Pierre-Yves Collart Dutilleul

Cunin

Keywords:

Adhésion, croissance et différenciation de cellules souches pulpaires sur silicium

Résumé en français:

Le silicium poreux est un biomatériau prometteur pour l'ingénierie tissulaire car il est non toxique et biorésorbable. Des modifications de surface permettent de controler sa vitesse de dégradation et peuvent favoriser l'adhésion cellulaire. Les cellules souches de la pulpe dentaire (DPSC) sont des cellules souches mésenchymateuses retrouvées dans la pulpe dentaire, à l'intérieur des dents, et constituent une source accessible de cellules souches. Regrouper les capacités de prolifération et différenciation des DPSC avec les propriétés morphologiques et biochimiques du pSi représente une approche intéressante pour des applications thérapeutiques de médecine régénératrice. Dans cette thèse, nous avons étudié le comportement de DPSC humaines sur des supports de pSi, avec des pores variant de quelques nanomètres à plusieurs centaines de nanomètres. Nous avons travaillé sur différentes fonctionalisations chimiques afin d'optimiser l'adhésion cellulaire et de stabiliser le matériau: oxydation thermique, silanisation et hydrosilylation. L'adhésion, la prolifération et la différenciation osseuse ont été évaluées par microscopie à fluorescence, microscopie électronique à balayage, activité enzymatique, tests de prolifération (activité mitotique), immunofluorescence et spectroscopie FTIR. Le pSi avec des pores de 30 à 40 nm de diamètre s'est révélé être le plus approprié pour l'adhésion, la prolifération cellulaire et la différenciation ostéoblastique. De plus, la structure nanométrique et le relargage d'acide silicique par le pSi a démontré un effet positif sur l'induction osseuse et la formation d'une matrice minéralisée. Le pSi est donc apparu comme un matériau prometteur pour l'adhésion de cellules souches mésenchymateuses, que ce soit pour une transplantation immédiate in vivo ou pour expansion et différenciation in vitro.

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Abstract in english:

Porous silicon (pSi) is a promising biomaterial for tissue engineering as it is both non-toxic and bioresorbable. Moreover, surface modification can offer control over the degradation rate of pSi and can also promote cell adhesion. Dental pulp stem cells (DPSC) are mesenchymal stem cells found within the teeth and constitute a readily source of stem cells. Coupling the good proliferation and differentiation capacities of DPSC with the textural and chemical properties of the pSi substrates provides an interesting approach for therapeutic use. In this thesis, the behavior of human DPSC is analyzed on pSi substrates presenting pore of various sizes, from few to hundreds nanometers. We investigated different chemical surface treatments, in order to enhance cell adhesion and stabilize the material: thermal oxidation, silanization and hydrosilylation. DPSC adhesion, proliferation and further osteodifferentiation were followed for up to 3 weeks by fluorescence microscopy, scanning electron microscopy (SEM), enzymatic activity assay, BrdU assay for mitotic activity, immunostaining and FTIR spectroscopy. Porous Silicon with pore size ranging from 30 to 40 nm was found to offer the best adhesion, the fastest growth rate for DPSC and the highest osteoinductive effect. Moreover, the pSi nanostructure and the release of silicic acid had a positive effect on precursor cells osteodifferentiation and mineralized matrix formation. Porous silicon appeared to be an appropriate biomaterial for mesenchymal stem cells adhesion and immediate in vivo transplantation, or for long term in vitro culture, for stem cells proliferation and differentiation.

Table of contents

Chapter 1 Porous Silicon Scaffolds for Stem Cells

Culture and Bone Tissue Engineering

Numerous studies have been carried out to verify if stem cells could become a source of stable differentiated cells capable of inducing tissue formation, including mineralized hard tissue. Stem cells for hard tissue formation has considerably increased attention of researchers as these cells can be an interesting source of stable differentiated cells, able to induce bone formation ie secretion of an organic matrix and hydroxyapatite crystal growth. Stem cells of various origins are able to produce in vitro calcified nodules and to form a mineralized tissue after transplantation in vivo. The property to form mineralized tissue is found in normal calcification system and also in pathological conditions and involves many cell types [START_REF] Huitema | What triggers cell-mediated mineralization[END_REF]. Supporting scaffolds for bone tissue engineering are often required to lead cell implantation, growth and differentiation (D' [START_REF] Aquino | Human postnatal dental pulp cells co-differentiate into osteoblasts and endotheliocytes: a pivotal synergy leading to adult bone tissue formation[END_REF]Graziano et al., 2008). These scaffolds can be inductive or instructive to direct stem cells down specific lineage pathways and increase the therapeutic effect. Thus, directed stem cells have the potential for rapid bone regeneration and repair. Since the advent of tissue engineering, bone has received particular interest as it is one of the tissue with most regenerative abilities in the human body and because of it apparent simplicity.

Porous silicon (pSi) is a promising biomaterial for bone tissue engineering, as it is non-toxic, it is bioresorbable and it can be inductive for osteodifferentiation. It has the ability to degrade in aqueous solutions into non-toxic silicic acid and surface modifications offer control over the degradation rate of pSi [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF].

Moreover, this tunable, biocompatible and resorbable material has been reported for the growth of hydroxyapatite, inferring the possible bone implantibility of the material (Cahnam, 1995).

Stem cells for bone tissue engineering

The first cell-based strategy for bone repair was autologous connective tissue progenitors from the iliac crest. This autograft procedure was accomplished without any in vitro manipulation, and was thus considered as part of the surgical procedure (Conolly et al., 1991). However, bone graft is not always possible, as it injures the donor site, and the success of grafting is entirely dependent on the transfer of sufficient numbers of progenitor cells. In addition, this approach may be least applicable in situations where it is most needed, such as in the case of ageing or disease with a reduction of osteogenic progenitors. Stem cell therapy involves in vitro sorting and expansion, before cell grafting. Cells can be of autologous or allogenic origin.

Autologous grafts offer the optimal compatibility, but require available cells from the patients. Therefore, allogenic grafts with cells from alternative allogenic source are of great interest, as they could provide "ready for use" products for bone treatments. were able to differentiate into a number of mesenchymal cell types including osteoblasts, chondrocytes and adipocytes. These cells were called "mesenchymal stem cells (MSCs)" in reference to their high self-renewing properties, clonogenicity, and ability to form cartilage and bone (Figure 1.1). They were suggested to be responsible for the normal turnover and maintenance of adult mesenchymal tissues. These MSC have been defined to be plastic-adherent when maintained in standard culture conditions; they must express CD105, CD73 and CD90, and lack expression of CD45 and CD34 (hematopoietic markers) surface molecules; and they must also differentiate to osteoblasts, adipocytes and chondroblasts in vitro (Dominici et al., 2006). MSC's therapeutic potential has generated significant excitement in the field of regenerative medicine, as they can be found in various niches within human body. The ability of these cells to self-renew and differentiate into multiple tissues makes them an attractive cell source for cell-based regenerative therapies. MSCs have considerable potential for the treatment of musculoskeletal disorders owing to their expansion capacity, immunosuppressive properties and ability to differentiate into bone and cartilage (Ma, 2010). The most studied source of adult MSCs has been the bone marrow, as it was recognized early that its stroma contained stem cells capable of forming bone and cartilage. Later, other sources of adult stem cells have been described, such as adipose tissue, dental pulp, skeletal muscle, umbilical cord blood and others (Marolt et al., 2010).

Bone marrow mesenchymal stem cells (BMSCs)

BMSCs are commonly recovered from pure bone marrow aspirates and isolated according to their adherence and culture on tissue culture plastics. They can be also selected using specific cell surface markers, and then expanded in culture. The number

A B C

of stem cells varies between different patients. Age and marrow aspiration volume influence also the number of isolated stem cells. BMSCs can, however, be expanded in vitro to large numbers and keep their osteogenic potential even in older patients (Bianco et al., 2001).

1. 1.1.1.2 Adipose-derived stem cells (ADSCs) [START_REF] Zuk | Human adipose tissue is a source of multipotent stem cells[END_REF] have been the first to describe and characterise ADSCs from human lipoaspirates. They showed that a multi-lineage stem cell population could be isolated from the stromo-vascular fraction of adipose tissue. These ADSCs are able to differentiate into osteoblasts, can be recovered from raw lipoaspirates consecutive through liposuction under anaesthesia, and represent an important niche of adult stem cells. They have been extensively studied in vitro for their ability to form various tissues, including bone-like tissue.

1. 1.1.1.3 Dental pulp stem cells (DPSCs) Stem cells within the dental pulp have been first described by Gronthos et al. (2000). They have then been studied and characterised as an adult stem-cell population that possesses the properties of high proliferative potential, the capacity of self-renewal and multi-lineage differentiation. Dental pulp can be easily collected from adult (DPSC) or deciduous teeth (stem cells from human exfoliated deciduous teeth) after dental extraction, when teeth have to be removed (Figure 1.2). The most frequent case is the collection of normal human third molar extracted for orthodontic reasons (Gronthos et al., 2002). Although they show great capacities for cell therapy and tissue engineering, these particular cells create a delicate ethical debate, as they induce destruction of human embryos. 1.1.1.3 Neonatal stem cells from the umbilical cord Neonatal tissues constitute a stem cell source that can be accessed in a noninvasive and rapid manner during and just after birth. The main source is the umbilical cord, covered by an amniotic epithelium that protects a gelatinous matrix called "Wharton's jelly". Stem cells can be recovered during birth, in utero or ex utero, from both cord blood and cord matrix (Wharton's jelly) (Forraz and McGuckin, 2011). Cord blood can be collected using a collecting needle connected to an anticoagulantcontaining bag. Wharton's jelly is obtained by collecting the whole umbilical cord.

With the average umbilical cord measuring 50 cm, this source of MSCs might become more and more clinically relevant.

Stem cells osteodifferentiation and bone formation

Bone tissue is composed of matrix and cells: bone matrix is mainly built of type I collagen (90%), with non-collagenous proteins (such as osteocalcin, osteonectin, bone sialoproteins). These scarce non-collagenous proteins participate in matrix maturation and mineralization. Osteoblasts and osteoclasts are responsible for bone formation and bone resorption. With progenitor cells, they mediate the mechanism of bone remodelling. During bone formation phase, the osteoblasts are recruited from MSC in bone marrow, since osteoclasts are derived from haematopoietic cells. Thus, stem cell therapy for bone regeneration will focus on MSC for osteoblastic differentiation, since host cells will provide osteoclasts during vascularisation process.

Molecular control of osteogenesis

Bone repair consists in inflammation, intra membranous ossification, chondrogenesis, endochondral ossification and remodelling. Li et al. (2007) have described molecular events that regulate osteogenesis: they observed an upregulation of genes related to cell cycle and cell-to-cell signalling, confirming the fact that both cell This differentiation is induced through addition of various molecules and chemicals, such as dexamethasone, ascorbic acid and β-glycerophosphate. Osteoblast-specific genes expression and proteins demonstrate this in vitro differentiation. However, in vitro assays have limitations and there is always a need to verify osteoblast differentiation potential in vivo. Thus, it has been shown that osteoprogenitor cells can only form bone tissue when transplanted in vivo, while MSCs can form both bone and bone microenvironment supporting haematopoesis in vivo. Moreover, MSCs demonstrate self-renewal and maintenance of stemness capacity during in vivo implantation (Arvidson et al., 2011). 1.1.2.3 Stem cells potential for scaffold-based bone tissue engineering applications ESCs and adult (or neonatal) MSCs are the main types of stem cells usable for scaffold-based tissue engineering. However, ESCs have a limited applicability to human cell therapy due to ethical concerns, but also immunological incompatibilities, potential for malignant tumor growth and insufficient understanding and control over differentiation. Thus, adult MSCs seem to be the cell type of choice for scaffold-based therapeutic applications. MSC delivery to the site of bone tissue defects requires scaffolds adapted to cell culture and activity in endogenous osteogenesis, including correct biochemical and mechanical stimuli.

Stem cells and mineralization process

Biomineralization is a general process by which inorganic calcium phosphate crystals are produced in biological environments. For bone formation, deposition of calcium phosphate salts occurs on an extracellular organic matrix by means of a heterogeneous nucleation process, which constitutes the first stage of hard tissue mineralization (Ngankan et al., 2000).

Biomineralization process: Ions and nucleation sites

Bone biominerals are composed of complex structures, having precise architectural order over several length scales. They rely on their biological constituents for regulation of mineralization, forming an organic-inorganic hybrid structure. The formation mechanisms of biomineralization involving several possible pathways remain unclear. Biomineralization has been described to start with matrix vesicle-mediated mineral initiation. It also involves heterogeneous nucleation of mineral crystals with an important role played by non-collagenous glycoproteins that bind calcium and collagen.

Calcium phosphate crystals can be formed via stable pre-nucleation clusters, with aggregation into an amorphous precursor phase, then transformation of this phase into a crystal [START_REF] Cölfen | Biomineralization: A crystal-clear view[END_REF]. Nudelman et al. (2010) demonstrated that collagen forms the structural matrix and has an active role in the mineralization of apatite for bone formation. Indeed, collagen fibrils have positively charged regions, and these nanosized regions can be used for mineral infiltration and charge-charge attractions. This initial mineral deposition leads to a dense network of pre-nucleation clusters, to transformation into amorphous calcium phosphate and to oriented crystalline hydroxyapatite inside the fibrils (Figure 1.4, next page). Even though the biomineralization process has been mainly studied with in vitro models (and in vivo bone formation may follow different rules), observations in small animals and tooth formation suggest that at least part of the mineralization mechanism of the collagen model is relevant for in vivo systems (Beniash et al., 2009).

Collagen and calcium-phosphate deposition during osteodifferentiation

Calcium and inorganic phosphate are the two main constituents of hydroxyapatite, the bone mineral that give mechanical resistance to the organic matrix.

The Ca/P mass ratio in bone is 2.2, and the initial step of Ca-P crystal nucleation takes place within matrix vesicles that appears from the plasma membrane of osteogenic cells and migrates into the extracellular skeletal compartment. Phosphate accumulates inside these matrix vesicles by a transport system, followed by an influx of calcium ions. This process permits the formation of hydroxyapatite crystal, associated with organic matrix collagen fibrils. Collagen is synthesized and assembled by osteoprogenitors, following a classical pathway for a secreted protein. The collagen chains are synthesized as precursors, and undergo a series of maturation reaction such as glycosylation and hydroxylation in the endoplasmic reticulum and Golgi complex. They are then assembled to form a triple helix. After processing and assembly, this type I procollagen is secreted into the extracellular space, where it will mature and polymerize into normal fibrils. Thus, collagen, calcium and phosphate are provided to the extracellular space through osteoprogenitors. During osteodifferentiation MSCs will turn into progenitor cells, then osteoblasts, expressing type I collagen genes.

Porous silicon scaffolds

Porous silicon (pSi) is a nanostructured form of the chemical element silicon that has introduced nanoporous holes in its microstructure, rendering a large surface to volume ratio. It is a semiconductor material created by electrochemical techniques, having electrical conductivity to a degree between that of metal and that of glass (insulator). Porous silicon was accidentally discovered in the 50's as an undesired product of experimental methods for polishing and shaping silicon surfaces for use as semiconductors: under certain conditions the combination of acid etching and anodic current resulted in the formation of a thick black, red or brown film on the material's surface. This was regarded as a failure for the purposes of polishing until the 70's when a interests arose because the high surface area of porous silicon was found to be useful as a model of the crystalline silicon surface in spectroscopic studies and as a dielectric layer in capacitance-based chemical sensors [START_REF] Sailor | Porous Silicon in Practice[END_REF]. In the 90's, pSi was evaluated as biomaterial, especially as bone implantable material: the huge surface area of silicon exposed by the pores relative to the volume, coupled with the covalently bonded hydrogen on the surface, renders pSi soluble in water and aqueous solutions (Cahnam, 1995).

Porous silicon fabrication and degradation

Fabrication

Porous silicon is generated by etching crystalline silicon in aqueous or nonaqueous electrolytes containing hydrofluoric acid (HF). Under appropriate electrochemical conditions, the silicon wafer doesn't dissolve uniformly but instead, fine holes appears, propagating primarily in the <100> (x;y;z) direction in the wafer.

The most common method is through anodization with a platinum cathode and a silicon wafer anode, both immersed in a hydrofluoric acid electrolyte. A direct current, passed through the set-up, corrodes the anode wafer, producing a porous layer (Figure 1.5, next page). The pore diameter and depth can be controlled by HF concentration, with longer and narrower pores produced by concentrated HF solutions. Once the wafer is fully etched, the created pSi can be rinsed in ethanol then dried under nitrogen airflow. Pore width is classified into 3 sizes, according to the pore diameters: less than 2 nm as nanoporous, from 2 to 50 nm as mesoporous and larger than 50 nm as macroporous [START_REF] Sailor | Porous Silicon in Practice[END_REF].

Degradation in aqueous solutions

Due to the large surface area of Si-H on the surface and interior space of the material, pSi is soluble in water and aqueous solutions, dissolving along the mechanism:

Si + 4 H 2 O Si(OH) 4 + 2 H 2
Orthosilicic acid Si(OH) 4 is liberated from the surface of the decomposing pSi in addition to hydrogen gas. Silicic acid is the form of bioavailable silicon in human body, under physiological conditions. It is mainly found in foods rich in fiber and whole grains, with beer being one of the main sources of this element in the diet [START_REF] González-Muñoz | Role of beer as a possible protective factor in preventing Alzheimer's disease[END_REF]. Surface treatment resulting in accessible porosity with chemically stable inner surface is a key step to prevent rapid hydrolysis and degradation of the pSi in aqueous solutions.

The most common and simple surface treatment is oxidation, which can be performed by either ozone, aging, thermal or chemical treatments and creates Si-OH bounds on the surface. Amine-terminated modifications as silanization with aminopropyl trimethoxysilane or triethoxysilane improve pSi chemical stability [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF].

Thermal hydrosilylation can also be used to graft chemical species, such as dodecene, undecenoic acid or oligoethylene glycol (Alvarez et al., 2009).

Biological activity and biocompatibility

Silicon is an very abundant element, constituting the majority of mineral found on earth, as sand, quartz, agate, amethyst, rock crystal, flint, chalcedony, jasper and opal in combination with other elements. But despite this abundance, silicon is rarely found in biology, forming only 0.02% of the human body by weight. However, silicon plays an important role in connective tissue, such as bone and cartilage, for the formation of organic matrix (collagen and glycosaminoglycan). Silicon is a major ion of osteogenic cells and participates in the biochemistry of subcellular enzyme-containing structures, highlighting its metabolic role in connective tissue [START_REF] Carlisle | Silicon as a trace nutrient[END_REF].

Soluble silicic acid, the biologically active form of silicon, is mainly absorbed for man from silicic acid-containing drinking water, beverages made from phytolithscontaining plants (mainly beer) and some food (mainly banana and green bean). In western countries diet, beer and bananas form the main source of bioavailable silicon for men, while the main source for women are bananas and green beans [START_REF] Jugdaohsingh | Dietary silicon intake and absorption[END_REF].

For biomedical applications, porous silicon biocompatibility and immunogenicity have already been demonstrated under different conditions [START_REF] Ainslie | In vitro immunogenicity of silicon-based micro-and nanostructured surfaces[END_REF]. Furthermore, pSi-based scaffolds have been investigated for orthopaedic (Whitehead et al., 2008) and ophtalmic implants [START_REF] Cheng | Intravitreal properties of porous silicon photonic crystals: a potential self-reporting intraocular drug-delivery vehicle[END_REF], for controlling the adhesion and proliferation of different cell types [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF] (Alvarez et al., 2009[START_REF] Torres-Costa | Engineering of silicon surfaces at the micro-and nanoscales for cell adhesion and migration control[END_REF], and even as intravenously injectable particles for imaging tumors in vivo (Park et al., 2009) or retina pressure actuators [START_REF] Noval | Design and characterization of biofunctional magnetic porous silicon flakes[END_REF].

Influence of pore size, nanoroughness and chemical surface treatment

To mimic the in vivo cell behavior for in vitro culture, one should consider scaffold chemistry, mechanical properties and architecture (Yang et al., 2012). An extra-cellular matrix is composed of various proteins within a 3D network. Thus, a scaffold for tissue engineering should provide optimal mechanical properties and environment for cell culture. PSi can be designed to mimic the properties of bone tissue, and coatings can impart proper surface chemistry, adapting various pore size, porosity and nanoroughness. Indeed, nanotopography can affect protein adsorption, cytoskeletal arrangement, and cell adhesion and spreading.

Influence of pSi pore size on MSC adhesion

The effect of pore size and porosity on cell growth is of particular relevance, as it will direct the fabrication of a porous biomaterial. Additionally, its biocompatibility depends on its porosity and pore size. These pore dimensions can be precisely controlled and are highly tunable during pSi electrochemical anodization or etching. A variety of pore sizes can be produced: from micropores (< 2 nm), mesopores (2-50 nm) to macropores (> 50 nm) depending on the preparation conditions [START_REF] Sailor | Porous Silicon in Practice[END_REF].

Although microscale topography modulates cellular behavior in vitro, it is important to consider that cells in vivo make contact with nanoscale as well as microscale topographical features. Even if cells are typically tens of microns in diameter, the dimension of subcellular structures tends to the nanometer scale. Furthermore, extracellular supporting tissues present an intricate network of cues at the nanoscale composed of a complex mixture of nanometer-size (5-200 nm) pits, pores, protrusions, and fibers (Biggs et al., 2010). The main mean to evaluate the influence of pore size on MSCs adhesion and proliferation is the use of pSi gradients, allowing cells growth on surface with a pore range from few nanometers to few micrometers: this method involves electrochemical etching of crystalline Si in a hydrofluoric acid (HF) solution to fabricate pSi films displaying a gradient of pore sizes in an asymmetrical arrangement.

The size range of these pores can be controlled by modifying the applied current and the HF (Collins et al., 2012;Clements et al., 2012). Different cell types respond differently to topographical features, such as nanoscale pores and studies demonstrated that MSCs show better attachment to pSi with pore size of 20-30 nm in diameter (Wang et al., 2012). Pore size in this range is similar to the pore size of biological systems, such as pores in basement membranes, and corresponds to the size of biological pits, pores, protrusions and fibers (Figure 1.6). While producing pSi, pore depth can also vary.

However, cells are not sensitive to changes in pore depth of the range observed with various type of pSi [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF]. MSCs were shown to respond to both topographical and chemical characteristics. However, chemical surface treatment and grafting of proteins seem to have a higher impact than changes in pore size during short-term culture (Clements et al., 2012). PSi surface treatment is necessary not only to improve cell adhesion, but firstly to permit accessible and stable pores: the inner walls of the porous matrix must be protected from an excessive degradation in the aqueous cell culture environment without eliciting any undesirable effects on the cells. Moreover, surface modification by turning pSi from hydrophobic to hydrophilic promotes cell adhesion and growth. The most common and simple treatment is oxidation, which can be done by ozone, aging, or thermal or chemical treatment. PSi biocompatibility has been extensively studied with various cell lines cultured on chemically modified pSi samples. Few studies have been done with primary MSC cultures to determine an optimal surface treatment (Hajj- et al., 2011;Clements et al., 2012). Moreover, it has been shown that stem cells growth can be modulated by factors mimicking extracellular matrix, such as amino or RGD residues. Cell behavior is influenced by changing the integrin and focal adhesion assembly, leading to changes in cytoskeletal organization and cell mechanical properties [START_REF] Yim | Nanotopography-induced changes in focal adhesions, cytoskeletal organization, and mechanical properties of human mesenchymal stem cells[END_REF]. Thus, amine-terminated samples are expected to enhance MSC adhesion and proliferation as it does for other cell lines: silanisation with aminopropyl trimethoxysilane and triethoxysilane improves pSi stability and enhances cell adhesion in comparison to oxidized pSi. The chemistry greatly improves the stability of pSi in contact with cultured cells while allowing cell coverage levels comparable to standard culture preparations on tissue culture polystyrene [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF]. Thermal hydrosilylation is also used to graft chemical species to generate a stable substrate for cell adhesion and proliferation. Among these species are: dodecene, undecylenic acid and oligoethylene glycol (Alvarez et al., 2009).

Hassan

In addition to chemical surface treatment, a coating of proteins can be realized onto the pSi surface to mimic cell adhesion and relative responses in an in vivo condition with surface presenting biocompatible contact sites in a nanoscale dimension (topographic scale of adhesive proteins such as collagen fibers, laminine and fibronectin is present on the order of nanometers). The cell adhesion process depends upon cell surface receptors interacting with proteins belonging to the extracellular matrix. Several proteins have been considered for pSi coating: collagen, fetal calf serum and the tripeptide arginine-glycine-aspartate (RGD). Collagen is a component of the extracellular matrix and can be used to coat surfaces to stimulate cell adhesion, as well as foetal calf serum, which contains several cell adhesive proteins. Collagen and fetal calf serum have been successfully used to promote cell attachment and cell spreading [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF]. The tripeptide RGD is an integrin-binding domain present within many matrix proteins including fibronectin and vitronectin. Clements et al. (2012) showed that RGD immobilized on pSi is still a bioactive peptide and that cells adhere more when the concentration of RGD grafted increases.

Composite materials, such as poly(ε-caprolactone) / pSi, are also used for cell culture and tissue engineering. Highly porous scaffolds composed of mesoporous Si and commonly-used biopolymers, such as polycaprolactone, are electrically conductive composite materials exhibiting an interfacial behavior that promotes calcification and sustains the in vitro stability and proliferation of connective tissue cells in implant surfaces. Poly(ε-caprolactone) has been shown to have biocompatible and biodegradable character that is widely exploited in guided bone regeneration and drug delivery, and the resorbable pSi component sensitively accelerates the necessary calcification process. This type of composite presents a relevant material to orthopedic tissue engineering with tunable properties as mechanical strength, controlled release of molecules, and rate of resorption of the composite to the host surroundings (Coffer et al., 2005). This pSi-containing polymer scaffold enhances calcification, can be considered nontoxic to cells, and support the proliferation, viability, attachment, and differentiation of bone precursor cells, suggesting that materials of this type of composition have potential merit as biomaterials (Whitehead et al., 2008).

Another type of pSi membrane has also been described using porous nanocrystalline Si, which is a new type of Si nanomaterial, with potential applications in cell culture and tissue engineering. This pSi material is a 15 nm thick nanoporous membrane made with scalable Si manufacturing, permeable to small solutes. This nanoporous and ultrathin material can serve as a cell culture substrate, and is of interest to separate tissues during surgical grafting, such as bone tissue, bone substitute, periosteum or epithelium (Agrawal et al., 2010).

Roughness and surface energy

Several molecular mechanisms by which cells perceive and respond to their surrounding environments have been studied, including integrin-ligand interactions, contact surface hydrophilicity and extracellular matrix topography and architecture. The in vitro control of the cellular environment is a challenge for manipulating both cell development and scaffold substitute function in the field of bioengineering. However, the differences existing between natural and experimental surface characteristics may confer some of the discrepancies reported between in vivo and in vitro results. Several resembling the in vivo conditions (Yang et al., 2010). However, the nanotopography plays a relevant role in the final wettability properties of surfaces, where hydrophobic surfaces prevent cell adhesion. PSi can be arranged with nanoscale patterns but freshly formed nanostructured pSi shows hydrophobic properties. Surface oxidation turns hydrophobic surfaces into hydrophilic: freshly etched pSi is hydrophobic and controlled ozone or thermal oxidation can drop the water contact angle to below 5°. This is attributed to the formation of a polar Si-OH capped surface after oxidation [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF]. 1.3.2.4 Influence of pore size, stiffness and surface treatment on osteodifferentiation Cell attachment on pSi substrate has been extensively studied, but few studies evaluated the influence of topography on cell differentiation. Recently, Wang et al. (2012) explored the influence of nanoscale surface topography on cell behavior. Using rat MSCs, they found that surface topography influences cell differentiation, but not proliferation. Osteogenesis was enhanced by porous topography with a ridge roughness lower than 10 nm with an increase in osteodifferentiation when pores size decreased.

Thus, osteodifferentiation was comparable between flat Si and pSi with 100-200 nm, and was clearly enhanced when pore sizes decreased to 10-30 nm. This impact of pore size on osteodifferentiation is of high importance for the development of scaffold materials that can stimulate stem cell differentiation into osteoblasts in absence of chemical treatment without compromising material properties. Dalby et al. (2007) already demonstrated the use of nanoscale disorder to stimulate human MSCs to produce bone mineral in vitro in the absence of osteogenic supplements. Their approach had similar efficiency to that of cells cultured with osteogenic media. Nanocues, such as pSi pores, may prove to be important in tissue-specific stem cell differentiation. These findings are supported by various observations of chemical stimulation, chemical patterning and material hardness having roles in stem cell differentiation [START_REF] Zemel | Optimal matrix rigidity for stress-fibre polarization in stem cells[END_REF]. The effect of surface treatment of pSi on stem cells has been investigated mainly for cell adhesion, however, several studies have shown that the immobilization of RGD on the substrate not only enhanced cell adhesion but also modulated the intercellular mechanisms of cell proliferation and differentiation [START_REF] Hu | Porous polymer scaffolds surface-modified with arginine-glycine-aspartic acid enhance bone cell attachment and differentiation in vitro[END_REF]Cavalcanti-Adam et al., 2007;Lagunas et al., 2011). Thus, RGD peptide coated surfaces enhance osteogenic differentiation when present in sufficient concentration (Frith et al., 2011).

At last, the mechanical properties of the underlying substratum have previously been shown to affect a number of cellular processes including locomotion, proliferation and differentiation. Microenvironments appear important in stem cell lineage specification, but can be difficult to control with pSi. MSCs are shown to specify lineage and commit to phenotypes with extreme sensitivity to tissue-level elasticity. Soft matrices that mimic brain are neurogenic, stiffer matrices that mimic muscle are myogenic, and comparatively rigid matrices that mimic collagenous bone, like pSi, prove to be osteogenic. After several weeks in culture, cells commit to the lineage specified by matrix elasticity (Engler et al., 2006).

Growth factors delivery and Si effects on osteodifferentiation

Stem cells grown in biomaterials with bone growth factors constitute a source of transplantable material. Indeed, mesenchymal stem cells differentiation into elements of the skeleton can be enhanced by the addition of bone growth factors. PSi is a potential carrier scaffold to deliver growth factors to the site of defect, where regeneration is needed.

