
EDC+NHS/COOH ratio combined with low carboxylation degree suggest a low grafting
density following the immobilization, leading to an uncharged diffuse layer, but extended
enough to shield the plasma surface. However, at high electrolyte concentration, analysis of
the interaction forces between a silica sphere and CMD layers made under condition no. 5
showed that there was adhesion upon retraction and small attraction was sometimes observed
on approach. This was not observed with layers made under condition no. 2. As previously
mentioned, immobilization of CMD layers made using condition no. 2 involved the use of a
very low salt concentration, allowing the adsorption of oriented and extended molecules,
while those made using condition no. 5 involved the use of quite high ionic strength, probably
forming a randomly packed layer, with a high surface area occupied per molecule, hence, a
reduced amount of adsorbed polysaccharide despite the high CMD solution concentration.
Both final layers presented extended chains but in a different fashion: layers produced using
condition no. 2 were quite dense and uniform in term of chain extension and spatial
distribution (at the molecular scale) while those produced using condition no. 5 were not
uniformly covering the surface, with random distribution of long loops and tails, and possibly
some pinholes.
Low pinning density (resulting from a low EDC+NHS/COOH ratio) leads to an
incomplete coverage of the surface with formation of mostly irregular, inhomogeneous layers,
similar to adsorbed layers.5,11 Yet, immobilization in absence of salt induced adsorption of
extended and oriented molecules on surface, improving the efficiency of immobilization; the
homogeneous and extended formed layers provide steric hindrance from the surface.
3.5.5 CMD spot homogeneity and CMD fouling from serum measured by SPR
microscopy
Surface plasmon resonance (SPR) spectroscopy is a widely used method for the
determination of optical thickness of thin layers. When p-polarized light is shined through the
glass slide and onto the gold surface at angles near the so-called “surface plasmon resonance”
59
condition, the optical reflectivity of the gold changes very sensitively with the presence of a
thin coating on the gold.28 SPR measures a certain combination of thickness and refractive
index and the shift of the plasmon angle is used to determine the optical thickness.
Reflected intensity is collected by a CCD camera through a zoom lens, and at plasmon
angle, spots appeared as uniform dark surfaces distinct from HApp background (Fig. 3.6a).
CMD spots seemed homogeneous in optical density suggesting that CMD immobilization
occurred uniformly across the whole spot area. As round spots were observed from an oblique
angle, they appeared as ellipses.
Figure 3.6: Protein adsorption on spots of CMD graft layers evaluated by SPR microscopy. a) Image of CMD layers produced in conditions no. 2 and no. 5 obtained at plasmon angle θSPR (scale bars: 500µm). b) SPR angle shifts (°) resulting from FBS protein adsorption on CMD arrays. Condition no. 9 involves the use of 70kDa CMD, 25% carboxylation degree, high EDC+NHS/COOH ratio, and a 2mg/ml CMD solution concentration.
a
b
0
0,1
0,2
0,3
0,4
0,5
0,6
0,7
#1 #2 #3 #4 #5 #6 #7 #8 #9
CMD conditions
SP
R a
ngle
shift
(°)
60
CMD arrays were then exposed to a serum solution and analysed by SPR microscopy.
Serum is a biological fluid composed of many proteins (of various sizes, charge densities and
shapes), and was then chosen to test general layer resistance to biofouling (including protein
adsorption). CMD arrays were exposed to harsh conditions (incubation with 50% FBS, for 1
hour under agitation) first to enhance differences between CMD layers fouling levels, as our
intention is to correlate layers physico-chemical characteristics with biofouling (including
protein repellence), and second, to be more selective, to screen for the best fouling resistant
surface(s). Figure 3.6b presents a summary of the serum adsorption on the different CMD
spots tested here.
CMD layers made using conditions no. 4 and no. 7 were subject to very important non-
specific adsorption from the FBS solution. XPS and AFM data suggest that those layers did
not completely cover the surface and were not extended in solution. Plasma could shine
through the CMD graft layers, therefore allowing for more serum component adsorption.
CMD layers produced using conditions no. 2 and no. 5 presented a thicker, more extended
layer, hindering the underlying plasma surface from the silica colloidal probe. Yet, serum
components could largely adsorb: the absence of charge in the layers indicates the
immobilization of a diffuse layer with an open structure, allowing proteins to penetrate.29,30
CMD layers made using conditions no. 2 and no. 5 have mobile protruding chains, but too
separated, therefore, proteins could penetrate without disturbing the chain conformation and
the hydration layer.
On the contrary, when proteins approach dense layers with protruding loops such as
CMD layers made using conditions no. 1, no. 3, and no. 6, flexible polymer chains are
compressed, causing an entropic loss.13 Moreover, water molecules tightly bound to the
polysaccharide units would be expulsed, causing an increase in polymer free energy; both
phenomena would lead to better protein repulsion. No difference are noted between layers
made using conditions no. 1 and no. 3, suggesting that the extension range and charge density
have no major effect on protein rejecting ability. However, layers made under condition no. 6
seem to be more protein resistant. XPS data revealed that a higher amount of polysaccharide
was immobilized, most likely creating a denser and more repulsive layer. As previously
mentioned, this may be due to better molecule reorganization at the surface or to higher CMD
solution concentration during immobilization (Table 3.1). Besides, the quite rigid
61
conformation of layers made using condition no. 8, along with a high train conformation
density and reduced loop extension and mobility, would prevent chain compressibility under
protein contact and steric repulsion. Hence, density and thickness are important parameters on
protein repulsion: proteins could penetrate into loose layers like those made under conditions
no. 2 and no. 5, while the absence of compressibility (thin layers) observed in CMD layers
produced under conditions no. 4, no. 7 and no. 8 prevented steric-entropic repulsion.14 As
CMD are highly hydrated polymers, their ability to reject serum components (including
proteins) may also result from the strength of chains interactions with surrounding water
molecules. At interface of dense, uniform CMD layers, chains are close enough to enable
formation of a stable water network. Water molecules tightly bound to polysaccharide units at
the interface may form a stable hydration layer, preventing proteins to adsorb on underneath
layer.15 However, diffuse or irregular layer would not allow formation of a stable water
network at the polymer interface, proteins would then interact with the CMD layer without
disturbing chain hydration layer.
To resume, CMD layers showing some resistance to serum adsorption were extended
enough and dense and were characterized by a high EDC+NHS/COOH ratio, a low salt
concentration and seemed more repellent when produced using a higher CMD solution
concentration. Hence, an optimized immobilization condition (called no. 9) was defined from
the results presented above, corresponding to 70kDa CMD MW, 25% carboxylic residues
with a high EDC+NHS/COOH ratio and a 2mg/ml CMD solution concentration. This
condition was used to produce CMD arrays to be tested towards serum adsorption (Fig. 3.6b)
and showed a low level of serum adsorption, well below CMD layers produced using
conditions no. 1, no. 3, and no. 6. The combination of a higher CMD solution concentration
with a high pinning density must have led to the immobilization of a denser and extended
layer. However, adsorption is still non negligible, but the adsorption assay was conducted in
harsh conditions. Most studies reporting protein adsorption on solid surfaces from FBS use
10% (v/v) FBS in a buffer solution with no agitation. Agitation may have facilitated chains
dissociation and structure opening and disrupted the hydration layer. Nonetheless, this
experiment allowed to determine immobilization parameters involved in layers ability to
adsorb or reject proteins.
62
3.6. Conclusions
Carboxy-methyl-dextran (CMD) array technology was successfully developed and
validated to simultaneously study covalent immobilization modes of CMD layers and their
resulting CMD layer structures to relate them to the resistance towards serum adsorption.
XPS analysis was used to characterize the surface elemental composition, while AFM
force measurements were used to study surface roughness and layer structure and density.
XPS analyses revealed that polymer layers immobilized on arrays presented the same
characteristics as layers grafted on uniform surfaces. Hence, polymer chips may be useful in
surface sciences to study polymer structural diversity and relate their properties to their low-
fouling behaviour, in a high-throughput manner.
Screening of the CMD immobilization conditions revealed the importance of grafting
reaction conditions on layer properties. Electrolyte concentration during immobilization
influences polymer conformation upon initial adsorption, while EDC+NHS/COOH ratio and
carboxylation degree regulate final layer density, segments distribution in trains, loops and
tails and hence the layer thickness (Fig. 3.7 shows a résumé). Our results suggest that a lower
ionic strength induces adsorption of an extended and ordered layer, resulting in diffuse or
dense mobile layer depending on pinning density. The use of higher salt concentrations during
immobilization likely initially leads to shrinked, randomly adsorbed molecules on surface,
forming erratic layers, with random distribution of loop sizes at low pinning densities, or the
presence of many trains and reduced flexible loops in case of a high pinning density. The
present XPS, AFM, and SPR analyses revealed no significant effect of the CMD molecular
weight on the layer properties and on the level of fouling by serum.
Finally, serum adsorption experiments showed that diffuse, irregular and/or
constrained layers were not resistant to serum adsorption. Dense, uniformly extended and
highly hydrated layers have better protein repellence: minimum surface density and layer
thickness and presence of a tightly bound layer of water at the polymer interface may be
responsible for protein rejection.
63
Figure 3.7: Schematic picture (not to scale) of immobilized CMD molecules. Final layer conformation can be controlled by electrolyte concentration and coupling agent ratio (EDC+NHS/COOH) during immobilization. CMD graft layers structures are hypothesized based on the present XPS and AFM results.
3.7. Acknowledgement
This work was supported by the Canadian Foundation for Innovation through an On-
going New Opportunities Fund (project no. 7918), by an NSERC Discovery grant, and by the
Université de Sherbrooke.
3.8. References
(1) Ratner, B. D. J. Control Release 2002, 78, 211-218.
(2) Wisniewski, N.; Reichert, M. Colloids Surf. B Biointerfaces 2000, 18, 197-219.
(3) Vermette, P.; Meagher, L. Colloids and Surfaces B-Biointerfaces 2003, 28, 153-198.
(4) Martin, Y.; Vermette, P. Macromolecules 2006, 39, 8083-8091.
(5) Osterberg, E.; Bergstrom, K.; Holmberg, K.; Riggs, J. A.; Van Alstine, J. M.; Schuman, T. P.; Burns, N. L.; Milton, H. J. Colloids and Surfaces A: Phisicochemical
and Engineering Aspects 1993, 77, 159-169.
(6) Osterberg, E.; Bergstrom, K.; Holmberg, K.; Schuman, T. P.; Riggs, J. A.; Burns, N. L.; Van Alstine, J. M.; Harris, J. M. J. Biomed. Mater. Res. 1995, 29, 741-747.
(7) Griesser, H. J.; Hartley, P. G.; McArthur, S. L.; McLean, K. M.; Meagher, L.; Thissen, H. Smart Materials & Structures 2002, 11, 652-661.
(8) Holland, N. B.; Qiu, Y.; Ruegsegger, M.; Marchant, R. E. Nature 1998, 392, 799-801.
high EDC
low EDC
low salt concentration
high salt concentration high EDC
low EDC
64
(9) Gupta, A. S.; Wang, S.; Link, E.; Anderson, E. H.; Hofmann, C.; Lewandowski, J.; Kottke-Marchant, K.; Marchant, R. E. Biomaterials 2006, 27, 3084-3095.
(10) Morra, M.; Cassinelli, C. Langmuir 1999, 15, 4658-4663.
(11) Morra, M.; Cassinelli, C.; Pavesio, A.; Renier, D. J Colloid Interface Sci 2003, 259, 236-243.
(12) Morra, M. J. Biomater. Sci. Polym. Ed 2000, 11, 547-569.
(13) Jeon, S. I.; Lee, J. H.; Andrade, J. D.; DeGennes, P. G. J Colloid Interface Sci 1991, 142, 149-158.
(14) Halperin, A. Langmuir 1999, 15, 2525-2533.
(15) Harde, P.; Grunze, M.; Dahint, R.; Whitesides, G. M.; Laibinis, P. E. J. Phys. Chem. B 1998, 102, 426-436.
(16) Fleer, G. J.; Cohen Stuart, M. A.; Scheutjens, J. M. H. M.; Cosgrove, T.; Vincent, B. Polymers at Interfaces; Chapman & Hall: 1993.
(17) Kuge, T. Carbohydrate Research 1987, 160, 205-214.
(18) Martin, Y.; Boutin, D.; Vermette, P. Thin Solid Films, in press.
(19) Ducker, W. A.; Senden, T. J.; Pashley, R. M. Nature 1991, 353, 239-241.
(20) Cleveland, J. P.; Manne, S.; Bocek, D.; Hansma, P. K. Rev Sci Instrum 1993, 64, 403-405.
(21) Gengenbach, T. R.; Chatelier, R. C.; Griesser, H. J. Surface Interface Anal 1996, 24, 271-281.
