A Vallés-Lluch

E Costa

G Gallego Ferrer

M Monleón Pradas

M Salmerón-Sánchez

M Monleón Pradas

Structure and biological response of polymer/silica nanocomposites prepared by sol-gel technique

Keywords: Nanocomposite, hybrid, acrylate, silica, cellular response

Structure of P(EMA-co-HEA)/SiO 2 nanocomposites with silica content in the range from 0 to 30 %wt was correlated with cell behaviour on substrates of those compositions by making use of two different populations of primary human cells: articular cartilage chondrocytes and dental pulp cells. Substrates were prepared by the simultaneous copolymerization of the organic monomers and the sol-gel reaction of the silica precursor in different proportions, which led to weight fractions of the silica phase in the materials closely matching the stoichiometric ratios employed during the preparation, both in the bulk and at the material surface. The silica nanophase increases surface wettability and improves the mechanical properties of the base materials. Both chondrocytes and dental pulp cells were cultured on serum-coated nanocomposite substrates in the same conditions, but very different cellular responses were obtained.

While chondrocytes adhered and proliferated, dental pulp cells formed viable aggregates weakly adhered on the sample that were viable up to 11 days. The results suggest that these sol-gel derived nancomposites may be used as culture surfaces maintaining the dental pulp cell phenotype in vitro.

Introduction

Polymer hydrogels stand out among synthetic materials used as biomaterials because of their good biocompatibility due, mainly, to their ability to absorb large amounts of water. However, they have poor mechanical properties and, as a consequence, they are in need of reinforcement. There are several routes to reinforce a polymer hydrogel, including the copolymerization with a hydrophobic component and the preparation of interpenetrating polymer networks [START_REF] Sánchez | Influence of the hydrophobic phase on the thermal transitions of water sorbed in a polymer hydrogel based on interpenetration of a hydrophilic and a hydrophobic polymer network[END_REF][START_REF] Sánchez | The structure of poly(ethyl acrylate-co-hydroxyethyl methacrylate) copolymer networks by segmental dynamics studies based on structural relaxation experiments[END_REF]. A different strategy consists in the preparation of hybrid organic/inorganic nanocomposites, which include a silica phase obtained simultaneously with the polymer in a sol-gel process. [START_REF] Hernández | Properties of poly(2-hydroxyethyl acrylate)-silica nanocomposites obtained by the sol-gel technique[END_REF].

Apart from reinforcement, there is a great need of providing bioactivity to surfaces in applications where good interaction between the synthetic material and the living tissues in mandatory. Since the discovery of the 45S5 Bioglass® by Hench in 1971 [START_REF] Hench | The story of Bioglass[END_REF], which appeared to strongly adhere to bone tissue, various kinds of ceramics such as Na 2 O-CaO-SiO 2 -P 2 O 5 glasses, sintered hydroxyapatite and glass-ceramics containing apatite or wollastonite, are known to bond directly to living bone avoiding the formation of fibrous tissue [START_REF] Ben-Nissan | Bioactive glasses and glass ceramics[END_REF][START_REF] Kokubo | Bioactivity of glasses and glass-ceramics[END_REF][START_REF] Hench | Handbook of Biomaterials Properties[END_REF]. Some of them or their derivatives are being clinically used in artificial middle-ear bone implants, artificial vertebrae, intervertebral discs, iliac bones, fillers in bone and maxillofacial defects or to fill the gaps around the implants, etc. and have inspired new bioactive materials with varied compositions and structures to fulfill different applications [START_REF] Kokubo | How useful is SBF in predicting in vivo bone bioactivity?[END_REF][START_REF] Ohtsuki | Mechanism of apatite formation on CaO-SiO 2 -P 2 O 5 glasses in a simulated body fluid[END_REF][10][START_REF] Zhang | Effects of materials parameters on mineralization and degradation of sol-gel bioactive glasses with 3D-ordered macroporous structures[END_REF][START_REF] Jones | Optimising bioactive glass scaffolds for bone tissue engineering[END_REF][START_REF] Kim | The mechanism of biomineralization of bone-like apatite on synthetic hydroxyapatite: an in vitro assessment[END_REF]. However, the use of bioactive glasses for the stimulated regeneration of tissues is limited by their poor mechanical properties, since their brittleness makes difficult their use in load-bearing applications.