Moreover, controlled release of ionic products from PSi dissolution can stimulate and enhance bone-cell function: dissolved Si has been shown to enhance osteoblast differentiation (Gupta et al., 2010). Drug Administration for specific clinical cases (Bessa et al., 2008). Indeed pSi also exhibits a number of properties that make it an attractive material for controlled drug delivery applications. Its high surface area, coupled with the ability to control pore sizes allow pSi to be loaded with a range of bioactive species. Tunable properties of surface area, porous volume, and pore size allow controlling the amount and in vivo release rate of molecule payloads (Anglin et al., 2008). By promoting osteogenic development of MSCs, the association of pSi and BMPs offer interesting perspective for bone reparation and regeneration. et al., 1998;Park et al., 2009). The release of Si(OH) 4 due to the corrosion of pSi in plasma stimulates calcification and collagen growth, which could ideally accelerate new bone regeneration. Cell differentiation assays and ultrastructural measurements (TEM) have been carried out to explore the influence of Si on stem cell osteodifferentiation. It has been demonstrated that Si plays an important role in the expression of alkaline phosphatase, a specific biomarker for mesenchymal stem cell differentiation (Fan et al., 2012). Osteocalcin (OCN) and bone sialoprotein (BSP) (non-collagenic proteins produced by osteoblast and involved in Ca-P crystal formation) are significantly more expressed by stem cells cultured in Si-rich medium, confirming that Si enhances stem cell differentiation into osteoblast and that Si may directly be involved in OCN and BSP-mediated mechanisms (Gupta et al., 2010). Histological evaluation and histomorphometric quantification of none ingrowth and mineral apposition in vivo demonstrated the benefits to bone ingrowths and repair through incorporation of Si (Hing et al., 2006). These results show that pSi, by its degradation, plays a significant role in both stem cell osteodifferentiation and mineralization (see Figure 1.7). If it is clear that the dissolution of pSi in the form of Si(OH) 4 is a powerful mechanistic force behind the nucleation and growth of calcium phosphate, it also seems that Si surfaces provide nucleation sites for the formation of apatite deposits. This conclusion is based on mechanistic observations concerning surfaces formed of CaO-rich silica, noting that Si in these materials are responsible for directing the formation of apatite deposits in simulated plasma. Si intermediates as a nucleation site in directing the formation of calcium phosphate spherulites and growth of apatite on the scaffold (Whitehead et al., 2008). Bone cells attached onto the surface containing high Si content absorb significantly higher concentrations of calcium than cells attached onto the surface that contains low Si content. In conjunction with the absorption of high calcium concentration, attached bone marrow stem cells produce calcified nodules and mineralized extracellular matrix, indicating osteoblastic differentiation (El-Gahnnam et al., 2006). Thus, the mechanism of bone mineralization at the interface with bioactive surfaces is mainly cell mediated: a Si-rich phase provides a guided cell adhesion and tissue growth in vitro. The addition of Si to a scaffold affects surface topography, dissolution-reprecipitation rates and the bone apposition process (Porter, 2006).

Objectives of the investigation and future trends

In tissue engineering, developing synthetic grafts with sufficient combined mechanical strength, bioactivity and biodegradability for bone-tissue regeneration applications therefore remains an important challenge. A material that combines high bioactivity and porous structure (for good tissue growth) with high levels of mechanical strength (for load-bearing applications) will improve clinical outcomes and enable a significantly wider application of restorative materials [START_REF] Zreiqat | The incorporation of strontium and zinc into a calcium-silicon ceramic for bone tissue engineering[END_REF] The overall aim of the research presented here is to assess porous silicon as a potential bioactive scaffold, promoting both human stem cell adhesion and osteodifferentiation, for bone tissue engineering using primary human dental pulp stem cells (DPSC). A secondary aim is to further elucidate the influence of pore diameter (at the nanometer scale) on stem cells adhesion, proliferation and differentiation.

Therefore, the aims of this work can be presented as follow (each point outlines the themes for the following chapter):

Development and optimization of human dental pulp stem cells recovery, in order to have an efficient and accessible source of primary human stem cells.

Dental pulp stem cells adhesion and proliferation onto various porous silicon scaffolds, in order to determine the ideal porosity and functionalization of pSi.

Dental pulp stem cells growth and osteodifferentiation on porous silicon scaffold, to study the inductive effects of pSi for osteodifferentiation, due to the silicic acid released during pSi degradation and due to its porous structure. odontoblasts, along the inner surface of the dentin inside the pulp chamber, continue to deposit dentin matrix to form secondary dentin throughout life. In addition to secondary dentin, odontoblasts can form tertiary (reparative) dentin in response to several stimuli, such as mechanical, chemical, and/or bacterial stimulation. Even when odontoblasts have been damaged, the reparative dentin can be formed in the dental pulp to protect against further disruption of the pulp tissue. This reparative dentinogenesis has been thought to be mediated by newly generated odontoblasts arising from dental pulp tissue.

These findings have led to the speculation that odontogenic progenitor cells or stem cells may exist in dental pulp tissue [START_REF] Sonoyama | Multipotent Stem Cells in Dental Pulp[END_REF]. And in 2000, Dental Pulp

Stem Cells have been, for the first time, clearly identified, isolated and characterized (Gronthos et al., 2000). Thus, we can represent dental pulp as a soft connective tissue entrapped within the dental crown, and divided into four layers. The external layer is made up of odontoblasts producing dentin; the second layer is poor in cells and rich in collagen fibers; the third layer contains progenitor cells and undifferentiated cells, some of which are considered stem cells. From this layer, undifferentiated cells migrate to various districts where they can differentiate under different stimuli and make new differentiated cells and tissues. The innermost layer is the core of the pulp and comprises the vascular area and nerves (D' Aquino et al., 2009).

Dental pulp stem cells (DPSC)

A stem cell is defined as a cell that has the ability to continuously divide to [START_REF] Yamada | Promising cellbased therapy for bone regeneration using stem cells from deciduous teeth, dental pulp, and bone marrow[END_REF].

Dental tissue from human third molar represents an easily accessible and often discarded source for MSC harvesting. The first type of dental stem cell was isolated from the human pulp tissue and termed dental pulp stem cells (DPSC) (Gronthos et al., 2000). Subsequently, four more types of dental-MSC-like populations were identified: stem cells from exfoliated deciduous teeth (SHED) [START_REF] Miura | SHED: stem cells from human exfoliated deciduous teeth[END_REF], periodontal ligament stem cells (PDLSCs) [START_REF] Seo | Investigation of multipotent postnatal stem cells from human periodontal ligament[END_REF], stem cells from apical papilla (SCAP) [START_REF] Sonoyama | Mesenchymal stem cell-mediated functional tooth regeneration in swine[END_REF], and dental follicle precursor cells (DFPCs) [START_REF] Morsczeck | Isolation of precursor cells (PCs) from human dental follicle of wisdom teeth[END_REF]. Among them, all except SHED are from permanent teeth.

Stem cells that reside in dental pulp (DPSC) are a population of MSC, as they match the definition given by the Mesenchymal and Tissue Stem Cell Committee of the International Society for Cellular Therapy (Dominici et al., 2006). This committee proposed minimal criteria to define human MSC: first, cells must be plastic-adherent when maintained in standard culture conditions; second, they must express some specific surface molecules such as CD105, CD73 and CD90, and lack expression of Characterization revealed that DPSC possessed osteogenic, chondrogenic and adipogenic differentiation capacities. In addition, DPSCs were also found to undergo myogenic and neurogenic differentiation capacities in vitro, expressing respective gene markers and exhibiting neuronal-like cell morphologies. The plasticity and multipotential capability of DPSC can be explained by the fact that dental pulp is made of both ectodermic and mesenchymal components, containing neural crest-derived cells (D' Aquino et al., 2009). Concerning cell surface molecules, markers that have been shown positive for various markers, such as STRO-1, CD13, CD24, CD29, CD44, CD73, CD90, CD105, CD106, CD146, Oct4, Nanog and β2 integrin, and negative for CD14, CD45 and HLA-DR. The persistence of negativity for CD45 demonstrates that these cells are not derived from a hematopoietic source, although they are of mesenchymal origin (D' [START_REF] Aquino | Human postnatal dental pulp cells co-differentiate into osteoblasts and endotheliocytes: a pivotal synergy leading to adult bone tissue formation[END_REF]. Like all MSC, DPSC are also heterogeneous and the various markers listed here may be expressed differently by subpopulations of these stem cells (Huang et al., 2009). In 2005, Laino and co-workers described a selected subpopulation of CD34 + /CD45 -DPSC, which roughly represented 10% of dental pulp cells. These cells displayed a great capability of self-expanding and differentiating in pre-osteoblasts, able to self-maintain and renew for long time [START_REF] Laino | A new population of human adult dental pulp stem cells: a useful source of living autologous fibrous bone tissue (LAB)[END_REF]. Although MSC were described to be CD34(-), it seems that a subpopulation of DPSC expressed this cell surface antigen described for the most primitive stromal, other than hematopoietic, stem cells, and gradually lost after lineage committed progenitors differentiation [START_REF] Yoo | Endothelial progenitor cells from human dental pulp-derived iPS cells as a therapeutic target for ischemic vascular diseases[END_REF].

Potential applications of DPSC in regenerative medicine

Upon discovery of stem cells in the dental pulp, DPSC demonstrated their ability to regenerate a complex consisting of a mineralized matrix with odontoblasts and connective tissue containing blood vessels similar to that observed in normal human tooth (Gronthos et al., 2000). Nowadays, the use of DPSC concerns the repair and regeneration of liver tissue, heart tissue, bone, central nervous system, eye, and holds great potential in the field of regenerative medicine [START_REF] Graziano | Human CD34+ stem cells produce bone nodules in vivo[END_REF] (Gronthos et al., 2002). They even appear to have greater potential than the actual mesenchymal stem cells gold standard, the bone marrow-derived mesenchymal stem cells [START_REF] Alge | Donor-matched comparison of dental pulp stem cells and bone marrow-derived mesenchymal stem cells in a rat model[END_REF]. aminotransferase. Furthermore, the liver functions, observed by the levels of serum bilirubin and albumin, appeared to be improved following transplantation of DPSC [START_REF] Ikeda | Multipotent cells from the human third molar: feasibility of cellbased therapy for liver disease[END_REF]. These findings suggested that DPSC were one of the candidates for cell-based therapy to treat liver diseases and offered opportunities for developing therapies. These strategies of transplantation of DPSC could help halting malignant progression in hepatitis patients receiving antiviral treatment.

Myocardiac regeneration

DPSC, like bone marrow stem cells, can secrete vascular endothelial growth factor (VEGF) [START_REF] Shi | Comparison of human dental pulp and bone marrow stromal stem cells by cDNA microarray analysis[END_REF] and stem cell factor (SCF, corresponding to the cell surface receptor CD117) [START_REF] Gagari | Expression of stem cell factor and its receptor, c-kit, in human oral mesenchymal cells[END_REF]. These factors have proangiogenic, antiapoptotic, and cardio protective actions [START_REF] Tran-Hung | Role of human pulp fibroblasts in angiogenesis[END_REF]. Thus, in 2008, Gandia and co-workers investigated whether DPSC could be useful for cardiac repair, in a rat model of myocardial infarction. DPSC were injected intramyocardially in nude rats 7 days after induction of myocardial infarction by coronary artery ligation (Gandia et al., 2008). At 4 weeks, DPSC-treated animals showed an improvement in cardiac function, observed by left ventricular wall thickness and reduction in infarct size (Figure 2.3, next page). Moreover, angiogenesis was increased compared to control animals.

Thus, DPSC appeared to be a potential alternative therapeutics for cardiac repair after acute myocardial infarction. The woven bone tissue produced in vitro could even be remodeled into lamellar bone when transplanted in vivo (Graziano et al., 2008). A study on human subject was performed to assess the efficiency of DPSC to rebuild mandibular bone: DPSC seeded on a collagen sponge scaffold were grafted for bone tissue repair in patients requiring extraction of their third molars (wisdom teeth). Three months after autologous DPSC grafting, alveolar bone of patients had optimal vertical repair and complete restoration of periodontal tissue back to the second molars.

Histological observations highlighted the complete regeneration of bone at the injury site one year after grafting (d'Aquino et al., 2009). The stability and quality of the regenerated bone and vessel network was assessed 3 years after the grafting intervention: the regenerated tissue from the graft sites was composed of a fully compact bone with a higher matrix density than control human alveolar spongy bone from the same patient [START_REF] Giuliani | Three years after transplants in human mandibles, histological and in-line holotomography revealed that stem cells regenerated a compact rather than a spongy bone: biological and clinical implications[END_REF]. [START_REF] Sakai | Human dental pulp-derived stem cells promote locomotor recovery after complete transection of the rat spinal cord by multiple neuro-regenerative mechanisms[END_REF]. These works showed up that tissue regeneration after DPSC graft was related to DPSC proliferation and differentiation, and to enhancement of local progenitor cells recruitment and maturation.

Eye reparation

Limbal stem cells reside in the eye, between the cornea and the sclera. They are responsible of constant renewal of the corneal epithelium and can promote its regeneration when injured. In case of partial or total limbal stem cell deficiency, a conjunctival epithelium invasion can occur, leading to a significant loss of visual acuity.

DPSC have been shown able to rebuild the eye surface after total limbal stem cell deficiency in rabbits [START_REF] Monteiro | Human immature dental pulp stem cells share key characteristic features with limbal stem cells[END_REF]. Thus, it is suggested that DPSC share similar characteristics with limbal stem cells and might be used as a potential alternative source of cells for corneal reconstruction.

Aims of the study

Important knowledge about the characteristics of bone marrow stem cells has permitted to isolate putative stem cell populations from the dental pulp of human third molars (DPSC) and deciduous teeth (stem cells from human exfoliated deciduous teeth; SHED), which exhibit properties similar to those of bone marrow stem cells. DPSC were found to reside in a specific perivascular microenvironment, where they were quiescent and maintained their basic stem cell characteristics, including a self-renewal capacity and undifferentiated status [START_REF] Shi | Perivascular niche of postnatal mesenchymal stem cells in human bone marrow and dental pulp[END_REF]. Various procedures for the isolation, purified preparation, expansion, and tissue regeneration potential of DPSC have been described.

The work presented in this chapter is about development and optimization of human dental pulp stem cells recovery, in order to have an efficient and accessible source of primary human stem cells. We aimed to establish a clear and efficient way to recover primary cultures of human DPSC. These cells can be collected from dental pulp by means of a non-invasive practice that can be performed in the adult during life and in the young after surgical extraction of wisdom teeth. We wanted to develop an access to this source of MSC, for potential autologous or allogenic therapeutic applications. And as DPSC can be cryopreserved and stored for long periods without losing their multipotential differentiation ability [START_REF] Laino | A new population of human adult dental pulp stem cells: a useful source of living autologous fibrous bone tissue (LAB)[END_REF], we studied cell stemness after storage in liquid nitrogen.

Materials and methods

Patients

Pre operative step

We developed our protocol at the public hospital in Montpellier, in accordance to the local ethical committee (Comité de Protection des Personnes, Montpellier hospital, France) (see statement, in french, in Appendix). Once the protocol was approved, we submitted it to a private clinic, where oral surgeons are working. In public hospital or in the private clinic, human impacted third molar extracted for orthodontic reasons were recovered from healthy patients (15-30 years of age). Patients were considered as "healthy" when they were not known to have any chronic or acute infection or pathology. Pregnant women were not included in the protocol. The pre operative step also evaluated the feasibility to recover intact impacted teeth, by radiographic evaluation (see Figure 2.4, next page). Written informed consent was obtained from all the patients or their parents, after a discussion in which the project was explained and patients had the opportunity to ask questions about the program and to freely accept or refuse. The document signed by the patients (or parents) is shown in program conducted with their teeth.

Post operative step

After teeth collection, the refrigerated box was brought to the laboratory, where the dental pulp tissue was recovered, under sterile conditions. All the research work was then conducted. After 3 to 6 months, a short summary of the work was sent to the patients who asked for.

Cell extraction

Transport tubes were opened under a laminar flow hood. Teeth surfaces were cleaned using 2% chlorhexidine and cut around the cementum-enamel junction by using

A B C D E F
sterilized disc. The teeth were then broken into 2 pieces to reveal the pulp chamber. The pulp tissues were gently separated from the crown and root. The pulps were submitted to enzymatic digestion. This technique has been already widely described: extracellular matrix of the pulp tissue were digested in a solution of 3 mg/ml type I collagenase and 4 mg/ml dispase for 1 hour at 37°C. The solution was then filtered through 70 µm Falcon strainers and immersed in αMEM supplemented with 10% fetal bovine serum (FBS), 100 U/ml penicillin, 100 µg/ml streptomycin and placed in 75 ml flasks. Cells were incubated for 1 week at 37°C with 5% CO 2 (Gronthos et al., 2002). For each dental pulp, the total number of extracted cells was evaluated by counting cells in a Glasstic ® slide with grids, under optical microscopy. Cell viability was assessed by 0.4% Trypan Blue staining.

Cell sorting

Non-adherent cells were removed by a change of medium 24 h after cell seeding. The remaining cells were let in culture medium and followed for a week, to avoid them to reach confluence. Subconfluent cells were then either selected by Fluorescence-activated cell sorting (FACS) or directly analyzed by flow cytometry. Cell sorting by FACS was realized for the expression of cell surface markers CD146 and CD34 (see below).

Cell characterization

After 1 week, subconfluent cells were collected (after FACS or without sorting)

and analyzed for minimal criteria to define human mesenchymal stem cells, such as adherence to plastic, clonogenicity, competence to proliferate, expression of specific cell surface antigens and ability to differentiate into osteoblasts, adipocytes and chondroblasts in vitro [START_REF] Morsczeck | Isolation of precursor cells (PCs) from human dental follicle of wisdom teeth[END_REF] (Dominici et al., 2006).

Colony forming efficiency

To assess colony-forming efficiency (CFU-F, colony forming units fibroblast), dental pulp cells were seeded in a dilution of 5x10 4 cells/well in 6-well cell culture plates, immediately after cell recovery. On day 7, all cultures were fixed with 4% formaldehyde, and then stained with 0.1% toluidine blue. Aggregates of >50 cells were scored as colonies [START_REF] Seo | Investigation of multipotent postnatal stem cells from human periodontal ligament[END_REF].

Proliferation rate

The proliferation rate of sub-confluent cultures (first passage) of stem cells was assessed by bromodeoxyuridine incorporation for 24 h, with a Zymed Laboratories BrdU staining kit. BrdU was incorporated into proliferating cells (during S-phase of mitosis) and revealed using a highly sensitive streptavidin-biotin staining system [START_REF] Morsczeck | Isolation of precursor cells (PCs) from human dental follicle of wisdom teeth[END_REF].

Flow cytometry characterization

For flow cytometry, cells were dissociated from their substrate with 0.5 mM

Accutase for 1 min. Cells were then resuspended in 1% BSA in PBS, counted and placed in flow cytometry tubes (10 5 cells per tube). The tubes were centrifuged at 1400 rpm for 10 min, the supernatant discarded, and 5 µL of antibody solutions were placed on the centrifuged cells. The antigen profiles of cultured DPSC were analyzed by detecting the expression of the cell surface markers CD90, CD146, CD117, CD34, STRO-1 and CD45 using flow cytometry and conjugated monoclonal antibodies, presented in Table 2.1.

• CD90 is a widely accepted marker for mesenchymal stem cells. This surface protein plays a role in cell-cell interaction events, including intracellular adhesion and cell recognition during development, and is considered an important stem cell marker [START_REF] Kern | Comparative Analysis of Mesenchymal Stem Cells from Bone Marrow, Umbilical Cord Blood, or Adipose Tissue[END_REF]) [START_REF] Mercati | Expression of mesenchymal stem cell marker CD90 on dermal sheath cells of the anagen hair follicle in canine species[END_REF].

• CD146 is a marker expressed in perivascular mesenchymal stem cells, also known as MCAM (Melanoma Cell Adhesion Molecule) or cell surface glycoprotein MUC18. Notably, CD146 is not expressed by bone marrow hematopoietic stem cells or their progenitors and has been linked to various cellular processes including cell adhesion, cytoskeletal reorganization, cell shape, migration, and proliferation through transmembrane signaling [START_REF] Shi | Perivascular niche of postnatal mesenchymal stem cells in human bone marrow and dental pulp[END_REF]. Several studies demonstrated that CD146 positive cells from connective tissue exhibited MSC potential (Kemoun et al., 2011).

• CD117 is a receptor expressed on the surface of hematopoietic stem cells as well as other cell types. CD117, also known as c-Kit, is a tyrosine kinase receptor that binds to the Stem Cell Factor (SCF), forming a dimer that activates its intrinsic tyrosine kinase activity. Signaling through CD117 plays a role in cell survival, proliferation, and differentiation (Li et al., 2003[START_REF] El-Sayed | Periodontal regeneration employing gingival margin-derived stem/progenitor cells: an animal study[END_REF].

• CD34 is a marker selectively expressed by primitive myeloid and lymphoid progenitor. However, CD34 is not restricted to hematopoietic progenitors, but is expressed by vascular endothelial cells and in basement membrane structures [START_REF] Simmons | CD34 expression by stromal precursors in normal human adult bone marrow[END_REF]. CD34+ cells are normally found in the umbilical cord and bone marrow as hematopoietic cells, and can be primitive stromal stem cells [START_REF] Gordon | Characterization and clinical application of human CD34+ stem/progenitor cell populations mobilized into the blood by granulocyte colony-stimulating factor[END_REF].

• STRO-1 is a marker of pre-osteogenic, pre-chondrogenic or pre-adipogenic populations. Its expression is progressively lost during cell proliferation and differentiation. Thus, STRO-1 seems to be an early marker of mesenchymal stem cells, and is often used in combination with cell adhesion molecules (such as CD106

or CD146) to isolate purified bone marrow stem cell populations [START_REF] Gronthos | Bone marrow stromal stem cells for tissue engineering[END_REF].

• CD45 is a marker of hematopoietic cells, mainly myeloid progenitors. It is used in this study as a control to verify the stromal origin of dental pulp cells and the absence of hematopoietic precursor contamination (Bakopoulou et al., 2011). At the end of the 3 weeks of induced differentiation, cultures were fixed and assessed for mineralization by Alizarin Red staining, for adipogenic differentiation by

Red O Oil staining and for chondrogenic differentiation by Alcian blue staining of thin section.

Use and conservation

Recovered primary DPSC were either used directly for various experiments (cell interactions with porous silicon scaffolds) or cryopreserved in liquid nitrogen. For storage, 10 6 cells were resuspended in 10% dimethylsulfoxide (DMSO) in FBS in a cryotube and frozen in liquid nitrogen. Storage was prolonged for up to 1 year. At the end of the storage period, cells were quickly thawed in a water bath at 37°C, then added to 10 ml of αMEM supplemented with 10% FBS. Cells were centrifuged at 1,400 rpm for 10 min, the supernatant was removed and fresh medium was added to the tube. Cells were then placed in T-75 flasks and cultured at 37°C in a 5% CO 2 humid atmosphere.

Results

Cells recovery

The pulp tissue enzymatic digestion permitted to extract an important amount of cells. This quantity was mainly depending on the tooth formation: the more immature the tooth, the bigger the pulp tissue available. We worked only with discarded wisdom teeth, extracted for orthodontic reasons, mainly impacted immature teeth, with a number of recovered cells varying between 5x10 5 to 1.5x10 6 : for all the evaluated teeth, we found an average amount of 1,068,301 ± 501,722 cells per pulp (mean value ± standard error of the mean).

Cells characterization

Adherence, clonogenicity and proliferation

We removed all non-adherent cells by changing the culture medium after 24 h.

We found the presence of a clonogenic cell population from the dental pulp tissue, and cells within each colony were characterized by a typical fibroblast-like morphology (Figure 2.6). The frequency of colony-forming cells derived from dental pulp tissue was of 20 to 65 colonies for 5x10 4 cells plated in 6-well plates. The number of proliferating cells, assessed for recovered dental cell cultures after 7 days, was of 63% BrdU-positive cells (62.8 ± 12.3 %).

Differentiation assays

Recovered DPSC were tested for their ability to differentiate into osteoblast, chondrocyte and adipocyte. After 3 weeks of in vitro differentiation assays, DPSC cultures formed calcium deposits in osteogenic medium (stained with Red Alizarin), lipid droplets in adipogenic medium (stained with Red O Oil) and proteoglycans in chondrogenic medium (stained with Alcian Blue) (see Figure 2.7, next page).

OM+ AM+ CM+

OM-AM-CM-

Flow cytometry assays

Dental pulp cells were analyzed by flow cytometry for several cell surface markers, such as CD90, CD146, CD117, CD34, STRO-1 and CD45. This characterization was realized immediately after cells recovery (without in vitro expansion), after 24 h in culture medium (to remove non adherent cells), after 7 days in culture (subconfluent cells) and after cryopreservation (6-12 months in liquid nitrogen).

The first flow cytometry analysis was realized immediately after dental extraction, tooth opening, pulp tissue recovery and cells extraction. These assays showed a widely heterogeneous population, mainly CD90 positive with cells varying from high to low positivity. We found this cell population to be 36% CD117+, 11%

CD34+ and 2% CD45+. No CD146+ cells were found within the recovered cells Side Scatter Forward Scatter

(Figure 2.8).

Fluorescence Activated Cell Sorting (FACS)

We found a distinguished high-positive CD34 population after 24 h and a clear CD146+ population after 7 days. Thus, after 3 days in culture, we characterized and sorted dental pulp cells by FACS for CD34 and CD146 positivity, and controlled for CD45 negativity. We obtained a pure population of CD34+/CD146+/CD45-stem cells.

We kept this purified population in culture in vitro for 3 weeks (passage 3), controlling the expression of CD146 and CD34 positivity after passage 3. It appeared that these sorted DPSC kept their CD146 positivity for up to 3 weeks. However, the high positive CD34 population couldn't be kept in culture: the CD34+ cells expressed weakly this marker after 3 weeks in culture (Figure 2.12)

Discussion

As already shown in numerous studies, several populations of stem cells have been isolated from different parts of the human tooth and all of them have been shown

to have generic mesenchymal stem cell-like properties [START_REF] Shi | Comparison of human dental pulp and bone marrow stromal stem cells by cDNA microarray analysis[END_REF] (Gronthos et al., 2002) [START_REF] Shi | Perivascular niche of postnatal mesenchymal stem cells in human bone marrow and dental pulp[END_REF] [START_REF] Laino | A new population of human adult dental pulp stem cells: a useful source of living autologous fibrous bone tissue (LAB)[END_REF]) (Papaccio et al., 2006) [START_REF] Aquino | Human postnatal dental pulp cells co-differentiate into osteoblasts and endotheliocytes: a pivotal synergy leading to adult bone tissue formation[END_REF] (Graziano et al., 2008) (d'Aquino et al., 2009). We demonstrated here the presence of an accessible and reliable source of Dental Pulp Stem Cells (DPSC) within the dental pulp of discarded wisdom teeth. Indeed, following the criteria of the International Society for Cellular Therapy (Dominici et al., 2006), we showed that we could recover mesenchymal stem cells from the dental pulp, even without cell sorting:

(1) We kept only plastic adherent cells by changing culture medium after 24 h.

(2) We confirmed the clonogenicity and proliferation capacities by counting CFU-F (fibroblast-like colony forming units) and measuring the rate of proliferative BrdUpositive cells. Our results were consistent with previous studies (Gronthos et al., 2000) ( [START_REF] Miura | SHED: stem cells from human exfoliated deciduous teeth[END_REF] [START_REF] Seo | Investigation of multipotent postnatal stem cells from human periodontal ligament[END_REF] [START_REF] Sonoyama | Mesenchymal stem cell-mediated functional tooth regeneration in swine[END_REF].

(3) We performed flow cytometry experiments on the recovered cells, from the immediate cells extraction until cells after cryopreservation. We confirmed the expression of CD90 and CD117 by the recovered cells. The expression of CD146 marker appeared after several days in vitro, when cells started to proliferate and migrate. The expression lack of CD45 showed that the recovered cells were not hematopoietic cells, but originated from the dental pulp. ( 4) We followed cells differentiation in vitro, to confirm that the recovered cells were able to differentiate into osteoblasts, adipocytes and chondroblasts.

The overall purpose of our whole project was to develop an efficient complex composed of resorbable biomaterials (porous silicon) covered by stem cells, for bone tissue engineering. As it had been demonstrated that different cell lines had different adhesion and growth characteristics, we wanted to work with primary human stem cells, usable for therapeutic applications. The third molars appeared to be one of the most accessible source of dental stem cells, because their extraction could be currently performed for orthodontic purpose and, according to the french regulation, these extracted teeth were considered to be medical waste. Because the third molar is the last tooth to develop, it was often, when extracted, in an early stage of development. Thus, wisdom teeth yielded an optimum quantity of dental pulp tissue for the isolation of DPSC (Gandia et al., 2008) [START_REF] Atari | Dental pulp of the third molar: a new source of pluripotent-like stem cells[END_REF]. Our aim was not to discover a new subpopulation of dental stem cells, but to work out an efficient technique to recover stem cells. Thus, we could proceed all experiments with fresh primary culture of human stem cells.

At the beginning of the study, we focused on the CD34+/CD146+ subpopulation because the characteristics of these adherent CD34+ cells were supposed to identify them as a separate putative stem/progenitor cell population [START_REF] Gordon | Characterization and clinical application of human CD34+ stem/progenitor cell populations mobilized into the blood by granulocyte colony-stimulating factor[END_REF], and because most of the cell population in our preliminary studies was found to express CD146 marker. In addition, CD34 has been showed as an epitope commonly expressed on human hematopoietic stem cells and on some primitive stromal stem cells from bone marrow (Garcia-Pacheco et al., 2001) and it was demonstrated that the expression of CD146 in the dental pulp was restricted to blood vessel areas, localized in the perivascular region of the pulp tissue [START_REF] Shi | Perivascular niche of postnatal mesenchymal stem cells in human bone marrow and dental pulp[END_REF]. The expression of CD34 antigen contrasted with previous studies (Gronthos et al., 2002) [START_REF] Miura | SHED: stem cells from human exfoliated deciduous teeth[END_REF] and with the general definition of MSC (Dominici et al., 2006). However, the targeted CD34+/CD146+ population lacked expression of the hematopoietic cell marker CD45 suggesting a mesenchymal origin of these cells and the absence of hematopoietic precursor contamination. And this finding was in agreement with previously published studies that identified a CD117+/STRO-1+/CD34+/CD45-population inside the dental pulp (Laino et al., 2006) (Bakopoulou et al., 2011). We aimed to isolate and purify a primitive stem cells population from mesenchymal tissue, with the ability to proliferate and migrate. However, we could not maintain the purified CD34+/CD146+ population in vitro. Cultured cells kept their CD146 positivity for several weeks (as did cells without cell sorting), but lost high CD34 expression after passage 1 or passage 2. Even though it had been reported that DPSC could keep CD34 expression for up to 4 weeks, we didn't manage to reproduce such results [START_REF] Aquino | Human postnatal dental pulp cells co-differentiate into osteoblasts and endotheliocytes: a pivotal synergy leading to adult bone tissue formation[END_REF]. For these reasons, we chose to conduct our tissue engineering experiments with non-purified DPSC, without cell sorting, following the minimal criteria for defining mesenchymal stem cells.