(22) Houseman, B. T.; Mrksich, M. Chem. Biol. 2002, 9, 443-454.
(23) Wang, D.; Liu, S.; Trummer, B. J.; Deng, C.; Wang, A. Nat. Biotechnol. 2002, 20, 275-281.
(24) McLean, K. M.; Johnson, G.; Chatelier, R. C.; Beumer, G. J.; Steele, J. G.; Griesser, H. J. Colloids Surf. B Biointerfaces 2000, 18, 221-234.
(25) Meagher, L.; Pashley, R. M. Langmuir 1995, 11, 4019-4024.
(26) Israelachvili, J. Intermolecular & Surface Forces; 2nd Ed.; London, 1991.
(27) Xu, F.; Persson, B.; Lofas, S.; Knoll, W. Langmuir 2006, 22, 3352-3357.
(28) Knoll, W. Annu. Rev. Phys. Chem. 1998, 49, 569-638.
65
(29) Piehler, J.; Brecht, A.; Gauglitz, G.; Maul, C.; Grabley, S.; Zerlin, M. Biosens.
Bioelectron. 1997, 12, 531-538.
(30) McArthur, S. L.; McLean, K. M.; Kingshott, P.; St John, H. A. W.; Chatelier, R. C.; Griesser, H. J. Colloids and Surfaces B-Biointerfaces 2000, 17, 37-48.
66
Chapitre 4
Étude des mécanismes de résistance à l’adhésion
cellulaire par utilisation de puces
d’un dérivé du dextran
4.1. Résumé
Afin d’étudier les interactions à l’interface cellule-matériau, des revêtements
polymériques anti-adhérents prévenant toute interaction non-spécifique sont utilisés. Des
biosignaux peuvent ensuite être greffés sur ces surfaces anti-adhérentes pour induire des
réponses biologiques spécifiques. Une technologie de puce de polymères a été développée
pour étudier la diversité structurelle de couches de carboxy-méthyl-dextran (CMD) en relation
avec leur capacité à résister à l’adhésion cellulaire. Des puces de CMD ont été exposées à des
fibroblastes pour analyser la résistance des couches en fonction des conditions
d’immobilisation utilisées pour les greffer. Le cytosquelette des cellules, la déposition et la
réorganisation de la matrice de fibronectine ont été marqués lors de tests d’adhésion à court
terme (4 et 12 heures) et à long terme (3 jours de confluence). Les résultats suggèrent que les
couches de CMD résistantes à l’adhésion cellulaire sont denses, flexibles et présentent une
surface hydratée régulière, sans défaut. Les couches de CMD résistantes à l’adhésion
cellulaire préviennent la déposition et l’assemblage de la matrice, affectant les adhésions
cellule-substrat et l’organisation du cytosquelette. Enfin, une couche de CMD optimisée est
déterminée et est aussi résistante qu’une couche de poly(éthylène glycol) (PEG).
67
Study of cell adhesion resistance mechanisms using
arrays of dextran-derivative layers
Emmanuelle Monchaux, Patrick Vermette
4.2. Abstract
To control interactions at the cell-material interface, low-fouling polymer coatings are
used to prevent non-specific interactions. Subsequent biosignals may be grafted on these low-
fouling layers to induce specific biological responses. A polymer array technology was
developed to study structural diversity of carboxy-methyl-dextran (CMD) grafted layers in
connection with their cell repulsion ability. Arrays of CMD layers were exposed to fibroblasts
to screen for cell resistant layers according to the immobilization conditions used to produce
these surfaces. Cell cytoskeleton and fibronectin matrix deposition and reorganization were
labeled in short-term (4 and 12 hours) and long-term (3 days of cell confluence) cell adhesion
assays. Results suggest that CMD layers that were resistant to cell adhesion were dense,
flexible, and presented a regular (i.e., defect-free) hydrated surface. Cell-resistant CMD layers
prevented cell matrix deposition and assembly, affecting cell-substrate adhesion and
cytoskeletal organization. Finally, an optimized CMD layer was chosen and proved to be as
resistant as a poly(ethylene glycol) (PEG) layer.
Keywords: cell responses to surfaces; polymer chips; dextran derivatives; carboxy-methyl-
dextran (CMD); surface properties affecting cell attachment and cell responses.
Funded by: On-going New Opportunities Fund, Canadian Foundation for Innovation, Grant
Number: 7918; Discovery grant, NSERC, Grant Number: 250296; Université de Sherbrooke
68
4.3. Introduction
Control of cell-material interactions is of great importance in biomaterials science and
tissue engineering as these interactions dictate success or failure of a medical device
implantation in an organism.1 To improve biomaterial implant performance, one needs to
control cellular interactions at the tissue-biomaterial interface using well defined surfaces. In
this way, the purpose of bioactive surfaces is to promote specific cell and tissue responses,
including selective cell adhesion. Controlled cell interactions could be achieved firstly, by
minimizing non-specific protein adsorption and cell adhesion and secondly, by chemically
grafting specific cell ligands on cell-resistant biomaterial surfaces.
Among surfaces and coatings lowering non-specific interactions, the most popular are
those made of poly(ethylene oxide) (PEO) and poly(ethylene glycol) (PEG). 2-4 Another
category of potential low-fouling molecules includes polysaccharides such as alginic acid,5
hyaluronic acid,6 dextran and its derivatives,7-12 to name a few. These molecules are highly
hydrated and flexible natural polymers, and the high concentration of hydroxyl groups along
their backbone allows incorporation of active groups, without significantly affecting their
hydrophilicity, and subsequent immobilization of bioactive molecules. Hence, polysaccharide-
covered surfaces can be used, to delay non-specific adsorption and, also, to enhance specific
responses from the host organism if activated with bioactive molecules.
Polymer layers ability to resist protein adsorption and subsequent cell colonization in
an organism results from interactive forces between the surface and molecules from biological
fluids. These interactions may arise from various origins, such as electrostatic, van der Waals
or hydration forces and steric-entropic effects.13 They are all dependent on the composition
and structure of the surface coating,2,13-15 which in turn depend on layer immobilization
conditions.
In a previous study, arrays of thin graft layers made of a dextran derivative i.e.,
carboxy-methyl-dextran (CMD), were developed to examine how immobilization conditions
affect CMD layers structure and protein repellence.12 Solution ionic strength during
immobilization influenced polymer conformation upon initial adsorption, while coupling
agent concentration and carboxylation degree regulated final layer density and thickness.
69
Moreover, protein adsorption experiments showed that minimum surface density and layer
thickness were associated with a better protein rejection.
To further investigate how immobilization conditions used to produce CMD graft
layers influence the cell repulsion ability of these layers, and how such repulsion can be
optimized, CMD arrays were exposed to cells. Arrays of acrylate-based polymers were
previously tested for their effects on human embryonic stem cell adhesion, growth and
differentation.16 In the present study, low-fouling polymer arrays were tested for their ability
to reject cell adhesion. Short-term and long-term cell culture assays were conducted with
human fibroblasts i.e., arrays were exposed to non-confluent and confluent cells. CMD graft
layers cell resistance and cell behavior on these coatings were studied and related to
immobilization conditions and layer structure. Finally, the cell resistance of an optimized
CMD layer and that of a PEG layer were compared.
4.4. Materials and Methods
4.4.1 Carboxy-methyl-dextran (CMD) synthesis
Carboxy-methyl-dextrans (CMDs) with different ratios of carboxyl groups to sugar
residues were prepared as follows. Five grams of dextran (70 and 500 kDa MW, Amersham
Biosciences, Uppsala, Sweden, cat. no. 17-0280-01 and no. 17-0320-01, respectively) were
dissolved in 25ml of a 1M NaOH solution. Then, bromoacetic acid (Lancaster Synthesis Inc.,
Pelham NH, cat. no. 3016) was added to a final concentration of either 0.125, 0.5 or 1M. The
solution was stirred overnight and dialyzed against Milli-Q water, 1M HCl, and finally Milli-
Q water for 24 hours each. Solutions were then lyophilized and stored at -20°C. The
carboxylation degree was determined by 1H NMR: carboxylation degrees were 5%, 25%, and
50% for 70-kDa dextran treated with 0.125M, 0.5M and 1M bromoacetic acid, respectively,
and 5% and 40% of carboxylated residues for 500-kDa dextran treated with 0.125M and 1M
bromoacetic acid, respectively.
4.4.2 Fabrication of arrays of CMD graft layers
To investigate cell resistance of CMD graft layers, CMD preparations of different
compositions (MW, %COOH, solution concentration) with the use of different immobilization
70
conditions (NaCl/sugar unit and EDC+NHS/COOH ratios) were printed on surfaces
previously covered with an amine-bearing surface. Conditions studied are presented in Table
4.1.
Details of the fabrication procedures of these arrays of CMD graft layers have been
reported elsewhere.12 Briefly, CMD arrays were made on borosilicate glass (Chemglass,
Vineland, NJ) washed in 35% nitric acid overnight, rinsed with Milli-Q water, and dried with
0.2-µm filtered air. Borosilicate glass substrates were surface-modified by plasma
polymerization of n-heptylamine (99.5% purity, Sigma-Aldrich, Oakville, Canada, cat. no.
126802) in a custom-built plasma reactor.17 Conditions of plasma polymerization were an
initial monomer pressure of 0.040 Torr, an ignition power of 80 W, an excitation frequency of
50 Hz, a deposition time of 45 sec, and a distance between the electrodes of 10 cm. The
resulting n-heptylamine plasma polymer (HApp) layers expose functional amine groups that
were used to graft CMD layers using carbodiimide chemistry.17
TABLE 4.1: CMD immobilization conditions used to produce arrays of CMD graft layers.
Conditions no. 1 no. 2 no. 3 no. 4 no. 5 no. 6 no. 7 no. 8 no. 9
MW (kDa) 70 500 70 500 70 500 70 500 70
%COOH 5 5 50 40 5 5 50 40 25
Concentration (mg/ml) 1 1 1 1 3 3 3 3 2
NaCl:sugar unit 10 1 1 10 10 1 1 10 5
EDC+NHS:COOH 10 0.1 10 0.1 0.1 10 0.1 10 5
10M NaOH a a a a
µI (mM) 60 6 70 150 185 20 80 365 60
µI stands for solution final ionic strength. a indicates when 10M NaOH was added to help dissolution of CMD solutions (pH was between 5.5 and 6).
NaCl was added to CMD solutions at varying ratios. In some samples, 10M NaOH was
added to help dissolution (see Table 4.1 for conditions). CMD solutions were first filtered
with 0.45-µm pore size membrane to limit CMD aggregates and then activated with (1-ethyl-
3-(3-dimethylaminopropyl)carbodiimide (EDC, Sigma-Aldrich, cat. no. E1769) and N-
hydroxysuccinimide (NHS, Sigma-Aldrich, cat. no. H7377) for 10 minutes. EDC/NHS ratio
was fixed to 1 for all CMD immobilization conditions. It should be noted that upon addition of
EDC/NHS to CMD solutions (even those in which NaOH was added), pH reaches values of
71
5.5-6.0. Solutions were then dispensed in duplicate with a non-contact robot designed to
produce arrays (BioChip Arrayer, PerkinElmer) on HApp-coated surfaces. The piezo-tips
were set to dispense 50 drops of solution per spot with 350 pL of sample per drop, resulting in
ca. 600-µm diameter spots. The reaction was allowed to proceed overnight in a chamber
saturated with humidity to maintain the substrates at dew point. CMD arrays were rinsed with
1M NaCl solution (2x12 hours), then with Milli-Q water (2x12 hours). They were autoclaved
and stored in sterile Milli-Q water at 4°C until needed.
As a comparison, a PEG solution at cloud-point conditions was also printed to produce
PEG graft layers with excellent low-fouling properties. Cloud point conditions induce
molecules packing, forming a dense layer.4 Boc-NH2-poly(ethylene glycol)-NHS of 3400-Da
MW (Nektar Therapeutics, San Carlos, CA, USA, cat. no. 4M530F02) was attached on the
HApp layers under cloud point conditions, using a 20% Na2SO4 solution at room temperature
to yield a final PEG concentration of 1mg/ml.4
4.4.3 Testing arrays of CMD graft layers towards cell responses
Human foreskin fibroblasts were cultured in sterile Dulbecco’s modified eagle’s
medium (Gibco, Grand island, NY, cat. no. 12100) with 10% fetal bovine serum (Sigma, cat.
no. F1051), 100 units/ml penicillin and 100 µg/ml streptomycin (Gibco, cat. no. 15140). Cells
were incubated at 37°C in a 5% CO2 incubator. Fibroblasts were chosen to perform cell
adhesion assays as these cells are known to easily adhere on most solid substrates.