One approach to combine bioactivity and enhanced mechanical properties in the same material is the preparation of organic/inorganic hybrids, in which the inorganic phase is incorporated into an organic polymeric matrix. This type of structural organization to a certain extent mimics the composition of mineralized tissues, where the hydroxyapatite (HAp) inorganic phase is dispersed in a collagen matrix with which it interacts at a molecular level yielding a natural composite. Various hybrid systems have been explored employing as inorganic phases bioactive glasses [START_REF] Fang | Processing of hydroxyapatite reinforced ultrahigh molecular weight polyethylene for biomedical applications[END_REF][START_REF] Li | Preparation and in vitro investigation of chitosan/nano-hydroxyapatite composite used as bone substitute materials[END_REF], HAp [START_REF] Jie | Tissue engineering scaffold material of nano-apatite crystals and polyamide composite[END_REF][START_REF] Eglin | Comparative study of the in vitro apatite-forming ability of poly(ε-caprolactone-silica sol-gels using three osteoconductivity tests (static, dynamic, and alternate soaking process)[END_REF][START_REF] Zhao | Preparation and characterization of macroporous chitosan/wollastonite composite scaffolds for tissue engineering[END_REF] or pure silica [START_REF] Kawai | In vitro apatite formation on polyamide containing groups modified with silanol groups[END_REF], in the form of particles or fibres. Porous scaffolds of hybrid composites mimicking natural bone structures have been prepared and investigated to serve as a support and guide new tissue in-growth and regeneration [START_REF] Meretoja | Crosslinked poly(ε-caprolactone/D,L-lactide)/bioactive glass composite scaffolds for bone tissue engineering[END_REF][START_REF] Rezwan | Biodegradable and bioactive porous polymer/inorganic composite scaffolds for bone tissue engineering[END_REF][START_REF] Huang | Structural characteristics and properties of silica/poly(2-hydroxyethyl methacrylate) (PHEMA) nanocomposites prepared by mixing colloidal silica or tetraethyloxysilane (TEOS) with PHEMA[END_REF][START_REF] Hajji | Synthesis, structure and morphology of polymer-silica hybrid nanocomposites based on hydroxyethyl methacrylate[END_REF]. In the recent past many studies have been devoted to organic/inorganic nanocomposites with the aim of combining the properties of both phases at the nanoscale level. Polymer/silica nanocomposites have been prepared by physically mixing silica nanoparticles with polymers [START_REF] Liu | Poly(methylmethacrylate)-silica nanocomposites films from surface-functionalized silica nanoparticles[END_REF][START_REF] Kawai | In vitro apatite formation on polyamide containing groups modified with silanol groups[END_REF] or by copolymerization of the organic polymers with surfacefunctionalized silica nanoparticles [START_REF] Oyane | Apatite formation on ethylene-vinyl alcohol cpopolymer modified with silanol groups[END_REF][START_REF] Kawashita | Apatite-forming ability of carboxyl group-containing polymer gels in a simulated body fluid[END_REF]. Various surface modification techniques have also been proposed aiming to enhance the bioactivity of polymer surfaces. These methods are based on the introduction of hydrophilic polar groups, such as phosphate, hydroxy groups, Si-OH or Ti-OH, and carboxy and carboxylate groups, which have been found to be effective for apatite nucleation, onto more hydrophobic less bioactive substrates [START_REF] Rhee | Preparation of a bioactive and degradable poly(caprolactone)/silica hybrid through a sol-gel method[END_REF][START_REF] Rhee | Effect of molecular weight of poly(e-caprolactone) on interpenetrating network structure, apatite-forming ability, and degradability of poly(caprolactone)/silica nano-hybrid materials[END_REF][START_REF] Costa | Apatite formation on poly(2-hydroxyethyl methacrylate)-silica hybrids prepared by sol-gel process[END_REF]. Nonetheless, the idea of nanohybridizing biocompatible polymers with silica by this simultaneous in situ sol-gel polymerization to improve their bioactivity is a very recent one. Few published works address it. Rhee et al. [START_REF] Rhee | Preparation of a bioactive and degradable poly(caprolactone)/silica hybrid through a sol-gel method[END_REF][START_REF] Rhee | Effect of molecular weight of poly(e-caprolactone) on interpenetrating network structure, apatite-forming ability, and degradability of poly(caprolactone)/silica nano-hybrid materials[END_REF] polymerized poly(ε-caprolactone) simultaneously via sol-gel with tetraethoxysilane (TEOS) as silica precursor and calcium nitrate tetrahydrate to dope the nanohybrid with calcium ions to facilitate the apatite nucleation. Costa et al. [START_REF] Costa | Apatite formation on poly(2-hydroxyethyl methacrylate)-silica hybrids prepared by sol-gel process[END_REF] polymerized poly(2hydroxyethyl methacrylate) by the same procedure with tetramethoxysilane (TMOS) as silica precursor and characterized the hydroxyapatite layer formed.