The dental pulp is, in many ways, one of the most interesting adult tissues for recovery, therapeutic use and storage of stem cells. The properties of DPSC have been demonstrated to be remarkable for their multipotency as well as their ability to proliferate after cryopreservation. Currently, stem cells that are most used in medicine are stem cells from bone marrow. Comparing the properties of DPSC to those of stem cells from the bone marrow, it seems that DPSC could be a new gold standard for the use of multipotent stem cells in regenerative medicine, for both autologous or allogenic use (Huang et al., 2009) (Collart Dutilleul et al., 2012).

Chapter 3 DPSC Adhesion and Proliferation on surface modified porous silicon scaffolds

Introduction

In regenerative medicine, tissue engineering applications are based on the development of biological substitutes that can restore, maintain or even improve tissue functions: tissue engineering combines cells and bioactive factors in a defined microenvironment created by a biomaterial scaffold. A key component for tissue engineering is the biomaterial scaffold that, ideally, should support cell attachment, proliferation and differentiation, and be biocompatible and biodegradable at a controlled rate [START_REF] Wang | Stem cell-based tissue engineering with silk biomaterials[END_REF]. Stem cell-based therapy often requires a scaffold to carry cells to the injured site. Porous silicon (pSi) appears to be a promising biomaterial for tissue engineering as it is both non-toxic and bioresorbable under physiological conditions, and it dissolves progressively into nontoxic silicic acid [START_REF] Dutilleul | Porous Silicon for Biomedical Applications[END_REF].

Its dissolution rate is dependent on the pore geometry and surface chemical properties, and these two factors influence cell adhesion [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF] [START_REF] Khung | Using continuous porous silicon gradients to study the influence of surface topography on the behaviour of neuroblastoma cells[END_REF].

Moreover, this tunable, biocompatible, and resorbable material has been reported to favor the growth of hydroxyapatite, suggesting the possible bone implantability of the material (Canham, 1995). Its biocompatibility and immunogenicity has already been demonstrated under different conditions and it additionally offers useful photoluminescence properties [START_REF] Ainslie | In vitro immunogenicity of silicon-based micro-and nanostructured surfaces[END_REF]. Furthermore, pSi-based scaffolds have been investigated for orthopedic (Whitehead et al., 2008) and ophthalmic implants [START_REF] Cheng | Intravitreal properties of porous silicon photonic crystals: a potential self-reporting intraocular drug-delivery vehicle[END_REF], for controlling the adhesion and proliferation of different cell types [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF] (Alvarez et al., 2009[START_REF] Torres-Costa | Engineering of silicon surfaces at the micro-and nanoscales for cell adhesion and migration control[END_REF], and even as intravenously injectable particles for imaging tumors in vivo (Park et al., 2009) or retina pressure actuators [START_REF] Noval | Design and characterization of biofunctional magnetic porous silicon flakes[END_REF].

Surface treatment resulting in accessible porosity with chemically stable inner surface is a key step to prevent rapid hydrolysis and degradation of the porous matrix in aqueous environments, without eliciting any undesirable effects on the cells. Moreover, surface chemical modification by turning pSi from hydrophobic to hydrophilic promotes cell adhesion and growth [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF][START_REF] Torres-Costa | Engineering of silicon surfaces at the micro-and nanoscales for cell adhesion and migration control[END_REF]. The most common and simple surface treatment is oxidation, which can be performed by either ozone, aging, thermal or chemical treatments. Amine-terminated modifications as silanization with aminopropyl trimethoxysilane or triethoxysilane improve pSi chemical stability and enhance cell adhesion in comparison to oxidized pSi [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF].

Thermal hydrosilylation was also used to graft chemical species to generate a substrate for cell adhesion and proliferation, such as dodecene, undecenoic acid or oligoethylene glycol (Alvarez et al., 2009).

Substrate topography has been known for a long time to affect crucial cell functions, namely adhesion, proliferation, migration, and differentiation. Moreover, material surface morphology has strong effects on cell cytoskeleton and morphology [START_REF] Dalby | Nucleus alignment and cell signaling in fibroblasts: response to a micro-grooved topography[END_REF]. Furthermore cellular organization requires that a cell assess its relative location, taking in multiple cues from its microenvironment i.e. the extracellular matrix in vivo, and the supporting scaffold in vitro [START_REF] Buxboim | Matrix elasticity, cytoskeletal forces and physics of the nucleus: how deeply do cells "feel" outside and in?[END_REF]. Cells respond to topographic surfaces in a wide variety of ways, which depend upon cell type, pore size, as well as the physicochemical properties of the substrate material. The effect of pore size and porosity of pSi on cell growth is of particular relevance. Indeed, its pore dimensions can be precisely controlled and are highly tunable during pSi electrochemical anodization. A variety of pore sizes can be produced: from micropores (< 2 nm), mesopores (2-50 nm) to macropores (> 50 nm) depending on the preparation conditions [START_REF] Sailor | Porous Silicon in Practice[END_REF]. Although microscale topography modulates cellular behavior in vitro, it is important to consider that cells in vivo make contact with microscale as well as nanoscale topographical features. Diverse topographical features have been assessed for different cell types, at the micrometer and submicrometer scale, and some recent studies on pSi have focused specifically on pore geometry influence on cell adhesion and proliferation [START_REF] Sapelkin | Interaction of B50 rat hippocampal cells with stain-etched porous silicon[END_REF] [START_REF] Khung | Using continuous porous silicon gradients to study the influence of surface topography on the behaviour of neuroblastoma cells[END_REF] (Wang et al., 2012) (Clements et al., 2012) [START_REF] Gentile | Differential Cell Adhesion on Mesoporous Silicon Substrates[END_REF]. Pore geometry was clearly shown to affect the cellular response, but each cell type responded differently. Rat hippocampal neurons were observed to preferentially adhere on macroporous surfaces, with a pore size ranging between 50 and 100 nm, rather than on flat silicon surfaces [START_REF] Sapelkin | Interaction of B50 rat hippocampal cells with stain-etched porous silicon[END_REF]. Neuroblastoma cells cultured over continuous porous gradient substrates were more likely to develop on surface topography with feature sizes of <20 nm, and substrates with an average pore size of a few hundreds of nanometers restricted cell adhesion and proliferation [START_REF] Khung | Using continuous porous silicon gradients to study the influence of surface topography on the behaviour of neuroblastoma cells[END_REF]. Primary human endothelial cells, mouse fibroblasts, mouse neuroblastoma cells, and human cortical neuron cell lines adhered and proliferated more on mesoporous silicon than on flat silicon, with a tendency to proliferate more on pSi with an average pore size of ≈5 nm, rather than ≈20 nm [START_REF] Gentile | Differential Cell Adhesion on Mesoporous Silicon Substrates[END_REF]. It is of interest to note that this was, so far, the only study assessing human primary cells. Rat mesenchymal stem cells (rMSC) adhesion was enhanced as pore size decreased, with a maximum proliferation for an average pore size of ≈20 nm but responded more strongly to surface chemical changes during short-term culture (Clements et al., 2012) and had a high proliferation rate also on flat silicon (Wang et al., 2012).

Coupling the auspicious capacities of human adult mesenchymal stem cells (hMSC) with the unique properties of pSi substrates provides a promising approach for therapeutic application in regenerative medicine. Indeed, mesenchymal stem cells therapeutic potential has generated significant excitement in the field of regenerative medicine, as they can be found in various niches within human body. The ability of these cells to self-renew and differentiate into multiple tissues makes them an attractive cell source for cell-based regenerative therapies. MSCs have considerable potential for the treatment of musculoskeletal disorders owing to their expansion capacity, immunosuppressive properties and ability to differentiate into bone and cartilage (Ma, 2010). Adult stem cells constitute the source of differentiated cells for the regeneration of tissues that are diseased or injured. In adults, these cells are accessible from different origins, such as bone marrow, adipose tissue or dental pulp (Marolt et al., 2010).

In the study presented in this chapter, we investigated for the first time the behavior of primary culture of hMSC on pSi scaffolds. Mesenchymal stem cells from the dental pulp, named Human Dental Pulp Stem Cells (DPSC) (Gronthos et al., 2000), were studied for adhesion and proliferation on oxidized mesoporous and macroporous pSi substrates, presenting pore size of 10 ± 2 nm, 36 ± 4 nm and 1.0 ± 0.1 µm. We also evaluated the influence of pSi substrate surface modification with CO 2 H, NH 2 or NHCONHNH 2 groups on the cell behavior. After characterization of the silicon substrates, cell adhesion and proliferation behavior were studied using fluorescence and electron microscopy, enzymatic proliferation assays and bromodeoxyuridine (BrdU) incorporation for the detection of proliferating cells.

Materials

Human Dental Pulp Stem Cells (hDPSC)

Human impacted third molar extracted for orthodontic reasons were recovered from healthy patients (15-18 years of age). Written informed consent was obtained from the parents of the patients. This protocol was approved by the local ethical committee (Comité de Protection des Personnes, Montpellier hospital, France). Tooth surfaces were cleaned using 2% chlorhexidine and cut around the cementum-enamel junction by using sterilized disc. The teeth were then broken into 2 pieces to reveal the pulp chamber. The pulp tissue was gently separated from the crown and root and then digested in a solution of 3 mg/mL collagenase type I and 4 mg/mL dispase for 1 hour at 37°C. The solution was then filtered through 70 µm Falcon strainers and immersed in αMEM supplemented with 10% fetal bovine serum (FBS), 100 U/mL penicillin, 100 µg/mL streptomycin and placed in 75 ml flasks. Cells were incubated for 1 week at 37°C with 5% CO2. Non-adherent cells were removed by a change of medium 24 h after cell seeding.

DPSC characterization

After 1 week, sub-confluent cells were collected and analyzed for minimal criteria to define human mesenchymal stem cells, such as adherence to plastic, expression of cell surface antigens and ability to differentiate into osteoblasts, adipocytes and chondroblasts in vitro (Dominici et al., 2006). The antigen profiles of cultured DPSC were analyzed by detecting the expression of the cell surface markers CD90, CD146, CD117 and CD45 using flow cytometry. CD90 is a widely accepted marker for mesenchymal stem cells, CD146 is a marker expressed in perivascular mesenchymal stem cells, CD117 is the receptor of stem cell factor, and CD45 is a marker of hematopoietic cells, mainly myeloid progenitors [START_REF] Coppe | Characterization of primary dental pulp cells in vitro[END_REF] (Shi et al., 2003). The latter has been used to demonstrate the absence of contamination by CD45+ hematopoietic progenitors. Cells were controlled for multipotency with in vitro osteogenic, adipogenic and chondrogenic differentiation following a previously described protocol (Kemoun et al., 2007). Briefly, for osteogenic and adipogenic differentiation, dental pulp cells were seeded at a density of 105 cells/cm2 and maintained in αMEM supplemented with 10% FBS until confluence. Cells were then cultured in osteogenic medium (αMEM supplemented with 15% FBS, dexamethasone, L-ascorbate phosphate, and β-glycerophosphate) or in adipogenic medium (αMEM supplemented with 10% FBS, dexamethasone, insulin, and isobutylmethylxanthine) for 21 days. Control cultures were maintained without osteogenic or adipogenic supplements. To induce chondrogenic differentiation, cells were cultured for 21 days as pellets in a serum-free expansion medium containing dexamethasone, L-ascorbate phosphate, L-proline, sodium pyruvate, ITS-Premix, and TGF-β3. Control pellet cultures were maintained in basal medium.

Porous silicon scaffolds preparation

P++ type boron-doped crystalline silicon wafers with 0.0008 -0.0012 Ωcm resistivity were obtained from Siltronix®. Wafers were etched in a custom-made Teflon cell at a constant current density of either 30 mA/cm 2 for 10 min or 300 mA/cm 2 for 2 min 15 s, in a hydrofluoric acid (HF) solution in ethanol (3:1 HF:ethanol solution, volume ratio). To create pSi with larger pores (1 µm), p-type <100> wafers with 7 -21 Ωcm resistivity were etched at a constant current density of 4.5 mA/cm 2 for 10 min, in a 50% HF solution in dimethylformamide (DMF). Etched wafers were oxidized at 800°C for 1 hour. The wafers were cut into 0.5 cm 2 pieces. Some of the samples etched in a 300 mA/cm 2 current were also submitted to various chemical treatments: silanization with aminopropyltriethoxysilane (APTES) after thermal oxidation (as described above), hydrosilylation with undecenoic acid without pre-oxidation treatment, and hydrosilylation with semicarbazide without pre-oxidation treatment. PSi surface modifications with APTES, and with undecenoic acid were described elsewhere [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF] (Alvarez et al., 2009). The covalent attachment of the semicarbazide to hydrogen-terminated pSi surfaces by thermal hydrosilylation was realized according to recently published procedure using tert-Butyl-2[(allylamino)carbonyl]hydrazine carboxylate [START_REF] Secret | Antibody-functionalized porous silicon nanoparticles for vectorization of hydrophobic drugs[END_REF]. The removal of the protecting group yields a semicarbazide-terminated monolayer, [START_REF] Coffinier | Semicarbazide-Functionalized Si(111) Surfaces for the Site-Specific Immobilization of Peptides[END_REF]. For use in cell culture, the wafers were sterilized with 70% ethanol (volume ratio) for 10 min before drying under sterile airflow.

Surface characterization

The topography of the surface modified pSi samples was analyzed by environmental scanning electron microscopy (SEM) (Analytic FEI Quanta FEG 200) to determine the pore size, and by atomic force microscopy (AFM) (Asylum MFP-3D, Asylum Research, Santa Barbara, CA) to determine the surface roughness. For SEM, an acceleration voltage of 20.00 kV was used in a pressure of 0.5 Torr. For AFM, goldcoated silicon nitride rectangular cantilevers were used with a typical spring constant of 30 pN/nm-1, tip radius ~30 nm (BL-RC150 VB-C1: Bio-Lever A, Olympus Optical Co., Ltd., Tokyo, Japan). The spring constant for each cantilever was determined by thermal noise method within the supplied software. Height images were recorded in tapping mode in liquid at room temperature. Typically, 512 x 512 point scans were taken at a scan rate of 1 Hz per line.

Surface tension was determined using sessile drop contact angle measurements, using a custom-made set-up consisting of a syringe dispenser, a sample stage, a macro lens and a CCD camera. 5µl of MilliQ water was spotted onto the surface at room temperature. Images of the drop profiles were captured and ImageJ® software was used to measure contact angles on both sides of the droplet. A minimum of four replications was conducted for each sample surface.

Cell adhesion

DPSC attachment was monitored by fluorescence microscopy. Cells were seeded onto the surface of sterilized pSi at a cell density of 5x10 4 cell/ml. Flat silicon (non etched silicon wafers) and glass coverslip were used as controls. Cells were incubated for 4h, 24h, 48h and 72h at 37°C with 5% CO2, in a humidified incubator, in ultra low adherence 24 well plates (Corning, NY, USA) which inhibit cell attachment on the tissue culture plate, allowing DPSC to attach only to the pSi. After the incubation time, the cells were fixed with 2.5% glutaraldehyde and stained with 50 µg/mL of 4,6diamidino-2-phenylindole (DAPI) for 30 min before being washed with PBS to remove any non-adherent cells. Cells were observed under fluorescence microscopy at an excitation wavelength of 290 nm. Controls were cells cultured on glass coverslip and flat silicon, cut into 0.5 cm 2 squares. Cell counts were conducted at five different locations on the surface of each sample (four peripheral and one central) in areas of 1400 µm x 1050 µm.

Proliferation assays

Cell proliferation was first measured via quantification of acid phosphatase activity using paranitrophenylphosphate phosphatase test (pNPP) [START_REF] Werner | The effect of microstructured surfaces and laminin-derived peptide coatings on soft tissue interactions with titanium dental implants[END_REF]. DPSC proliferation was assessed after 4 h, 24 h, 48 h and 72 h incubation. At the end of each experimental time, the cells were washed three times with PBS and lysed with 500 mL of acid phosphatase lysis buffer (0.1 M sodium acetate, 0.1% Triton X-100, pH 5.5), supplemented with 1 mg/mL of pNPP (Sigma-Aldrich). After 1 h incubation at 37 °C, the reaction was stopped by the addition of 10 µL of 1 N NaOH. The supernatant was removed, placed in a new 24 wells plate and the yellow colorimetric reaction was measured by a microtiter plate reader at 405 nm. All proliferation experiments were performed at least in triplicate and results were normalized according to the glass coverslip control at 72 h, for which 100% of proliferation was attributed.

The proliferation rate of DPSC on pSi was also assessed by bromodeoxyuridine (BrdU) incorporation for 24 h [START_REF] Horner | Proliferation and differentiation of progenitor cells throughout the intact adult rat spinal cord[END_REF]. DPSC were seeded on the various pSi scaffolds at a density of 105 cells/ml in αMEM supplemented with 10% FBS with 1:100 diluted BrdU labeling (Invitrogen). After 24 h incubation at 37°C, samples were rinsed 3 times in PBS, fixed in 4% paraformaldehyde, rinsed again 3 times in PBS and put in 1.5M HCl (Sigma) for 30 min at room temperature to dissociate DNA strands.

Samples were washed with PBS, incubated in PBS with 0.1% Triton X-100 and 1% BSA for 1 h at room temperature, then incubated with mouse anti-BrdU primary antibody (Miltenyi) overnight at 4°C. After immunostaining, cells were washed with PBS/1% BSA, incubated with FITC conjugated rabbit anti-mouse secondary antibody for 30 min and counterstained with 2µg/ml Hoechst 33342 (Sigma) for nucleus staining.

Samples were observed under fluorescence microscopy at an excitation wavelength of 290 nm for nuclei (blue staining) and 490 nm for BrdU (green staining). All experiments were performed at least in triplicate and the number of BrdU-positive cells was expressed as a percentage of the total number of cells, counted at 5 different points per sample.

Cellular morphology and viability

Fluorescein diacetate (FDA) and propidium iodide (PI) staining was used to observe cell morphology and to distinguish viable and dead cells. FDA, which enters viable cells by energy-dependent endocytosis, yield a bright green color; while PI, which interacts with RNA and DNA of cells having disrupted cytoplasmic and nuclear membranes, produces a red color [START_REF] Cao | Cell adhesive and growth behavior on electrospun nanofibrous scaffolds by designed multifunctional composites[END_REF]. Living cells were stained with 25 µg/mL of fluorescein diacetate (FDA) and 20 µg/ml of propidium iodide (PI), and incubated for 3 min at 37°C. After staining, samples were washed with PBS before being observed under a Nikon TE2000-E microscope equipped with a Nikon digital camera at an excitation wavelengths of 480 nm for FDA and 630 nm for PI.

Observations were conducted at five different locations on the surface of each sample (four peripheral and one central) at magnifications x20 and x40. All experiments were made in triplicate.

SEM evaluation of DPSC morphology and spreading

The cells were cultured on the various pSi scaffolds and control for 24 h under normal conditions, as described above. After 24 h incubation the cells were washed twice with PBS buffer and fixed with 2.5% glutaraldehyde for 1 h at room temperature.

After washing the specimens were dehydrated in graded ethanol solutions from 50% to 100%, and in hexamethyldisilazane (HMDS, Ted Pella, USA). The samples were then sputter-coated with platinum. Scanning electron microscopy (SEM) was performed on an Analytic FEI Quanta FEG 200 microscope with an acceleration voltage of 15 kV in a pressure of 10 -5 Torr. From the SEM images, two parameters for characterizing the cell morphology were considered: the cell area (in µm 2 ), defined as the area covered by the cell projected over the substrate and the cell circularity, defined as the ratio between the shorter and the longer axis of the cell (value between 0 for elongated cells and 1.00 for round cells). All experiments were made in triplicate.

Statistical analysis

All data were evaluated with a Shapiro-Wilk normality test. The cell attachment evaluations (cell counts) were plotted as mean ± standard error of the mean and statistical analyses were performed using a parametric Student t-test. The results of proliferation experiments with pNPP enzymatic assays were normalized according to the glass coverslip control at 72 h, for which 100% of proliferation was attributed. The data were plotted as mean ± standard error of the mean and statistical analyses were performed using a parametric Student t-test. The results of BrdU proliferation assay, cell surface area and cell circularity were analyzed with a non-parametric one-way Tukey ANOVA test (SigmaStats, SPSS Inc., Chicago, IL, USA). A p-value of <0.05 was considered to be significant.

Results

pSi samples

PSi substrates were generated from bulk boron-doped p-type silicon wafers via anodization [START_REF] Sailor | Porous Silicon in Practice[END_REF]. Substrates with various pore sizes were obtained by modulating the etching conditions and the doping level of the silicon wafer. Figure 3.1 shows representative SEM and AFM images of the pSi substrates; distinct textural features are depicted when comparing the different pSi substrates. PSi substrates obtained from silicon wafers with a resistivity of 0.0008 -0.0012 Ωcm by applying a constant current density of 30 mA/cm2 and 300 mA/cm2 had an average pore diameter of 10 ± 2 nm (pSi 10 nm) and 36 ± 4 nm (pSi 36 nm), respectively. The pSi layers produced from silicon wafers with 7 -21 Ωcm resistivity, by applying a current density of 4.5 mA/cm2, had an average pore diameter of 1020 ±100 nm (pSi 1 µm). The pSi substrates were quite uniformly porosified. Topographic images of pSi were obtained using AFM, and the calculated statistical roughness (RMS) of the substrate were found at 1.9 ± 0.7 nm for "pSi 10 nm", 7 ± 2 nm for "pSi 36 nm" and 54 ± 20 nm for "pSi 1 µm". Surface chemical treatments realized on "pSi 36 nm" substrates are hydrosilylation with undecenoic acid or semicarbazide, and silanization (after thermal oxidation) with APTES. The reaction schemes of the chemical treatments are presented in Figure 3.1. The sample pSi 36 nm was chosen because it displayed the best results in term of cell adhesion, compared to pSi 10 nm and pSi 1 µm (Figure 3.2). Water contact angle measurements performed after surface functionalization revealed that all the substrates were hydrophilic (Table 3.1). Oxidized pSi hydrophilicity was attributed to the formation of a polar Si-OH capped surface after oxidation. By analyzing SEM micrographs, the surface porosities (ratio between total area of the pores and the area of the considered region of interest) were found to be as follows: 40 ± 4% for pSi 10nm, 45 ± 6% for pSi 36nm and 24 ± 4% for pSi 1µm.

Flat Si pSi 10 nm pSi 36 nm pSi 1 µm

Oxidized 26° ± 4 16° ± 2 15° ± 3 33° ± 5 Undecenoic acid 35° ± 5 APTES 37° ± 6
Semicarbazide 12° ± 3 The experimental results show that initial cell adhesion and growth were very similar on glass coverslip, Flat Si, APTES-treated pSi and semicarbazide-treated pSi, after 4 h and 24 h. Initial cell adhesion on oxidized pSi (pSi 10nm, pSi 36nm and pSi 1µm) and undecenoic acid-treated pSi were significantly lower compared to glass coverslip (p<0.001, p=0.003, p<0.001, and p<0.001 respectively). However, DPSC started to proliferate on pSi 10 nm and pSi 36 nm after 24 h following the same growth pattern observed on glass coverslip, while the number of cells slightly decreased on pSi 1 µm and on undecenoic acid-treated pSi after 24h and 48 h respectively. After 48 h, the numbers of cells were similar on glass coverslip and on flat Si, significantly higher on APTES-treated pSi (p<0.001) and significantly lower on semicarbazide-treated pSi The experiments show that DPSC clearly proliferated at all time points on glass coverslip, flat Si, pSi 10 nm, pSi 36 nm and APTES-treated pSi, while the proliferation was limited on pSi 1 µm and undecenoic acid-treated pSi. For semicarbazide-treated pSi, the proliferation rate was higher compared to glass coverslip until 24 h, and dropped dramatically after 48 h. After 24 h of incubation, the proliferation rate was equivalent on glass coverslip and on flat Si. At 24 h, APTES-treated pSi and semicarbazide-treated pSi improved DPSC proliferation of 27 % and 24 % (p=0.042 and p=0.039, respectively). The proliferation rate was lower on pSi 10 nm, pSi 1 µm and undecenoic acid-treated pSi (p=0.012, p=0.004 and p=0.011, respectively). On pSi 36 nm, this rate was slightly lower compared to glass coverslip and slightly higher compared to pSi 10 nm, but without statistical significance (p=0.106 and p=0.094, respectively). Interestingly, we found that, after 48 h and 72 h, the proliferation rate on pSi 36 nm was significantly higher than the rate on pSi 10 nm (p=0.006 at 48 h and p=0.003 at 72 h), revealing the tendency for DPSC to attach and proliferate more on pores of about 36 nm than on pores of about 10 nm. After 72 h of incubation, the proliferation rate was equivalent on flat Si and glass coverslip, higher for 34 % on APTES-treated (p=0.005), and lower on pSi 10 nm, and pSi 36 nm.

To confirm the proliferation capacities of DPSC on pSi, we investigated their The ratios of BrdU-positive cells were remarkably enhanced on semicarbazidetreated pSi (29.2 ± 6%) compared to glass coverslip (14.7 ± 4%, p<0.001), Flat Si (15.6 ± 6%, p<0.001) and APTES-treated pSi (19.5 ± 7%, p=0.011). The BrdU-positive cells rate was 14.7 ± 4% on glass coverslip. In comparison, this rate was similar on flat Si (15.6 ± 6%, p=0.756), slightly higher on APTES-treated pSi (19.5 ± 7%, p=0.142), and slightly lower on pSi 10 nm (11.8 ± 5%, p=0.331), on pSi 36 nm (13.0 ± 5%, p=0.555), on pSi 1 µm (9.6 ± 4%, p=0.075) and on undecenoic acid-treated pSi (10.9 ± 4%, p=0.189). To gain a more quantitative understanding of cell surface adhesion, the cell surface area and circularity were determined on the different substrates (Figure 3.8). Compared to glass coverslip, the mean cell surface area was similar on flat Si, pSi 10 nm and semicarbazide-treated pSi; it was higher on pSi 36 nm (p=0.040) and APTES-treated pSi (p<0.001); and it was lower on pSi 1 µm (p<0.001) and undecenoic acid-treated pSi (p<0.001). APTES-treated pSi had a significantly higher cell surface area compared to all other surfaces (Figure 3.8A). It is interesting to note that, on all surfaces except pSi 1 µm and undecenoic acid-treated, DPSC were spread and covered a very large area (between 1290 ± 316 µm2 on glass coverslip and 1958 ± 314 µm2 on APTES-treated pSi), which is consistent with the mesenchymal stem cells characteristics [START_REF] Krishna | Integrin-mediated adhesion and proliferation of human MSCs elicited by a hydroxyproline-lacking, collagen-like peptide[END_REF]. We also considered cell circularity which value varies between 0 (elongated cell) and 1.0 (perfect circle). Comparing directly circularity of cells on the various pSi substrates, morphology was found to vary significantly on 3 groups of substrates: DPSC had a similar circularity on glass coverslip (0. As expected according to results presented above, DPSC reached confluence on glass coverslip, flat Si and APTES-treated pSi, growing as clonogenic cell clusters (CFU-F: colony-forming units-fibroblastic) with a high proliferation rate. On pSi 10 nm and pSi 36 nm, DPSC were forming few CFU-F with a moderate proliferation rate, while DPSC were hardly growing on pSi 1 µm and undecenoic acid-treated pSi. The semicarbazidetreated pSi was clearly degraded, with DPSC attached only on the porous remaining substrate (Figure 3.9H), highlighting the fast resorbability of the pSi scaffold.

Discussion

The aim of this work was to elucidate the influence of pore morphology and chemical status of different pSi substrates on immediate human mesenchymal stem cell adhesion and on their proliferation rate. Therefore we compared the attachment, spreading, morphology and proliferation of DPSC seeded on different porous surfaces.

Our major finding was that pore geometry and surface chemical surface modification strongly influenced the subsequent behavior of DPSC that adhered and proliferated on pSi substrates. It was previously observed that cells could sense nanoscopic features on pSi substrates down to just a few nanometers (≈ 5 nm) and reacted differently to distinctive nanotopographic cues [START_REF] Gentile | Differential Cell Adhesion on Mesoporous Silicon Substrates[END_REF]) [START_REF] Buxboim | Matrix elasticity, cytoskeletal forces and physics of the nucleus: how deeply do cells "feel" outside and in?[END_REF] (Khung et al., 2003).

Considering topographical influence, our results showed a higher adhesion and higher proliferation rate for DPSC on pSi substrates with typical pore sizes around 36 nm, as compared to the other pore sizes, at all the considered time points. In term of morphology and spreading, DPSC seemed to react in a similar way on pSi 10 nm and pSi 36 nm as on other substrates such as glass coverslips, and flat Si wafers, with a wider surface area on pSi 36 nm indicating an enhanced cell spreading. On pSi 1 µm, DPSC were not able to proliferate after adhesion and to cover the surface. Even if flat surfaces (glass and flat Si), used as controls in this work, allowed a better cell adhesion compared to porous surfaces, the same proliferation profile was obtained for simple oxidized pSi surface as for the flat surfaces, with the advantage for the porous surfaces to be bioresorbable (Figure 3.2 and 3.4). Recently, it has been demonstrated, for four distinct cell types (endothelial cells, mouse fibroblasts, mouse neuroblastoma cells, and human cortical neurons), a higher adhesion on pSi substrates with pore size of 10 nm, as compared to flat silicon or pSi with pore size of 20 nm [START_REF] Gentile | Differential Cell Adhesion on Mesoporous Silicon Substrates[END_REF]. Our results clearly portrayed a different behavior for human mesenchymal stem cells and confirmed the fact that each cell line had different adhesion characteristics on various pSi surfaces at different time points [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF]. As mesenchymal cells, DPSC are involved in structural and fibrous tissues, and their preferential adhesion and proliferation on 36 nm pores could be related to the formation of focal adhesion complex: it has been shown that focal adhesions exhibited a complex multiscale architecture where nanoscopic, doughnut-shaped complexes (≈25 nm in diameter and spaced at ≈45 nm intervals) were distributed within the adhesion area [START_REF] Patla | Dissecting the molecular architecture of integrin adhesion sites by cryo-electron tomography[END_REF]. More recently, it has been reported that a 34 nm lateral spacing of adhesion peptides increased the formation of mature focal adhesions point (Frith et al., 2012). Thus, we hypothesize that DPSC, as other hMSC, have the ability to form mature focal adhesion on pSi with pores around 36 nm, explaining the differences observed in their attachment onto pSi 36 nm, pSi 10 nm and pSi 1 µm.