Short-term cell adhesion assay
Cells were trypsinised (trypsine-EDTA, Gibco, cat. no. 25200), suspended in complete
culture medium and seeded at 15,000 cells/cm² on autoclaved CMD arrays. Autoclaving has
been shown to have no significant effect on the surface properties of CMD graft layers
(unpublished results). Cell incubation was stopped after either 4 or 12 h: for each incubation
time, two arrays were fixed with 3.7% formaldehyde during 15 min for further labeling.
72
Long-term cell adhesion assay
Cells were seeded at 10,000 cells/cm² on autoclaved arrays made of CMD graft layers.
Cells were observed every 24h and once 90% confluence was reached (72h, day 3),
experiments were allowed to proceed for another two days before cells were fixed with 3.7%
formaldehyde. Images of each spot were taken using phase contrast microscopy (Nikon
Eclipse TE2000-S) at 100x magnification during the 3 days of confluence (days 3, 4 and 5).
Spot area uncovered by cells was measured using SigmaScan Pro software (SPSS, Chicago,
IL). The area was outlined manually using the software.
Cell labeling
All antibodies used were purchased from Sigma. Fixed cells were rinsed with PBS and
permeabilized with 0.5% Triton X-100 for 5 min and finally blocked either with 2% bovine
serum albumin (BSA, cat. no. A7906) for vinculin and actin labelling or with 5% goat serum
(cat. no. G9023) for fibronectin labeling. For actin filament labeling, samples were incubated
with phalloidin-TRITC (1:400 dilution, cat. no. P1951) for 1h at room temperature. For focal
adhesion and secreted fibronectin staining, samples were incubated for 1h at room temperature
with the primary antibody: mouse anti-vinculin (1:25, cat. no. V4505) and mouse anti-human
fibronectin (1:400, cat. no. F0916), respectively, Then, samples were labeled with a FITC-
goat-anti-mouse secondary antibody (1:250, cat. no. F5262). Samples were rinsed and
mounted on glass slides and observed with an epifluorescence microscope (Nikon Eclipse).
4.5. Results
In the present study, previously characterized CMD arrays12 were exposed to cells: i)
first, to study parameters influencing resistance of CMD layers towards cell attachment, with
the aim to identify optimal immobilization conditions to produce CMD layers with excellent
cell-resistance properties; ii) second, to examine cell adhesive behavior when exposed to
CMD surfaces with various cell-resistance properties.
73
4.5.1 Initial cell interaction with CMD graft layers: short-term assays
Arrays of CMD graft layers were first seeded with cells for a short-time interval to
investigate early responses of cells exposed to various CMD layers. In vitro, initial cell
attachment on a substrate usually results in cell spreading, formation of focal adhesions and
cytoskeletal reorganization.18-20 Focal adhesions are peripheral cellular structures in which
cells interact with the extracellular matrix (ECM) through integrin receptors. One of the most
prominent structural component of focal adhesions is vinculin.21 Formation of focal adhesion
sites and actin cytoskeletal organization into stress fibers are interdependent and regulate each
other.20,21
Cell morphology and secreted fibronectin matrix at 4 h
Arrays of CMD layers were seeded with a high density of fibroblasts (i.e., 15,000
cells/cm²), and cells were fixed at 4 and 12 h. Four hours following cell seeding, fibroblasts
outside CMD spots i.e., those attached on the HApp layer, appeared well spread with filopodia
extensions. Few isolated ones looked small and rounded. The very few cells present on CMD
surfaces were small and rounded (Fig. 4.1a,b).
Figure 4.1: Optical microscope images of fibroblasts seeded on CMD arrays following 4h (a,b) and 12h (c,d) cell seeding on CMD spots made using condition no. 8 (a,c) and condition no. 3 (b,d).
4h
12h
200µm
CMD-Conditions #8 CMD-Conditions #3
a b
c d
74
Phalloidin actin-labeling revealed that, on HApp layers, cells had a quite organized
actin cytoskeleton with the presence of stress fibers associated with focal adhesions that were
characteristic of spread cells anchored to the surface (Fig. 4.2a,b). Actin fibers of cells on the
extreme periphery of CMD spots already tended to align along the spot curvature. However,
the few small cells present on cell-resistant CMD graft layers appeared asymmetrical and
showed no central stress fibers but most had bundles and a network of actin filaments at
periphery (Fig. 4.2d), usually associated with lamellipodia formation for cell spreading.19
Vinculin staining of focal adhesions was not visible on these cell-resistant CMD layers (Fig.
4.2e).
Figure 4.2: Initial cell behavior on CMD arrays. Cells were fixed 4h following cell seeding. Cells on the HApp layer (a,b,c) and on a CMD spot (d,e,f: condition no.1 and g,h,i: condition no. 7).
CMD spot edge
HApp layer 50µm
a b c
d f e
g h i
Actin Vinculin Fibronectin
CMD #1
CMD #7
HApp
75
As cytoskeletal architecture and ECM organization are dynamically linked through
integrin signaling, human fibronectin deposition and reorganization were evaluated at this
early stage of cell interaction with CMD surfaces. Fibronectin forms fibrils when activated by
integrins present in fibrillar adhesions and when stress fibers exert tension on fibronectin
matrix through focal adhesions.21,22 Fibronectin fibrils formation is essential as it affects the
composition, organization and stability of the ECM and also the stability of cell-ECM fibrillar
adhesion sites.23,24 On HApp layers, the presence of central fibronectin fibrils indicated that
fibroblasts had the ability to deposit and remodel the fibronectin matrix (Fig. 4.2c). For cells
present on the CMD graft layers, no fibronectin structure, neither granular nor fibrillar, was
distinguished (Fig. 4.2f). Early cell interaction with CMD films induced limited spreading,
rather cell rounding, associated with an absence of stress fibers and focal adhesions and CMD
surfaces also altered cells ability to deposit fibronectin matrix.
Nonetheless, on some CMD graft layers and at the edge of some CMD spots, cells
found in group exhibited thin actin stress fibers and weak focal adhesions (Fig. 4.2g,h). They
could remodel fibronectin matrix (Fig. 4.2i). As cell-surface and cell-cell adhesion points both
interact with the actin cytoskeleton, the formation of cell-cell adhesions also triggers
reorganization of the actin cytoskeleton.25,26 Hence cell-cell adhesions modulate cell spreading
and cell-surface inside-out signaling through integrins, favoring cell integrity and cell
survival.
Cell morphology and fibronectin matrix deposition at 12 h
Figure 4.1c,d shows that 12h following cell seeding, cells on HApp surfaces were
more spread and emitted less filopodia than after 4h (Fig. 4.1a,b). The presence of stress fibers
associated with long focal adhesions indicates that cells strongly adhered to the HApp surfaces
(Fig. 4.3a,b). More fibronectin fibrils were observed at 12h than at 4h, suggesting advanced
fibronectin matrix assembly. Contrary to cell responses observed on CMD arrays at 4h, three
types of CMD spots could be distinguished at 12h. First, on CMD spots showing good
resistance towards cell attachment i.e., CMD layers made under condition nos. 1, 3, and 9,
very few or no rounded cells were observed (Fig. 4.1d). Second, on less cell-resistant CMD
spots i.e., those made under condition nos. 2, 6, and 8, some cells were present, rounded or
spread (Fig. 4.1c). These CMD graft layers can be referred to as semi-resistant. The few cells
76
found on cell-resistant and semi-resistant CMD spots exhibited disrupted networks of actin
cytoskeleton and vinculin, while fibronectin was not detected under the few cells (Fig. 4.3d-f).
Third, on CMD layers made using condition nos. 4, 5, and 7, cells appeared quite spread, with
stress fibers associated with focal adhesions (Fig. 4.3g,h). Fibronectin labeling revealed that
isolated cells could deposit fibronectin in a granular structure, whereas grouped cells exposed
some fibrils (Fig. 4.3i). Twelve hours following cell seeding, fibroblasts were able to deposit
proteins on those CMD layers and to establish cell-surface contacts.
Figure 4.3: Initial cell behavior on CMD arrays. Cells were fixed 12h following cell seeding. Cells on the HApp layer (a-c), on cell-resistant CMD spots (d-f, condition no. 9), and on CMD layers made using condition no. 4 (g-i).
Actin Fibronectin Vinculin
50µm
a b
d e f
g h i
c
CMD #4
CMD #9
HApp
77
4.5.2 CMD resistance following 3 days of cell confluence
CMD spot size evolution
To further test CMD graft layers resistance towards cell attachment, arrays of CMD
surfaces were exposed to confluent cells for three days. Preceding results showed that grouped
cells having established cell-cell adhesions seemed less dependent on cell-surface signaling.
Thus, exposing CMD graft layers to confluent cells constitute harder conditions to study cell
resistance of thin films.27,28 CMD arrays were exposed to fibroblasts, at high seeding density
(i.e., 10,000 cells/cm²). Once 90% confluence was reached on day 3, another 48h was elapsed
before cells were fixed. The size of each spot (up to eight spots per condition) was measured
on days 3, 4 and 5 and compared to the size measured on day 3 to quantify surface coverage
by cells (Fig. 4.4).
Figure 4.4: CMD spots surface coverage following 3-day exposition to confluent cells. Spot size was measured each day and compared to initial spot size. (*) Some spots were completely covered on Day 3, therefore the average initial size was not 1, resulting in larger standard deviations.
Figure 4.5 shows representative images of long-term cell adhesion assays. The three
classes of CMD layers defined earlier regarding their cell-resistance are still observed. Some
spots were partially covered by cells on days 3 and 4 and disappeared on the third day of
confluence (CMD surfaces made using condition nos. 4, 5, and 7; Fig. 4.5a-c), while some
showed good cell resistance on days 3 and 4, but were invaded by cells on day 5 (CMD
surfaces made using condition nos. 2, 6, 8; Fig. 4.5d-f). Finally, CMD spots made using
-0,2
0
0,2
0,4
0,6
0,8
1
1,2
1,4
#1 #2 #3 #4 #5 #6 #7 #8 #9
CMD immobilization conditions
No
rma
lize
d s
po
t a
rea
Day 3
Day 4Day 5
* *
*
78
condition nos 1, 3, 9 resisted cell adhesion over the 3 days, but their size progressively
decreased, with the expansion of the surrounding cell sheet (Fig. 4.5g-i).
Figure 4.5: CMD spots exposed to confluent fibroblasts for a period of 3 days. CMD layers made using condition nos. 5 (a-c), 2 (d-f), and 9 (g-i).
Cell morphology and fibronectin matrix on CMD array after 3 days of confluence
To shed more light on cell behavior observed on CMD surfaces presenting various
resistance levels, cell labeling was carried out on day 5.
Phalloidin actin-labeling of cells fixed on CMD arrays following 3 days of confluence
revealed that outside CMD spots, that is, on HApp layers, cells were arranged in a bilayer.
They had well developed actin stress fibers and peripheral focal adhesions (Fig. 4.6a,b). They
were surrounded by a dense fibronectin fibrillar network. Cells found on covered CMD spots
made using condition nos. 4, 5, and 7 presented stress fibers and were arranged in a bilayer as
on HApp film (Fig. 4.6j). On these CMD spots, fibronectin matrix was organized into a
fibrillar network spanning several cells (Fig. 4.6l). Except for cell density, almost no
Day 5 Day 3 Day 4
c
f
a
d
g
b
e
h i
CMD #5
CMD #2
CMD #9
200µm
79
Figure 4.6: Fibroblast actin cytoskeleton (a,d,g,j), focal adhesion formation (b,e,h,k) and human fibronectin deposition and reorganization (c,f,i,l) after 3 days of confluence on CMD spots. Cells found on HApp surfaces (a,b), at the edge of CMD spots made using condition nos. 1, 3, and 9 (c-f), on CMD spot made using condition no. 6 (g,h,i,k), and on CMD spot made using condition no. 5 (j,l).
distinction could be noted between cells on HApp layers and cells on CMD spots made using
condition nos. 4, 5 and 7 at day 5.
On partially covered semi-resistant CMD spots (made using condition nos. 2, 6, 8),
invading cell extensions presented a weak actin staining: they had few stress fibers when
a- HApp b- HApp c- CMD #9
d- CMD #1 e- CMD #9 f- CMD #3
g- CMD #6 h- CMD #6 i- CMD #6
l- CMD #5 k- CMD #6 j- CMD #5 50µm
Actin Vinculin Fibronectin
80
directly on CMD surfaces (Fig. 4.6g) and few focal adhesions at extremities of isolated cells
(Fig. 4.6h), while focal adhesions were more numerous at the extremities of group of cells
(Fig. 4.6k). Some actin filaments appeared out of focus in Figure 4.6g: they correspond to
cells forming a bridge over the spot. These cells leaned on surrounding cells and did not seem
to interact with the CMD surface.