In previous works we addressed the structure, physicochemical properties and hydroxyapatite nucleation ability of poly(ethyl methacrylate-co-hydroxyethyl acrylate)/silica, P(EMA-co-HEA)/SiO 2 , nanohybrids with silica contents comprised between 0 and 30 %wt as potential candidates for the synthetic matrix of scaffolds for bone or dentin regeneration [START_REF] Hernández | Properties of poly(2-hydroxyethyl acrylate)-silica nanocomposites obtained by the sol-gel technique[END_REF][START_REF] Vallés-Lluch | Biomimetic apatite coating on P(EMA-co-HEA)/SiO 2 nanocomposites[END_REF][START_REF] Vallés-Lluch | Surface modification of P(EMA-co-HEA)/SiO 2 nanohybrids for faster hydroxyapatite deposition in simulated body fluid[END_REF]. Here we analyze the effect of the presence of silica on the in vitro biological performance of this series of substrates making use of two different human primary cells: articular cartilage chondrocytes and dental pulp cells. This work is a first attempt to understand the interaction of these specific sol-gel derived nanocomposites with human cells in vitro. We employed two cell populations stemming from human tissues with very different properties and lineage. Chondrocytes are mature cells which are in need of adequate substrates for their in vitro expansion while maintaining their phenotype; dental pulp cells are immature, multipotent cells able to differentiate to osteoblasts. Our results show that sol-gel derived nanocomposites might be useful for maintaining the characteristic dental pulp phenotype in vitro, minimizing interactions with the artificial substrate while keeping the cell population viable.

Materials and methods

Synthesis of P(EMA-co-HEA)/SiO 2 nanohybrids

Nanocomposites of poly(ethyl methacrylate-co-hydroxyethyl acrylate), p(EMA-co-HEA), with fixed EMA/HEA weight ratio of 70/30 %wt and with varying proportions of silica, SiO 2 : 0, 5, 10, 15, 20 and 30 %wt, were obtained in the form of sheets of 0.8 mm in thickness similarly as in [START_REF] Hernández | Properties of poly(2-hydroxyethyl acrylate)-silica nanocomposites obtained by the sol-gel technique[END_REF]. Briefly, the procedure consisted in preparing an organic mixture of the organic monomers ethyl methacrylate, EMA (99%, Aldrich), and hydroxyethyl acrylate, HEA (96%, Aldrich), with a 0.5 %wt of ethylene glycol dimethacrylate, EGDMA (98%, Aldrich), as crosslinking agent and a 2 %wt of benzoyl peroxide, BPO (97%, Fluka), as thermal initiator, relative to monomer weight.