After investigating the pore size influence, pSi with 36 nm pores were used to investigate the influence of various surface chemical treatments, including silanization (after thermal oxidation) and hydrosilylation. Hydrosilylation with undecenoic acid offered a limited adhesion and proliferation rate. As already reported in previous works, with other cell types [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF]) (Alvarez et al., 2010), we showed that silanization with APTES encouraged cell attachment, spreading and proliferation, with a significant number of DPSC attached to the pSi surface and a proliferation rate higher than on glass coverslip or flat Si. We also demonstrated here, for the first time, the efficiency of semicarbazide grafted by hydrosilylation for DPSC adhesion and early proliferation. Indeed, semicarbazide-treated pSi permitted cell adhesion, spreading and proliferation similar to those found on APTES-treated pSi, with an even higher rate of mitotically active cells after 24 h. These results are in accordance with studies showing that amine-capped surfaces promoted cell attachment [START_REF] Faucheux | Self-assembled monolayers with different terminating groups as model substrates for cell adhesion studies[END_REF]. However, after 48 h, semicarbazide-treated pSi degraded and only few cells remained attached on the surface. In contrast, thermal oxidation, by creating Si-OH bonds at the pSi surface, significantly stabilized pSi in aqueous solution and functionalization with APTES further reduced the rate of hydrolytic dissolution. Thus, the semicarbazide grafting approach on pSi appears not to be used for long-term cell culture in vitro, but offers promising perspectives for cell attraction and immediate transplantation in vivo, as stem cell therapy often requires a scaffold to carry stem cells to the injured site in the body.

We have investigated three different pore sizes and three different chemical treatments to evaluate the use of pSi scaffolds for the adhesion and proliferation of primary culture of human mesenchymal stem cells from the dental pulp. We have identified, for the first time, two efficient amino-grafted pSi scaffolds for human mesenchymal stem cells adhesion and growth, with optimized pore diameter, for in vitro proliferation (and further differentiation), with an interesting potential for in vivo transplantation.

DPSC on pSi 36 nm were observed to have a better adhesion and a faster growth compared to pSi with smaller (10 nm) or larger (1 µm) pore size, in particular after silanization with APTES. Hydrosilylation with semicarbazide led to a new chemical modification favoring cell adhesion and proliferation, especially mitosis after cell adhesion. As this modified pSi surface was stable for only 24 to 48 h, it appeared to be potentially usable for stem cells adhesion and immediate in vivo transplantation, while APTES-treated pSi was more suitable for long term in vitro culture, for stem cells proliferation and differentiation.

More studies are on course to investigate: ( 1 effective if they are inert; they are now also expected to elicit a response from the body in order to favor their integration within the host bone [START_REF] Campion | Microstructure and chemistry affects apatite nucleation on calcium phosphate bone graft substitutes[END_REF]. Scaffoldbased bone tissue engineering aims to regenerate bone defects, using osteogenic/progenitor cells and a bioresorbable scaffold to create a scaffold-cell hybrid.

The combination of living cells, biologically active molecules and a structural scaffold to form a tissue-engineered construct to promote the regeneration of tissues is the most common concept underlying tissue engineering. The scaffold component is expected to support cell colonization, migration, growth and differentiation, so that it guides the development of the required tissue. In parallel with tissue formation, the scaffold may undergo degradation through the release of products that are either biocompatible or that can be excreted or subjected to metabolism [START_REF] Hutmacher | State of the art and future directions of scaffold-based bone engineering from a biomaterials perspective[END_REF]. Porous silicon shows great promise as a biointerface material due to the ability to precisely adjust its textural features, and due to its ability to degrade into non-toxic silicic acid in aqueous solution. In this chapter, we present the growth of primary cultures of DPSC, plated at confluence, for up to 21 days in culture with or without osteogenic induction. Our aim was to culture DPSC onto pSi in order to generate a mineralized matrix, mimicking bone formation. We followed induced or spontaneous osteodifferentiation onto porous silicon scaffolds and assessed the influence of silicic acid release and porosity.

Calcium-Phosphate formation onto pSi

In vivo grafting of a biomaterial initiates various inflammatory reactions. Bioinert materials are characterized by minimal corrosion and fibrous encapsulation in the surrounding tissues. On the contrary, pSi is considered to be bioactive as this biomaterial elicits physiological responses when implanted in living tissues [START_REF] Canham | Calcium phosphate nucleation on porous silicon: factors influencing kinetics in acellular simulated body fluids[END_REF] [START_REF] Low | Generation of reactive oxygen species from porous silicon microparticles in cell culture medium[END_REF] (Park et al., 2009). For bone tissue engineering, the ability of pSi to form direct bond with bone and soft tissues could reduce the inflammation due to implant micro-movements and promote integration within the body [START_REF] Canham | Calcium phosphate nucleation on porous silicon: factors influencing kinetics in acellular simulated body fluids[END_REF]). An essential requirement for such tissue-bonding ability is that pSi could induce precipitation of calcium phosphate (CaP) on its surface. Hydroxyapatite growth is a function of both composition and textural properties such as surface area, pore dimensions, density and porosity [START_REF] Campion | Microstructure and chemistry affects apatite nucleation on calcium phosphate bone graft substitutes[END_REF]. Canham and co-workers demonstrated that it was possible to grow hydroxyapatite on pSi surfaces in vitro over periods of days to weeks. Without defining an optimum porous silicon structure, they

showed that low porosity films could promote hydroxyapatite nucleation, whereas bulk silicon and macroporous silicon didn't grow any apatite nucleation [START_REF] Canham | Apatite Nucleation on Low Porosity Silicon in Acellular Simulated Body Fluids[END_REF]. Calcium-Phosphate deposits on pSi are shown in Figure 4.1. It was also reported that high porosity pSi structures (>70%) were rapidly corroded, thus preventing CaP nucleation, while low porosity films (<50%) were sufficiently stable to promote nucleation. Thus, both surface morphology and resorbability appeared to be two important factors to take in account for mineral Ca/P growth and further integration in the body.

Mineralization process

Mineralization for bone and dentin takes place within an organized matrix of type I collagen fibers, and the crystalline phase consists of aggregates of individual nanocrystals of hydroxyapatite, oriented and located specifically by the collagen fibril structure [START_REF] Tsai | The ionic products of bioactive glass particle dissolution enhance periodontal ligament fibroblast osteocalcin expression and enhance early mineralized tissue development[END_REF]. More generally, simple biominerals are composed of complex structures, which often have precise architectural order over several length scales. This architecture contributes to their superior strength and toughness. These biominerals rely heavily on their biological constituents, such as biomacromolecules, for regulation of mineralization [START_REF] Cölfen | Biomineralization: A crystal-clear view[END_REF]. Even in the case of bone and teeth, the details of biomineralization remain unclear. Recent works demonstrated that calcium-phosphate biominerals were formed via stable pre-nucleation clusters, with aggregation into an amorphous precursor phase and, subsequently, transformation of this phase into a crystal [START_REF] Cuisinier | High resolution electron microscopy study of crystal growth mechanisms in chicken bone composites[END_REF] [START_REF] Houllé | High resolution electron microscopy: structure and growth mechanisms of human dentin crystals[END_REF]) [START_REF] Dey | The role of prenucleation clusters in surface-induced calcium phosphate crystallization[END_REF].

These pre-nucleation clusters could be stabilized then turned into apatite with an active role of collagen, in the case of bone and tooth formation. These structures are mobile and negatively charged, while collagen fibrils exhibiting positively charged regions.

The nanosized, positively charged regions can be used for mineral infiltration, leading to the deposition of a dense network of pre-nucleation clusters and their subsequent transformation into amorphous calcium phosphate and, finally, oriented crystalline hydroxyapatite inside the fibrils (Nudelman et al., 2010). Thus, a disordered amorphous calcium phosphate phase (ACP) is the major mineral phase in the newly formed mineralized tissues, and this disordered amorphous phase is a precursor of mature crystalline hydroxyapatite (Cuisinier et al., 1993) (Houllé et al., 1998) [START_REF] Mahamid | Amorphous calcium phosphate is a major component of the forming fin bones of zebrafish: Indications for an amorphous precursor phase[END_REF]. Amorphous calcium phosphates (ACP) represent a unique class of calcium orthophosphate salts, in which there is no orientational ordering of the atomic positions.

They are described as the first solid phases precipitated in aqueous solutions containing Ca 2+ and PO 4 3-(Dorozhkin, 2010).

Silicon effect on osteodifferentiation

The dissolution behavior of implantable scaffolds in physiological solutions is of biological and clinical importance as biomaterials dissolution leads to graft resorption.

This dissolution is responsible of the material bioactivity, and depends on the physicochemical properties, porosity, and matrix ultra structure. Biomaterials resorption in vivo and in vitro is mediated by both a cell and solution phenomenon. Cellular activity affects the dissolution reactions and the material degradation can, in turn, affect the cell function (Gupta et al., 2010). Porous silicon, as a resorbable material, degrades in physiological solutions into silicic acid, the soluble form of silicon (Si). It was shown that silicon deficiency in animal led to abnormally shaped bones and defective cartilaginous tissue, with reversible effects when soluble Si was restored to the diet.

This knowledge led to the suggestion that Si may play an important role in connective tissue metabolism [START_REF] Carlisle | Silicon: an essential element for the chick[END_REF]) [START_REF] Schwarz | Growth-promoting effects of silicon in rats[END_REF]. Its main effect in bone and cartilage was thought to be on matrix synthesis rather than on mineralization, although its influence on calcification might be an indirect phenomenon through its effects on matrix components [START_REF] Seaborn | Effects of germanium and silicon on bone mineralization[END_REF]. More recently, it was observed that human osteoblasts cultured in media treated with ionic dissolution products from bioactive glass exhibited significantly upregulated IGF-II gene expression and protein synthesis and dissolved Si (silicic acid, in physiological concentrations) was shown to increase type 1 collagen synthesis and enhance osteoblast differentiation [START_REF] Reffitt | Orthosilicic acid stimulates collagen type 1 synthesis and osteoblastic differentiation in human osteoblast-like cells in vitro[END_REF]. However, the precise cellular mechanism by which silicon affected type 1 collagen synthesis was not elucidated. It was then shown that silicon-rich phases could provide a guided cell adhesion and tissue growth in vitro, and bone-cells attached to silicon-rich materials could absorb higher concentrations of calcium (El-Ghannam et al., 2006). Another recent study investigated the effects of dissolved silicon on osteoblast differentiation in vitro. Through the investigation of osteocalcin and osteopontin expression, it was demonstrated that Si played an important role in enhancing bone-cell differentiation in vitro (Gupta et al., 2010). This was confirmed by the significant role in the kinetics of mineralization and stromal cells differentiation played by resorbable porous silicon particles embedded in electrospun polycaprolactone. The presence of the pSi structures in such polycaprolactone composites could accelerate mesenchymal stem cell differentiation and the corresponding calcification process (Fan et al., 2010).

Aims of the study

Taking in account that silicic acid could enhance osteoblastic differentiation and that Ca/P formation was influenced by surface chemical composition and morphology, we investigated the osteodifferentiation of DPSC onto pSi surface. We aimed to elucidate whether DPSC could attach, proliferate and differentiate on pSi scaffolds for up to 3 weeks, until producing a mineralized extra cellular matrix. We wanted to evaluate the impact of Si release on osteogenic differentiation and on mineralization.

Finally, we focused on porous surface influence on Ca/P formation.

Materials

Human dental pulp stem cells

Human impacted third molar extracted for orthodontic reasons were recovered from healthy patients (15-30 years of age). Written informed consent was obtained from the patients or their parents. This protocol was approved by the local ethical committee (Comité de Protection des Personnes, Montpellier hospital). Tooth surfaces were cleaned using 0.12% chlorhexidine and cut around the cementum-enamel junction by using sterilized disc. The teeth were then broken into 2 pieces to reveal the pulp chamber. The pulp tissue was gently separated from the crown and root and then digested in a solution of 3 mg/ml type I collagenase and 4 mg/ml dispase for 1 hour at 37°C. The solution was then filtered through 70 µm Falcon strainers and immersed in αMEM supplemented with 10% fetal bovine serum (FBS), 100 U/ml penicillin, 100 µg/ml streptomycin and placed in 75 ml flasks. Cells were incubated for 1 week at 37°C with 5% CO 2 . Non-adherent cells were removed by a change of medium 24 h after cell seeding. After 1 week, sub-confluent cells were collected and analyzed for minimal criteria to define human mesenchymal stem cells.

Flow cytometry assays

The antigen profiles of cultured DPSC were analyzed by detecting the expression of the cell surface markers CD90, CD146, CD117 and CD45 using flow cytometry. CD90 is a widely accepted marker for mesenchymal stem cells, CD146 is a marker expressed in perivascular mesenchymal stem cells, CD117 is the receptor of stem cell factor, and CD45 is a marker of hematopoietic cells, mainly myeloid progenitors. The latter has been used to demonstrate the absence of contamination by CD45+ hematopoietic progenitors. For cell surface antigen detection, DPSC were detached with accutase, then incubated with fluorescence-conjugated antibody or the isotype control IgG for 1 h at 4°C in PBS containing 1% BSA. Antibodies used were the following mouse anti-human antibodies: CD90-APC conjugated, CD45-PerCP conjugated, CD117-APC conjugated and CD146-FITC conjugated. Analysis of 15,000 events was performed in an Epics XL4C (Beckman-Coulter). Data were analyzed with the FCS Express software.

Cell differentiation assays

Cells were controlled for multipotency with in vitro osteogenic, adipogenic and chondrogenic differentiation assays, following protocols described in chapter 2. Briefly, for osteogenic and adipogenic differentiation, dental pulp cells were seeded at a density of 10 5 cells/cm 2 and maintained in αMEM supplemented with 10% FBS until confluence. Cells were then cultured in osteogenic medium (αMEM supplemented with 15% FBS, dexamethasone, L-ascorbate phosphate, and β-glycerophosphate) or in adipogenic medium (αMEM supplemented with 10% FBS, dexamethasone, insulin, and isobutylmethylxanthine) for 21 days. Control cultures were maintained without osteogenic or adipogenic supplements. To induce chondrogenic differentiation, cells were cultured for 21 days as pellets in a serum-free expansion medium containing dexamethasone, L-ascorbate phosphate, L-proline, sodium pyruvate, ITS-Premix, and TGF-β3. Control pellet cultures were maintained without the addition of TGF-β3.

Porous silicon scaffolds preparation

P++ type boron-doped crystalline silicon wafers with 0.0008 -0.0012 Ωcm resistivity were obtained from Siltronix®. Wafers were etched in a custom-made Teflon cell at a constant current density of 300 mA/cm 2 for 2 min 15 s, in a hydrofluoric acid (HF) solution in ethanol (3:1 HF:ethanol solution, volume ratio). Etched wafers were oxidized at 800°C for 1 hour. The wafers were cut into 0.5 cm 2 pieces. We also worked with pSi presenting a gradient of porosity: these pSi gradients were prepared by anodization of highly doped p ++ type silicon wafers (0.0005-0.001 Ω cm, <100>, boron doped, obtained from Siltronix®) using a 1:1 HF:ethanol solution in a custom-built Teflon etching cell. A current density of 115 mA/cm 2 was applied for 60 s using a. For surface gradient preparation, the platinum electrode was placed perpendicular to the substrate and at one end of the etching cell (Clements et al., 2012) (Collins et al., 2002).

For a more precise description of the set-up, see chapter 5 (Figure 5.1). PSi gradients were thermally oxidized for 1 h at 800 °C, and silanized with aminopropyltriethoxysilane (APTES) 4 % in dry toluene at room for 1 h with gentle shaking. For use in cell culture, all the wafers were sterilized with 70% (vol/vol) ethanol for 10 min before drying under sterile air flow.

Surface characterization

The topography of surface modified pSi samples was analyzed by scanning electron microscopy (Analytic FEI Quanta FEG 200) to determine pore size, with an acceleration voltage of 20 kV, in a pressure of 0.5 Torr.

To determine surface tension, sessile drop contact angle measurements were conducted in a custom-made set-up consisting of a syringe dispenser, a sample stage, a macro lens and a CCD camera connected to a video-acquisition card. At room temperature, 3µl of MilliQ water was spotted onto the surface. Images of the drop profiles were captured and ImageJ® software was used to measure contact angles on both sides of the droplet. Solutions were sprayed through a high solid type nebulizer as plasma into a thermoelectrically controlled spray chamber. The results of the analysis were then compared with a standard for identification and quantification of Si concentration. The average and standard deviation of the Si concentrations per set of samples were reported as absolute concentrations, expressed in ppm. The Si concentrations were then calculated as mg/ml.

Osteodifferentiation onto pSi

DPSC were seeded at a density of 10 5 cells/cm 2 on porous silicon scaffolds and on oxidized unetched flat silicon wafers, and maintained in αMEM supplemented with 10% FBS for 5 days. Osteogenic induction was done as described in chapter 2, for 21 days. Flat silicon and culture plates were used as control. All experiments were conducted either in osteogenic medium (OM) or in basal medium (MEM). The osteogenic differentiation of DPSC was followed by the quantification of alkaline phosphatase (ALP) after 3 days, 7 days, 10 days and 14 days. Mineralization was assessed by the quantification of calcium deposits with Alizarin Red staining and by energy dispersive X-ray spectroscopy (EDX) after 21 days. The same experiments were conducted with bare pSi scaffolds, incubated without cells in culture medium (osteogenic medium or basal medium) at 37°C to reflect surface chemical reactions that occur in vivo.

Quantification of alkaline phosphatase (ALP) activity

ALP is a widely used marker, mainly expressed in differentiated cells producing mineralized matrix. After 3, 7, 10 and 14 days in differentiation medium, DPSC layers were rinsed with PBS, and scraped off into a 0.1% Triton X-100 solution. ALP activity was determined by the hydrolysis of p-nitrophenyl phosphate in 2-amino-2-methyl-1propanolbuffer, pH 10.4, at 37°C for 30 min. Absorbance was measured at 504 nm with a microplate reader. Total protein was quantified with the Bio-Rad protein assay kit.

ALP activity was expressed as nanomole of paranitrophenol per microgram of protein.

The experiments were performed in triplicate. The results were normalized according to the total protein contents. Statistical analyses were performed using a parametric Student t-test, and a p-value of <0.05 was considered to be significant.

Mineralization assays

After 21 days in differentiation medium, DPSC layers were rinsed with PBS.

Calcium deposits were stained with 40 mM pH 4.2 Alizarin red (AR) solution. Cells were rinsed three times with PBS and then fixed in 3.7% formaldehyde for 15 min.

Alizarin red was added to the wells for 1 h and rinsed five times to avoid non-specific staining. Cells were treated with 10% (w/v) cetylpyridinium chloride in 10 mM sodium phosphate pH 7.0 for 15 min at room temperature to remove the bound dye. The calcium concentration was determined by absorbance measurements at 562 nm with a spectrophotometer. Values were normalized to the total protein content of DPSC cultured under the same conditions, in a parallel assay. Results were presented as optical densities (OD) per microgram of total cellular proteins. Statistical analyses were performed using a parametric Student t-test, and a p-value of <0.05 was considered to be significant.

Calcium deposits within bone nodule-like colonies were assessed by calcein blue (CB) staining. This experiment was conducted for DPSC incubated on pSi gradients after 3 weeks of incubation in either osteogenic medium (OM) or basal medium (MEM). Samples were stained using a 30 µM calcein blue solution overnight in the dark. After staining, samples were rinsed with PBS, followed by fixation with 3.7% formaldehyde for 15 min. Cells were then counterstained with 50 µg/ml propidium iodide (PI) for 3 min at 37°C. The percentage of CB-positive colony area was defined by the area of CB-positive colony area divided by the total colony area in each image.

Images were taken by fluorescence microscopy, with excitation wavelength The colony area and CB-positive colony area were calculated using ImageJ software. Statistical analyses were performed using a parametric Student t-test, and a p-value of <0.05 was considered to be significant.

Results

Dental pulp stem cells characterization

A complete characterization of the recovered dental pulp cells is presented in chapter 2 (section 2.3 Results). Briefly, the cells that were extracted from human dental pulp matched all the criteria to be defined as mesenchymal stem cells. Thus, we could assume that we conducted our experiments with Dental Pulp Stem Cells (DPSC), as preciously named (Gronthos et al., 2000). Indeed, the recovered DPSC were plastic adherent cells, clonogenic and able to proliferate and differentiate into osteoblast, adipocyte and chondroblast. And they expressed, with various degrees of positivity, the cell surface markers CD90, CD117 and CD90, and they were negative for CD45 antigen (Figure 4.2).

pSi scaffolds characterization

PSi scaffolds were generated from bulk p-type silicon wafers (resistivity of 0.0008 -0.0012 Ωcm) by applying a constant current density of 300 mA/cm 2 . SEM examination showed that the porous substrates had an average pore diameter of 36 ± 4 nm, and were quite uniformly porosified. The surface porosity (ratio between total area of the pores and the area of the considered region of interest) was 45 ± 6%. SEM images of the pSi surface are shown in Figure 4.3, with top view and cross section. PSi gradients were prepared from highly boron-doped p ++ type silicon wafers (resistivity of 0.0005-0.001 Ωcm) by applying a constant current density of 115 mA/cm 2 was applied for 60s in a custom-built Teflon etching cell, with the platinum electrode placed perpendicular to the substrate, at one end of the set-up. SEM evaluation showed that the pore diameters were different along the scaffold, with largest pores under the electrode (electropolished area) and the smallest pores at the other extremity of the scaffold, in continuity with the flat silicon (unetched bulk wafer). The pore diameters were measured along the gradient, every 2 mm, starting from the electropolished area.

The pore diameter measurements are presented in Table 4

DPSC adhesion and viability on pSi

Cells adhesion and viability were assessed using vital FDA/PI staining. FDA enters into viable cells and stains their cytoplasm in bright green. PI interacts with RNA and DNA of dead cells, producing a red color [START_REF] Cao | Cell adhesive and growth behavior on electrospun nanofibrous scaffolds by designed multifunctional composites[END_REF]. Thus we could observe cell adherence, cytoplasmic spreading and vitality. As already developed in chapter 3, we showed that DPSC could adhere and spread on oxidized pSi scaffolds, with initial cell adhesion and growth very similar compared to the controls on flat Si. Morphologies of DPSC on pSi scaffolds are presented in The cell counts after 4 h, 24 h and 48 h confirmed that DPSC could attach and grow on oxidized pSi surfaces, as well as on flat Si. Initial cell adhesion on pSi was

A1 A2 A3 B1 B2 B3
slightly lower compared to flat Si and the number of DPSC counted on pSi was significantly lower after 24 h and 48 h, compared to flat Si (p=0.025 and p=0.002, respectively). However, DPSC started to proliferate on pSi after 24 h following the same growth pattern as the one observed on flat Si (Figure 4.5). More information and investigations about DPSC adhesion, spreading and proliferation on pSi are presented in chapter 3.

DPSC osteodifferentiation on pSi

To investigate the kinetics of DPSC osteoblastic differentiation on pSi scaffolds compared to flat Si, we assessed alkaline phosphatase (ALP) activity during the 2 first weeks of osteogenic induction. After only 3 days in osteogenic medium (OM), ALP activity of DPSC on pSi was slightly increased compared to cultures on flat Si. At day 7, ALP activity was again greater on pSi than values obtained for cells on flat Si. After 10 days, and even more after 14 days, ALP activity was significantly greater on pSi than on flat Si. These data are presented as histograms in Figure 4.7. It is interesting to note that DPSC differentiation was greater on pSi, while their original adhesion and proliferation was delayed compared to flat Si. This experiment was also conducted without osteogenic induction, in basal medium (MEM). A low ALP activity was detectable after 3, 7 and 10 days and ALP activity was slightly greater on pSi than on flat Si. After 14 days, this increased ALP activity of DPSC on pSi (in MEM) was statistically significant (Figure 4.7). the Ca deposits produced by DPSC were more abundant on pSi than on flat Si (p=0.004). These results were consistent with the increased DPSC osteodifferentiation on pSi scaffolds. We also investigated the formation of calcium deposits in the absence of cells, by incubating pSi scaffolds and flat Si without cells in culture medium (OM and MEM) for 3 weeks. We found the presence of a significant amount of Ca deposits on bare pSi when it was incubated in OM, while no Ca deposits was noticeable on bare

Calcein blue staining

In order to evaluate the influence of pore diameter on DPSC differentiation and on Ca deposits, DPSC were incubated on pSi gradients for 3 weeks, in OM. Calcein blue was chosen instead of alizarin red to be able to follow the fluorescence on the Psi gradients. Calcein blue staining indicated the extent of the mineralization at each stage of the gradient. Some CB-positive colonies were found along the whole gradient.

However, we demonstrated that there was a gradual increase in frequency and colony size along the gradient as the pore diameters decreased, with an exception on the area of 37 nm pores where CB-positive colonies (46 %) were more abundant than on areas of 24 nm (36 %) and 18 nm pores (44 %). The fraction of CB-stained cells was 40.3 ± 8.7

% on the flat Si, which value was comparable with that on the areas with 24 nm pore diameter and 70 nm pore diameter. The percentage of CB-positive colonies is presented in Table 4.2, and a summary of these results is shown in Figure 4.9. To investigate the influence of pore diameter on DPSC differentiation and mineralization process, we used pSi gradients as scaffolds for DPSC growth and further osteodifferentiation. As described above, DPSC were incubated for 3 weeks on pSi gradients, in either osteogenic medium (OM) or basal medium (MEM). After the 3 weeks, samples were fixed and dehydrated, then observed under SEM and analyzed with EDX spectroscopy. EDX analyses were realized on the whole corresponding SEM images (Figure 4.12). We discovered that Ca and P distribution were not homogeneous along the gradient: there was a gradual increase in amount of Ca and P (in atomic %)

along the gradient as the pore diameters decreased, except on 37 nm pores where Ca and P deposits were more abundant than on 24 nm and 18 nm pores. The lowest deposits were found on flat Si. The results were normalized according to the quantity of cells at each stage of the gradient, represented by the amount of nitrogen (N from the proteins, in atomic %). The same experiments were reproduced without cells, on bare pSi gradients incubated in OM, in order to investigate the influence of pore size on Ca-P deposits. As observed on bare pSi scaffolds (with 36 nm pore diameter), some mineralized structures were observable on pSi gradients. EDX spectroscopy confirmed the Ca-P composition of these structures. Ca and P distributions were not homogeneous along the gradient: as observed with DPSC on pSi gradient, there was a gradual increase in amount of Ca and P (in atomic %) along the gradient, as the pore diameters decreased (Figure 4.13). But unlike the previous experiment with DPSC on pSi gradients, CaP deposits were not more abundant on 37 nm pores than on 24 nm and 18 nm pores. In this work, we managed to grow DPSC for up to 3 weeks showing that oxidized pSi, without additional functionalization, was sufficiently stable for long-term cell growth, and could promote CaP precipitation. DPSC could attach and grow on oxidized pSi surfaces, as well as on flat Si, except that cell adhesion on pSi was lower compared to flat Si. However, DPSC started to proliferate on pSi after 24 h as they did on flat Si, with just a delayed growth (Figure 4.3 and Figure 4.4). As we wanted to explore pores influence on cell differentiation and mineralization, we didn't functionalize the pSi surface, except with thermal oxidation to stabilize the surface.

Chemical functionalization could enhance cell adhesion and proliferation [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF]) (Alvarez et al., 2009), and such functionalization strategies are presented in chapter 3. On pSi gradients, we silanized the surface with APTES because the scaffold was not stable enough to support cell growth for 3 weeks. After APTES treatment, pSi gradients were able to support cell adhesion and differentiation.

Si effects on DPSC osteodifferentiation

Cell differentiation was assessed in terms of ALP activity of DPSC cultured on pSi scaffolds at 3, 7, 10 and 14 days. ALP activity increased from day 3 to day 14 on pSi and flat Si in OM and after 10 days the ALP activity was significantly higher on pSi. At 10 and 14 days after induction, pSi scaffolds had significantly higher ALP activity compared to flat Si (p < 0.05). Even in MEM, we found a limited ALP activity for DPSC, with a significantly higher ALP activity on pSi after 14 days. PSi dissolution induced the presence of silicic acid (the soluble form of Si) in culture medium, while Si was hardly detectable when flat Si was incubated in culture medium. This is confirmed by the fact that ALP activity for flat silicon is of the same order than on plastic (data not shown). Previous studies have already shown that silicon-rich phase could provide guided cell adhesion and tissue growth (El-Ghannam et al., 2006) and it was suggested that Si could enhance bone-cell differentiation (Gupta et al., 2010). As alkaline phosphatase is a specific biomarker for mesenchymal stem cell differentiation, the levels of ALP expression and the correlated Si content recorded in our study demonstrate that pSi have the ability to enhance osteodifferentiation. The enhanced expression of ALP on pSi, even without osteogenic supplement (DPSC in MEM), confirmed that pSi dissolution was sufficient to stimulate osteodifferentiation in the absence of exogenous supplements. This finding is in accordance with a recent study that investigated and confirmed the ability of silicic acid to accelerate the differentiation of mouse stromal cells (Fan et al., 2011).

Concerning the resorption rate of pSi, we didn't find any evidence that cells interfered with pSi dissolution. Indeed, there was no difference between bare pSi and psi incubated with DPSC when measuring the Si content of culture media after incubation. This finding confirm a previous observation where bone marrow cells did not significantly affected the dissolution behavior of silica-calcium phosphate composite in vitro (El-Ghannam et al., 2006).

Mineralization and CaP structure

We postulate that high ALP activity related to DPSC osteodifferentiation could lead to early calcium incorporation. Indeed, it had been already reported that increased expression of ALP and osteocalcin led to earlier incorporation of calcium mineralized tissues [START_REF] Tsai | The ionic products of bioactive glass particle dissolution enhance periodontal ligament fibroblast osteocalcin expression and enhance early mineralized tissue development[END_REF]. Alizarin Red staining and EDX analysis demonstrate significantly more Ca and P deposits on pSi than on flat Si with or without OM.