Finally, on the edge of cell-resistant CMD spots (made using conditions nos. 1, 3 and
9), cells were piled up (i.e., stacked), showed actin stress fibers uniformly aligned on the spot
curvature (Fig. 4.6d) and a fibrillar fibronectin network. Some cells on the extreme edge
showed no extracellular fibronectin (Fig. 4.6c) probably owing to fibronectin inability to
adsorb on these CMD surfaces. However, few cells dived in some cell-resistant CMD spots,
grouped in a single layer. They showed some focal adhesions and granular fibronectin at their
extremity inside the CMD spot (Fig. 4.6e,f). Those invading cells seemed to need a support
from neighboring cells settled on the HApp layer. Besides, as illustrated in Figures 4.5h and
4.5i, on day 4, cells were already stacked along the spot edge and, some cells were able to get
over the CMD barrier on day 5, leaning on the cell sheet.
Validation of an optimal condition towards cell resistance
In order to determine an optimal cell-resistant CMD surface, CMD spots were made on
HApp layers using a 2 mg/ml CMD solution concentration, 70-kDa CMDs, an
EDC+NHS/COOH ratio of 5:1, with no added electrolyte, and with CMD molecules having
5%, 25% or 50% carboxylic residues. These conditions were selected based on results of the
present study and on those of a previous work reporting the relationship between physico-
chemical properties of CMD layers and the level of fouling from serum on these same CMD
layers.12 These CMD surfaces, referred as to CMD-5, CMD-25 and CMD-50, were exposed to
fibroblasts. Cell resistance of these CMD spots were compared to that of PEG layers made
under cloud point conditions. These PEG layers were used as a reference, because PEG is one
of the most popular low-fouling polymer and it has been shown that these PEG layers
produced under cloud point conditions had no detectable protein adsorption using quartz
crystal microbalance.4 CMD-5 and CMD-50 layers spot size decreased progressively while
CMD-25 was as resistant as PEG layers towards cell attachment and invasion (Fig. 4.7). For
81
some CMD-25 and PEG spots, spot size even increased between the second and third day,
indicating that cells may have detached or retracted.
Figure 4.7: Determination of an optimal cell-resistant CMD surface. CMD surfaces immobilized using optimized conditions and various carboxylation degrees (referred to as CMD-5, -25 and -50) and PEG surfaces made under cloud point conditions (PEG-CP) were exposed to confluent fibroblasts for a period of 3 days. A: spot size evolution. B: optical microscope images of spots over the 3 days.
PEG-CP CMD-5 CMD-25 CMD-50
No
rma
lize
d s
po
t a
rea
0,0
0,2
0,4
0,6
0,8
1,0
1,2 Day 3
Day 4
Day 5
A
B Day 3 Day 4 Day 5
PEG-CP
CMD-25
a b c
d e f 200µm
82
4.6. Discussion
Polysaccharides6-11,27,29 and poly(ethylene glycol) (PEG)3,4,28,30,31 layers are widely
studied for their capacity to resist protein and cell adhesion. Layers low-fouling property
depends on their chemical composition but also on their structure (layer density, thickness and
conformation, for instance), which in turn depends on immobilization conditions.6,30,32,33
Low-fouling properties of thin polymer layers may be explained first by steric-entropic
repulsion effects. This mechanism is believed to explain repellant abilities of dense and thick
enough polymer layers. Polymer chains compression and desolvation upon interaction with
approaching proteins could result in an entropic loss, unfavorable to protein adsorption.2 A
second mechanism regarding hydrated polymers relies on the interactions at the interphase
between water molecules and hydrophilic polymer moieties.14 The formation of a stable water
network involving tightly bound water molecules at polymer interface depends on molecules
orientation and on the strength of interaction.34 This stable solvation shell may form a barrier
preventing proteins or molecules to interact with the hydrated layer.
To find optimal low-fouling coatings, arrays of CMD layers immobilized under
various conditions were developed, and their structure was analyzed in relation to their protein
adsorption resistance in a first step.12 These results are reported in Table 4.2. CMD
immobilization was assessed by X-ray electron spectroscopy (XPS) and the layers interactions
with a silica sphere were analyzed by atomic force microscopy (AFM) force measurements.
The results suggested that the ionic strength influenced polymer conformation upon initial
adsorption; the presence of electrolytes induces the collapse of the CMD molecules. The
coupling agent concentration (EDC+NHS) and the carboxylation degree rather regulated
pinning density, controlling final surface coverage, layer density and thickness. Moreover,
protein adsorption measured by surface plasmon resonance microscopy (SPR) showed that,
better protein rejection was obtained for uniform, dense and thick CMD layers.12
In a second step, that is, in the present study, CMD arrays were exposed to cells to
further screen CMD coatings cell resistance and to observe cell behavior towards layers with
various structures.
83
TABLE 4.2: Physico-chemical characterization of CMD layers and their resistance to protein adsorption and cell adhesion.
Conditions no. 1 no. 2 no. 3 no. 4 no. 5 no. 6 no. 7 no. 8 no. 9
O/C ratio a 0.19 0.20 0.19 0.13 0.13 0.24 0.13 0.18 0.20
Separation distance 10 mM b 63 48 44 19 50 65 20 10 60
Separation distance 150 mMb 33 44 22 nd 43 33 nd nd 38
Protein adsorption c + ++ + +++ ++ + +++ ++ +/-
Cell coverage d + ++ + +++ +++ ++ +++ ++ +/-
a oxygen/carbon (O/C) atomic ratio determined by XPS analysis. b Separation distance in nm observed for AFM force curves obtained between a silica sphere and layers in 10 and 150 mM PBS solutions. c Protein adsorption measured by SPR microscopy. a, b, c Results are derived from Ref 12. d CMD layers coverage by cells after 3 days of confluence.
Four hours after seeding CMD arrays with fibroblasts, few cells were present on CMD
spots and those cells were rounded, with an undeveloped actin cytoskeleton, no focal
adhesion, and they were not able to deposit matrix fibronectin. Like most polysaccharide
coatings, all CMD layers offered some resistance to cell adhesion, impeding cell adhesion
and/or spreading.6,9,27,29,35 Previous studies of protein adsorption on dextran-derivative graft
layers showed that protein adsorption was lower on these polysaccharide layers than on amine
surfaces.10-12,36 The hydrated polysaccharide layer may hide, at least partially, the adhesive
HApp surface. Fibronectin and other matrix proteins, either secreted by the cells or contained
in culture media, may not adsorb in sufficient amount on the CMD surfaces following a 4h
assay. By limiting protein adsorption, CMD surfaces may impair formation of cell-surface
contacts, affecting cytoskeleton organization as well as ECM assembly, thus decreasing cell
adhesion and cell survival.
However, 12h after seeding, cells on CMD spots had either disappeared (possibly by
detaching) or begun to spread. Observation of the various conditions allowed defining three
classes of CMD layers depending on their resistance to cell adhesion; the same distinct
categories were observed after exposing CMD arrays to confluent cells for 3 days. The
difference in cell behavior towards the various CMD surfaces may be explained by a failure of
the resistance barriers offered by CMD layers towards protein adsorption and cell adhesion.
Immobilization conditions and layer structure may account for the various responses.
84
Layers made of CMD condition nos. 4, 5 and 7 supported cell adhesion and spreading
at 12h, associated with deposition of matrix fibronectin, formation of focal adhesions and
organization of actin into stress fibers. Exposition of these CMD layers to confluent cells
resulted in a rapid and complete coverage, with a cell organization similar to that on HApp
layers. Protein adsorption assays revealed that these CMD surfaces presented the least
resistance towards protein adsorption.12 These CMD layers were immobilized with a low
coupling agent concentration (resulting in a low pinning density) and with electrolyte.
Physico-chemical analysis suggested that these CMD layers were not dense (low O/C ratio)
and were irregular (random size distribution of loops and tails) and possibly presented
pinholes. Polysaccharide layers immobilized with a low pinning density, either by polymer
adsorption or end-on-grafting have been shown to be more subject to protein adsorption and
cell adhesion than covalently immobilized polysaccharides in side-on configuration.6,7,33 They
did not form a sufficiently dense protective barrier and the surface coverage may not
completely hide the underlying surfaces, which are subject to non-specific protein adsorption
and cell adhesion.
The second class of CMD layers is composed of condition nos. 2, 6 and 8, they are
called semi-resistant. At 12h, some cells were still present on those CMD layers, although
neither deposition of fibronectin, nor focal adhesion formation nor actin cytoskeleton was
observed. After exposition to confluent cells, semi-resistant CMD spots were quite resistant to
cell adhesion for 2 days, but they then were suddenly covered on the third day: few cells
established contacts with the CMD layer, while some formed bridges above the surface,
leaning on settled cells. Besides, observation of CMD spots covered at day 5 revealed that
while cells tended to be aligned in bundles on cell-covered CMD spots made using condition
nos. 4, 5, and 7 (Fig. 4.5c), similarly as on HApp surfaces, cells on covered semi-resistant
CMD spots formed a lattice (Fig. 4.5f), indicative of a different invasion mechanism. Layers
made using condition nos. 2, 6 and 8 were immobilized either with a high EDC+NHS/COOH
ratio (layers 6 and 8) or with almost no added electrolyte (layer 2), improving immobilization
efficiency despite a low EDC+NHS/COOH ratio.12 XPS and AFM force measurements
suggested that those layers have a higher surface density and that they were uniform, pinhole
free, probably hiding the underlying HApp surface, which explains their better resistance to
cell adhesion. Despite a uniform surface coverage, CMD layers made under condition nos. 2
85
and 8 were also subject to protein adsorption.12 Layers made under condition no. 2 were
extended, due to a very low electrolyte concentration, but they were diffuse, due to a low
pinning density. These layers did not constitute a steric barrier as their open structure allows
in-diffusion of proteins.37 Layers made under condition no. 8 were dense but presented short
extensions with limited mobility and compressibility, due to the immobilization in the
presence of a high electrolyte concentration and a high pinning density, limiting steric-
entropic repulsion of proteins. As previously stated, to prevent non-specific protein adsorption
and cell adhesion, polymers should be immobilized as dense and compressible layers i.e.,
thick enough and uniformly covering the surface.
Surprisingly, layers made using condition no. 6 were as resistant to protein adsorption
as layers made under condition nos. 1, 3 and 9,12 yet some cells were able to adhere and
spread on CMD layer 6. CMD surfaces made under condition nos. 1, 3, 6 and 9 all resulted in
dense, extended and compressible layers, hindering the access of proteins to the underlying
plasma layer by forming a steric barrier. However, high MW dextrans (e.g., 500-kDa used in
condition no. 6) have an increased polydispersity index and few long branches.38 The degree
of intra-molecular freedom is decreased and affects CMD molecules ability to rearrange on
surfaces. At lower MW, CMD is linear with very few to no long branches, therefore having a
greater conformational mobility and ability to reorient more effectively to optimize packing at
the interface (70-kDa used in condition nos. 1, 3 and 9). Layers made using condition no. 6
completely covered the HApp layer,12 but may present some surface defects with
heterogeneous thickness due to limited possibility of molecules reorganization. AFM imaging
of thiolated dextran39 and CMD (unpublished results)∗ of various molecular weights showed
presence of bulges for 500-kDa dextran derivative layers while 70-kDa layers looked quite
smooth. The interfacial solvation layer made of water molecules tightly bound to the CMD
layer, and potentially preventing protein adsorption,14,15 may be disturbed by protruding loops
and tails, locally exposing sites for protein adsorption and establishment of cell-surface
contacts. This may explain why previous experiments assessing total protein adsorption on
dextran and CMD surfaces did not reveal any significant effect of MW7,12,40 while cells were
able to establish contact with underlying 500-kDa CMD layer 6.
∗ some images are presented in Annexe B
86
Finally, graft layers made using CMD condition nos. 1, 3, 9 with a high pinning
density, low electrolyte concentration and 70-kDa CMD, formed uniform, dense and extended
hydrated layers12 providing a steric-entropic barrier to protein adsorption. Indeed, layers made
using condition nos. 1, 3 and 9 showed some protein adsorption but to lower extent than other
layers (except those made using condition no. 6). As they were made from 70-kDa CMD, they
were supposed to present a homogeneous surface, possibly favoring the formation of a stable
water layer, an additional protection towards protein adsorption and cell adhesion. Almost no
additional cells were observed on these layers after 12h. When exposing CMD arrays to
confluent cells, spots made using those conditions were progressively invaded by cells but
were still apparent on the last day of the experiment. As mentioned earlier, few cells resting
on fibroblasts settled at the spot edge could adhere on the CMD spots. It is noteworthy to
mention that such a mechanism of cell adhesion and cell spreading may not occur on a
uniformly covered cell-resistant CMD surface tested in standard 2D cell culture assays, as this
study reports the cell behavior at the interface between an amine-bearing plasma polymer
layer and CMD spots on a CMD array. Definitely, the immobilization of 70-kDa CMD instead
of 500-kDa CMD, with a high pinning density and low electrolyte concentration, provides an
additional protective barrier towards cell adhesion.