Separately, tetraethoxysilane, TEOS, was mixed with distilled water and hydrochloric acid (37%, Aldrich) in the molar ratio 1:2:0.0185, respectively (inorganic mixture).

After 30 min of separate stirring, both solutions were mixed, stirred for another 30 min and injected into glass moulds. The hybrid mixture was polymerized in an oven at 60ºC for 21 h and post-polymerized at 90ºC for 18 h, rinsed in boiling distilled water/ethanol mixture 50/50 %vol for 24 hours to eliminate monomer residues and finally dried in a vacuum desiccator at 80ºC until constant weight. Materials with silica contents of 0, 5, 10, 15, 20 and 30 %wt were obtained by calculating the (EMA+HEA)/TEOS ratios assuming that the sol-gel reaction was complete. Hereafter, the hybrids will be referred to as Hx, x being the percentage of silica. Besides, poly(ethyl methacrylate), PEMA, and poly(hydroxyethyl acrylate), PHEA, homopolymers were prepared following the same procedure as reference systems.

Fourier-transform infrared spectroscopy (FTIR) analyses

Fourier-transform infrared (FTIR) spectra were collected in a Thermo Nicolet Nexus FTIR spectrometer (Thermo Fischer Scientific Inc., Waltham, MA, USA), in the attenuated total reflection mode (ATR), to elucidate the surface compositions. The spectra resulted from averages of 128 scans at 4 cm -1 resolution, between 650 and 4000 cm -1 .

Quantification of SiO 2 by energy dispersive x-ray spectroscopy (EDS)

The percentage of silica at the surfaces and in the interior of the nanohybrids was quantified by energy dispersive x-ray spectroscopy (EDS) in an Oxford Instruments spectrometer, attached to a JSM-6300 scanning electron microscope (JEOL Ltd., Tokyo, Japan). Exterior and inner (fracture) surfaces were analyzed. Samples were previously sputter-coated with carbon under vacuum. Spectra were taken at 10 kV of acceleration voltage and 15 mm of working distance. Silicon was employed as optimization standard.

Contact angle measurements

In order to determine the surface wettability of the different substrates, contact angle measurements were conducted in air atmosphere with the sessile drop method, employing water (extra pure, Scharlau) as drop liquid. A Dataphysics OCA instrument (DataPhysics Instruments GmbH, Filderstadt, Germany) was used for this purpose.

Glass Pasteur pipettes were employed as droppers. A minimum of 10 drops of water was analyzed for each composition.

Compression tests

Mechanical compression tests were performed in a Seiko TMA/SS6000 device (Seiko Instruments Inc., Chiba, Japan), from 0.05 to 150 g at 10 gmin -1 , at room temperature.

Specimens were disk-shaped (3 mm in diameter, 0.8 mm-thick) and each composition set consisted of five replicas.

Chondrocytes

Human articular cartilage from the knee of a patient undergoing total knee arthroplasty was processed for chondrocyte isolation. Briefly, the cartilage tissue was aseptically dissected from the joint, minced, and washed with Dulbecco's modified Eagle's medium (DMEM, Life Technologies). Then, the cartilage was incubated for 30 min with a 0.5 mg/mL hyaluronidase (Sigma-Aldrich) solution and for 1 h with a 1 mg/mL pronase (Merck) solution in a shaking water bath at 37ºC. After that, the cartilage fragments were washed with DMEM and incubated with a 0.5 mg/mL collagenase-IA (Sigma-Aldrich) solution in a shaking water bath at 37ºC overnight. The resulting cell suspension was filtered with a 70 µm cell strainer (BD Biosciences) to remove any undigested tissue and collagenase was rinsed off with DMEM containing 10% FBS (Invitrogen SA). Finally, the cell suspension obtained was transferred in 15 mL DMEM supplemented with 10% FBS and 50 µg/mL ascorbic acid (Sigma-Aldrich) to a 75 cm 2 tissue culture flask (Nunc) and maintained at 37ºC, in a humidified atmosphere under 5% CO 2 . The culture medium was replaced every 2 days and cells were allowed to grow until subconfluence. Then, the cells were harvested by trypsinisation and counted with a hemacytometer for the experiment.