However, such a difference was also reported when the experiments were conducted without cells, on bare scaffolds: there were significant Ca deposits (as assessed by Alizarin Red staining) and CaP deposits (as assessed by EDX spectroscopy) on pSi incubated in osteogenic medium, while only few deposits were found on flat Si. Images taken with SEM showed abundant mineralized formations on pSi, and not on flat Si, confirming the CaP nucleation role of the porous surface. It had already been demonstrated that pSi, unlike flat Si, had the ability to promote CaP (and even hydroxyapatite) nucleation, without elucidating the calcification mechanisms involved (Canham, 1995) [START_REF] Canham | Calcium phosphate nucleation on porous silicon: factors influencing kinetics in acellular simulated body fluids[END_REF]. We confirmed here that mesoporous silicon could behave as nucleation sites for CaP minerals. The driving force of nucleation is the supersaturation representing the free energy gains [START_REF] Kashchiev | Nucleation Basic Theory with Applications[END_REF]. The critical supersaturation for the heterogeneous nucleation of Ca/P on bare silica for hydroxyapatite is 7.6 mM (Ngankam et al., 2000). Porous silicon is used as an effective nuleator for protein crystals [START_REF] Chayen | Porous silicon: an effective nucleation-inducing material for protein crystallization[END_REF]. It is assumed that nucleation rate is dependent of relation between pore radius and protein 'stockes' radii [START_REF] Chayen | Experiment and theory for heterogeneous nucleation of protein crystals in a porous medium[END_REF]. The presence of nanoscopic pores on certain polymer surfaces led to order-ofmagnitude faster aspirin nucleation rates when compared with surfaces without pores [START_REF] Diao | Surface design for controlled crystallization: the role of surface chemistry and nanoscale pores in heterogeneous nucleation[END_REF]. Biocompatible polymer surfaces imprinted with nanopores were utilized as heteronucleants to influence the nucleation kinetics (Lopez-Mejias et al., 2013). A deep investigation on of the effect of pores on nucleation of carbonated hydroxyapatite on porous silica demonstrated that pore diameter is critical and that the surface chemistry (silanol and hydroxyl surface groups) are not involved [START_REF] Pereira | Mechanisms of hydroxyapatite formation on porous gel-silica substrates[END_REF]. This was confirmed by Canham et al., who studied hydroxyapatite nucleation on pSi in simulate body fluid [START_REF] Canham | Calcium phosphate nucleation on porous silicon: factors influencing kinetics in acellular simulated body fluids[END_REF]. Even if this work was confirmed by phenomenological studies [START_REF] Chadwick | Hydroxyapatite formation on metallurgical grade nanoporous silicon particles[END_REF]) [START_REF] Pramatarova | Porous silicon as a substrate for hydroxyapatite growth[END_REF], the influence of the length scale of the pore was never investigated as is was done for polymeric substraste [START_REF] Diao | Surface design for controlled crystallization: the role of surface chemistry and nanoscale pores in heterogeneous nucleation[END_REF]) (Lopez-Mejias et al., 2013).

We wanted to further investigate the nature of the CaP deposits, focusing at Ca/P molar ratio. We compared the Ca/P ratio with theoretical values of calcium phosphate.

According to the Ca/P ratio, the calcium phosphate particles formed on pSi without cells were octacalcium phosphate (OCP) (Ca/P = 1.341), while on pSi with DPSC they were closer to tricalcium phosphate (Ca/P = 1.418) [START_REF] Dorozhkin | Amorphous calcium (ortho)phosphates[END_REF]. Previous studies of calcium phosphate nucleation on Psi have used simulated body fluid with a different calcium phosphate ratio than the osteogenic media used.

Pore diameter effects on mineralization

After demonstrating the influence of Si content on differentiation and of pSi surface on mineralization, we wanted to assess the influence of pore diameter on mineralization. We reproduced the previously described experiments on pSi gradients, with pore diameters ranging from few nanometers to 500 nm and more. We seeded DPSC on pSi gradients (and flat Si as control) and followed differentiation and mineralization for up to 3 weeks. We assessed CaP deposits at each stage of the gradient by fluorescence microscopy (Calcein Blue staining) and SEM (EDX spectroscopy).

With fluorescence microscopy, we observed that there was a gradual increase in frequency of CB-positive colonies along the gradient as the pore diameters decreased, with an exception on the area with 37 nm pores diameter (Figure 4.8). A very similar pattern was found with the mean of EDX spectroscopy: there was a gradual increase in amount of Ca and P (in atomic %) along the gradient as the pore diameters decreased, with an exception on the area with 37 nm pores diameter. The maximum amounts of Ca and P deposits, assessed by both techniques, were recovered on the areas with pores diameter ranging from 8 ± 5 nm to 14 ± 7 nm. Deposits were the lowest on flat Si and larger pores (assessed by EDX spectroscopy). And deposits were more abundant on 37 nm pores than on 18 nm and 24 nm pores, which is not consistent with the global tendency of gradual increase along the gradient as pores diameters decrease (Figure 4.11).

To further investigate the effect of pores and of cells, we reproduced the experiments without cells, by incubating bare pSi gradient in OM for 3 weeks. EDX spectroscopy revealed a gradual increase in amount of Ca and P (in atomic %) along the gradient as the pores diameters decreased. The maximum amounts of Ca and P deposits were recovered on the areas with pores diameter ranging from 8 ± 5 nm to 14 ± 7 nm (Figure 4.12).

The area with pores diameter of 37 ± 18 nm was of particular interest as it showed a specific behavior only when pSi gradients were incubated with DPSC. During the 3 weeks of osteogenic induction, DPSC differentiate into osteoblast-like cells, likely to produce collagen extra cellular matrix. The collagen fibrils form a scaffold for a highly organized arrangement of oriented apatite crystals [START_REF] Orgel | Microfibrillar structure of type I collagen in situ[END_REF]. In the periodic 67 nm cross-striated pattern of the collagen fibrils, the less dense 40 nm long gap zone is implicated as the place where apatite crystals nucleate from an amorphous phase, and subsequently grow [START_REF] Mahamid | Mapping amorphous calcium phosphate transformation into crystalline mineral from the cell to the bone in zebrafish fin rays[END_REF]. Thus, we hypothesize that pSi Indeed cells with large adhesion areas (focal adhesion plaques) containing increased levels of adhesion molecules (such as talin and vinculin) show usually less migration and proliferation and express more differentiation markers [START_REF] Mann | Cell adhesion peptides alter smooth muscle cell adhesion, proliferation, migration, and matrix protein synthesis on modified surfaces and in polymer scaffolds[END_REF]. The extent and strength of cell adhesion (and subsequent cell proliferation and differentiation) depend on the physical and chemical properties of the biomaterial surface, such as surface morphology and micropatterning of the surface with celladhesive domains [START_REF] Bacakova | Modulation of cell adhesion, proliferation and differentiation on materials designed for body implants[END_REF]. Substrate topography has been shown to be an important factor for the regulation of cellular responses [START_REF] Tsai | The ionic products of bioactive glass particle dissolution enhance periodontal ligament fibroblast osteocalcin expression and enhance early mineralized tissue development[END_REF], and cell attachment, migration, proliferation and differentiation have been demonstrated to be effectively regulated by topographic features [START_REF] Wang | Modulation of alignment, elongation and contraction of cardiomyocytes through a combination of nanotopography and rigidity of substrates[END_REF], including pore diameter (Wang et al., 2012), with dimensions ranging from the nano to the microscale [START_REF] Lim | The regulation of integrin-mediated osteoblast focal adhesion and focal adhesion kinase expression by nanoscale topography[END_REF].

Cellular responses to nanostructured pSi have been investigated both in vitro and in vivo [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF] [START_REF] Sapelkin | Interaction of B50 rat hippocampal cells with stain-etched porous silicon[END_REF] (Whitehead et al., 2008) (Alvarez et al., 2009) (Fan et al., 2011). Porosity and pore size can be tuned by varying the etching conditions to generate surfaces with pore size ranging from a few nanometers to micrometer [START_REF] Sailor | Porous Silicon in Practice[END_REF] and pSi surface can be modified by oxidation or various chemical treatments, such as silanization or hydrosilylation (surface functionalizations are described in Chapter 1 and Chapter 3). We investigated the effect of pore diameter on DSPC adhesion, proliferation and further differentiation. As proliferation and differentiation are related, we aimed to determine optimal pores diameter for DPSC proliferation and for osteodifferentiation. For this purpose, we used pSi gradients with pore diameters ranging from <10 nm to >500 nm. These pSi films, containing a controlled distribution of pore sizes, are produced by using an inhomogeneous etching technique (Collins et al., 2002). These pSi gradients have been used to assess cell attachment, with various cell types, such as neuroblastoma cell lines, for which attachment and morphology were shown to critically depend on pore size [START_REF] Khung | Using continuous porous silicon gradients to study the influence of surface topography on the behaviour of neuroblastoma cells[END_REF] and rat mesenchymal stem cells, for which pore size of 20-30 nm enhanced cell attachment (Clements et al., 2011) (Clements et al., 2012). However, different cell types respond differently to topographical features [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF]. In addition, nanotopography influence cell proliferation through conformation of adhesive proteins and cell focal adhesion [START_REF] Popat | Osteogenic differentiation of marrow stromal cells cultured on nanoporous alumina surfaces[END_REF], and human mesenchymal stem cell differentiation (Dalby et al., 2007). Therefore we investigated primary human stem cells adhesion, proliferation and differentiation on such pSi gradients. DPSC were isolated as previously described (Gronthos et al., 2002). Briefly, tooth L-glutamine, 100 IU/ml penicillin, and 100 µg/mL streptomycin), then incubated at 37°C in 5 % CO 2 .

Materials and Methods

Human

Porous Silicon Gradients

Porous silicon (pSi) gradients were prepared by anodization of highly doped p ++ type silicon wafers (0.0005-0.001 Ω cm, <100>, boron doped, obtained from Siltronix®) using a 1:1 HF:ethanol solution in a custom-built Teflon etching cell using aluminium foil to contact the back of the silicon wafer and a Pt mesh as the counter electrode. A current density of 115 mA.cm -2 was applied for 60 s using a Keithley 2425 sourcemeter. For surface gradient preparation, the platinum electrode (1 mm diameter, Sigma-Aldrich) was placed perpendicular to the substrate and at one end of the etching cell (Clements et al., 2012) (Collins et al., 2002); the etching set up is presented in

Surface characterization

Gradient pSi surfaces were characterized by means of scanning electron microscopy (SEM) and interferometric reflectance spectroscopy (IRS), to determine the pore diameter and the porosity of the layer along the surface of the gradient. Samples were observed after preparation, before any cell experiment, and after 21 days in culture medium (αMEM) to assess their stability in aqueous solution. The SEM analysis was conducted using a FEI Helios SEM (10 kV accelerating voltage, under a pressure of 1.3

x 10 -3 Pa). For IRS, a bifurcated optical cable was used to direct white light from a tungsten lamp source to the pSi gradient at normal incidence and the reflected light was collected and transmitted through the other arm of the bifurcated fiber optic cable to a CCD spectrometer.

Atomic Force Microscopy (AFM)

For atomic force microscopy (AFM) cell imaging, we used an Asylum MFP-3D

AFM system, mounted on an Olympus inverted microscope. Gold-coated silicon nitride rectangular cantilevers were used with a typical spring constant of 30 pN/nm -1 , tip radius ~30 nm. The spring constant for each cantilever was determined by thermal noise method within the supplied software. DPSC were examined after 24 h of in vitro incubation on pSi with pores diameter of 36 ± 4 nm or on flat Si. Cells were fixed in 2.5

% glutaraldehyde for 1 h, and then rinsed 3 times in PBS. AFM topographic images were obtained in contact or tapping mode in liquid (PBS) at an average temperature of 30°C.

Cell adhesion

DPSC attachment was monitored by fluorescence microscopy. Cells were seeded onto the surface of sterilized pSi at a cell density of 5 x 10 4 cell/mL. Flat silicon (non etched silicon wafers) was used as control. Cells were incubated for 4 h, 24 h, 48 h and 72 h at 37 °C with 5 % CO 2 , in a humidified incubator. For cell counts, samples were fixed with 2.5 % glutaraldehyde for 1 h, and stained with 50 µg/ml of Hoechst 33342 for 30 min to stain nuclei. Samples were then washed with PBS buffer to remove any non-adherent cells and unbounded dye, and observed under fluorescence microscopy at an excitation wavelength of 360 nm. Cell counts were conducted every 2 mm along the pSi gradients, starting from the area under the electrode (electropolished area). Three measurements were conducted in each region over an area of 1400 µm x 1050 µm. All experiments were performed in triplicate, and statistical differences along the gradients were determined using the Student t-test with p < 0.05 deemed as statistically significant.

Cell morphology and spreading

To evaluate cell morphology and spreading, DPSC were incubated at a density of 1x10 5 cell/mL for 24 h on pSi gradients, and stained for actin, nuclei, and vinculin.

After 24 h of incubation, cells were rinsed three times with PBS, fixed with 4 % paraformaldehyde for 20 min, then washed three more times to remove excess paraformaldehyde. DPSC were permeabilized with 0.5 % Triton X-100 in PBS at 4 °C for 15 min. In order to reduce non-specific background, samples were blocked with PBS/BSA 1 % for 10 min at 37 °C. After blocking, PBS/BSA was aspirated and the samples were first incubated for 1 h with TRITC-labeled phalloidin (1:200) at 37 °C in the dark. Mouse anti-vinculin antibody, at a dilution of 1:100 in 1 % BSA/PBS, was added to each sample for 1 h at 37 °C in the dark, then the samples were incubated for 30 min with secondary rabbit anti-mouse antibody at a dilution of 1:200 at room temperature. Finally, samples were incubated with Hoechst 33342 for 10 min at room temperature, and washed twice with deionized water. Samples were observed under fluorescence microscopy at an excitation wavelength of 360 nm for Hoechst staining (nucleus staining), 490 nm for vinculin and 630 nm for phalloidin (actin staining).

After staining the actin, vinculin and nuclei, fluorescent images were analyzed using imageJ software. Actin staining was used for identification of cell shapes into either round, branched or spindle cell morphology. The circularity of the cells was determined by using the length/width ratio. Round cells were defined to have a length/width ration of > 0.8, while spindle cells had length/width ratio below 0.8.

Branched cells were characterized by more than two branches in their morphology. Cell areas were also measured for each type of cell morphology, giving insights into the spreading of cells on surfaces. The number of focal points was counted per cell for each type of cell morphology and normalized by the cell area for each cell morphology.

Proliferation assays

The proliferation rate of DPSCs on pSi gradients was assessed by bromodeoxyuridine (BrdU) incorporation for 24 h, 48 h and 72 h. BrdU incorporation is used as a marker for mitotically active cells. BrdU is a synthetic nucleotide analogue of thymidine, incorporated into the newly synthesized DNA of replicating cells (during the S phase of the cell cycle), substituting for thymidine during DNA replication [START_REF] Horner | Proliferation and differentiation of progenitor cells throughout the intact adult rat spinal cord[END_REF]. DPSC were seeded on the scaffolds at a density of 1x 10 5 cells/mL in αMEM supplemented with 10 % FBS with 1:100 diluted BrdU labeling (Invitrogen). After 24 h, 48 h and 72 h incubation at 37 °C, samples were rinsed 3 times in PBS, fixed in 4 % paraformaldehyde, rinsed again 3 times in PBS and put in 1.5 M HCl (Sigma) for 30 min at room temperature to dissociate DNA strands. Samples were washed with PBS, incubated in PBS with 0.1 % Triton X-100 and 1 % BSA for 1 h at room temperature, then incubated with mouse anti-BrdU primary antibody overnight at 4 °C. After immunostaining, cells were washed with PBS/1 % BSA, incubated with FITC conjugated rabbit anti-mouse secondary antibody for 30 min and counterstained with 2 µg/ml Hoeschst 33342 for nucleus staining. Samples were observed under fluorescence microscopy at an excitation wavelength of 290 nm for nuclei (blue staining) and 490 nm for BrdU (green staining). All experiments were performed at least in triplicate and the number of BrdU-positive cells was expressed as a percentage of the total number of cells.

Cell differentiation

DPSCs were seeded at a density of 1 x 10 5 cells/cm 2 on pSi gradients and maintained in αMEM supplemented with 10 % FBS for 72 h, until they reached confluence. Osteogenic differentiation was then induced by culture in osteogenic medium (αMEM supplemented with 15 % FBS, dexamethasone, L-ascorbate phosphate, and β-glycerophosphate) for 21 days. Control cultures were maintained without osteogenic supplements and flat silicon was used as control. The osteoblastic differentiation of DPSC was followed, after 7 days, by the quantification of alkaline phosphatase (ALP) and osteocalcin (OC) by immunostaining. After 21 days, the osteodifferentiation was assessed by type I collagen immunostaining and by infrared spectroscopy (FTIR).

FTIR microscopy

After 21 days in osteogenic medium, pSi gradients were characterized with infrared microscopy (FTIR), to assess DPSC osteodifferentiation along the gradients.

FTIR studies were performed with a Bruker Hyperion 1000 IR microscope operating with a Bruker Vertex 80 IR spectrometer. Along the whole gradient, reflectance spectra were collected over 64 scans, fram 650 to 4000 cm -1 , with a resolution of 4 cm -1 . All spectra were background corrected with an unetched silicon wafer of the same type.

Spectra were recovered with OPUS 6.5 software and analyzed with Matlab software for averaging spectra and MagicPlotPro software for deconvolution.

Data analysis

Data was reported as the mean ± standard error. All post-acquisition image processing was conducted with ImageJ software. Statistical analysis was performed using SigmaStat. All data were checked for normalization. Data were analyzed using parametric Student t-test, and for ANOVA, non-parametric Tukey test was performed.

For all tests, significance was determined to be p < 0.05

Results

Surface characterization

The pSi gradients were fabricated by anodization at a current density of 115 mA.cm -2 , using an asymmetric electrode configuration and a 1:1 HF/ethanol electrolyte solution, as previously described (Clements et al., 2012) The topography and the average pore size along the pSi gradients were characterized by means of scanning electron microscopy (Figure 5.3). The average pore size measured at 2 mm intervals decreased with a logarithmic profile from the region near the electrode to the other end of the surface. The non-linear nature of this profile was attributed to the non-linear decrease in current density across the substrate, in accordance with previous studies (Clements et al., 2012) (Collins et al., 2002). The pore sizes along the gradients are presented in Table 5.1: in the region closest to the Pt electrode, the largest pores were observed with a pore size of 1022 ± 353 nm. The pore size decreased to 500 ± 208 nm at a position 2 mm from the electrode, 119 ± 44 nm at a position 6 mm and 37 ± 18 nm at a position 10 mm from the electrode. At the extreme 5.1 (next page). Both porosity and pore depth decreased with increasing distance from the electrode. After 21 days in αMEM, pores were still visible on pSi gradients, demonstrating the stability of the porous layer for up to 3 weeks, in the conditions of the in vitro experiments (Figure 5.5). From the fluorescence images, we assessed cell spreading by considering the cell surface area (area covered by the cell, in µm 2 ) and the cell circularity (ratio between the shorter and the longer axis of the cell, between 0 for elongated cells and 1.00 for round cells). As observed with fluorescence images, cell surface area dramatically decreased on large pores (pores diameter ≥ 120 µm) and cell circularity was significantly higher on these large pores. It was interesting to note that cell surface area gradually increased while pores diameter decreased and cell circularity gradually decreased while pores diameter increased; with an important exception in the middle of the gradient, at 10 mm from the electrode (pores diameter = 37 ± 18 nm). On this specific area, with pores diameter roughly ranging from 20 to 50 nm, cell spreading was enhanced compared to other regions with smaller or larger pores (Figure 5.8).

AFM: Filopodia and lamellipodia formation

We were interested in filopodia formation for DPSC growing on pSi samples, as this gives information about cell behavior since these cellular projections contribute to intercellular communication, cell adhesion and motility [START_REF] Mattila | Filopodia: molecular architecture and cellular functions[END_REF]. After 24 h on flat Si, DPSC developed short filopodia at their apical pole, spreading from lamellipodia. After 24 h on pSi samples, DPSC developed such filopodia spreading from lamellipodia, but they also developed long and thin protrusions, growing directly from the cellular body. These findings were controlled by fluorescence microscopy with actin staining; the long and thin protrusions were formed by actin, confirming them as filopodia. These results are presented in Figure 5.10 (next page). BrdU assay revealed an increased proliferation rate on areas with pores diameter ranging from 18 to 70 nm. These areas (18 to 70 nm) were only partially matching with the areas where we counted the more cells after 72 h (flat Si and pores ≤ 37 nm) (Figure 5.13, next page). This apparent contradiction is quite difficult to interpret, but reinforce the fact that cells are influenced by the pores diameter at the nanometer scale, for early attachment, adhesion process and proliferation. On large pores (diameter ≥ 250 nm), adhesion and proliferation were clearly disadvantaged, compared to smaller pores. After 21 days in osteogenic medium, DPSC osteodifferentiation along the pSi gradients was assessed with infrared spectroscopy (FTIR). Spectra acquisitions were performed with a microscope that scanned the whole surface of the pSi gradients, permitting to create a mapping of the surface. Figure 5.18 shows single representative FTIR spectra corresponding to various stages of the pSi gradients (freshly prepared scaffold). On bare scaffolds (pSi gradients before any cell experiment), we described 3 main peaks. The major one (between 1070 to 1090 cm -1 ) corresponds to asymetric stretching of Si-O-Si groups. Its usual wavenumber have been described between 1000

to 1100 cm -1 , according to the crystal phase of silicon dioxide [START_REF] Pisciella | FTIR spectroscopy investigation of the crystallisation process in an iron rich glass[END_REF], and as pSi is a crystalline material, this peaks was found here to be closer to 1100 cm -1 .

The second peak, at 810 cm -1 , corresponds to symetric stretching of Si-O-Si groups. The third peak, at 1250 cm -1 , represents C-N stretching that corresponds to APTES grafting on the pSi surface. Concerning proteins peaks, several absorption bands were present, corresponding to carbonate bands (at 870 cm -1 , 1420 cm -1 and 1445 cm -1 ) and amide bands (C-N and C=O vibrations at 1250 cm -1 , 1420 cm -1 , 1445 cm -1 , 1540 cm -1 and 1650 cm -1 ) with the most intense peak at 1650 cm -1 [START_REF] Figueiredo | Characterization of Bone and Bone-Based Graft Materials Using FTIR Spectroscopy[END_REF][START_REF] Dutilleul | Root canal hydrophobization by dentinal silanization: Improvement of silicon-based endodontic treatment tightness[END_REF].

Vibration bands and corresponding molecules are presented in table 5.2. From the FTIR spectra, we selected the area ranging from 960 to 1025 cm -1 for phosphate and from 1250 to 1650 cm -1 for proteins, and we created a 3D graph representing the amount of proteins and the amount of phosphate on the different regions of the gradients. The proteins distribution was quite homogeneous on the whole gradients, indicating that we had a regular cell density along the gradient after 21 days, with cells covering all surface (Figure 5.22). The phosphate distribution revealed a maximum in the middle of the gradients, in the area at 8 to 10 mm from the electrode (pores diameter ranging from 37 to 70 nm) (Figure 5.23). showing that the amount of phosphate deposition is maximized on the middle of the gradient.

Discussion

In this work asymmetric anodization give gradient of porosity used to study the influence of topography (pores diameter) on cell behavior. The dispersion of pore sizes, was consistent with other published studies (Collins et al., 2002) (Clements et al., 2012) from few nanometer to 1 µm. Upon thermal oxidation, the surface chemistry of the pSi is similar to the chemistry of glass, and glass is a common surface for cell culture studies [START_REF] Khung | Using continuous porous silicon gradients to study the influence of surface topography on the behaviour of neuroblastoma cells[END_REF]. However, after thermal oxidation, pSi gradients were not stable enough to support cell growth for more than 3 to 5 days. Surface treatment by silanization (grafting of APTES) stabilized the scaffolds and allowed long-term cell growth.

DPSC adhesion and spreading according to pores diameter

It has been described that porosity played a regulatory role in determining cell spreading and migration by changing the access to the binding sites on the substrate [START_REF] Miron-Mendoza | The differential regulation of cell motile activity through matrix stiffness and porosity in three dimensional collagen matrices[END_REF]. Cytoskeleton staining and fluorescence microscopy showed that DPSC adhered and spread well on smaller pores while they were round or less spread on larger pores. In the middle of the gradient (pores diameter = 37 ± 18 nm) cell spreading was enhanced compared to regions with smaller or larger pores. This pores diameter increase cell organization as assessed by cell spreading and corresponds to biological structures in human body, such as the pore size recovered in glomerular basement membrane [START_REF] Gagliardini | Imaging of the porous ultrastructure of the glomerular epithelial filtration slit[END_REF] or the gap region between collagen fibrils in type 1 collagen [START_REF] Shoulders | Collagen Structure and Stability[END_REF].

To evaluate the role of adhesion molecules in nanopores-induced cell attachment and further focal adhesion formation, immunostaining specific to vinculin was performed. Vinculin is an integrin that supports adhesion to the ECM, stabilizes protrusions through structural connections by bridging focal adhesions to the cytoskeleton, and activate migration-related signaling molecules [START_REF] Huttenlocher | Integrins in Cell Migration[END_REF]. Mature focal adhesions were recovered on flat Si and pSi with pores diameters ≤ 70 nm, while they were hardly visible on larger pores (Figure 5.9). These findings are in accordance with a recently published study investigating the effects of pores on the growth and function of fibroblasts [START_REF] Pan | The spatial and temporal control of cell migration by nanoporous surfaces through the regulation of ERK and integrins in fibroblasts[END_REF]. They evaluated fibroblast growth on a stainless steel porous material with pores ranging from 40 to 210 nm: vinculin was widely distributed within cells grown on flat surfaces and the on 40 and 75 nm nanopores surfaces and decreased in cells grown on 185 and 210 nm nanopores surfaces. Focal adhesion, cell spreading and subsequent proliferation are closely related to the surface characteristics, even though the detailed mechanics through which the geometry and size of pores enhance cell attachment and proliferation are still unclear [START_REF] Zhu | Effects of topography and composition of titanium surface oxides on osteoblast responses[END_REF]. We hypothesize that nanopores may affect cell adhesion by promoting an interaction between vinculin/integrin anchoring sites and the edges of pores for cell attachment. Thus, focal adhesions that were formed on pSi with pores >70 nm were aberrant, whereas we recovered mature focal adhesions on pores ≤ 70 nm. This hypothesis is supported by previous studies reporting that focal adhesions were enhanced when cells were cultured on pore surfaces with critical feature sizes smaller than 50 nm [START_REF] Hu | Cell culture on AAO nanoporous substrates with and without geometry constrains[END_REF] and the negative cellular response on larger pores up to 70 nm was ascribed to a lack of effective integrin clustering, which therefore inhibits the formation of focal adhesions and actin fibre networks (Huang et al., 2009).

As we clearly demonstrated higher adhesion and spreading on pores ranging from 20 to 50 nm, we investigated cell spreading on this specific region by AFM. Long and thin filopodia growing from the cell body were observed, with size related to pores diameter. Cells attach to and pull on the matrix through focal adhesion and cytoskeleton during migration and filopodia act as cellular tentacles [START_REF] Mattila | Filopodia: molecular architecture and cellular functions[END_REF]. It has been found that actin-dependent retraction of the filopodia occurred with discrete steps with an average size of 36 ± 13 nm with a myosin motor involved [START_REF] Kress | Filopodia act as phagocytic tentacles and pull with discrete steps and a loaddependent velocity[END_REF].

Therefore we could hypothesize that cell adhesion and spreading was more efficient on 20-50 nm pores than on smaller pores due to the correspondence with filopodia retraction steps: at each retraction step, the integrin-binding complexes are displaced from a pore edge to the next edge, to form more easily a mature focal adhesion.

DPSC proliferation

The highest rate of proliferative cells (BrdU positive cells) was recovered in the middle of the gradient, on the region with 37 ± 18 nm. The correlation of the switch between the proliferation and differentiation programs and the maturation of the focal adhesion remains to be fully elucidated [START_REF] Bacakova | Modulation of cell adhesion, proliferation and differentiation on materials designed for body implants[END_REF]. Considering the important role of integrin-mediated focal adhesion and intracellular signaling in anchorage-dependent cell function, a previous study evaluated the influence of nanoscale pit texture. This study suggested that nanostructured physical signals could regulate cell function [START_REF] Lim | The regulation of integrin-mediated osteoblast focal adhesion and focal adhesion kinase expression by nanoscale topography[END_REF]. Thus, we assume that the mature focal adhesions and higher cell spreading recovered for pores ranging from 20 to 50 nm led to a higher cell proliferation in this region, by favoring cell function.

DPSC osteodifferentiation on pSi gradients

After 3 weeks in osteogenic medium, DPSC gradually differentiated into osteoblast-like cells. ALP staining revealed that DPSC differentiated earlier on 37 ± 18 nm pores. High cell density might increase the availability of secreted bioactive molecules and result in higher differentiation potential [START_REF] Tang | The regulation of stem cell differentiation by cell-cell contact on micropatterned material surfaces[END_REF] and this is in this region that higher spreading and cell density was noted (see previous sections).

However initial cell density was similar after 3 days on flat Si and pSi with pores diameter < 70 nm and this shows that differences of osteodifferentiation are due to pore diameter. Collagen immunostaining and calcium phosphate accumulation confirmed the enhanced osteodifferentiation on the 37 ± 18 nm pores.

Chapter 6 General discussion and perspectives

General discussion

Regenerative medicine, and more specifically tissue engineering, is based on the use of cells preferably stem cells, bioactive molecules and scaffolds. Designing a material that induces a favorable response for stem cells is important in biomaterials research. In the thesis, short-term and long-term performances of human MSC from the dental pulp on porous silicon fabricated by electrochemical etching were investigated.

Dental Pulp Stem Cells have been extensively studied in vitro and in vivo and they represent a promising source of stem cells. This is because they are easily accessible when a tooth is extracted, without any supplementary invasive procedures or ethical issues. We developed an efficient stem cell recovery system where porous silicon scaffolds supported DPSC adhesion, proliferation, and viability while at the same time functionalization with APTES and semicarbazide allowed for enhanced cell adhesion and early proliferation. Semicarbazide treatment led to a highly resorbable scaffold while APTES-treated pSi was stable for long-term culture. Stem cell based tissue engineering is considered according to two strategies: in vitro expansion and differentiation before grafting, or immediate grafting of the cell-scaffold complex. The two-studied functionalization permitted to address the two strategies.

In addition, we cultured DPSC on pSi surfaces for a period of 3 weeks to investigate the long-term effect of silicon nano-architecture on their osteogenic potential. Cells cultured on pSi surfaces demonstrated higher ALP activity, suggesting a faster osteodifferentiation compared to flat Si. Furthermore, the calcium and phosphorous concentrations were higher compared to flat Si, confirming that DPSC differentiated faster and produced more mineralized matrix. This enhanced differentiation was attributed to silicic acid release and porous structure: silicic acid induced a faster cell differentiation and the porous surface enhanced CaP nucleation compared to flat surface (higher CaP deposits even without cells).

In order to demonstrate the influence of pore size, we used pSi gradients as a screening tool. We investigated short-term DPSC attachment, spreading, proliferation and further differentiation on these gradients of porosity. This investigation aimed to design biomaterials with optimal porosity, able to enhance DPSC adhesion and guide their differentiation. We demonstrated that pSi with pores ranging from 20 to 50 nm enhanced cell adhesion, proliferation and osteodifferentiation. Unlike the belief that biological mineralization is mainly cell mediated, we demonstrated that nanoporosity was of primary importance for CaP nucleation, with or without osteoblast-like cells.

Without cells, the maximum amounts of Ca and P deposits were recovered on the areas with pores diameter ranging from 8 ± 5 nm to 14 ± 7 nm. With DPSC, the pores diameter of 37 ± 18 nm led to the highest osteodifferentiation and mineralized matrix formation. ACP nucleation started with nanometer-sized CaP complexes, and these prenucleation clusters aggregated to form amorphous calcium phosphate, such as OCP.