CMD layers made using condition nos. 1 and 3 were comparable and showed less
resistance to protein adsorption and cell adhesion than those made under condition no. 9. The
latter performance could be attributed either to the use of a higher CMD concentration during
the immobilization or to an intermediate carboxylation degree. Higher concentration would
result in a denser layer,41 offering a better steric-entropic barrier, while carboxylation degree
regulates pinning density, layer thickness via loops and tails size distribution, and charge
density. Hence, three conditions with various carboxylation degrees and a concentration of 2
mg/ml were tested to determine the optimal immobilization conditions leading to a cell
repulsive layer. They were immobilized with a high EDC+NHS/COOH ratio to increase
surface coverage and density, with no added electrolyte to favor chains extension, and they
were made of CMD 70-kDa, to allow the formation of a homogeneous surface. CMD-25 was
resistant towards cell adhesion while CMD-5 and CMD-50 were progressively covered by
cells, though they were immobilized with the same concentration. Martwiset et al. studied the
effect of the modification degree of oxidized dextran layers on protein adsorption and they
87
found better protein repulsion for dextran layers made with an intermediate degree of
oxidation (25%).36 Modification degree first controls pinning density; dextran derivatives
immobilized with lower modification degrees such as CMD-5 are more diffuse.10,32 However,
although those layers were immobilized with a higher concentration than layers made under
conditions #1, they were less repulsive. It should be noted than CMD-5 was grafted in total
absence of added electrolytes therefore molecules can be immobilized on the HApp surface in
a more extended conformation, presenting less carboxylic groups for covalent linkage to the
amine surface, thus reducing pinning density. On the contrary, CMD-50 was immobilized
with a high pinning density, resulting in short loops and tails with a probably reduced
mobility; it must have been immobilized with a higher density due to a higher concentration.
It was also noted that dextran derivatives with higher modification degrees had reduced
dissolution (personal observation and Ref. 36). Reactive groups added confer a different
molecular structure to the polymer chains, which may alter the interactions with the
surrounding water molecules and the structure of the interfacial water layer. Finally, in the
case of CMD, increasing carboxylation degree and surface density lead to an increased charge
density,32 generally associated with protein adsorption via electrostatic interactions.11 Hence,
layers ability to reject protein adsorption and cell adhesion results from a combination of
different types of interactions such as i) electrostatic interactions, which should be minimized
by decreasing surface charge density, ii) steric-entropic effects, maximum for dense and
flexible layers, and iii) molecular interactions between water molecules and polymer moieties
at the interphase, which may provide a protective shell.
CMD layers were compared to PEG surfaces as they are the most popular and most
studied low-fouling polymers. PEG was immobilized close to CP conditions, as this
immobilization mode yields the highest chain density and the best low-fouling property.4,31,42
CMD-25 and PEG-CP layers both resisted cell adhesion while exposed to confluent cells for
three days; they can be considered as non-fouling layers. Then, CMD layers immobilized with
optimal conditions can provide a good alternative to PEG layers, being as cell-resistant, less
expensive, and derived from a natural source. CMD has also the advantage of being
multivalent: carboxylic groups present in its structure can be easily modified yielding an
interface useful for the development of bioactive surfaces on which biosignals can be
covalently incorporated.
88
4.7. Conclusions
Arrays of CMD graft layers were exposed to fibroblasts for short-term and long-term
cell adhesion assays. Behavior of cells exposed to the different CMD layers was studied and
CMD layers cell resistance was screened in relationship with the immobilization conditions
used.
Cell-CMD layer interactions were time-dependent. Non-resistant CMD layers,
characterized by a non-uniform surface coverage, allowed matrix protein adsorption and the
subsequent formation of focal contacts, cell adhesion and cell spreading. On semi-resistant
CMD layers, cells encountered some resistance and most of them relied on interactions with
attached neighboring cells for spot invasion. Only the dense and thick CMD graft layers made
using 70-kDa CMD could partially resist cell adhesion after exposition to confluent cells. By
minimizing protein adsorption, resistant CMD layers impaired the formation of cell-surface
adhesion sites, preventing cell adhesion and cell survival.
Finally, screening of CMD layer cell resistance led to the synthesis of an optimized
CMD layer, which was dense, thick, and homogeneous. This optimized CMD layer was as
cell-resistant as PEG layers produced under CP conditions and represent a good alternative.
4.8. Acknowledgements
The authors wish to thank Dr. Charles J. Doillon for providing cells.
4.9. References
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92
Chapitre 5
Puces bioactives immobilisées sur des surfaces anti-
adhérentes pour étudier l’adhésion spécifique des
cellules endothéliales
5.1. Résumé
Afin d’étudier comment contrôler l’adhésion spécifique et le comportement de cellules
endothéliales sur les surfaces de biomatériaux, des puces bioactives ont été développées et
validées. Ces puces sont faites d’une surface anti-adhérente qui prévient l’adhésion cellulaire
non-spécifique et portent des molécules bioactives à des positions données exposant des
ligands spécifiques pour les récepteurs cellulaires. Les puces de molécules bioactives
(peptides de séquence RGD, REDV, SVVYGLR et le facteur de croissance vascular
endothelial growth factor (VEGF)) ont été immobilisées sur une couche anti-adhérente de
carboxy-méthyl-dextran (CMD) et ont été exposées à des cellules endothéliales et des
fibroblastes humains pour étudier l’effet de la composition moléculaire de chaque “spot”
bioactif sur l’adhésion cellulaire. Seules les cellules endothéliales ont été sensibles au peptide
RGD co-immobilisé avec les séquences REDV ou SVVYGLR : celles-ci induisent un
étalement réduit et la perte de fibres de stress d’actine. La co-immobilisation de RGD avec le
VEGF résulte aussi en la réorganisation des filaments d’actine et des points focaux chez les
cellules endothéliales. La combinaison de RGD avec ces molécules sélectives pour les cellules
endothéliales n’induit pas un phénotype associé à une adhésion forte mais plutôt
caractéristique de cellules en mouvement.
93
Bioactive microarrays immobilized on
low-fouling surfaces to study specific endothelial
cell adhesion
Emmanuelle Monchaux and Patrick Vermette
5.2. Abstract
With the aim to study how to modulate the specific endothelial cell patterning and
responses on biomaterials surfaces, bioactive micro-arrays were developed and validated for
specific cell patterning. These micro-arrays were made of low-fouling surfaces, that prevent
nonspecific cell adhesion, bearing bioactive molecules at given known locations by presenting
specific ligands to cell receptors. Arrays of bioactive molecules (RGD, REDV, and
SVVYGLR sequences and vascular endothelial growth factor (VEGF)) were immobilized on
a carboxy-methyl-dextran low-fouling surface and were exposed to human endothelial cells
and fibroblasts to screen for the effect of bioactive spot molecular composition on cell
adhesion. Endothelial cells only were sensitive to RGD peptide co-immobilized with REDV
or SVVYGLR sequences: they induced a reduction in cell spreading and a loss of actin stress
fibers. RGD co-immobilized with VEGF also resulted in the reorganization of actin filaments
and focal points in endothelial cells. Combination of RGD with these endothelial cell-selective
biomolecules did not elicit a strong adhesion phenotype but rather one characteristic of
migrating cells.
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5.3. Introduction
Control over endothelial cell responses at the biomaterial interface is important in
applications such as vascular prosthesis endothelialization and tissue engineering.1 The
formation of a continuous monolayer of endothelial cells on the luminal surface of vascular
prosthesis helps to resolve the problem of graft thrombogenicity.2 On the other hand, to
construct or regenerate organs, invasion of a scaffold by endothelial cells and the subsequent
formation of a capillary network is essential for tissue survival and growth.3,4
Endothelial cell adhesion to a biomaterial surface and subsequent cell survival,
proliferation, or migration are mediated by cell-material interactions via cell adhesion
receptors such as integrins.5 Control of cell-material interactions, and thus cell behavior, may
first be achieved by preventing nonspecific protein adsorption and cell adhesion and, second,
by exposing specific cell ligands or bioactive molecules to optimize the adhesion, migration,
and/or proliferation of desired cells.
Micro-arrays of extracellular matrix (ECM) proteins and growth factors have been
used to screen molecules impact on cell adhesion, proliferation, and differentiation.6-9 In the
present study, arrays of bioactive molecules known to be specific for endothelial cells were
fabricated to analyze their effect on endothelial cell adhesion. Bioactive molecules were
immobilized on a low-fouling template layer to be able to directly observe only the effect of
molecule-cell receptor binding on cell responses by efficiently eliminating nonspecific
interactions. Layers made of carboxy-methyl-dextran (CMD), a dextran derivative, were used
because they highly resist nonspecific protein adsorption and cell adhesion10-14 and possess
reactive groups convenient for the grafting of bioactive molecules such as peptides and
proteins.15
RGD is the most effective and most often used peptide sequence to promote cell
adhesion on synthetic surfaces.16 The RGD sequence is present in many ECM proteins such as
fibronectin, vitronectin, and collagen, and RGD is able to address more than one integrin
receptor. The RGD peptide was immobilized on its own and in combination with other
bioactive molecules specific for endothelial cells.
The REDV sequence, found within the alternatively spliced CS5 fibronectin domain, is
specifically recognized by the integrin α4β1.17 REDV peptides grafted on synthetic surfaces
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were shown to induce selective adhesion of endothelial cells,18,19 and the REDV sequence
mediates endothelial cell migration on fibronectin via its α4β1 receptor.20 The SVVYGLR is a
cryptic sequence exposed after osteopontin thrombin cleavage and is principally recognized
by integrin α9β1.21 Grafted SVVYGLR induces endothelial cell adhesion, migration, and
enhances angiogenesis in its soluble form.22-24 REDV and SVVYGLR sequences are believed
to be selective for endothelial cells as α4β1 and α9β1 expression is limited to a small number
of cell types.25,26
Vascular endothelial growth factor (VEGF) is an endothelial specific growth factor.
VEGF-induced effects are principally mediated by interaction with its receptor VEGF-R2: it
enhances endothelial cell proliferation, migration, and survival and is a potent angiogenic
factor.27,28 In vivo, endothelial cells respond to both soluble VEGF and immobilized VEGF,
bound to the ECM. In vitro, immobilized VEGF induces endothelial cell migration and
survival.29,30
In this study, micro-arrays made of bioactive molecules were exposed to human
endothelial cells and human fibroblasts to investigate cell adhesion, spreading, cytoskeletal
organization, and focal adhesion assembly with regard to bioactive spots molecular
composition.
5.4. Materials and methods
5.4.1. Carboxy-methyl-dextran layers.
Detailed fabrication procedures of an optimized low-fouling CMD graft layer have
been reported elsewhere.12 CMD was prepared as follows. Five grams of 70 kDa dextran
(Amersham Biosciences, Uppsala, Sweden, cat. no. 17-0280- 01) were dissolved in 25 mL of
a 1 M NaOH solution. Then, bromoacetic acid (Lancaster Synthesis Inc., Pelham, NH, cat. no.
3016) was added to a final concentration of 0.5 M. The solution was stirred overnight and
dialyzed against Milli-Q water (Millipore Canada, Nepean, Canada), 0.1 M HCl, and finally
Milli-Q water, for 24 h each. The solution was then lyophilized and stored at -20 °C. The
carboxylation degree was determined by 1H NMR to be 27.5%.
Borosilicate glass substrates (Assistent, Sondheim, Germany) were washed in
Sparkleen solution (Fisher Scientific Canada, Ottawa, Canada), thoroughly rinsed with Milli-
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Q water, and dried with 0.2 µm filtered air. Clean substrates were surface-modified by plasma
polymerization of n-heptylamine (99.5% purity, Sigma-Aldrich, Oakville, Canada, cat. no.
126802) in a custom-built plasma reactor.31 Conditions of plasma polymerization were an
initial monomer pressure of 0.040 torr, an ignition power of 80 W, an excitation frequency of
50 Hz, a deposition time of 45 s, and a distance between the electrodes of 10 cm. The resulting
n-heptylamine plasma polymer (HApp) layers exposed functional amine groups that were
used to graft CMD layers using carbodiimide chemistry.