Dental pulp cells

Human dental pulp cells were isolated from third molar teeth extracted due to odontological reasons from healthy patients. First, pulp was extracted and washed with Hank's balanced solution (Gibco-Invitrogen, USA) supplemented with penicillin/streptomycin (100 U/ml and 100 µg/ml respectively, Gibco-Invitrogen). Pulp was minced in small fragments and incubated with a 2 mg/mL collagenase-IA (Sigma-Aldrich) solution in a shaking water bath at 37ºC for 90 min. Cellular suspension was centrifuged (400 g for 7 min) and the resulting pellet was seeded on a to a 75 cm 2 tissue culture flask (Nunc) in DMEM containing 10% FBS and antibiotics and maintained at 37ºC, in a humidified atmosphere under 5% CO 2 . The culture medium was replaced every 2 days and cells were allowed to grow until subconfluence. Then, the cells were harvested by trypsinisation and counted with a hemacytometer for the experiment.

Cell culture

Films pre-sterilized with 25 kGy gamma radiation were placed in a 96-well tissue culture plate and were soaked in culture medium for 72 h before cell seeding. Then, 40 µL (10 4 cells) of the cell dispersion were placed onto the polymer films and were incubated at 37ºC under 5% CO 2 condition for 1 h. After this time, 600 µL of fresh medium were added to each well. Samples were maintained at 37ºC, in a humidified atmosphere under 5% CO 2 for 10 days. The culture medium used was DMEM supplemented with 10% FBS and 50 µg/mL ascorbic acid and it was renewed every 2-3 days. Contrast phase light microscopy was carried out to observe morphological changes in cell shape after different time cultures. Neutral red staining was employed as a vital staining. Each experiment was performed in triplicate and TCPS served as the control substrate.

Results and discussion

Figure 1 shows the FTIR spectra of the nanocomposites and both homopolymers. The CH x asymmetric and symmetric stretching peaks appear at 2962 and 2888 cm -1 , respectively, in all spectra, and the broad band between 3700 and 3100 cm -1 , as well as the shoulder at 1650 cm -1 , characteristic of the hydroxy groups, only appears in the PHEA spectrum (data not shown). The well-defined C=O stretching peak (1700 cm -1 ) decreases with the percentage of silica. Between 1700 and 650 cm -1 the FTIR curves show the complex spectrum of the copolymer blurring the silica fingerprint region. The intensity of the peaks at 1100 cm -1 and 800 cm -1 (Si-O-Si asymmetric and symmetric stretching, respectively) and 950 cm -1 (Si-OH stretching), characteristic of the silica phase [START_REF] Rhee | Preparation of a bioactive and degradable poly(caprolactone)/silica hybrid through a sol-gel method[END_REF][START_REF] Rhee | Effect of molecular weight of poly(e-caprolactone) on interpenetrating network structure, apatite-forming ability, and degradability of poly(caprolactone)/silica nano-hybrid materials[END_REF], increases with the silica content. No evidence of hybrid Si-O-C bonds from heterocondensation reactions can be found by FTIR, because their characteristic peak (1120-1080 and 836 cm -1 ) [START_REF] Rhee | Effect of molecular weight of poly(e-caprolactone) on interpenetrating network structure, apatite-forming ability, and degradability of poly(caprolactone)/silica nano-hybrid materials[END_REF] overlaps the absorption interval of the Si-O-Si bonds. The absence of the broad band characteristic of the OH groups of the silanols or the OH vibrations of HEA can be explained by an imperfect contact of the sample and the measuring device, caused by the rigidity of the samples at room temperature. This band appeared only after immersion of the samples in water to render them softer (data not shown).