We hypothesized that pSi with 37 ± 18 nm pores favor ACP transformation and crystal growth during extra cellular matrix formation. This research showed that tuned pSi scaffolds had a significant influence on the DPSC differentiation into mature osteoblast-like cells and on the kinetics of mineralization. Stem cell performance could be significantly improved using controlled nano-topography. These pSi scaffolds showed interesting potential applications in biomedicine: the nanoscale pore surfaces were able to influence the expression of integrin molecules and directed cell behaviors such as spreading, proliferation and focal adhesion formation. Silicic acid release and nano-patterned surfaces enhanced DPSC differentiation and induced calcium phosphate deposits to nucleate in vitro, thus promoting biomineralization process.

The complex mechanisms of calcification remain partially unknown but our results suggest that pSi could improve bonding between grafted biomaterial and tissue in vivo. However, the in vivo situation is probably much more complex than suggested by the in vitro experiments. The problem of Ca and PO 4 delivery to the mineralization front, the grafted stem cell survival, the scaffolds degradation kinetics and the possibility to use pSi as a drug delivery carrier are all issues that should be studied in in vivo models.

Perspectives

Stem cell therapies represent a new approach for the treatment of various diseases. Mesenchymal stem cells can operate through many different mechanisms and have generated great therapeutic expectations. However, their use is limited by the way to bring cells to the site of interest. When directly injected to the injured site, it is difficult to evaluate if cells remain in the correct site. When injected intravenously, cells cannot be followed and directed to their target site, with the risk of being stuck in pulmonary capillary vessels. Thus, it is admitted that tissue engineering requires a scaffold to carry cells to the injured site. The principle of bone tissue engineering is the fabrication of a complex (scaffold and stem cells) that matches the physical and biological properties of the natural bone tissue, without the disadvantages of autologous or allogenic bone grafts. Mesenchymal stem cells represent an advantageous source of stem cells for tissue regeneration, due to their favorable properties (regeneration of damaged tissues, minimal donor-site morbidity and low risk of autoimmune rejection).

As biomaterial, pSi has been successfully investigated in vitro and hold promises for biomedical applications. Its compatibility and degradation have been already assessed in vivo. This biomaterial has now to be tested in vivo to carry stem cells for regenerative medicine, as chips, microparticles or nanoparticles. The porous structure allows the loading of the scaffold with bioactive molecules, and the optical properties permit to follow the grafted complex in vivo. Further loading with iron nanoparticles could even permit live tissue imaging of the grafted complex with magnetic resonance imaging or or embryonic stem cells (ESC) with the textural and chemical properties of the PSi substrates provides an interesting approach for therapeutic use. Moreover, the PSi nanostructure and the release of silicic acid have positive effect on precursor cell osteodifferentiation and mineralized matrix formation.

Key words: porous silicon, mesenchymal stem cells, cell adhesion, osteodifferentiation, mineralization.

Introduction

Numerous studies have been carried out to verify that stem cells could become a source of stable differentiated cells capable of inducing tissue formation, including mineralized hard tissue. Stem cells for hard tissue formation have received considerably increased attention from researchers, as these cells can be an interesting source of stable differentiated cells, capable of inducing bone formation and control hydroxyapatite crystal growth. Stem cells of various origins are able to produce in vitro calci ed nodules and to form a mineralized tissue after transplantation in vivo . Supporting scaffolds for bone tissue engineering are often required to lead cell implantation, growth and differentiation (D [START_REF] Aquino | Human postnatal dental pulp cells co-differentiate into osteoblasts and endotheliocytes: a pivotal synergy leading to adult bone tissue formation[END_REF]Graziano et al. , 2008 These scaffolds can be inductive or instructive to direct stem cells down speci c lineage pathways and increase the therapeutic effect. Thus, directed stem cells have the potential for rapid bone regeneration and repair. Since the advent of tissue engineering, bone has received particular interest, since it is one of the tissues with most regenerative abilities in the human body.

Porous silicon (PSi) is a promising biomaterial for bone tissue engineering, as it is non-toxic, bioresorbable, and can be inductive for osteodifferentiation. It has the ability to degrade in aqueous solutions into non-toxic silicic acid and surface modi cations offer control over the degradation rate of PSi [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF]. Moreover, this tunable, biocompatible and resorbable material has been reported for the growth of hydroxyapatite, inferring the possible bone implantability of the material (Canham, 1995).

Stem cells for bone tissue engineering: adult, neonatal and embryonic stem cells (ESCs)

The rst cell-based strategy for bone repair was autologous connective tissue progenitors from the iliac crest. This autograft procedure was accomplished without any in vitro manipulation, and was thus considered part of the surgical procedure (Connolly et al ., 1991). However, bone grafting is not always possible, as it injures the donor site, and the success of graft is entirely dependent on the transfer of suf cient numbers of progenitor cells. In addition, this approach may be least applicable in situations where it is most needed, such as in the case of ageing or disease with a reduction of osteogenic progenitors. Stem cell therapy involves in vitro sorting and expansion before cell grafting. Cells can be of autologous or allogenic origin. Autologous grafts offer the optimal compatibility, but require available cells from the patients. Therefore, allogenic grafts with cells from alternative allogenic source are of great interest, as they could provide 'ready for use' products for bone treatments.

Adult mesenchymal stem cells: bone marrow, adipose tissue and dental pulp-derived stem cells

It has been shown that plastic-adherent human cells derived from bone marrow were able to differentiate into a number of mesenchymal cell types including osteoblasts, chondrocytes and adipocytes. These cells were called 'MSCs' in reference to their high self-renewing properties, clonogenicity, and ability to form cartilage and bone (Fig. 20.1). They were suggested to be responsible for the normal turnover and maintenance of adult mesenchymal tissues. These MSCs have been de ned to be plastic-adherent when maintained in standard culture conditions; they must express CD105, CD73 and CD90, and lack expression of CD45 and CD34 (haematopoietic markers) surface molecules; and they must also differentiate to osteoblasts, adipocytes and chondroblasts in vitro (Dominici et al. , 2006). MSCs' therapeutic potential has generated signi cant excitement in the eld of regenerative medicine, as they can be found in various niches within human body. The ability of these cells to self-renew and differentiate into multiple tissues makes them an attractive cell source for cell based regenerative therapies. MSCs have considerable potential for the treatment of musculoskeletal disorders owing to their expansion capacity, immunosuppressive properties and ability to differentiate into bone and cartilage (Ma, 2010).

The most studied source of adult MSCs has been the bone marrow, as it was recognized early that its stroma contained stem cells capable of forming bone and cartilage. Later, other sources of adult stem cells have been described, such as adipose tissue, dental pulp, skeletal muscle, umbilical cord blood and others (Marolt et al. , 2010).

Bone marrow mesenchymal stem cells (BMSCs)

BMSCs are commonly recovered from pure bone marrow aspirates and isolated according to their adherence and culture on tissue culture plastics. They can be also selected using speci c cell surface markers, and then expanded in culture. The number of stem cells varies between different patients. Age and marrow aspiration volume in uence also the number of isolated stem cells. BMSCs can, however, be expanded in vitro to large numbers and keep their osteogenic potential even in older patients (Bianco et al. , 2001).

Adipose-derived stem cells (ADSC) [START_REF] Zuk | Human adipose tissue is a source of multipotent stem cells[END_REF] have been the rst to describe and characterize ADSCs from human lipoaspirates. They showed that a multi-lineage stem cell population could be isolated from the stromo-vascular fraction of adipose tissue. These ADSCs are able to differentiate into osteoblasts, can be recovered from raw lipoaspirates consecutive through liposuction under anaesthesia, and represent an important niche of adult stem cells. They have been extensively studied in vitro for their ability to form various tissues, including bone-like tissue.

Dental pulp stem cells (DPSC)

Stem cells within the dental pulp have been rst described by Gronthos et al. (2000). They have then been studied and characterized as an adult stem cell population that possesses the properties of high proliferative potential, the capacity of self-renewal and multi-lineage differentiation. Dental pulp can be easily collected from adult DPSC or deciduous teeth (stem cells from human exfoliated deciduous teeth) after dental extraction, when teeth have to be removed. The most frequent case is the collection of normal human third molar extracted for orthodontic reasons (Gronthos et al. , 2002). proliferate in vitro while maintaining an undifferentiated state, and are capable of differentiating into many somatic cell types. Although they show great capacities for cell therapy and tissue engineering, these particular cells create a delicate ethical debate, as they induce destruction of human embryos.

Embryonic stem cells

Neonatal stem cells from the umbilical cord

Neonatal tissues constitute a stem cell source that can be accessed in a noninvasive and rapid manner during and just after birth. The main source is the umbilical cord, covered by an amniotic epithelium that protects a gelatinous matrix called 'Wharton's jelly' . Stem cells can be recovered during birth, in utero or ex utero , from both cord blood and cord matrix (Wharton's jelly) (Forraz and McGuckin, 2011). Cord blood can be collected using a collecting needle connected to an anticoagulant-containing bag. Wharton's jelly is obtained by collecting the whole umbilical cord. With the average umbilical cord measuring 50 cm, this source of MSCs might become more and more clinically relevant.

Stem cells osteogenic differentiation and bone formation

Bone tissue is composed of matrix and cells: bone matrix is mainly built of type I collagen (90%), with non-collagenous proteins (such as osteocalcin (OCN), osteonectin, and bone sialoproteins). These scarce non-collagenous proteins participate in matrix maturation and mineralization. Osteoblasts and osteoclasts are responsible for bone formation and bone resorption. With progenitor cells, they mediate the mechanism of bone remodelling. During bone formation phase, the osteoblasts are recruited from MSC in bone marrow, since osteoclasts are derived from haematopoietic cells. Thus, stem cell therapy for bone regeneration will focus on MSC for osteoblastic differentiation, since host cells will provide osteoclasts during vascularization process.

Molecular control of osteogenesis

Bone repair consists in in ammation, intra membranous ossi cation, chondrogenesis, endochondral ossi cation and remodelling. Li et al. (2007) have described molecular events that regulate osteogenesis: they observed an upregulation of genes related to cell cycle and cell-to-cell signalling, con rming the fact that both cell division and communication are essential to initiate bone healing.

Osteoblastic differentiation in vitro and in vivo

MSCs are capable of osteodifferentiation in vitro under appropriate conditions. This differentiation is induced through addition of various molecules and chemicals, such as dexamethasone, ascorbic acid and β -glycerophosphate. Osteoblast-speci c genes expression and proteins demonstrate this in vitro differentiation. However, in vitro assays have limitations and there is always a need to verify osteoblast differentiation potential in vivo . Thus, it has been shown that osteoprogenitor cells can only form bone tissue when transplanted in vivo , while MSCs can form both bone and bone-microenvironment supporting haematopoiesis in vivo . Moreover, MSCs demonstrate self-renewal and maintenance of stemness capacity during in vivo implantation (Arvidson et al. , 2011).

Stem cells potential for scaffold-based bone tissue engineering applications

ESCs and adult (or neonatal) MSCs are the main types of stem cells usable for scaffold-based tissue engineering. However, ESCs have a limited applicability to human cell therapy due to ethical concerns, but also to immunological incompatibilities, potential for malignant tumour growth, and insuf cient understanding and control over differentiation. Thus, adult MSCs seem to be the cell type of choice for scaffold-based therapeutic applications. MSC delivery to the site of bone tissue defects requires scaffolds adapted to cell culture and activity in endogenous osteogenesis, including correct biochemical and mechanical stimuli.

Stem cells and mineralization process

Biomineralization is a general process by which inorganic calcium phosphate crystals are produced in biological environments. For bone formation, deposition of calcium phosphate salts occurs on an extracellular organic matrix by means of a heterogeneous nucleation process, which constitutes the rst stage of hard tissue mineralization (Ngankam et al. , 2000). Biomineralization process: ions and nucleation sites Bone biominerals are composed of complex structures, having precise architectural order over several length scales. They rely on their biological constituents for regulation of mineralization, forming an organic-inorganic hybrid structure. The formation mechanisms of biomineralization involving several possible pathways remain unclear. Biomineralization has been described to start with matrix vesicle-mediated mineral initiation. It also involves heterogeneous nucleation of mineral crystals with an important role played by non-collagenous glycoproteins that bind calcium and collagen. Calcium phosphate crystals can be formed via stable pre-nucleation clusters, with aggregation into an amorphous precursor phase, then transformation of this phase into a crystal (C ö lfen, 2010). Nudelman et al. (2010) demonstrated that collagen forms the structural matrix and has an active role in the mineralization of apatite for bone formation. Indeed, collagen brils have positively charged regions, and these nanosized regions can be used for mineral in ltration and charge-charge attractions. This initial mineral deposition leads to a dense network of pre-nucleation clusters, to transformation into amorphous calcium phosphate and to oriented crystalline hydroxyapatite inside the brils. Even though the biomineralization process has been mainly studied in in vitro models and in vivo bone formation may follow different rules, observations in small animals and tooth formation suggest that at least part of the mineralization mechanism of the collagen model is relevant for in vivo systems (Beniash et al. , 2009).

Collagen and calcium phosphate deposition during osteodifferentiation

Calcium and inorganic phosphate are the two main constituents of hydroxyapatite, the bone mineral that gives mechanical resistance to the organic matrix. The Ca/P mass ratio in bone is 2.2, and the initial step of Ca-P crystal nucleation takes place within matrix vesicles that appear from the plasma membrane of osteogenic cells, and migrates into the extracellular skeletal compartment. Phosphate accumulates inside these matrix vesicles by a transport system, followed by an in ux of calcium ions. This process permits the formation of hydroxyapatite crystal, associated with organic matrix collagen brils. Collagen is synthesized and assembled by osteoprogenitors, following a classical pathway for a secreted protein.

The collagen chains are synthesized as precursors, and undergo a series of maturation reactions, such as glycosylation and hydroxylation in the endoplasmic reticulum and Golgi complex. They are then assembled to form a triple helix. After processing and assembly, this type I procollagen is secreted into the extracellular space, where it will mature and polymerize into normal brils. Thus, collagen, calcium and phosphate are provided to the extracellular space through osteoprogenitors. During osteodifferentiation, MSCs will turn into progenitor cells, then osteoblasts, expressing collagen type I genes.

Influence of pore size, nanoroughness and chemical surface treatment

To mimic the in vivo cell behaviour for in vitro culture, one should consider scaffold chemistry, mechanical properties and architecture (Yang et al. , 2012). An extracellular matrix is composed of various proteins within a 3D network. Thus, a scaffold for tissue engineering should provide optimal mechanical properties and environment for cell culture. PSi can be designed to mimic the properties of bone tissue, and coatings can impart proper surface chemistry, adapting various pore sizes, porosity and nanoroughness. Indeed, nanotopography can affect protein adsorption, cytoskeletal arrangement, and cell adhesion and spreading.

Infl uence of PSi pore size on MSC adhesion

The effect of pore size and porosity on cell growth is of particular relevance, as it will direct the fabrication of a porous biomaterial. Additionally, its biocompatibility depends on its porosity and pore size. These pore dimensions can be precisely controlled and are highly tunable during PSi electrochemical anodization or etching. A variety of pore sizes can be produced: from micropores (<5 nm), mesopores (5 -50 nm) to macropores (>50 nm) depending on the preparation conditions [START_REF] Sailor | Porous Silicon in Practice[END_REF].

Although microscale topography modulates cellular behaviour in vitro , it is important to consider that cells in vivo make contact with nanoscale as well as microscale topographical features. Even if cells are typically tens of microns in diameter, the dimensions of subcellular structures tend to the nanometre scale. Furthermore, extracellular supporting tissues present an intricate network of cues at the nanoscale composed of a complex mixture of nanometre-size (5 -200 nm) pits, pores, protrusions, and bres (Biggs et al. , 2010). The main means to evaluate the in uence of pore size on MSCs adhesion and proliferation is PSi gradients, allowing cells growth on surface with a pore range from few nanometres to few micrometres: this method involves electrochemical etching of crystalline Si in a hydro uoric acid (HF) solution to fabricate PSi lms displaying a gradient of pore sizes in an asymmetrical arrangement. The size range of these pores can be controlled by modifying the applied current and the HF (Collins et al. , 2002;Clements et al. , 2012). Different cell types respond differently to topographical features, such as nanoscale pores, and studies have demonstrated that MSCs show better attachment to PSi with pore size of 20 -30 nm in diameter (Wang et al. , 2012). Pore size in this range is similar to the pore size of biological systems, such as pores in basement membranes, and corresponds to the size of biological pits, pores, protrusions and bres (Fig. 20.3). While producing PSi, pore depth can also vary. However, cells are not sensitive to changes in pore depth in the range observed with various type of PSi [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF].

Optimal surface treatment and coating for cell adhesion

MSCs were shown to respond to both topographical and chemical characteristics. However, chemical surface treatment and grafting of proteins seem to have a higher impact than changes in pore size during short-term culture (Clements et al. , 2012). PSi surface treatment is necessary not only to improve cell adhesion, but also to permit accessible and stable pores: the inner walls of the porous matrix must be protected from excessive degradation in the aqueous cell culture environment without eliciting any undesirable effects on the cells. Moreover, surface modi cation by turning PSi from hydrophobic to hydrophilic promotes cell adhesion and growth.

The most common and simple treatment is oxidation, which can be done by ozone, ageing, or thermal or chemical treatment. PSi biocompatibility has been extensively studied with various cell lines cultured on chemically modi ed PSi samples. Few studies have been done with primary MSC cultures to determine an optimal surface treatment (Hajj-Hassan et al ., 2011;Clements et al. , 2012). Moreover, it has been shown that stem cell growth can be modulated by factors mimicking extracellular matrix, such as amino or arginine-glycine-aspartate (RGD) residues. Cell behaviour is in uenced by changing the integrin and focal adhesion assembly, leading to changes in cytoskeletal organization and cell mechanical properties [START_REF] Yim | Nanotopography-induced changes in focal adhesions, cytoskeletal organization, and mechanical properties of human mesenchymal stem cells[END_REF]. Thus, amine-terminated samples are expected to enhance MSC adhesion and proliferation as it does for other cell lines: silanization with aminopropyl trimethoxysilane and triethoxysilane improves PSi stability and enhances cell adhesion in comparison to oxidized PSi. The chemistry greatly improves the stability of PSi in contact with cultured cells while allowing cell coverage levels comparable to standard culture preparations on tissue culture polystyrene [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF]. Thermal hydrosilylation is also used to graft chemical species to generate a stable substrate for cell adhesion and proliferation. Among these species are dodecene, undecylenic acid and oligoethylene glycol (Alvarez et al. , 2009).

In addition to chemical surface treatment, a coating of proteins can be realized onto the PSi surface to mimic cell adhesion and relative responses in an in vivo condition with a surface presenting biocompatible contact sites in a nanoscale dimension (topographic scale of adhesive proteins such as collagen bres, laminin and bronectin is present on the order of nanometres). The cell adhesion process depends upon cell surface receptors interacting with proteins belonging to the extracellular matrix. Several proteins have been considered for PSi coating: collagen, foetal calf serum and the tripeptide RGD. Collagen is a component of the extracellular matrix and can be used to coat surfaces to stimulate cell adhesion, as well as foetal calf serum, which contains several cell adhesive proteins. Collagen and foetal calf serum have been successfully used to promote cell attachment and cell spreading [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF]. The tripeptide RGD is an integrin-binding domain present within many matrix proteins including bronectin and vitronectin. Clements et al. (2012) showed that RGD immobilized on PSi is still a bioactive peptide and that cells adhere more when the concentration of RGD grafted increases.

Composite materials, such as poly( ε -caprolactone)/PSi, are also used for cell culture and tissue engineering. Highly porous scaffolds composed of mesoporous Si and commonly-used biopolymers, such as polycaprolactone, are electrically conductive composite materials exhibiting an interfacial behaviour that promotes calci cation and sustains the in vitro stability and proliferation of connective tissue cells in implant surfaces. Poly( ε -caprolactone) has been shown to have a biocompatible and biodegradable character that is widely exploited in guided bone regeneration and drug delivery, and the resorbable PSi component sensitively accelerates the necessary calci cation process. This type of composite presents a relevant material to orthopaedic tissue engineering, with such tunable properties as mechanical strength, controlled release of molecules, and rate of resorption of the composite to the host surroundings (Coffer et al. , 2005). This PSi-containing polymer scaffold enhances calci cation, can be considered non-toxic to cells, and supports the proliferation, viability, attachment, and differentiation of bone precursor cells, suggesting that materials of this type of composition have potential merit as biomaterials (Whitehead et al. , 2008).

Another type of PSi membrane has also been described using porous nanocrystalline Si, a new type of Si nanomaterial with potential applications in cell culture and tissue engineering. This PSi material is a 15 nm thick nanoporous membrane made with scalable Si manufacturing, permeable to small solutes. This nanoporous and ultrathin material can serve as a cell culture substrate, and is of interest to separate tissues during surgical grafting, such as bone tissue, bone substitute, periosteum or epithelium (Agrawal et al. , 2010).

Roughness and surface energy

Several molecular mechanisms by which cells perceive and respond to their surrounding environments have been studied, including integrin-ligand interactions, contact surface hydrophilicity and extracellular matrix topography and architecture. The in vitro control of the cellular environment is a challenge for manipulating both cell development and scaffold substitute function in the eld of bioengineering. However, the differences existing between natural and experimental surface characteristics may confer some of the discrepancies reported between in vivo and in vitro results. Several approaches have been used to generate biomimetic in vitro environments. These experiments explore cell interaction, either with the substrate's topography (physical domain) or with the substrate's chemical nature (chemical domain). The effects of nanotopographic features on cell behaviours have been investigated using a variety of nanoscale patterns, such as columns, sharp tips, pores and dots. Therefore, to biomimic cell adhesion and relative responses in an in vivo condition, cell adhesion surfaces are required to present biocompatible contact sites in a nanoscale dimension with periodicity matching with cell surface proteins. Nonetheless, the physical nature of the interactions between the cell and the surfaces in controlling speci c cell behaviour remains unclear. Several results suggest that nanometre-sized topographical biocompatible substrates potentially provide a distinct in vitro environment for cells, resembling the in vivo conditions (Yang et al. , 2010). However, the nanotopography plays a relevant role in the nal wettability properties of surfaces, where hydrophobic surfaces prevent cell adhesion. PSi can be arranged with nanoscale patterns, but freshly formed nanostructured PSi shows hydrophobic properties. Surface oxidation turns hydrophobic surfaces into hydrophilic: freshly etched PSi is hydrophobic, and controlled ozone or thermal oxidation can drop the water contact angle to below 5 ° . This is attributed to the formation of a polar Si-OH capped surface after oxidation [START_REF] Low | Evaluation of mammalian cell adhesion on surface-modi ed porous silicon[END_REF].

Infl uence of pore size, stiffness and surface treatment on osteodifferentiation

Cell attachment on PSi substrate has been extensively studied, but few studies evaluated the in uence of topography on cell differentiation. Recently, Wang et al. (2012) explored the in uence of nanoscale surface topography on cell behaviour. Using rat MSCs, they found that surface topography in uences cell differentiation, but not proliferation. Osteogenesis was enhanced by porous topography with a ridge roughness lower than 10 nm with an increase in osteodifferentiation when pore sizes decreased. Thus, osteodifferentiation was comparable between at Si and PSi with 100 -200 nm, and was clearly enhanced when pore sizes decreased to 10 -30 nm. This impact of pore size on osteodifferentiation is of high importance for the development of scaffold materials that can stimulate stem cell differentiation into osteoblasts in absence of chemical treatment without compromising material properties. Dalby et al . (2007) had already demonstrated the use of nanoscale disorder to stimulate human MSCs to produce bone mineral in vitro in the absence of osteogenic supplements. Their approach had similar efciency to that of cells cultured with osteogenic media. Nanocues, such as PSi pores, may prove to be important in tissue-speci c stem cell differentiation. These ndings are supported by various observations of chemical stimulation, chemical patterning and material hardness having roles in stem cell differentiation [START_REF] Zemel | Optimal matrix rigidity for stress-fibre polarization in stem cells[END_REF]. The effect of surface treatment of PSi on stem cells has been investigated mainly for cell adhesion; however, several studies have shown that the immobilization of RGD on the substrate not only enhanced cell adhesion but also modulated the intercellular mechanisms of cell proliferation and differentiation [START_REF] Hu | Porous polymer scaffolds surface-modified with arginine-glycine-aspartic acid enhance bone cell attachment and differentiation in vitro[END_REF]Cavalcanti-Adam et al. , 2007;Lagunas et al. , 2012). Thus, RGD peptide coated surfaces enhance osteogenic differentiation when present in suf cient concentration (Frith et al. , 2012). Finally, the mechanical properties of the underlying substratum have previously been shown to affect a number of cellular processes including locomotion, proliferation and differentiation. Microenvironments appear important in stem cell lineage speci cation, but can be dif cult to control with PSi. MSCs are shown to specify lineage and commit to phenotypes with extreme sensitivity to tissue-level elasticity. Soft matrices that mimic brain are neurogenic, stiffer matrices that mimic muscle are myogenic, and comparatively rigid matrices that mimic collagenous bone, such as PSi, prove to be osteogenic. After several weeks in culture, cells commit to the lineage speci ed by matrix elasticity (Engler et al. , 2006).

Growth factors delivery and Si effects on osteodifferentiation

Stem cells grown in biomaterials with bone growth factors constitute a source of transplantable material. Indeed, mesenchymal stem cells differentiation into elements of the skeleton can be enhanced by the addition of bone growth factors. PSi is a potential carrier scaffold to deliver growth factors to the site of defect, where regeneration is needed. Moreover, controlled release of ionic products from PSi dissolution can stimulate and enhance bone-cell function: dissolved Si has been shown to enhance osteoblast differentiation (Gupta et al. , 2010).

Drug delivery devices for osteogenic induction

Tissue engineering combines precursor cells from the patient with scaffolding matrices and the stimulus of growth factors. BMPs are probably the most important growth factors in bone formation and healing. These cytokines have been extensively studied during recent decades and, nowadays, recombinant human BMPs (rhBMP) are widely used in several tissue engineering products that might serve for the complete regeneration of bone or cartilage. Current applications include rhBMPs loaded in delivery systems made of synthetic or natural polymers and the differentiation of transplanted stem cells from the patient with rhBMPs for later body implantation. BMPs have a strong effect on bone and cartilage growth during early skeletal formation, and thus have promising potential for clinical bone and cartilage repair, working as powerful bone-inducing components in diverse tissue engineering products. PSi may also be used as carrier for the delivery of BMPs. BMP-2 and BMP-7 have received approval from the US Food and Drug Administration for speci c clinical cases (Bessa et al. , 2008). Indeed, PSi also exhibits a number of properties that make it an attractive material for controlled drug delivery applications. Its high surface area, coupled with the ability to control pore sizes, allows PSi to be loaded with a range of bioactive species. Tuneable properties of surface area, porous volume, and pore size allow controlling the amount and in vivo release rate of molecule payloads (Anglin et al. , 2008). By promoting osteogenic development of MSCs, the association of PSi and BMPs offers an interesting perspective for bone reparation and regeneration.

Silicic acid and Si effects on bone formation

Since the discovery that bioactive ceramics bond to bone without brous encapsulation, many studies have focused on the mechanism of bioactivity in vitro and in vivo . These studies have indicated that controlled dissolution of bioactive ceramics was of prime importance for stimulation of bone cell function and tissue regeneration (El-Ghannam et al. , 2006). The main advantage of PSi over other biomaterials is its ability to degrade, in aqueous solutions, into non-toxic silicic acid [Si(OH) 4 ], which is the biologically signi cant form of Si that is in blood plasma. Silicic acid is vital for normal bone and connective tissue homeostasis. In vivo , PSi releases Si(OH) 4 with a negligible in ammatory response before being excreted in the urine (Popplewell et al. , 1998;Park et al. , 2009). The release of Si(OH) 4 due to the corrosion of PSi in plasma stimulates calci cation and collagen growth, which could ideally accelerate new bone regeneration. Cell differentiation assays and ultrastructural measurements transmission electron microscope (TEM) have been carried out to explore the in uence of Si on stem cell osteodifferentiation. It has been demonstrated that Si plays an important role in the expression of alkaline phosphatase, a speci c biomarker for MSC differentiation (Fan et al. , 2011). OCN and bone sialoprotein (BSP) (non-collagenic proteins produced by osteoblast and involved in Ca-P crystal formation) are signi cantly more expressed by stem cells cultured in Si-rich medium, con rming that Si enhances stem cell differentiation into osteoblast, and that Si may directly be involved in OCN and BSPmediated mechanisms (Gupta et al. , 2010). Histological evaluation and histomorphometric quanti cation of bone ingrowth and mineral apposition in vivo demonstrated the bene ts to bone ingrowth and repair through incorporation of Si (Hing et al. , 2006). These results show that PSi, by its degradation, plays a signi cant role in both stem cell osteodifferentiation and mineralization (Fig. 20.4). If it is clear that the dissolution of PSi in the form of Si(OH) 4 is a powerful mechanistic force behind the nucleation and growth of calcium phosphate, it also seems that Si surfaces provide nucleation sites for the formation of apatite deposits. This conclusion is based on mechanistic observations concerning surfaces formed of CaO-rich silica, noting that Si in these materials is responsible for directing the formation of apatite deposits in simulated plasma. Si intermediates as a nucleation site in directing the formation of calcium phosphate spherulites and growth of apatite on the scaffold (Whitehead et al. , 2008). Bone cells attached onto a surface containing high Si content absorb signi cantly higher concentrations of calcium than cells attached onto a surface containing low Si content. In conjunction with the absorption of high calcium concentration, attached bone marrow stem cells produce calci ed nodules and mineralized extracellular matrix, indicating osteoblastic differentiation (El-Ghannam et al. , 2006). Thus, the mechanism of bone mineralization at the interface with bioactive surfaces is mainly cell mediated: an Si-rich phase provides a guided cell adhesion and tissue growth in vitro . The addition of Si to a scaffold affects surface topography, dissolution-reprecipitation rates and the bone apposition process (Porter, 2006).

Conclusions and future trends

In tissue engineering, developing synthetic grafts with suf cient combined mechanical strength, bioactivity and biodegradability for bone-tissue regeneration applications therefore remains an important challenge. A material that combines high bioactivity and porous structure (for good tissue growth) with high levels of mechanical strength (for load-bearing applications) will improve clinical outcomes and enable a signi cantly wider application of restorative materials [START_REF] Zreiqat | The incorporation of strontium and zinc into a calcium-silicon ceramic for bone tissue engineering[END_REF]. Most materials currently used in medical implants are bioinert and, as such, do not interact with living systems. PSi has the property to be used as a scaffold, which could be well vascularized and integrated within the host skeleton, to stimulate the formation of new living bone. It provides a controlled biophysical and biochemical microenvironment, possibly directing the differentiation of stem cells into a speci c lineage. Its porous, tunable and loadable structure allows cell adhesion and maintenance of cell function.