A 2 mg/mL CMD solution was filtered with 0.45 µm pore size membrane to limit
CMD aggregates and then activated with 17 mM 1-ethyl-3-(3-dimethylaminopropyl)
carbodiimide (EDC, Sigma-Aldrich, cat. no. E1769) and N-hydroxysuccinimide (NHS,
Sigma-Aldrich, cat. no. H7377) for 10 min. Substrates freshly covered by HApp layers were
rapidly immersed into the activated CMD solution, and reaction was allowed to proceed for 2
h under agitation. CMD surfaces were then rinsed with 1 M NaCl solution (2 x 12 h), with
Milli-Q water (2 x 12 h), and autoclaved to remove any unbound molecules left. They were
dried with 0.2 µm filtered air and stored in a desiccator until needed.
5.4.2. Fabrication of arrays of bioactive molecules.
Peptide sequence G-R-G-D-S and its inactive control G-R-G-E-S, named RGD and
RGE (American Peptide Company, Sunnyvale, CA, cat. nos. 44-0-23 and 44-0-51,
respectively), were dissolved in 150 mM PBS (pH 8.5) to a final concentration of 25 µg/mL
(solutions R25 and RGE25). In some samples, G-R-E-D-V-D-Y or G-D-S-V-V-Y-G-L-R
(named REDV and SVVYGLR, Celtek Bioscience, Nashville, TN) was added with a final
concentration of 1 µg/mL to R25 solutions (solutions RE1 and SV1). Also, in other samples,
recombinant human VEGF165 (Peprotech, Rocky Hill, NJ, cat. no. 100-20) was added to a
final concentration of 2 µg/mL to R25, RE1, or SV1 solutions (solutions R25V, RE1V, and
SV1V, respectively).
CMD surfaces were activated by a 200 mM EDC + NHS solution for 8 min, rapidly
rinsed in Milli-Q water, and dried with 0.2 µm filtered air. Biomolecule solutions were placed
in a 384-well plate and three individual spots of each biomolecule solution were deposited
with a 1200 µm pitch on activated CMD surfaces using a non-contact dispensing robot
(BioChip Arrayer, Perkin-Elmer Life and Analytical Sciences, Boston, MA) with arraying
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capabilities such as those used to manufacture DNA and protein chips. The piezo-dispensers
were set to dispense 50 drops of solution per spot with 350 pL of sample per drop, resulting in
ca. 700 µm diameter spots. The reaction was allowed to proceed overnight in a chamber filled
with air saturated with humidity to maintain the substrates at dew point. Bioactive molecule
arrays were rinsed in PBS (pH 8.5) coupling buffer (2 x 3 h) and then with PBS at pH 7.4.
They were finally immersed overnight in sterile PBS containing antibiotics (100 units/mL
penicillin and 100 µg/mL streptomycin, Invitrogen, Burlington, Canada, cat. no. 15140) and
stored at 4 °C until needed.
5.4.3. Testing bioactive arrays toward cell responses.
Human umbilical vein endothelial cells (HUVECs) from Promocell (Heidelberg,
Germany, cat. no. C-12200) were cultured in Medium 199 (M199, Sigma, cat. no. M5017)
supplemented with 10% fetal bovine serum (FBS, Sigma, cat. no. F1051), 100 units/mL
penicillin, and 100 µg/ mL streptomycin, 90 µg/mL heparin (Sigma, cat. no. H1027), 2 mM L-
glutamine (Invitrogen, cat. no. 25030), and 20 µg/mL endothelial cell growth supplement
(ECGS, Sigma, cat. no. E2759). Human foreskin fibroblasts harvested from human biopsies
with collagenase were maintained in Dulbecco’s modified Eagle’s medium (DMEM,
Invitrogen, cat. no. 12100) with 10% FBS and antibiotics. Cells were incubated at 37 °C in a
5% CO2 incubator.
5.4.4. Cell adhesion assay.
Near confluence cells were harvested by a short trypsin-EDTA treatment (Invitrogen,
cat. no. 25200), washed, and resuspended in serum-free culture media (M199 for HUVECs
and DMEM for fibroblasts), incubated at 37 °C for 30 min, and seeded at 7500 cells/cm2 on
sterilized bioactive molecule arrays. Six hours after cell seeding, arrays were washed with
PBS to remove non-adherent cells and were fixed with 3.7% formaldehyde during 15 min for
further labeling. Images of each spot were taken using phase contrast microscopy (Nikon
Eclipse TE2000-S) at 100x magnification. Cell number per spot was counted manually. The
cell projected area was outlined manually using SigmaScan Pro software (SPSS Inc., Chicago,
IL). Each biomolecule solution was deposited in triplicate on an array, and the assay was
carried out in duplicate for each cell type.
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5.4.5. Cell labeling.
Fixed cells were rinsed with PBS, permeabilized with 0.5% Triton X-100 for 5 min,
and blocked with 2% bovine serum albumin (BSA, Sigma, cat. no. A7906). Samples were
incubated for 1 h at room temperature with the primary mouse anti-vinculin antibody (1:25,
Sigma, cat. no. V4505), rinsed with PBS, and incubated with Alexa Fluor 488 goat anti-mouse
secondary antibody (1:1000, Invitrogen, cat. no. A11001) mixed with phalloidin-TRITC
(1:300, Sigma, cat. no. P1951) for 1 h at room temperature. The samples were rinsed and
mounted on glass slides and observed with an epifluorescence microscope (Nikon Eclipse)
equipped with oil immersion objectives (60x and 100x).
5.4.6. Statistical analysis.
The adhesion results were presented as mean values ± standard deviations. Statistical
analysis was performed using Student’s t test for paired samples to compare bioactive spots to
the RGD spots (R25). A p-value <0.05 was considered statistically significant.
5.5. Results
Surfaces bearing arrays of bioactive molecules were made by covalent immobilization
of bioactive molecules on a low-fouling CMD layer (Figure 5.1). They were exposed to
HUVECs and human fibroblasts to examine how grafted bioactive molecules affect cell
adhesion, spreading, cytoskeletal organization, and focal adhesion assembly.
Figure 5.1. Reaction scheme for the grafting of carboxy-methyl-dextran (CMD) to the HApp-modified surface and subsequent bioactive molecule (peptide or growth factor) immobilization.
99
5.5.1. RGD peptide is required to initiate cell adhesion.
Arrays were seeded with HUVECs or fibroblasts, and cell density was measured on
each spot after 6 h. For both cell types, no cell adhesion was observed on the bare CMD layer
or on RGE spots. Cells adhered and spread on spots bearing covalently immobilized RGD
(Figure 5.2, parts A and B). However, no cell attached on spots where REDV, SVVYGLR
(0.1-100 µg/mL), or VEGF (2-8 µg/mL) were grafted individually (data not shown). The co-
immobilization of these molecules with RGD was necessary to induce HUVEC and fibroblast
adhesion. Bioactive spots then supported cell adhesion and growth as illustrated in Figure
5.2C. Cells remained within the spot limits even after a few days in culture in the presence of
serum.
Figure 5.2. Microscopy images of phalloidin-actin labeling of (A) human umbilical vein endothelial cells (HUVECs) and (B) human foreskin fibroblasts adhering on RGD spots, 6 h after cell seeding, and of (C) confluent HUVECs cultured on a RGD spot for 5 days with 10% serum. Scale bar is 250 µm.
5.5.2. REDV, SVVYGLR, and VEGF specifically affect endothelial cell adhesion when
co-immobilized with RGD.
Figure 5.3A reveals that co-immobilization of the RGD peptide with REDV (RE1) did
not induce any effect on HUVEC attachment 6 h after seeding. However, the co-
immobilization of RGD or RGD + REDV with VEGF (R25V and RE1V, respectively) leads
to a significant increase in endothelial cell density. RGD co-immobilized with SVVYGLR
(SV1) also significantly enhanced HUVEC adhesion. Addition of VEGF to this combination
(SV1V) did not elicit an increase in endothelial cell attachment. On the other hand, the co-
immobilization of RGD with either endothelial cell-selective peptides or VEGF induced no
effect on fibroblast attachment (Figure 5.3B).
A B C
100
Figure 5.3. (A) Human umbilical vein endothelial cells (HUVECs) and (B) human foreskin fibroblasts adhesion on bioactive spots 6 h after cell seeding. Significant difference compared with RGD alone (R25) at *p< 0.05 or ** p< 0.01.
In vitro, initial cell attachment results in cell spreading. Projected cell area was
measured for cells adhering on each bioactive spot, and results presented in Figure 5.4 suggest
that HUVECs attached on spots exposing either REDV or SVVYGLR showed a reduction in
spreading, while immobilized VEGF did not induce any effect 6 h after seeding. The
molecular composition of bioactive spots did not affect fibroblasts spreading (data not shown).
Figure 5.4. Spreading levels measured for human umbilical vein endothelial cells (HUVECs) on bioactive spots 6 h after cell seeding.
R25 R25V RE1 RE1V SV1 SV1V RGE25
cell n
um
ber/
mm
²
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40
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80
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A
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30
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B
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101
5.5.3. REDV, SVVYGLR, and VEGF affect actin filaments organization and focal
adhesion assembly in endothelial cells.
In vitro, cell adhesion is usually followed by formation of focal adhesions and
reorganization of the actin cytoskeleton into stress fibers. Cells adhering to bioactive spots
were fixed 6 h after cell seeding and labeled for their actin cytoskeleton and for vinculin, one
of the major structural components of focal adhesions.32 HUVECs adhering on spots made of
RGD displayed a well-organized actin cytoskeleton with actin stress fibers associated with
thick focal adhesions (Figure 5.5, parts E and F). However, cells on surfaces exposing REDV,
SVVYGLR, or VEGF presented a different cytoskeletal organization. HUVECs on RGD +
REDV (RE1) spots (Figure 5.5I) or RGD + SVVYGLR (SV1) spots (Figure 5.5M) displayed
actin organized into thin filaments, but few actin stress fibers were observed; focal contacts
were small (Figure 5.5, parts J and N). HUVECs adhering on spots bearing VEGF (R25V,
RE1V, and SV1V) showed no stress fibers, and actin filaments were rather organized into
cortical networks associated with membrane ruffling (Figure 5.5, parts G, K, and O). Focal
contacts were small and mostly present at the cell edges (Figure 5.5, parts H, L, and P). In
comparison, covalently immobilized VEGF did not seem to affect fibroblast actin
cytoskeleton organization into stress fibers or focal adhesion assembly (Figure 5.5A-D).
5.6. Discussion
Arrays of RGD and bioactive molecules specific for endothelial cells were exposed to
endothelial cells and fibroblasts to screen for their effect on cell adhesion, spreading, actin
cytoskeleton organization, and focal adhesions assembly. Bioactive molecules were covalently
immobilized on a low-fouling CMD layer, which prevents nonspecific protein adsorption and
cell adhesion12,14 and which possesses multiple functional carboxylic groups convenient for
peptides and proteins grafting.15
HUVECs and human fibroblasts both adhered on bioactive spots bearing the RGD
adhesion peptide, whereas no cell attached to the background CMD layer and to the RGE
inactive peptide. This is in good agreement with a previous study in which RGE peptide did
not support cell adhesion.33 As expected, CMD layer properties prevented any unspecific cell
adhesion. The introduction of a peptide containing charged groups may have altered surface
102
Figure 5.5. Actin filaments labeled with phalloidin (A, C, E, G, I, K, N, O) and focal adhesions labeled with anti-vinculin (B, D, F, H, J, L, N, P) in human foreskin fibroblasts and human umbilical vein endothelial cells (HUVECs) on R25 (A, B, E, F), R25V (C, D, G, H), RE1 (I and J), RE1V (K and L), SV1 (M and N), and SV1V (O and P) 6 h after cell seeding. Scale bars are 25µm.
properties, such as charge and hydrophilicity, and induced nonspecific cell adhesion. The
absence of any cell adhesion on the inactive grafted RGE peptide proves that cell adhesion
was directly mediated by specific interactions between the immobilized RGD sequence and its
cell receptors. Moreover, the co-immobilization of RGD with either REDV, SVVYGLR, or
VEGF molecules, believed to be rather specific for endothelial cells as their receptors are
expressed by reduced number of cell types,25,26 elicited changes in HUVEC adhesion,
spreading, and cytoskeletal organization, whereas no change could be detected in fibroblasts
behavior. It is then hypothesized that molecules immobilized on a low-fouling CMD layer are
E F G H
I L K J
M N O P
A B C D
HUVECs
Human fibroblasts
103
specifically recognized by their receptors and that the observed cell responses directly result
from specific interactions between cell receptors and these immobilized bioactive molecules.
Grafting molecules on a CMD layer may be a way to control elicited signals and to design
bioactive surfaces that can specifically control cell responses and patterning in co-culture
systems, for instance.