The experimental silica contents, SiO 2 EDS , listed in Table 1 were calculated from the silicon average contents obtained by EDS, making use of the silica and silicon molar masses and neglecting the amount of hydrogen in the copolymer and in the silanol groups, undetectable by this technique and the carbon coating of the samples for performing the EDS measurements. The surface and bulk silica percentages are very similar and correlate quite well with the nominal silica contents of the hybrids, despite the considerations concerning the detection of hydrogen and carbon. This is an important issue since the biological media will interact with the components on the surface of the sample. The larger standard deviations of the silica contents obtained in the bulk are due to the limitations of this technique when working with fractured samples of variable height. Both the FTIR spectra and the silica quantifications by EDS corroborate that TEOS was efficiently hydrolyzed and condensed to silica during the sol-gel process, and silica was found to be homogeneously distributed throughout the polymeric matrix both in the bulk and at the surface.

The average water contact angles, θ w , are also listed in Table 1. The water contact angle of the H05 surface is identical to that of the copolymer, lying between those of the homopolymers. It then decreases for intermediate silica contents (10-20 %wt) and increases again for the H30 surface, reaching a value close to that of the copolymer. The compressive elastic moduli, E, were obtained from the initial slopes of the stress-strain curves, and are also listed in Table 1. The modulus of the copolymer lies between those of the homopolymers, and maintains similar values up to 15 %wt of silica, from where it starts increasing up to a value similar to that of PEMA for the H30 composite.

As explained in a previous work [START_REF] Vallés-Lluch | Biomimetic apatite coating on P(EMA-co-HEA)/SiO 2 nanocomposites[END_REF], different inorganic phase morphologies are obtained within the copolymeric matrix depending on the silica content. Below a 15 %wt of silica the inorganic phase is uniformly dispersed in the nanocomposites as porous nanoparticles which form larger aggregates leaving spaces of tens of nanometers. Higher silica contents produce silica networks which percolate and eventually extend continuously throughout the samples, totally interpenetrated with the organic network. These different phase morphologies influence the physicochemical properties of the surfaces. The number of free Si-OH groups is larger the larger the boundary interface of the silica phase, and thus decreases as percolation occurs and with the development of the continuous silica network. This is reflected in the wettability of the nanocomposites with high silica contents. Besides, the reinforcing effect of the interpenetrated silica phase is observed in the increase of the compressive elastic modulus of the nanohybrids.

Cells respond to different kinds of surface parameters: chemical, topographical and mechanical, through the intermediacy of extracellular matrix (ECM) proteins. Here we investigated the role of increasing amounts of silica of the material surface on the biological performance of the substrate. Cellular behavior at the biomaterials interface can be understood in terms of the natural adhesive interaction of cells with the ECM [START_REF] Spie | Tissue engineering and reparative medicine[END_REF]. Cells do not interact directly with foreign materials, but they attach to the adsorbed layer of proteins on the material surface, such as fibronectin (FN), vitronectin (VN), fibrinogen (FNG), representing the so called soluble matrix proteins in the biological fluids [START_REF] Spie | Tissue engineering and reparative medicine[END_REF]. Cells recognize these matrix proteins via integrins -a family of cell surface receptors -that provide trans-membrane links between the ECM and the actin cytoskeleton [START_REF] Hynes | Integrins: bidirectional, allosteric signaling machines[END_REF]. When integrins are occupied they cluster and develop focal adhesion complexes that actually anchor the cells to the surface and trigger the subsequent cellular response [START_REF] Hynes | Integrins: bidirectional, allosteric signaling machines[END_REF]. Thus, the cell-material interaction is a complex process consisting of early events, such as adsorption of proteins, followed by cell adhesion and spreading, and later events, related to cell growth, differentiation, matrix deposition and cell function.