MSCs have a great potential for cell replacement, relying on their capacities to differentiate into several cell lineages, to modulate immune response and to secrete biological soluble factors. They have been assessed in vivo by direct transplantation to injured sites. However, new strategies involve the generation of an engineered construct combining biocompatible scaffolds with MSCs. Several pilot clinical trials are currently in process (http://www. clinicaltrials.gov).

PSi has been extensively studied in vitro for its capacity to support cell growth and differentiation, and in vivo for drug delivery applications. Its ef ciency in bone tissue engineering has now to be con rmed in experimental research in vivo and in clinical trials. Zemel, A., Rehfeldt, F., Brown, A.E.X., Discher, D. E. and Safran, S.A. (2010) 'Optimal matrix rigidity for stress ber polarization in stem cells', Nat Phys , 6 (6),(468)(469)(470)(471)(472)(473)H.,Ramaswamy,Y.,Wu,C.,Paschalidis,A.,Lu,Z.,James,B., Le silicium poreux est un biomatériau prometteur pour l'ingénierie tissulaire car il est non toxique et biorésorbable. Des modifications de surface permettent de controler sa vitesse de dégradation et peuvent favoriser l'adhésion cellulaire. Les cellules souches de la pulpe dentaire (DPSC) sont des cellules souches mésenchymateuses retrouvées dans la pulpe dentaire, à l'intérieur des dents, et constituent une source accessible de cellules souches. Regrouper les capacités de prolifération et différenciation des DPSC avec les propriétés morphologiques et biochimiques du pSi représente une approche intéressante pour des applications thérapeutiques de médecine régénératrice. Dans cette thèse, nous avons étudié le comportement de DPSC humaines sur des supports de pSi, avec des pores variant de quelques nanomètres à plusieurs centaines de nanomètres. Nous avons travaillé sur différentes fonctionalisations chimiques afin d'optimiser l'adhésion cellulaire et de stabiliser le matériau: oxydation thermique, silanisation et hydrosilylation. L'adhésion, la prolifération et la différenciation osseuse ont été évaluées par microscopie à fluorescence, microscopie électronique à balayage, activité enzymatique, tests de prolifération (activité mitotique), immunofluorescence et spectroscopie FTIR. Le pSi avec des pores de 30 à 40 nm de diamètre s'est révélé être le plus approprié pour l'adhésion, la prolifération cellulaire et la différenciation ostéoblastique. De plus, la structure nanométrique et le relargage d'acide silicique par le pSi a démontré un effet positif sur l'induction osseuse et la formation d'une matrice minéralisée. Le pSi est donc apparu comme un matériau prometteur pour l'adhésion de cellules souches mésenchymateuses, que ce soit pour une transplantation immédiate in vivo ou pour expansion et différenciation in vitro. ______________________________________________________________________ Abstract: Porous silicon (pSi) is a promising biomaterial for tissue engineering as it is both non-toxic and bioresorbable. Moreover, surface modification can offer control over the degradation rate of pSi and can also promote cell adhesion. Dental pulp stem cells (DPSC) are mesenchymal stem cells found within the teeth and constitute a readily source of stem cells. Coupling the good proliferation and differentiation capacities of DPSC with the textural and chemical properties of the pSi substrates provides an interesting approach for therapeutic use. In this thesis, the behavior of human DPSC is analyzed on pSi substrates presenting pore of various sizes, from few to hundreds nanometers. We investigated different chemical surface treatments, in order to enhance cell adhesion and stabilize the material: thermal oxidation, silanization and hydrosilylation. DPSC adhesion, proliferation and further osteodifferentiation were followed for up to 3 weeks by fluorescence microscopy, scanning electron microscopy (SEM), enzymatic activity assay, BrdU assay for mitotic activity, immunostaining and FTIR spectroscopy. Porous Silicon with pore size ranging from 30 to 40 nm was found to offer the best adhesion, the fastest growth rate for DPSC and the highest osteoinductive effect. Moreover, the pSi nanostructure and the release of silicic acid had a positive effect on precursor cells osteodifferentiation and mineralized matrix formation. Porous silicon appeared to be an appropriate biomaterial for mesenchymal stem cells adhesion and immediate in vivo transplantation, or for long term in vitro culture, for stem cells proliferation and differentiation.

1. 1 . 1

 11 Adult, neonatal and embryonic stem cells 1.1.1.1 Adult mesenchymal stem cells: Bone Marrow, Adipose tissue and Dental Pulpderived stem cells It has been shown that plastic adherent human cells derived from bone marrow

Figure 1 . 1 :

 11 Figure 1.1: Proliferation and differentiation model of mesenchymal stem cells, illustrating their high self-renewing properties, clonogenicity, and ability to form various tissues, such as bone, cartilage, adipose tissue and other connective tissues. Images A, B and C present optical microscopy images of mesenchymal stem cells in vitro, proliferating (A) to form progenitor cells colonies (B) and starting to differentiate when reaching confluence (C).

Figure 1 . 2 :

 12 Figure 1.2: Human mesenchymal stem cells from dental pulp. A: Stem cells outgrow from pulp tissue explants. B, C, D: stem cells colonies formed after digestion and being plated in culture dishes. Colonies formed in 3 and 5 days (A and B respectively), cells reaching confluence after 10 days (D). Scale bar: 300µm. From Collart Dutilleul et al. (2014)

  division and communication are essential to initiate bone healing. MSCs proliferate and differentiate under the influence of a number of molecular pathways, but MAPK (mitogen-activated protein kinase) signalling plays a pivotal role in osteogenic induction. MAPK pathway is activated by a variety of growth factors such as FGF (fibroblast growth factors), PDGF (platelet derived growth factors), TGF-β (transforming growth factor beta), BMP (bone morphogenic proteins) and IGF (insulinlike growth factors). These extracellular signals activate a signalling cascade of MAP kinase inside the cell, resulting in cell proliferation and osteodifferentiation. Another pathway involved in osteodifferentiation is the Wnt signalling pathway: Wnt represents a family of secreted molecules that are involved in various aspects of cellular biology, including cell growth, differentiation and death. The binding of Wnt proteins permits the translocation of β-catenin into the nucleus, resulting in an enhanced expression and activity of the osteogenic promoter Runx2. Wnt signalling and Hedgehog signalling cooperate in controlling bone formation, especially at the early stage of osteoblast formation, initiating cellular differentiation. Cell-cell interactions are also required for activating some signalling cascade: Notch signalling and Ephrin signalling are the main cell-cell interaction pathways. Notch receptors and ligands are membrane proteins able to up-regulate bone related genes to favour the proliferation of immature osteoblasts, while limiting osteoblast maturation by inhibiting the Runx2 transactivation function. Ephrin ligands and receptors are also membrane proteins. They are constituted by two classes of ephrins: the B class ephrin ligands and receptors (EphB-Tyrosine Kinase receptor) induce osteogenic transcription factors, and inhibit osteoclast differentiation, leading to bone formation. EphrinA/EphA receptor interaction facilitates bone remodelling by suppressing osteoblast differentiation and activating osteoclast differentiation (Figure 1.3, next page).

Figure 1 . 3 : 2

 132 Figure 1.3: Schematic representation of the principal signalling networks and transcription factors regulating osteoblast differentiation and osteogenesis. Scanning electron microscopy of a mesenchymal stem cell (MSC, original magnification x2000) and of an osteoblast differentiated in vitro from MSC (original magnification x2000). BMP: bone morphogenic protein, TGFβ: transforming growth factor, PDGF: platelet derived growth factor, FGF: fibroblast growth factors, IGF: insulin-like growth factor, MAPK: mitogen-activated protein kinase. From Collart Dutilleul et al.(2014)

Figure 1 . 4 :

 14 Figure 1.4: Schematic representation of biomineralization process, describing matrix vesicle-mediated mineral initiation. (A) Mineralization of a collagen fibril, with calcium phosphate mineral droplets (red) binding to collagen fibers and entering the fibrils. Once inside the collagen, the mineral in a liquid state diffuses through the interior of the fibril and solidifies into an amorphous phase (black). Finally, directed by the collagen, the amorphous mineral transforms into oriented apatite crystals (yellow). (B) Two-dimensional cryo TEM image of a mineralized collagen fiber. Crystals are visible on insets 1 and 2 (white arrows). Scale bar = 100 nm (C) SEM image of mineralized collagen fibers (dentin tissue from a human tooth). Scale bar = 2 µm. (D) Surface-directed mineralization of calcium phosphate. (1) Aggregates of pre-nucleation clusters are in equilibrium with ions in solution. (2) Pre-nucleation clusters approach and aggregate near the functionalized surface. (3) Further aggregation causes densification near the surface. (4) Nucleation of amorphous spherical particles occurs at the surface. (5) Crystallization occurs in the region of the amorphous particles directed by the surface. Adapted from Cölfen et al. (2010), Nudelman et al. (2010) and Collart[START_REF] Dutilleul | Root canal hydrophobization by dentinal silanization: Improvement of silicon-based endodontic treatment tightness[END_REF]

Figure 1 . 5 :

 15 Figure 1.5: Schematic representation of an electrochemical set-up used to make porous silicon. (A) Schematic side view of the custom made set-up. Silicon is the anode and platinum electrode is the cathode. (B) Schematic side view of the silicon wafer after etching, with a superficial porous part (porous silicon) and an unetched part (bulk silicon). (C) Top view photograph of a freshly etched silicon wafer. The black area is the porous part while the mirror area is the unetched part. The etched area is round because of the O-ring in the set-up. Samples are cut in square to remove the unetched surfaces. Scale bar = 1 cm. (D) SEM image of pSi surface. Scale bar = 500 nm. (E) and (F) SEM images of pSi: cross section of the wafer. E: Scale bar = 10 µm. F: Scale bar = 300 nm. Adapted from Alvarez et al. (2009) and personal data.

Figure 1 . 6 :

 16 Figure 1.6: Epifluorescence microscopy and SEM of MSC culture on pSi, after 24H of culture. A: mean pore size = 30 nm. A1 and A2: epifluorescence microscopy and SEM, respectively. B: mean pore size = 1 µm. B1 and B2:epifluorescence microscopy and SEM, respectively. Cells present a better adhesion on pSi with small pore size, compared to pSi with large pores: cells recover a more physiological shape (A1) and spread more (A2) on 30nm-pSi, while cells are more bulging and spread less on 1 µm-pSi (B1 and B2). From Collart[START_REF] Dutilleul | Porous Silicon for Biomedical Applications[END_REF]

  approaches have been used to generate biomimetic in vitro environments. These experiments explore cell interaction either with the substrate's topography (physical domain) or with the substrate's chemical nature (chemical domain). The effects of nanotopographic features on cell behaviors have been investigated using a variety of nanoscale patterns, such as columns, sharp tips, pores and dots. Therefore, to biomimic cell adhesion and relative responses in an in vivo condition, cell adhesion surfaces are required to present biocompatible contact sites in a nanoscale dimension with periodicity matching with cell surface proteins. Nonetheless, the physical nature of the interactions between the cell and the surfaces in controlling specific cell behavior remains unclear. Several results suggest that nanometer-sized topographical biocompatible substrates potentially provide a distinct in vitro environment for cells,

1. 3 . 3 . 1

 331 Drug delivery devices for osteogenic induction Tissue engineering combines precursor cells from the patient with scaffolding matrices and the stimulus of growth factors. Bone morphogenetic proteins (BMPs) are probably the most important growth factors in bone formation and healing. These cytokines have been extensively studied during recent decades and, nowadays, recombinant human BMPs (rhBMPs) are widely used in several tissue-engineering products that might serve for the complete regeneration of bone or cartilage. Current applications include rhBMPs loaded in delivery systems made of synthetic or natural polymers and the differentiation of transplanted stem cells from the patient with rhBMPs for later body implantation. Bone morphogenetic proteins have a strong effect on bone and cartilage growth during early skeletal formation, and thus have promising potential for clinical bone and cartilage repair, working as powerful bone-inducing components in diverse tissue-engineering products. PSi may also be used as carrier for the delivery of BMPs. BMP-2 and BMP-7 have received approval by the US Food and

Figure 1 . 7 :

 17 Figure 1.7: A and B: SEM of MSC osteodifferentiation and mineralized matrix deposits. A: on psi membrane (pore diameter = 30 nm); B: on culture plate. On pSi, osteoblast-like cells are embedded in a mineralized matrix, while cells are more clearly visible on polystyrene, with less extracellular matrix. C: quantity of mRNA bone sialoprotein (BSP) and osteocalcin (OCN) expressed by osteoblast-like cells on Si composite and on culture plate, after 2 days of culture in basal medium MEM(-) or in osteogenic medium MEM(+). Mean concentration of [Si] in culture medium containing Si composite is 6.5mg/l. Cells express more of osteogenic marker when cultured in the presence of [Si]. Adapted from Gupta et al. (2010) and Collart Dutilleul et al. (2014).

  CD45, CD34, CD14, CD19 and HLA-DR surface molecules; third, they must differentiate to osteoblasts, adipocytes and chondroblasts in vitro (Figure2.2, next page).

Figure 2 . 2 :

 22 Figure 2.2: Minimal criteria to define MSC: differentiation assays and flow cytometry analysis. A-B: Optical microscopic images representing results of differentiation assays. (A) Alizarin Red S-stained monolayer of osteogenic MSC. (B) Oil Red O-stained monolayer of adipogenic MSC. (C) Sections of chondrogenic pellets stained with Toluidine Blue. (D) Flow Cytometry analysis of MSC for various cell surface markers. MSC are negative for CD34, CD45, CD150, CD18 and CD14. They are positive for the other markers, with various intensities. Adapted from Sonoyama et al. (2006 and 2007).

  MSC have also been investigated for DPSC: DPSC have been

2. 1 . 3 . 1

 131 Hepatic tissue regeneration DPSC transplantation in injured hepatic tissue in rat leaded to liver primary function recovery, after hepatocytes restoration. In 2008, Ikeda and co-workers examined DPSC potential by the transplantation into a carbon tetrachloride (CCl4)-treated liver injured rat, to determine whether these cells might be useful for cell-based therapy to treat liver diseases. The successful engraftment of the DPSC was demonstrated at 4 weeks after transplantation. They prevented the progression of liver fibrosis and contributed to the restoration of liver function, as assessed by the measurement of hepatic serum markers aspartate aminotransferase and alanine

Figure 2 . 3 :

 23 Figure 2.3: Effect of DPSC transplantation on infarct size. (A): Masson's trichrome-stained heart sections from infracted rats receiving either no cells (control) or DPSC. Arrows show the fibrotic area. (B): Thin sections and quantification of the Left Ventricular Wall thickness. Adapted from Gandia et al. (2008)

Figure 2 . 4 :Figure 2 . 5 :

 2425 Figure 2.4: Orthopantomogram (panoramic radiograph) of impacted wisdom teeth, before extraction for orthodontic purpose. (A) Favorable case, where teeth will probably remain intact during dental extraction. (B)Unfavorable case, where wisdom teeth will probably have to be cut in several pieces by the dental surgeon during dental extraction. From Collart[START_REF] Dutilleul | Dental pulp stem cells: characteristics, cryopreservation and therapeutic potentialities[END_REF]

Figure 2 . 6 :

 26 Figure 2.6: Optical microscopic images of DPSC stained with Toluidine Blue. Left picture: cells with typical fibroblast morphology. Central picture: colonies of DPSC, forming aggregates (CFU). Right pictures: DPSC reaching confluence.

Figure 2 . 7 :

 27 Figure 2.7: Optical microscopic images of differentiation assays: (OM+) & (OM-): Alizarin Red S staining of DPSC cultures after 21 days in osteogenic medium (OM+) or in basal medium (OM-). (AM+) & (AM-): Red-O-oil staining of DPSC cultures after 21 days in adipogenic medium (AM+) or in basal medium (AM-). (CM+) & (CM-): Alcian Blue staining of DPSC cultured ad pellet during 21 days in chondrogenic medium (CM+) or in basal medium (CM-).

Figure 2 . 8 :

 28 Figure 2.8: Representative flow cytometry analysis performed on cells immediately after their recovery, without in vitro step.

Figure 2 . 9 :

 29 Figure 2.9: Flow cytometry performed on dental pulp cells after 24 h in culture in vitro, and rinsing to remove non-adherent cells.

Figure 2 . 10 :

 210 Figure 2.10: Flow cytometry performed on dental pulp cells after one week of in vitro expansion.

Figure 2 . 11 :

 211 Figure 2.11: Flow cytometry performed on dental pulp cells after cryopreservation (1 year in liquid nitrogen). There was not any noticeable difference compared to flow cytometry preformed on fresh cells after 7 days in culture.

Figure 2 . 12 :

 212 Figure 2.12: FACS of CD34+/CD146+ cells. (A) Total population analyzed for CD34 and CD146 positivity. Double positive cells were selected and sorted. (B) Cell analysis after sorting, showing the CD34+/CD146+ population. (C) Histogram of CD34 expression, after sorting. (D) CD34 expression after 3 weeks in culture (Passage 3).

Figure 3 . 1 :

 31 Figure 3.1: Characterization of the porous silicon scaffolds. Surface topography of the silicon substrates imaged with (A) scanning electron microscopy and (B) atomic force microscopy. Porous silicon with mean pore diameter of 10 ± 2 nm, 36 ± 4 nm and 1.0 ± 0.1 µm are shown as A1-B1, A2-B2 and A3-B3 respectively. Surface chemical treatments are presented as: (1) Non-oxidized pSi hydrosilylated with semicarbazide. (2) Non-oxidized pSi hydrosilylated with undecenoic acid. (3) Oxidized pSi silanized with APTES.

  We examined DPSC attachment to various surface-modified pSi layers and compared the results with those of DPSC cultured on glass coverslips and on flat silicon wafers after a period of 4 h, 24 h, 48 h and 72 h. The results are shown in Figure3.2.

Figure 3 . 2 :

 32 Figure 3.2: DPSC attached on various type of pSi, after 4h, 24h, 48h and 72h of incubation. Glass coverslip and (non-porous) flat silicon were used as controls. Cells were counted per 1400 µm x 1050 µm area, in five areas per samples. After 72 h, cells couldn't be counted on glass, flat Si and APTES as they reached confluence (protruding bars).

  reached confluence on glass coverslip, flat Si and APTES-treated pSi. Cells were stained with the vital dye FDA to observe cell adherence and cytoplasmic spreading. The morphologies of DPSC on surface modified pSi at early attachment stage (after 4 h) are shown in Figure 3.3.

Figure 3 . 3 :

 33 Figure 3.3: Fluorescence microscopy of DPSC on the various scaffolds after 4 h incubation. Cells are stained with vital FDA staining (cytoplasma: green staining). A: glass coverslip, B: Flat Si, C: pSi 10nm, D: pSi 36nm, E: pSi 1µm, F: undecenoic acidtreated pSi, G: APTES-treated pSi, H: semicarbazide-treated pSi. Main pictures: x20, scale bar = 100 µm. Inserts: magnification x40, scale bar = 40 µm.

Figure 3 . 4 :

 34 Figure 3.4: DPSC proliferation at 4 h, 24 h, 48 h and 72 h on various pSi surfaces, assessed by acid phosphatase activity determination. Glass coverslip and oxidized nonporous silicon (flat Si) were used as control. In order to normalize the results, 100% adhesion was attributed to cells attached on glass coverslip after 72 h of cell seeding. All experiments were performed in triplicate. Statistical significance was determined by using the Student's t test (P < 0.05) for comparison with glass coverslip: (*)Statistical difference at 24 h. (**)Statistical difference at 72 h.

  mitotic activity just after cell seeding, by BrdU incorporation for 24 h. The percentages of BrdU positive cells, corresponding to mitotically active cells, are shown in Figure 3.5.

Figure 3 . 5 :

 35 Figure 3.5: Number of BrdU-positive cells after 24 h of incubation, expressed as a percentage of the total number of DPSC counted on the various pSi surfaces. Glass coverslip and oxidized non-porous silicon (Flat Si) were used as controls. All experiments were performed in triplicate. Statistical significance was determined by one-way ANOVA Tukey test (P < 0.05): (*) Comparison with glass coverslip. (**) Comparison with APTES

3. 3 . 4

 34 Cell morphology SEM was employed to study the effects of porosity and chemical surface treatments on the morphology of single adhering cells, and to determine cell circularity and cell surface area. Figure 3.6 shows representative images of DPSC growing on the different pSi substrates, on glass coverslip and flat silicon as controls, after 24 h of incubation.

Figure 3 . 6 :

 36 Figure 3.6: Scanning electron microscopy of DPSC after 24 h incubation. A: glass coverslip, B: flat Si, C: pSi 10nm, D: pSi 36nm, E: pSi 1µm, F: undecenoic acid-treated pSi, G: APTES-treated pSi, H: semicarbazide-treated pSi.

Figure 3 . 7 :

 37 Figure 3.7: Scanning electron microscopy of DPSC after 24 h incubation at magnification x 8000 (image A, B, C D and G), x 15000 (image F and H) or x 30000 (image E). A: glass coverslip, B: flat Si, C: pSi 10nm, D: pSi 36nm, E: pSi 1µm, F: undecenoic acid-treated pSi, G: APTES-treated pSi, H: semicarbazide-treated pSi.

Figure 3 . 8 :

 38 Figure 3.8: A: Mean cell surface area calculated from SEM images after 24 h (in µm2). B: Mean cell circularity, evaluated from the SEM images after 24 h. 0 represent a line and 1.0 a circle (non defined unit). Statistical significance was determined by one-way ANOVA Tukey test (P < 0.05): (*) Comparison with glass coverslip. (**) Comparison with APTES.

  23 ± 0.06), flat Si (0.22 ± 0.06), pSi 10 nm (0.20 ± 0.07) and pSi 36 nm (0.20 ± 0.05). Cell circularity was similar on APTES-treated (0.34 ± 0.07) and semicarbazide-treated pSi (0.36 ± 0.06), and it was also similar on pSi 1 µm (0.49 ± 0.08) and undecenoic acidtreated pSi (0.48 ± 0.08) (Figure3.8B).3.3.5 Cell morphology after 72 hoursWe followed cell growth for 72 h, until they reached confluence on the control surfaces (glass coverslip and flat Si). Cells were stained with the vital dye FDA to observe cell adherence and cytoplasmic spreading. The morphologies of DPSC on surface modified pSi after 72 h are shown in Figure3.9.

Figure 3 . 9 :

 39 Figure 3.9: Fluorescence microscopy of DPSC after 72 h incubation. Cells are stained with vital FDA. A: glass coverslip, B: Flat Si, C: pSi 10 nm, D: pSi 36 nm, E: pSi 1 µm, F: undecenoic acid-treated pSi, G: APTES-treated pSi, H: semicarbazide-treated pSi. Main pictures: x20, scale bar = 100 µm. Inserts: x40, scale bar = 40 µm.

  ) the ability of APTES-treated pSi as scaffold for stem cells differentiation in different lineage, as pore size might also influence cell differentiation, (2) the optimization of surface stabilization with the semicarbazide treatment, (3) the efficiency of semicarbazide-treated pSi as an immediate cell carrier for in vivo transplantation. Further studies will also elucidate the role played by the porosity on focal adhesion formation, as well as the implication on the cytoskeleton organization. -based bone tissue engineering Synthetic biomaterials designed for bone repair are no longer considered

Figure 4 . 1 :

 41 Figure 4.1: SEM images of porous silicon surfaces: (A) after 1 day in simulated body fluid (SBF) and (B) after 6 days in SBF. From Canham et al. (1997).

4. 2 . 6

 26 Cell adhesion and viabilityDPSC attachment and viability were monitored by fluorescence microscopy. DPSC were seeded onto pSi at a cell density of 5x10 4 cell/ml. Cell culture plates and flat unetched silicon wafers were used as controls. Cultures were incubated for 4 h, 24 h and 48 h at 37°C with 5% CO 2 , in a humidified incubator. To determine cellular morphology and viability on the surfaces, cells were stained with 25 µg/ml of fluorescein diacetate (FDA) and 50 µg/ml of propidium iodide (PI) for 3 min at 37°C. Samples were subsequently washed with PBS to remove non-adherent cells and unbounded dye, before being observed under fluorescence microscopy at excitation wavelengths of 480 nm for FDA and 630 nm for PI. Images were captured on a Nikon TE2000-E microscope equipped with a Nikon digital camera. For cell counts, samples were fixed with 2.5% glutaraldehyde for 1 hour, and stained with 100 ng/ml 4,6diamidino-2-phenylindole (DAPI) for 30 min to stain nuclei. Samples were then washed with PBS to remove any non-adherent cells and unbounded dye, and observed under fluorescence microscopy at an excitation wavelength of 360 nm. Cell counts were conducted at five different locations on the surface of each sample (four peripheral and one central) in areas measuring 1200x780 µm. All experiments were made in triplicate and statistical differences along the gradients were determined using the Student's t-test with p < 0.05 deemed as statistically significant.4.2.7 Quantification of Silicic acid release: ICP-MS analysis DPSC were cultured on pSi or flat Si, in osteogenic medium (OM) or in basal medium (MEM). The culture media were changed twice a week. The Si content in all the culture media after cell incubation was measured by means of inductively coupled plasma mass spectrometry (ICP-MS). ICP-MS analysis was performed by introducing extract aliquots into a mass spectrometer (Element XR ICP-MS, Thermo Scientific™).

Figure 4 . 2 :

 42 Figure 4.2: Characterization of DPSC. (A) Optical microscopy of adherent dental pulp cells forming colonies at 3 and 6 days, and reaching confluence 10 days after recovery. Scale bar = 100 µm. (B) Flow cytometry analysis of subconfluent dental pulp cells. Single-parameter histograms showing the expression of markers CD 90, CD 146, CD 117 and CD 45.

Figure 4 . 3 :

 43 Figure 4.3: Scanning electron microscopy of pSi scaffold and DPSC adhesion after 24 h in culture medium. (A) Top view and cross section of pSi scaffold (A1 and A2, respectively). (B) DPSC on pSi scaffold at magnification x2000 and x10000 (B1 and B2, respectively). (C) DPSC on flat Si at magnification x2000 and x8000 (C1 and C2, respectively).

Figure 4 . 3 (

 43 SEM after 24 h of incubation) and in Figure 4.4 (Fluorescence microscopy of stained living cells, after 4 h, 24 h and 48 h of incubation). The absence of red staining (PI) confirmed cell-compatibility of the substrate.

Figure 4 . 4 :

 44 Figure 4.4: Vital fluorescence microscopy of DPSC on pSi and flat Si. (A) DPSC on flat Si after 4 h, 24 h and 48 h of incubation (A1, A2 and A3, respectively). (B) DPSC on pSi after 4 h, 24 h and 48 h of incubation (B1, B2 and B3, respectively). Cells are stained with fluorescein diacetate (vital cells, green staining) and propidium iodide (dead cells, red staining). Scale bar: main image = 200 µm; inset = 30 µm.

Figure 4 . 5 :

 45 Figure 4.5: Cell counts after 4 h, 24 h and 48 h incubation. DPSC seeded on flat Si and pSi at a density of 5x10 5 cells/ml. After incubation, cells were fixed with 2.5% glutaraldehyde and stained with DAPI (nuclei blue staining). Cells counted on areas measuring 1200x780 µm. Experiment made in triplicate, with 5 measurements per sample. Results presented as mean ± standard deviation, with error bar representing standard deviation. Dotted line to guide eyes.

Figure 4 . 6 :

 46 Figure 4.6: Si concentration in culture media (OM or MEM) after incubation with pSi or flat Si during DPSC osteodifferentiation.

Figure 4 . 7 :

 47 Figure 4.7: DPSC osteodifferentiation during the 2 first weeks of experiment, assessed by alkaline phosphatase (ALP) quantification. Results were normalized to the total quantity of proteins. Experiments were performed in triplicate. (*) indicates a statistical difference between groups, with significance determined by using the Student t test (p < 0.05).

flat

  Si in OM. Similarly, Ca deposits were detectable on bare pSi in MEM while no Ca deposits were found on bare flat Si. These data are summarized in Figure4.8.

Figure 4 . 8 :

 48 Figure 4.8: Calcium deposits stained with Alizarin red: (A) DPSC on pSi after 3 weeks in OM and (B) DPSC on pSi after 3 weeks in MEM. (C) Quantification of calcium deposits after 21 days of differentiation, assessed by Alizarin red staining. Calcium concentration was determined by absorbance at 562 nm. Optical density (OD) values were normalized to the total protein content of DPSC cultured under the same condition. Experiments were performed in triplicate. (*) indicates a statistical difference between groups, with significance determined by using the Student's t test (p < 0.05).

Figure 4 . 9 :

 49 Figure 4.9: (A) Representative histograms of the percentage of CB-positive colonies along the pSi gradient. Abscissa indicates the corresponding pore diameters. (B)-(D) Representative fluorescence microscopy images of DPSC after 3 weeks in OM (red fluorescence) and Ca deposits stained by Calcein Blue (blue fluorescence), on flat Si, 8 nm pores and 37 nm pores (B, C and D, respectively). Scale bar = 100 µm.

4. 3 . 6 . 3

 363 SEM images and EDX spectroscopy DPSC were incubated for 3 weeks on pSi or flat Si, in either osteogenic medium (OM) or basal medium (MEM). After the 3 weeks, samples were fixed and dehydrated, then observed under SEM and analyzed with EDX spectroscopy. EDX analyses were realized on the whole corresponding SEM images (Figure4.10). We demonstrated that there were significantly more Ca and P on pSi, compared to flat Si, with DPSC cultured in OM. The same tendency was observed with DPSC cultured in MEM, where some Ca and P were found on pSi and not on flat Si. To get some information about the CaP structure, we consider the ratio Ca/P: the values are presented in Table4.3 (next page).

Figure 4 . 10 :

 410 Figure 4.10: SEM images of DPSC after 3 weeks of in vitro growth and differentiation, and corresponding EDX spectra. (A) DPSC cultured on pSi, in OM. (B) DPSC cultured on flat Si, in OM. (C) DPSC cultured on pSi, in MEM. (D) DPSC cultured on flat Si, in MEM. Scale bar = 50 µm. EDX analysis results are in atomic %.

  The same experiments were conducted without cells, on bare scaffolds incubated in either OM or MEM. Some mineralized structures were clearly observable, but only on pSi incubated in OM. EDX spectroscopy confirmed the Ca-P composition of these structures(Figure 4.11).

Figure 4 . 11 :

 411 Figure 4.11: SEM images of bare scaffolds after 3 weeks of in vitro incubation, and corresponding EDX spectra. (A) pSi in OM. (B) flat Si in OM. (C) pSi in MEM. (D) flat Si in MEM. Main image scale bar = 5 µm. Inset image scale bar = 1 µm. EDX analysis results are in atomic %.

Figure 4 .

 4 11 presents a summary of these data, with representative SEM images at different stage of the gradient.