When exposed to RGD, a peptide sequence recognized by integrins αvβ3 and α5β1
which are predominant in cell adhesion,16,32 endothelial cells adhered, spread, formed focal
adhesions and actin stress fibers indicative of a strong adhesion. However, REDV and
SVVYGLR peptides were not able to promote HUVEC adhesion on their own (i.e., without
the co-immobilization of RGD). The co-immobilization of RGD with either REDV or
SVVYGLR induced a reduction in HUVEC spreading and affected actin filaments
organization and focal adhesion assembly in adhering cells, with a reduced number of stress
fibers and smaller focal adhesions than in cells adhering to RGD-modified spots. The known
cell receptors for REDV and SVVYGLR are integrins α4β1 and α9β1, respectively.17,21
Integrin α4 and α9 subunits are the sole members of a structural subfamily of integrin α
subunits, and they share functional similarities as they both inhibit cell spreading and enhance
cell migration.34-36 α4 and α9 cytoplasmic domains directly bind to paxillin, a signaling
molecule, inducing inhibition of cell spreading. Moreover, paxillin binding to the α4
cytoplasmic domain leads to focal adhesion disassembly and stress fibers disappearance.34 As
α4 and α9 share structural and functional similarities, ligand binding to α9β1 integrin may
also elicit focal adhesion disruption and stress fibers disappearance. Similarities between α4
and α9 functions and the observed morphology of cells adhering to spots bearing REDV and
SVVYGLR (RE1 and SV1) suggest that there is a direct interaction between the immobilized
peptide sequences and their receptors α4β1 and α9β1, leading to an intermediate state of
adhesion.
Previous work on REDV and SVVYGLR sequences suggested that endothelial cells
were able to attach on surfaces bearing these peptides.18,19,22,23 The template layer used in this
study is non-fouling, i.e., absolutely no cell adhered on the CMD surface. Previous studies on
endothelial cell adhesion onto REDV- or SVVYGLR-modified surfaces usually revealed a
minimal nonspecific cell adhesion on control surfaces without peptide.18,22,23 This nonspecific
adhesion may have been sufficient to initiate cell attachment on REDV- or SVVYGLR
104
modified surfaces. Furthermore, although cells adhering on RE1 and SV1 spots displayed
similar changes in their spreading and cytoskeletal organization, only SVVYGLR co-
immobilized with RGD could enhance HUVEC adhesion. If α4 and α9 seem to trigger similar
signals involved in cell spreading and cytoskeletal organization, they may activate different
signaling pathways related to cell adhesion. In addition, although some previous reports
showed an enhanced endothelial cell adhesion on surfaces modified with the REDV
sequence,18,19 some were not able to observe any significant increase in endothelial cell
adhesion.37,38
VEGF also enhanced cell adhesion when co-immobilized with RGD or RGD + REDV
(RE1) but not with RGD + SVVYGLR (SV1). Surface-immobilized VEGF induces adhesion,
migration, and survival of endothelial cells via integrins αvβ3, α3β1, α9β1, and its receptor
VEGF-R2.29,30 When VEGF and SVVYGLR were co-immobilized, there might have been a
competition between the two α9β1 ligands, which may explain why combination of SV1 with
VEGF did not enhance HUVEC adhesion. Nonetheless, as previously noticed,29 endothelial
cells adhering on spots where VEGF was exposed displayed few actin filaments arranged in a
cortical network and peripheral focal points. RGD co-immobilized either with VEGF, REDV,
or SVVYGLR leads to a loss of stress fibers associated with membrane ruffling or
lamellipodia extension and focal adhesion rearrangement into small and/or peripheral ones in
adhering HUVECs. Immobilized VEGF has been shown to induce cell migration,29,30 and
integrins α4β1 and α9β1 also enhance endothelial cell motility.20,34,36 Cell migration requires
constant reorganization of the actin cytoskeleton and focal adhesions: extension of filopodia
and lamellipodia is followed by attachment to the matrix via focal adhesions and formation of
stress fibers to be able to contract the cell body to allow forward movement. Disruption of
focal adhesions at the front edge is then necessary for the cell to migrate.39 The observed
morphology of HUVECs adhering on bioactive spots exposing REDV, SVVYGLR, and/or
VEGF suggests that endothelial cell interaction with these bioactive molecules induces cell
migration rather than a strong anchorage to the surface. Endothelial cell migration is required
for in situ spontaneous endothelialization of vascular grafts or for construction or repair of a
vascularized tissue, for instance.2,3 Future work would then be to test the ability of these
bioactive coatings to promote endothelial cells migration on a 2D-modified surface and to
induce invasion of an engineered tissue substitute with formation of tubular structures.
105
SVVYGLR and VEGF should be the chosen candidates as they both selectively enhance
endothelial cell adhesion and their soluble forms have been shown to promote formation of
capillary-like structures in vitro and blood vessels in vivo.23,24,27
5.7. Conclusions
This study reveals that low-fouling CMD layers constitute a good template to develop
bioactive surfaces: their multivalence allows the immobilization of active biological ligands
while their high resistance to unspecific adhesion allows and promotes specific cell receptor-
ligand interactions. Bioactive micro-arrays were made by covalent grafting of biologically
active molecules specific for endothelial cells combined with RGD peptides. Arrays
exposition to human endothelial cells and human fibroblasts revealed that fibroblasts were not
affected by spot composition, whereas endothelial cells were.
While RGD initiated cell adhesion, the co-immobilization of RGD with the
SVVYGLR sequence or with VEGF selectively enhanced endothelial cell adhesion, whereas
the co-immobilization with either REDV or SVVYGLR induced a reduction in endothelial
cell spreading. RGD combination with any of the bioactive molecules tested here resulted in a
loss of stress fibers and rearrangement of focal adhesions, synonymous of a reduced strength
of attachment, but probably associated with cell migration.
Further studies of bioactive coatings exposing SVVYGLR or VEGF will be done in
regard to their ability to support invasion of a scaffold used for the (re)construction of a
vascularized tissue. These bioactive micro-arrays are also an excellent tool to investigate cell
patterning in co-culture assays.
5.8. Acknowledgment
The authors thank Dr. Charles J. Doillon for providing fibroblasts. This work was
supported by the Canadian Foundation for Innovation through an On-going New
Opportunities Fund (Project No. 7918), NSERC through a Discovery Grant (Project No.
250296), and by the Université de Sherbrooke.
106
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Conclusions générales et perspectives
L’objectif de ce travail de recherche était de développer et de caractériser une surface
bioactive modèle interagissant spécifiquement avec les cellules endothéliales. Il s’agissait
d’optimiser un revêtement anti-adhérent fait de carboxy-méthyl-dextran (CMD) résistant à
l’adsorption non-spécifique de protéines et à l’adhésion cellulaire. Ensuite, ce projet avait
comme objectif d’immobiliser sur cette surface des molécules bioactives spécifiques pour les
cellules endothéliales et d’étudier leur comportement face à ces surfaces. L’ensemble des
étapes du développement de la surface bioactive (caractérisation et optimisation des surfaces
de CMD, comportement cellulaire face aux surfaces bioactives) devait être réalisé par
l’utilisation de puces de polymères et de molécules bioactives afin de réaliser des analyses à
haut niveau de criblage.
Développement d’une surface de CMD anti-adhérente
Les résultats obtenus et les analyses présentées dans les Chapitres 3 et 4 montrent que
les caractéristiques du CMD et les conditions d’immobilisation employées influencent les
propriétés physico-chimiques et la structure des couches de CMD formées et ainsi, leur
capacité à prévenir toute interaction non-spécifique. La concentration d’agents de couplage et
le degré de carboxylation du polymère déterminent la densité de points de liaison avec le
substrat. La concentration de NaCl régule la conformation de la chaîne en solution ainsi que
l’extension de la couche immobilisée, et enfin, le poids moléculaire du CMD influencerait la
capacité des molécules à se réarranger et donc la régularité ou l’homogénéité de la surface.
Ces travaux démontrent que les couches de CMD résistent mieux à l’adsorption non-
spécifique de protéines si celles-ci sont assez denses et épaisses pour cacher le substrat sous-
jacent. De plus, il est suggéré qu’une couche de CMD homogène et faiblement chargée résiste
à l’adhésion cellulaire. Une condition optimale répondant à ces critères a été mise au point et
validée : cette couche de CMD résiste aussi bien à l’adhésion cellulaire que le poly(éthylène
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glycol) (PEG), polymère le plus couramment utilisé pour prévenir les interactions non-
spécifiques.
Afin d’étayer, voire confirmer les hypothèses émises au sujet de l’effet des paramètres
d’immobilisation sur la structure des couches de CMD (Figure 3.7 du Chapitre 3), différentes
analyses supplémentaires pourraient être conduites. Une analyse individuelle par QCM des
différentes couches fournirait des données sur les propriétés viscoélastiques de chacune ; cette
étude est en cours dans notre Équipe. Elle pourrait permettre de déterminer la contribution des
"trains" et des "boucles" dans la structure de la couche de CMD : une couche ayant une forte
densité de trains à la surface serait plus rigide qu’une couche avec une forte densité de
boucles ; une couche plus dense serait aussi plus rigide qu’une couche diffuse. Des premières
images AFM en milieu liquide présentées à l’Annexe B permettent de soutenir l’hypothèse
selon laquelle une couche de CMD de haut poids moléculaire formerait une couche
hétérogène, présentant des irrégularités de surface. Il serait intéressant de poursuivre cette
caractérisation par imagerie en milieu sec et en milieu liquide. Enfin, l’analyse de la
composition du mélange de protéines adsorbées sur les différentes couches de CMD
(protéines de différentes tailles et charges) pourrait fournir des informations sur le mécanisme
d’adsorption (adsorption par diffusion, par attraction électrostatique) et ainsi des informations
supplémentaires sur les propriétés de chaque couche de CMD. Par contre, peu de techniques
permettent ce type d’analyse avec un niveau de détection permettant d’identifier les protéines
adsorbées à partir un mélange complexe. Par exemple, l’identité des protéines adsorbées sur
les puces de CMD pourrait être connue soit par marquage immunochimique, soit par
spectroscopie de masse possédant une résolution spatiale adéquate (ex. : TOF-SIMS ou
MALDI-MS).
Surfaces bioactives spécifiques pour les cellules endothéliales
Le travail présenté dans le Chapitre 5 indique que la couche de CMD anti-adhérente
permet l’immobilisation covalente et la reconnaissance spécifique par les récepteurs
cellulaires de molécules bioactives. La réponse cellulaire observée est directement et
spécifiquement liée aux molécules immobilisées. Ainsi, le modèle de surface bioactive “CMD
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+ molécule(s) bioactive(s)” permet d’une part l’étude des interactions spécifiques entre la
cellule et la surface et d’autre part, le contrôle de la réponse cellulaire.
L’immobilisation de molécules dites spécifiques pour les cellules endothéliales,
REDV, SVVYGLR et VEGF, n’induit pas l’adhésion sélective de ces cellules puisque la
séquence peptidique RGD, aussi reconnue par les fibroblastes, est nécessaire pour initier
l’attachement des cellules. Cependant, les molécules spécifiques utilisées induisent un effet
sur l’organisation des adhésions et du cytosquelette uniquement pour les cellules
endothéliales. Cette réorganisation de la structure de la cellule leur confèrerait un phénotype
de migration.
Afin de compléter l’étude de l’effet des surfaces bioactives présentées sur le
comportement des cellules endothéliales, il serait intéressant d’identifier les récepteurs
cellulaires impliqués dans la réponse, par marquage immunochimique des intégrines α4β1 et
α9β1, par exemple. D’autre part, il faudrait vérifier que les changements induits dans la
morphologie cellulaire sont associés à une plus forte mobilité. Ces travaux sont maintenant
entamés par un autre étudiant au doctorat de notre Équipe de recherche.
La nature, la densité des molécules immobilisées mais aussi leur distribution et
notamment la formation de “clusters”, ou agglomération de ligands, influencent le
comportement cellulaire. Aussi, l’immobilisation d’un mélange de CMDs de différents degrés
de carboxylation (ratios de CMD-50:CMD-25 de 1:10, 1:5, …) pourrait affecter
l’immobilisation des molécules et ainsi la réponse cellulaire.
Enfin, SVVYGLR et VEGF étant des molécules pro-angiogènes, il est envisageable
d’immobiliser ces molécules (en combinaison avec RGD) dans des structures 3D et
d’observer leur influence sur la formation de structures de type capillaire.