Figure 4 . 12 :

 412 Figure 4.12: (A) Representative histograms of the EDX analysis along the pSi gradient, with results expressed in Atomic %. Abscissa indicates the corresponding pore diameters. (B)-(D) Representative SEM images of DPSC after 3 weeks in OM. Brighter spots correspond to Ca deposits. (B) flat Si. (C) 14 nm pores. (D) 37 nm pores. Scale bar = 50 µm.

Figure 4 . 13 : 1

 4131 Figure 4.13: (A) Representative histograms of the EDX analysis along the bare pSi gradient incubated in OM for 3 weeks, with results expressed in Atomic %. Abscissa indicates the corresponding pore diameters. (B)-(D) Representative EDX spectra at different stage of the pSi gradient surface, after 3 weeks in OM. (B) flat Si. (C) 14 nm pores. (D) 37 nm pores.

  with 40 nm pores favor ACP transformation and crystal growth during extra cellular matrix formation, by mimicking the 40 nm gap zone of the collagen fibrils. Such hypothesis is reinforced by the results obtained with pSi scaffolds (36 nm pores diameter): bare scaffolds incubated in OM induced the nucleation of OCP while scaffolds with DPSC permitted to CaP closer to β-TCP or HA.Biomaterials used as scaffold in tissue engineering must be tolerated by the cells, but are also expected to promote physiological cell responses: by acting as analogs of the natural extracellular matrix (ECM), they can regulate cell adhesion and extent through the binding of cell surface molecules to the materials surface[START_REF] Bacakova | Modulation of cell adhesion, proliferation and differentiation on materials designed for body implants[END_REF]. The cell extents and the strength of cell adhesion regulate the cell proliferation and the balance between proliferation and differentiation[START_REF] Mann | Cell adhesion peptides alter smooth muscle cell adhesion, proliferation, migration, and matrix protein synthesis on modified surfaces and in polymer scaffolds[END_REF] (See alsoChapter 1, Figure 1.1). The proliferation capacity and migration speed are dependent of both cell spreading and cell adhesion strength. Maximum cell proliferation and cell migration is achieved at an intermediate stage of cell extents and adhesion strength.

  Dental Pulp Stem Cells (DPSC) Mesenchymal stem cells derived from human teeth (dental pulp stem cells) were recovered from discarded normal human impacted third molars, collected with informed consent of patients undergoing routine extractions at the Dental Clinic of the University of Adelaide, under approved guidelines set by the University of Adelaide and Institute of Medical and Veterinary Science (IMVS) Human Subjects Research Committees.

  surfaces were cleaned and cracked open to reveal the pulp chamber. The pulp tissue was gently separated from the crown and root and then digested in a solution of 3 mg/mL collagenase type I and 4 mg/ml dispase for 1 hour at 37 °C. Single-cell suspensions were obtained by passing the cells through a 70 µm strainer. Cultures were established by seeding single-cell suspensions (1 -2 x 10 5 ) of dental pulp into T-25 flasks in growth media, (αMEM supplemented with 10% fetal calf serum, 100µM L-ascorbic acid, 2 mM

Figure 5 . 1 .

 51 Figure 5.1. In order to stabilize the scaffolds surface, pSi gradients were thermally oxidized for 1 h at 800 °C, and silanized with aminopropyltriethoxysilane (APTES) 4 % in dry toluene at room for 1 h with gentle shaking. For use in cell culture, the wafers were sterilised with 70% (vol/vol) ethanol for 10 min before drying under sterile airflow.

Figure 5 . 1 :

 51 Figure 5.1: Etching set up, with platinum electrode placed perpendicular to the surface, at one and of the pSi wafer. Adapted from Clements et al. (2012).

5. 2 . 9

 29 Immunofluorescence: ALP and type 1 Collagen Cell differentiation was assessed after 7 days by ALP staining and after 21 days by type I collagen staining. Samples were washed with PBS and fixed with 4 % paraformaldehyde for 20 min at room temperature. Cells were permealized with 0.5% Triton X-100 in PBS for 15 min at 4 °C then blocked with 1 % BSA in PBS for 1 hour at room temperature. Samples were incubated with either rabbit anti-ALP or goat anticollagen antibodies (1:100 in 1 % BSA) at 4 °C overnight. After three times of rinsing with PBS, samples were incubated with either PerCP-conjugated donkey anti-rabbit or FITC-conjugated donkey anti-goat secondary antibodies (1:200 in 1 % BSA) for 1 hour at room temperature. Finally, cell nuclei were stained with 50 µg/ml Hoechst 33342 for 30 min at room temperature. Samples were observed under confocal fluorescence microscopy at an excitation wavelength of 360 nm for Hoechst, 488 nm for type I collagen and 633 nm for ALP staining.

  . Electropolishing was observed near the Pt electrode and a conspicuous color change occurred towards the other extreme of the surface, indicating a change in pore architecture (Figure 5.2).

Figure 5 . 2 :

 52 Figure 5.2: Asymmetric configuration of the etching set-up, and photograph of a freshly etched pSi gradient. Color changes are visible along the gradient. The samples are 20 mm long. Schematic set-up adapted from Clements et al. (2011)

  end of the sample furthest away from the electrode (18 mm from the electrode), a pore size of 8 ± 5 nm was determined. Pore size gradients were reproducible and showed little sample-to-sample variation.

Figure 5 . 3 :

 53 Figure 5.3: Graph of the average pore size along the pSi gradients, according to the distance to the electrode. Images are representative SEM top view of the pSi at different stages of the gradient; the value on each image indicates the distance to the electrode.

Figure 5 . 4 :

 54 Figure 5.4: Porosity and thickness of the porous layer, assessed by IRS. The thickness corresponds to the pore depth. On the bottom of the figure are shown optical reflectance spectra showing the variation in the fringes along the length of a pSi gradient.

Figure 5 . 5 :Figure 5 . 6 :

 5556 Figure 5.5: Representative SEM images of pSi gradient after 21 days in culture medium, showing the persistency of pores. (A) image at 14 mm from the electrode. (B) image at 6 mm from the electrode. (C) image under the electrode. Scale bar = 1 µm.

Figure 5 . 7 :

 57 Figure 5.7: Fluorescence microscopy images of DPSC seeded on pSi grad, after 24 h of incubation. Cytoskeleton is stained in red (actin) and nuclei are stained in blue. The distance on each image indicates the distance to the electrode. Scale bar = 100 µm

Figure 5 . 8 :

 58 Figure 5.8: Graphical representation of cell surface area and cell circularity, assessed from fluorescence microscope images.

  cell spreading and adhesion were assessed for DPSC cultivated on pSi gradients, by cytoskeleton (actin) and focal adhesion (vinculin) staining. Adhesion-mediated cell spreading and focal adhesion was monitored after 24 h of incubation via immunostaining of F-actin (phalloidin staining) and vinculin (Figure 5.9). On flat Si and pSi with pores ≤ 70 nm, development of stress fibers and vinculin could clearly be seen, indicating well spread cells with mature focal adhesion. On large pores (pores diameter ≥120 µm), minimal cell spreading was observed, with actin observed only around the nuclei, and less organized focal adhesion.

Figure 5 . 9 :

 59 Figure 5.9: Fluorescence microscopy images of DPSC seeded on pSi grad, after 24 h of incubation, showing focal adhesion distribution. Cytoskeleton is stained in red (actin), nuclei are stained in blue and vinculin is stained in green. Focal adhesion (green staining) are fairly well organized on flat Si and pSi at 8 to18 mm from the electrode, and less organized (even not visible) on regions closer to the electrode. Scale bar = 100 µm

Figure 5 . 10 :

 510 Figure 5.10: AFM and fluorescence microscopy examination of DPSC after 24 h of incubation on pSi (pores 36 ± 4 nm) and flat Si. Representative AFM 3D reconstruction of DPSC on pSi (A1) or flat Si (A2) and corresponding deflection images (B1: pSi. B2: flat Si). (C) Fluorescence microscopy images of DPSC on pSi (C1) or flat Si (C2) with actin staining (red staining) showing the filopodia formations. Scale bar = 50 µm. The arrows highlight the long thin filopodia growing from the main cell body while the arrowheads show filopodia growing from lamellipodia.

For

  DPSC incubated on pSi scaffolds, we measured height and width of the filopodia growing from the main cell body by AFM: the size of these actin finger-like structures ranged from 30 to 70 nm of height and from 50 to 200 nm of width, with a length that could be of several micrometers(Figure 5.11).

Figure 5 . 11 :

 511 Figure 5.11: AFM images in contact mode of a DPSC cultured on pSi. (A) Deflection image; (B) Corresponding 3D reconstruction image; (C) Topography image; (D) Cross section along the line indicated in C.

  proliferation rate of DPSC seeded on pSi gradients, incorporation of bromodeoxyuridine (BrdU) for 24 to 72 h was performed, in order to label dividing cells. All cell nuclei were stained in blue with Hoechst 33342, and dividing cells nuclei were, in addition, stained in green (Alexa 488 conjugated secondary antibody). The proliferation rate corresponds to the number of BrdU stained cells over the total number of cells (Figure5.12).

Figure 5 . 12 :

 512 Figure 5.12: Fluorescence microscopy images of DPSC stained with Hoechst and anti-BrdU antibody. The percentage of BrdU positive cells is the ratio between BrdU stained cells (green nucleus) and total cells (blue nuclei).

Figure 5 . 13 :

 513 Figure 5.13: Percentage of BrdU positive cells along pSi gradients, assessed for up to 72 h after cell seeding. The black histograms show the total number of cells after 72 h.

Figure 5 . 14 :

 514 Figure 5.14: Representative confocal microscopy images of DPSC after 7 days in osteogenic medium. Nuclei are stained with Hoechst 33342 (blue staining) and ALP is stained with PE-conjugated antibodies (red staining). Images are treated by ImageJ software to enhance readability. (A) image at mm from the electrode. (B) image at 10 mm from the electrode. (C) image at 2 mm from the electrode. Scale bar = 200 µm.

Figure 5 . 15 :

 515 Figure 5.15: Quantification of ALP expression, assessed by Immunofluorescence. Data were normalized to the cell number on each area. (*) indicates a statistically significant difference compared to the ALP staining for the 37 nm pores (p < 0.05).

Figure 5 . 16 :

 516 Figure 5.16: Representative confocal microscopy images of DPSC after 21 days in osteogenic medium. Nuclei are stained with Hoechst 33342 (blue staining) and type 1 Collagen is stained with FITC-conjugated antibodies (green staining). Distance from electrode: (A) 18 mm, (B) 10 mm, (C) 2 mm. Scale bar = 400 µm.

Figure 5 . 17 :

 517 Figure 5.17: Quantification of type 1 Collagen expression, assessed by Immunofluorescence. Data were normalized to the cell number on each area. (*) indicates a statistically significant difference compared to the type 1 Collagen staining for the 37 nm pores (p < 0.05).

Figure 5 . 18 :

 518 Figure 5.18: Typical FTIR spectra of bare pSi from various positions on the gradients. (A) Spectrum on smaller pores (18 mm from the electrode). (B) Spectrum in the middle of the gradient (10 mm from the electrode). (C) Spectrum on larger pores (4 mm from the electrode). Si-O-Si bands come from the pSi scaffold and the siloxane bound to APTES. C-N band comes from APTES.

Figure 5 .

 5 Figure 5.19 shows single representative FTIR spectra corresponding to various stages of the pSi gradients after 3 weeks of cell differentiation. The spectra obtained portrayed two phosphate (P-O) peaks: one band at 960 cm -1 and one well-defined intense band at 1025 cm -1 , corresponding to symetric and asymetric P-O stretching, respectively. The Si-O-Si band at 1070 cm -1 is hardly distinguished from the P-O band.

Figure 5 . 19 : 2 :

 5192 Figure 5.19: Typical FTIR spectra of bare pSi from various positions on the gradients, showing the vibrational assignments of the most significant bands. (A) Spectrum on smaller pores (18 mm from the electrode). (B) Spectrum in the middle of the gradient (10 mm from the electrode). (C) Spectrum on larger pores (4 mm from the electrode).

Figure 5 . 20 :

 520 Figure 5.20: FTIR spectra and deconvolution. (A) Bare pSi spectrum, with a major peak at 1070 cm-1. (B) Region corresponding to phosphate; pSi scaffold with cells after 21 days of incubation; the major peak is at 1025 cm-1. (C) Region corresponding to proteins; pSi scaffold with cells after 21 days of incubation.

Figure 5 . 21 :

 521 Figure 5.21: Top view of a pSi gradient after 3 weeks in OM: screening of the surface along the pSi gradients. Each circle represents a spot where FTIR spectra were acquired.

Figure 5 . 22 :

 522 Figure 5.22: Reconstructed 3D image of the 1250 to 1650 cm -1 peaks distribution: presence of proteins along the gradient, indicating that cells were covering the whole gradient.

Figure 5 . 23 :

 523 Figure 5.23: Reconstructed 3D image of the 960 to 1025 cm -1 peaks distribution,

1

 1 Human MSCs from dental pulp. (a): Stem cells outgrow from pulp tissue explants. (b, c, d): Stem cells colonies formed after digestion and being plated in culture dishes. Colonies formed in 3 and 5 days (a and b, respectively), cells reaching confl uence after 10 days (d). Scale bar: 300 µ m.Porous silicon and stem cells for tissue engineering 489© Woodhead Publishing Limited, 2014

2

 2 Schematic representation of the principal signalling networks and transcription factors regulating osteoblast differentiation and osteogenesis. Scanning electron microscopy (SEM) of an MSC, original magnifi cation × 2000) and of an osteoblast differentiated in vitro from MSC (original magnifi cation × 2000). BMP: bone morphogenic protein, TGF β : transforming growth factor, PDGF: platelet derived growth factor, FGF: fi broblast growth factors, IGF: insulin-like growth factor, MAPK: mitogen-activated protein kinase, Osx: Osterix transcription factor.Porous silicon and stem cells for tissue engineering 493© Woodhead Publishing Limited, 2014

  Porous silicon and stem cells for tissue engineering 495© Woodhead Publishing Limited, 2014

3

 3 Epifl uorescence microscopy and SEM of MSC culture on PSi, after 24H of culture. A: mean pore size = 30 nm. A1 and A2: epifl uorescence microscopy and SEM, respectively. B: mean pore size = 1 µ m. B1 and B2: epifl uorescence microscopy and SEM, respectively. Cells present a better adhesion on PSi with small pore size, compared to PSi with large pores: cells recover a more physiological shape (A1) and spread more (A2) on 30 nm-PSi, while cells are more bulging and spread less on 1 µ m-PSi (B1 and B2). ( Source : Adapted from CollartDutilleul (nonpublished data.)

  Porous silicon and stem cells for tissue engineering 497© Woodhead Publishing Limited, 2014

  Porous silicon and stem cells for tissue engineering 499© Woodhead Publishing Limited, 2014

  a and b: SEM of MSC osteodifferentiation and mineralized matrix deposits. (a) on PSi membrane (pore diameter = 30 nm); (b) on culture plate (polystyrene). On PSi, osteoblast-like cells are embedded in a mineralized matrix, while cells are more clearly visible on polystyrene, with a less abundant extracellular matrix. (c) quantity of mRNA bone sialoprotein (BSP) and osteocalcin (OCN) expressed by osteoblast-like cells on Si composite and on culture plate (polystyrene), after 2 days of culture in basal medium mimimum essential medium (MEM( -)) or in osteogenic medium (MEM(+)). Mean concentration of[Si] in culture medium containing Si composite is 6.5 mg/L. mRNA concentration is normalized to the concentration of β -Actin. Cells express more of osteogenic marker when cultured in the presence of[Si]. ( Source : a and b: adapted from Collart Dutilleul (non-published data). c: adapted, with permission, fromGupta et al. (2010).)

  Porous silicon and stem cells for tissue engineering 503© Woodhead Publishing Limited, 2014

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  This is a great pleasure and honor for me to have you in my thesis jury. Your expertise in the field of porous silicon and its biological applications are of prime interest for me. I thank you for evaluating my work. Emeline Collart Dutilleul : Tu es si précieuse et formidable. J'ai bien fait de t'épouser ! A ma famille : mes chers parents, frères et belles-soeurs, soeur et beau-frère, et mes admirables et abondants neveux et nièces. « Pour toi, reste attaché à tout ce que tu as appris et reçu avec une entière conviction. Tu sais de qui tu l'as appris » Au laboratoire BioSanté et Nanoscience EA 4203, et ses collaborateurs du LCC-GES : Nous avons partagé 5 années de vie, tant professionnelle que personnelle. Sous la bienveillante supervision de Frédéric et Csilla, avec la modération de Bernard Levallois, mes chers amis, que je ne vais pas tous citer... De la Pompignane : vous êtes précieux pour nous. La fraternité existe, vraiment! De la bande à choco « j'allais me coucher, mais je me suis aperçu que je vous aimais trop » De la fac... De toujours...

		Remerciements personnels / Personal thanks
	•	#$,!"#)( 2222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222 >
	• () $ '+)$#(222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222222277
		Monsieur le Docteur Guy Ladam,
		Je vous suis tout particulièrement reconnaissant d'avoir accepté d'être rapporteur de
		cette thèse, alors que vous étiez déjà en charge d'un autre rapport. Vous m'avez déjà été
		présenté comme un modèle à suivre.
		Professor Nico Voelcker,

•

Professor Raùl Martìn Palma,

The four months spent in your lab have been a very good experience for me. Thank you for your disponibility in building and following research projects. We still have some works in progress. I wish I could work in your team again.

Monsieur le

Docteur Jean-Olivier Durand, Depuis le début de mon projet de thèse, vous avez été là. Je vous remercie pour votre présence discrète et votre expertise reconnue qui ont accompagné mes travaux. Vous avez contribué au point central de ma thèse, à savoir l'approche multidisciplinaire. l'expertise de Marta Martin et Thierry Cloitre, et l'encadrement de Hélène, Elodie, Catherine et Alban. Une pensée sincère pour mes co-thésards : Hamideh et Ivan. Et mes chaleureuses salutations à tout l'équipe : Marie-Alix, Amel, Farès, Satish (brother from another mother), Olivier Romieu, Chakib (pour nous, tu fait toujours partie de l'équipe), Wafa, Michel Fages, Jacques-Henri Torrès, Jean-Cédric, Bruno Jacquot, Matthieu, Yohan, Philippe et tous les autres, dont la présence est importante. Quelques remerciements plus spécifiques à Marta qui est à la fois mon encadrante et mon amie et à Cyril, à Emilie Secret qui a fabriqué de si nombreux échantillons de pSi, à Pierre Jacquemart qui m'a toujours accueilli au bloc, à Franck Chaubron, Marion, Severine et Catherine pour la collaboration avec l'Institut Clinident BioPharma, et à Christophe Hirtz avec qui le contact est resté intact. Warm greetings to the Voelcker laboratory, Mawson Institute, University of South Australia, with special thought to my dear friend Ruth Zoehrer, to Sandhya, Fran, Steve. Je remercie également l'école doctorale CBS2 et surtout l'Asso cbs2, qui arrive à donner une belle dynamique à notre formation. A Quelques remerciements plus spécifiques à Anne et Robby qui ont relu et corrigé l'anglais de ce manuscrit, à Monsieur Alban Desoutter qui m'a soutenu jusqu'au 42e km, et aux bons amis.

Chapter 2 Dental Pulp Stem Cells Recovery 2.1 Introduction 2.1.1 Dental pulp Teeth

  are formed with two main parts: the crown and the root that can be defined with anatomic or clinical criteria. They are linked, by the periodontal ligament, to the supporting alveolar bone, which is composed of both compact and trabecular bone (Figure2.1).

	Pore diameter influence on Dental Pulp Stem Cells Adhesion, Proliferation and
	Osteodifferentiation: use of porous silicon gradients to specifically determine an
	ideal pore diameter for DPSC use in bone tissue engineering.
	Dental pulp stem cells behavior on porous structure at the nanometer scale: atomic
	force microscopy assessment of DPSC spreading and filopodia formation on pSi
	scaffolds.

Figure 2.1: Schematic

  of a tooth, showing both internal structures and supporting ligament and bone. DPSC are recovered from the inner dental pulp.

	Adapted from the National Library of Medicine, National Institute of Health
	(http://www.nlm.nih.gov/medlineplus/ency/imagepages/1121.htm).
	The crowns of human teeth consist of enamel, dentin, and dental pulp tissue. During
	tooth growth and development, ameloblasts form enamel and odontoblasts generate
	primary dentin. After tooth eruption, ameloblasts disappear from the surface of the
	enamel; consequently, enamel formation ceases to occur naturally in vivo. In contrast,

Enamel Dentin Dental Pulp Trabecular Bone Compact Bone Ligament Cement Crown Root

Left Ventricular Wall thickness Fibrotic area (Left Ventricular) A B

	Control
	2.1.3.4 Neural tissue regeneration	Control	DPSC
	DPSC are thought to originate from migrating cranial neural crest cells. During
	DPSC embryonic development, these neural crest cells differentiate into a wide variety of cell
	types, including neurons of the peripheral nervous system. In 2008, Arthur and co-
	workers demonstrated that ex vivo expanded human adult DPSC responded to neuronal
	inductive conditions to form neuronal cells, and expressed neuronal specific markers,
	both in vitro and in vivo (Arthur et al., 2008). Thus, adult human DPSC could provide a
	readily accessible source of stem cells for therapeutic applications in neurological
	pathologies. Several experiments with DPSC grafted in mice and rats assessed the
	regeneration of injured neural tissue (brain or spinal chord injuries) (Huang et al., 2008)
	(Kiraly et al., 2009)

Table 2 .

 2

	Osteogenic medium	Adipogenic medium	Chondrogenic medium
	αMEM	αMEM	αMEM
	15% FBS	10% FBS	Serum free
	10 -8 M Dexamethasone	10 -5 M Dexamethasone	10

1: Antibodies for flow cytometry analysis

2.2.4.4 Differentiation assays

Recovered cells were controlled for multipotency with in vitro osteogenic, chondrogenic and adipogenic differentiation. For osteogenic and adipogenic differentiation, dental pulp cells were seeded at a density of 10 5 cells/cm 2 and maintained in αMEM supplemented with 10% FBS until confluence. Cells were then cultured in osteogenic medium (αMEM supplemented with 15% FBS, dexamethasone, L-ascorbate phosphate, and β-glycerophosphate) or in adipogenic medium (αMEM supplemented with 10% FBS, dexamethasone, insulin, and isobutylmethylxanthine) for 21 days, with medium changed twice a week. Control cultures were maintained without osteogenic or adipogenic supplements. To induce chondrogenic differentiation, cells were cultured for 21 days as pellets in a serum-free expansion medium containing dexamethasone, L-ascorbate phosphate, L-proline, sodium pyruvate, ITS-Premix, and TGF-β3. Control pellet cultures were maintained in basal medium

(Kemoun et al., 2007)

. The detailed differentiation media and supplements are presented in Table

2

.2 -7 M Dexamethasone 50 µg/mL L-Ascorbate Phosphate 50 µg/mL L-Ascorbate Phosphate 50 µg/mL L-Ascorbate Phosphate 5 mM β-Glycerophosphate 1 µg/mL Insulin 40 µg/mL L-Proline 1.8 mM Monopotassium Phosphate 0.5 mM Isobutylmethylxantine 100 µg/mL Sodium Pyruvate 50 mg/mL ITS-Premix 10 ng/mL TGF-β3

Table 2 .2: Composition

 2

of the various differentiation media.

Table 3 . 1 :

 31 Water contact angle of the various scaffolds (Flat Si and pSi)3.3.2 Cell adhesion and growth

pSi 10 nm pSi 36 nm pSi 1µm Undecenoic Acid APTES Semicarb Cell Counts

  , pSi 10 nm (p<0.001) and pSi 36 nm (p<0.001). It is remarkable that there were significantly more cells on pSi 36 nm than on pSi 10 nm after 48 h and 72 h (p=0.006 and p=0.004, respectively), and that the number of cells on semicarbazidetreated pSi increased during 24 h and decreased abruptly after 48 h. After 72 h, DPSC

	(p<0.001)
		4 h
	>1000	24 h 48 h
		72 h
	500
	400
	300
	200
	100
	Glass	Flat Si

Table 4 . 1 :

 41 .1. Pore diameters (nm) measured along pSi gradients

	Distance to Electrode (mm)	18	16	14	12	10	8	6	4	2
	Pore Diameter (nm)	8	14	18	24	37	70	120 248 500
	Standard Error	5	7	8	11	18	28	44	92	209

positive colonies 40.3 62.1 57.8 44.8 36.1 46.2 36.3 32.0 26.6 20.3

	Pore Diameter (nm)	Flat Si	8	14	18	24	37	70	120	248	500
	% of CB-Standard Error	8.7	7.7	8.7	10.5	8.7	12.5	6.7	7.8	7.4	6.6

Table 4 . 2 :

 42 Percentage of CB-positive colonies along the pSi gradient, with the corresponding pore diameters.

Table 4 . 3 :

 43 Ca/P ratio of the mineralized structures recovered on pSi and flat Si after 3 weeks of incubation with or without cells, in OM or MEM.

Table 5 . 1 :

 51 Pore diameter (Mean values and standard deviation of the mean), porosity and thickness of the porous layer at each stage of the pSi gradient.

	18	16	14	12	10	08	06	04	02	00	18	16	14	12	10	08	06	04	02	00
		Distance to electrode (mm)		electrode		Distance to electrode (mm)		Electrode

  Stem cells constitute the source of differentiated cells for the generation of tissues during development and for regeneration of tissues that are diseased or injured postnatally. Stem cell-based therapy often requires a scaffold to carry cells and/or growth factor to the injured site. Porous silicon (PSi) is a promising biomaterial for tissue engineering, as it is both non-toxic and bioresorbable. Surface modi cation can offer control over the degradation rate of PSi and can also impart properties to promote cell adhesion. Thus, coupling the good proliferation and differentiation capacities of either adult mesenchymal stem cells (MSC)

		20
	Porous silicon scaffolds for stem cells
	growth and osteodifferentiation
	P. Y. COLLART DUTILLEUL , D. DEVILLE DE P É RI È RE and F. J. CUISINIER, UFR Odontologie, France , F. CUNIN, Institut Charles Gerhardt Montpellier, France and i t e d C. GERGELY, Montpellier 2 University, France
	DOI: 10.1533/9780857097156.3.486 Abstract:	L i m
	P u b l i s h i n g
	confocal Raman microscopy. W o o d h e a d
	486
	© Woodhead Publishing Limited, 2014

  ).

	Porous silicon and stem cells for tissue engineering	487
		i t e d
	L i m
	P u b l i s h i n g
	W o o d h e a d
	© Woodhead Publishing Limited, 2014
	20_p486-506.indd 486	10/8/2013 7:00:17 PM

  Wnt represents a family of secreted molecules which are involved in various aspects of cellular biology, including cell growth, differentiation and death. The binding of Wnt proteins permits the translocation of β -catenin into the nucleus, resulting in an enhanced expression and activity of the osteogenic promoter Runx2. Wnt signalling and hedgehog signalling cooperate in controlling bone formation, especially at the early stage of osteoblast formation, initiating cellular differentiation. Cell-cell interactions are also required for activating some signalling cascade: notch signalling and ephrin signalling are the main cell-cell interaction pathways. Notch receptors and ligands are membrane proteins able to up-regulate bone related genes to favour the proliferation of immature osteoblasts, while limiting osteoblast maturation by inhibiting the Runx2 transactivation function. Ephrin ligands and receptors are also membrane proteins. They are constituted by two classes of ephrins: the B class ephrin ligands and receptors (EphB-Tyrosine Kinase receptor) induce osteogenic transcription factors, and inhibit osteoclast differentiation, leading to bone formation. EphrinA/EphA receptor interaction facilitates bone remodelling by suppressing osteoblast differentiation and activating osteoclast differentiation (Fig.20.2).

	Porous silicon and stem cells for tissue engineering	491
	pathway is activated by a variety of growth factors such as broblast growth
	factors (FGF), platelet derived growth factors (PDGF), transforming growth
	factor beta (TGF β ), bone morphogenic proteins (BMP) and insulin-like
	growth factors (IGF). These extracellular signals activate a signalling cas-
	cade of MAP kinase inside the cell, resulting in cell proliferation and osteo-
	differentiation. Another pathway involved in osteodifferentiation is the Wnt i t e d signalling pathway:
	L i m
	P u b l i s h i n g
	MSCs proliferate and differentiate under the in uence o o d h e a d W
	of a number of molecular pathways, but mitogen-activated protein kinase
	(MAPK) signalling plays a pivotal role in osteogenic induction. MAPK
	© Woodhead Publishing Limited, 2014

  Yim, E.K.F.,Darling, E.M., Kulangara, K., Guilak, F. and Leong, K.M. (2010) 'Nanotopography-induced changes in focal adhesions, cytoskeletal organization, and mechanical properties of human mesenchymal stem cells', Biomaterials , 31 , 1299-2306.

	504 506	Porous silicon for biomedical applications Porous silicon for biomedical applications
		i t e d i t e d
		L i m L i m
		P u b l i s h i n g P u b l i s h i n g
	W o o d h e a d W o o d h e a d
		© Woodhead Publishing Limited, 2014 © Woodhead Publishing Limited, 2014

  Birke, O., McDonald, M., Little, D. and Dunstan, C.R. (2010) 'The incorporation of strontium and zinc into a calcium-silicon ceramic for bone tissue engineering', Biomaterials , 31 (12), 3175-3184. Zuk, P. A., Zhu, M., Ashjian, P., De Ugarte, D.A., Huang, J.I., Mizuno, H., Alfonso, Z.C., Fraser, J.K., Benhaim, P. and Hedrick, M.H. (2002) 'Human adipose tissue is a source of multipotent stem cells', Mol Biol Cell , 13 , 4279-4295.

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ACS Paragon Plus EnvironmentACS Applied Materials & Interfaces

Acknowledgements

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ACS Applied Materials & Interfaces

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Investigation showed that ACP nucleation started with nanometer-sized CaP complexes, and these prenucleation clusters aggregated to form amorphous calcium phosphate, such as OCP [START_REF] Cuisinier | High resolution electron microscopy study of crystal growth mechanisms in chicken bone composites[END_REF] [START_REF] Houllé | High resolution electron microscopy: structure and growth mechanisms of human dentin crystals[END_REF] [START_REF] Habraken | Ion-association complexes unite classical and non-classical theories for the biomimetic nucleation of calcium phosphate[END_REF]. These works bring out that the nucleated clusters are of the nanometer-scale and nucleation rate is higher due to size similarities. Je sais que ma décision est révocable à tout moment, sans justification et sans que cela ne porte aucunement préjudice à la qualité des soins qui me seront prodigués. J'ai également été informé(e) que des données médicales me concernant peuvent être transmises au responsable de ces recherches scientifiques, dans le strict respect du secret médical.

Appendix

Si je le souhaite, les données me concernant, pourront m'être communiquées directement ou par l'intermédiaire d'un médecin de mon choix. Je dispose également du droit d'être informé(e) des résultats globaux de cette recherche.