Puces de polymères et de molécules bioactives
Pour chaque étape du projet de recherche présenté, la technique des puces a été
développée puis utilisée. L’utilisation des puces de polymères a rendu possible l’étude
simultanée de l’effet de cinq paramètres sur les propriétés physico-chimiques des couches et
leur caractère anti-adhérent. De plus, le robot autorise la fabrication de puces selon des
modèles variés adaptés à différentes techniques d’analyse : spectroscopie des photoélectrons-
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X (XPS), microscopie à force atomique (AFM, imagerie et mesure de force) et résonance des
plasmons de surface (SPR). Ainsi, les puces de polymères permettent une étude rapide et
comparative des propriétés de différentes surfaces, avec plusieurs techniques, et les résultats
analytiques obtenus avec les puces sont similaires à ceux obtenus pour des surfaces uniformes
classiques. Enfin, l’exposition des puces de polymères à des cellules confluentes constitue un
test sélectif fort et a permis d’observer des phénomènes d’invasion qui peuvent être non-
visibles sur des surfaces uniformes (i.e., recouvertes entièrement de la couche anti-adhérente).
Les puces de molécules bioactives ont été fabriquées pour l’étude du comportement
cellulaire relativement à des surfaces de composition moléculaire variée et présentes sur un
même substrat. Pour chaque puce, un test ou un marquage différent peut être réalisé et observé
soit au microscope, soit avec un lecteur de puces, permettant ainsi une comparaison directe.
De plus, l’utilisation de puces par rapport à des surfaces classiques nécessite une plus faible
quantité de réactifs, souvent coûteux, et de cellules pour chaque test ou marquage.
De façon générale, une surface modèle “CMD + molécule(s) bioactive(s)” sous forme
de puce peut être adaptée à d’autres systèmes pour l’identification de surfaces induisant
l’adhésion sélective d’un type cellulaire, la différentiation de cellules souches dirigée vers une
lignée donnée, ou le maintien d’un phénotype différencié in vitro, par exemple. De plus, le
robot permet de déposer des solutions de compositions variées à des positions déterminées sur
la surface de CMD anti-adhérente et ainsi de contrôler la position des cellules et leur
comportement. De cette façon, la formation de lignes ou de gradients de concentration dirige
la migration cellulaire, alors que l’apposition de “spots” favorisant l’adhésion de types
cellulaires différents permet la formation d’interactions cellulaires hétérotypiques. En effet, le
comportement cellulaire, notamment l’acquisition et le maintien d’un phénotype différencié,
dépend non seulement des interactions cellule-matrice (protéines et peptides de la matrice,
facteurs de croissance immobilisés), mais aussi des adhésions inter-cellulaires.
L’immobilisation à des positions données de différentes solutions de molécules bioactives, ou
“patterning” de la surface, permettrait donc de contrôler localement les interactions cellule-
substrat et cellule-cellule et d’étudier leurs effets respectifs sur le comportement cellulaire.
Ces informations sont essentielles pour la conception de surfaces induisant le développement
de tissus fonctionnels dans un environnement artificiel. Enfin, il est à souligner que les puces
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bioactives peuvent aussi être utilisées comme outils de diagnostic clinique et cette technologie
a fait l’objet d’un brevet PCT déposé en juin 2007 dont je suis l’une des co-inventrices.
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Annexe A
Synthèse du carboxy-méthyl-dextran (CMD)
Principe
Figure A.1: Principe de la réaction de carboxy-méthylation du dextran.
Protocole
Préparer une solution de NaOH 2M dans de l’eau Milli-Q.
Dissoudre 5g de dextran dans 20mL de cette solution sous agitation.
Ajouter de l’acide bromoacétique 98% : 0.44g, 1.77g ou 3.54g pour une concentration finale de 0.125M, 0.5M ou 1M, respectivement. L’acide bromoacétique est chauffé pour le dissoudre.
HOOC-CH2
CH2-COOH
Br-CH2-COOH
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CH2-COOH *
†
Compléter le volume à 25mL.
Laisser réagir toute la nuit, sous agitation.
Placer la solution dans une membrane à dialyse avec un « molecular cut-off » de 8000 rincée à l’eau Milli-Q.
Faire une première dialyse contre de l’eau Milli-Q, puis contre une solution de HCl 1M et enfin une dernière fois contre de l’eau Milli-Q.
Répartir la solution de CMD dans un grand plat de petri, congeler (de préférence à –80°C) et lyophiliser.
Détermination du degré de carboxylation par analyse de spectres de résonance magnétique nucléaire (RMN)
Un spectre RMN-proton dans l’eau deutérée (D2O) est réalisé pour chaque CMD synthétisé afin de déterminer le degré de carboxylation.
Figure A.2: Représentation schématique d’un monomère carboxylé de la chaîne de CMD.
Chaque unité glucosidique de la chaîne de CMD comprend un hydrogène anomérique (H†) et chaque groupe carboxy-méthyl ajouté à la chaîne comprend deux atomes d’hydrogène (H*) qui lui sont propres (Fig. A.2).
L’intensité des pics correspondant aux 2 atomes d’hydrogène H* présent dans chaque monomère carboxylé est divisée par 2 et est comparée à l’intensité de l’ensemble des pics correspondant à l’hydrogène anomérique H† présent dans chaque monomère (Fig. A.3 et A.4). Plusieurs pics correspondent à l’hydrogène anomérique et aux hydrogènes de carboxylation car chaque environnement (monomère carboxylé ou non, site de carboxylation) induit un déplacement différent.
Les CMDs obtenus avec un ajout de 0.125M, 0.5M ou 1M d’acide bromoacétique ont des degrés de carboxylation proches de 5%, 25% et 50%, respectivement.
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* †
†
Exemples de spectres
Figure A.3: Spectre H1-RMN dans D2O du dextran non modifié. †: hydrogène anomérique.
Figure A.4: Spectre H1-RMN dans D2O du CMD obtenu avec 1M d’acide bromoacétique. †: hydrogène anomérique, *: hydrogènes du groupe carboxyméthyl.
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Annexe B
Imagerie par microscopie à force atomique (AFM)
des puces de CMD hydratées
Protocole
Les images AFM des spots de CMD sont acquises en milieu liquide en mode contact, avec des cantileviers de Si3N4 ayant une pointe pyramidale intégrée et une constante de ressort de 0.06 N/m (Model DNP-S, Veeco NanoProbe Tips). Les puces de CMD sont immergées dans du PBS 150 mM à température ambiante, une heure avant l’imagerie.
Exemples d’images obtenues
Figure B.1: Image AFM de 1µm² de la surface HApp prise dans le PBS 150mM.
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Figure B.2: Images AFM de 1µm² prises dans le PBS 150mM des “spots” de CMD de 70kDa immobilisés selon les conditions #5 (a) et #9 (b).
Figure B.3: Images AFM de 1µm² prises dans le PBS 150mM des “spots” de CMD de 500kDa immobilisés selon les conditions #2 (a) et #4 (b).
a b
a b
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Annexe C
Marquages immunocytochimiques
Remarques générales
Les solutions de BSA 2% w/v, de formaldéhyde 3.7% v/v et de triton 0.5% v/v sont préparées avec du PBS 150mM pH 7.4, conservées à 4°C et utilisées pendant une période maximale d’une semaine.
Les anticorps sont dilués dans une solution de PBS/BSA 2%.
Les solutions mères et les dilutions d’anticorps ou de réactifs sont gardées dans la glace pendant le marquage et à l’obscurité si nécessaire.
Actine et vinculine
Ce marquage permet d’observer le cytosquelette d’actine et les adhésions focales (comprenant la vinculine) de la cellule.
Retirer le milieu de culture et rincer 3 fois les cellules avec du PBS.
Fixer les cellules au formaldéhyde 3.7% pendant 15 min à température ambiante.
Rincer avec du PBS.
Perméabiliser les cellules avec une solution de triton 0.5% pendant 5 min à température ambiante.
Rincer avec du PBS.
Bloquer avec une solution de BSA 2% pendant 30 min à température ambiante.
Rincer avec du PBS.
Incuber les cellules avec l’anticorps primaire de souris anti-vinculine dilué au 1/25 (Sigma, cat. # V4505).
Rincer avec du PBS.
Incuber avec un mélange comprenant l’anticorps secondaire chèvre anti-souris couplé à l’Alexa fluor 488 dilué au 1/1000 (Invitrogen, cat. # A11001), la phalloidine-TRITC diluée au 1/300 ( Sigma, cat. # P1951) et le Hoechst au 1/10,000 (Sigma, cat. # B2883) pendant une heure à température ambiante et à l’abri de la lumière.
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Rincer 3 fois avec du PBS, puis de l’eau milliQ pour retirer les cristaux de sel et finalement avec du PBS.
Monter entre lame et lamelle en utilisant du PBS pur ou dilué avec du glycérol (50/50 v/v).
Fibronectine
Ce marquage permet d’observer la fibronectine cellulaire humaine, c’est à dire d’examiner spécifiquement la fibronectine fabriquée et déposée par des cellules humaines en culture.
Retirer le milieu de culture et rincer 3 fois les cellules avec du PBS.
Fixer les cellules au formaldéhyde 3.7% pendant 15 min à température ambiante.
Rincer avec du PBS.
Perméabiliser les cellules avec une solution de triton 0.5% pendant 5 min à température ambiante.
Rincer avec du PBS.
Bloquer avec une solution de sérum de chèvre diluée au 1/20 dans du PBS (Sigma, cat. # G9023) pendant 30 min à température ambiante.
Rincer avec du PBS.
Incuber les cellules avec l’anticorps primaire de souris anti-fibronectine humaine dilué au 1/400 (Sigma, cat. # F0916).
Rincer avec du PBS.
Incuber avec un mélange comprenant l’anticorps secondaire chèvre anti-souris couplé à l’Alexa fluor 488 dilué au 1/1000 (Invitrogen, cat. # A11001) et le Hoechst au 1/10,000 (Sigma, cat. # B2883) pendant une heure à température ambiante et à l’abri de la lumière.
Rincer 3 fois avec du PBS, puis de l’eau milliQ pour retirer les cristaux de sel et finalement avec du PBS.
Monter entre lame et lamelle en utilisant du PBS pur ou dilué avec du glycérol (50/50 v/v).
VE-cadhérine
La vascular endothelial (VE)-cadhérine est une molécule d’adhésion intercellulaire spécifique aux cellules endothéliales. Son marquage permet donc d’identifier les cellules endothéliales différenciées.
Retirer le milieu de culture et rincer 3 fois les cellules avec du PBS.
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Fixer les cellules au formaldéhyde 3.7% pendant 15 min à température ambiante.
Rincer avec du PBS.
Bloquer avec une solution de sérum de chèvre diluée au 1/20 dans du PBS (Sigma, cat. # G9023) pendant 30 min à température ambiante.
Rincer avec du PBS.
Incuber les cellules avec l’anticorps primaire de souris anti-VE-cadhérine dilué au 1/25 (R&D, cat. # MAB9381).
Rincer avec du PBS.
Incuber avec un mélange comprenant l’anticorps secondaire chèvre anti-souris couplé à l’Alexa fluor 488 dilué au 1/1000 (Invitrogen, cat. # A11001) et le Hoechst au 1/10,000 (Sigma, cat. # B2883) pendant une heure à température ambiante et à l’abri de la lumière.
Rincer 3 fois avec du PBS, puis de l’eau milliQ pour retirer les cristaux de sel et finalement avec du PBS.
Monter entre lame et lamelle en utilisant du PBS pur ou dilué avec du glycérol (50/50 v/v).
von Willebrand factor (vWF)
Le vWF est un facteur thrombogène secrété par les cellules endothéliales. Il est retrouvé dans les corps de Weibel-Palade dans le cytoplasme de la cellule. Tout comme pour la VE-cadhérine, ce marquage permet d’identifier les cellules endothéliales différenciées.
Retirer le milieu de culture et rincer 3 fois les cellules avec du PBS.
Fixer les cellules au formaldéhyde 3.7% pendant 15 min à température ambiante.
Rincer avec du PBS.
Perméabiliser les cellules avec une solution de triton 0.5% pendant 5 min à température ambiante.
Rincer avec du PBS.
Bloquer avec une solution de sérum de lapin diluée au 1/20 dans du PBS (Biodesign, cat. # N66010R) pendant 30 min à température ambiante.
Rincer avec du PBS.
Incuber les cellules avec l’anticorps primaire de chèvre anti-vWF dilué au 1/500 (Haematologic Technologies Inc., cat. # PAHVWF-G).
Rincer avec du PBS.
Incuber avec un mélange comprenant l’anticorps secondaire lapin anti-chèvre couplé au FITC dilué au 1/100 (Chemicon, # AP106F) et le Hoechst au 1/10,000 (Sigma, cat. # B2883) pendant une heure à température ambiante et à l’abri de la lumière.
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Rincer 3 fois avec du PBS, puis de l’eau milliQ pour retirer les cristaux de sel et finalement avec du PBS.
Monter entre lame et lamelle en utilisant du PBS pur ou dilué avec du glycérol (50/50 v/v).
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