Figure 2 shows human articular chondrocytes after 1 day of culture on the different substrates. Cells are well spread and distributed throughout the surface of the samples independently of the amount of silica in the material (from 5 to 30 %). The fraction of the surface covered by cells is lower for the homopolymers (both PEMA and PHEA) than for the 70/30 %wt copolymer. This family of polymers represents a set of materials with very different mechanical properties (see table 1) that support chondrocyte expansion. From an applied perspective, the incorporation of silica within the polymer matrix represents a way of matching the properties of the material to those of articular cartilage; since the stiffness of the sample is known to influence the cell-material interaction [START_REF] Pelham | Cell locomotion and focal adhesions are regulated by substrate flexibility[END_REF][START_REF] Genes | Effect of substrate mechanics on chondrocyte adhesion to modified alginate surfaces[END_REF][START_REF] Ryan | Tissue Spreading on implantable substrates is a competitive outcome of cell-cell vs. Cell-substratum sensivity[END_REF], it is important in our case that this increased stiffness does not modify the biological performance of the substrate. Chondrocytes remain covering the surface of the material and start growing in layers (cell on cell) up to 10 days on culture (results not shown, microscopy images are similar to those shown in Figure 2 where a confluence monolayer is already formed).

The situation is completely different when the same family of materials is seeded with human dental pulp cells (Figure 3). After two days of culture the cells are spread on PEMA, PHEA and the 70/30 %wt copolymer. However, addition of silica to the system leads to a complete modification of the cell-material interaction, which changes dramatically as the amount of silica in the system increases. Cells tend to form clusters which adhere weakly to the surface of the material, in a growing number as the silica content increases. This is in agreement with the results found by Cousins et al. [START_REF] Cousins | The effect of silica nanoparticulate coatings on cellular response[END_REF] with L929 murine fibroblasts seeded on silica nanoparticulate coatings, which showed a weakly adhered rounded conformation in comparison with the same cells seeded on a plain glass surface, with spread flattened morphology. These authors hypothesized that the nanostructure of silica governs the serum proteins-substrate interactions, leading to localized focal adhesion sites, and thus to weakly adhered rounded cells. In our case, the presence of silica on the material substrate implies the presence of a higher concentration of OH groups on the material and a slightly negatively charged surface that leads to a more wettable surface (table 1).

After 5 days of culture the cells remained clustered on the compositions with the lower content of silica (5, 10 and 15 %wt) while good adhesion and proliferation occurred on the samples with 20 and 30 %wt silica (Figure 4). This behavior was previously found on TEOS based chitosan/silica membranes seeded with fibroblast-like cells. Cell growth rate, density and adhesion strength were lower for hybrid materials than for pure chitosan for lower silica contents (below 20 %), but similar or even increased values were obtained for a 50 %wt inorganic content [START_REF] Coradin | Sol-Gel Biopolymer/Silica Nanocomposites in Biotechnology[END_REF].

The pulp cells are shown to be viable after 5 days on the material as revealed by neutral red staining for the cell clusters (Figure 5). Besides, when these cells are collected from the nanocomposites and seeded again on a standard substrate they acquire the typical fibroblast-like morphology (Figure 5). This peculiar cell-material interaction behavior has interesting consequences for tissue engineering applications requiring phenotype maintenance.

It is interesting to stress that the same nanocomposite materials, coated with the same solution of proteins (FBS in our case, that consists mainly of fibronectin and vitronectin), has a very different cell response depending on the cell type seeded, which reveals the specificity of the cell-material interaction. It is known that human chondrocytes tend to spread on material surfaces and adopt an extended morphology as a consequence of the production of collagen and expression of the α 2 β 1 integrin that mediates interaction with this protein [START_REF] Brodkin | Chondrocyte phenotype on different extracellular matrix monolayers[END_REF]. However, dental pulp cells are in need of fibronectin and vitronectin adsorption on this set of substrates to adhere via the α 5 β 1 or α v β 3 receptors [START_REF] Zhu | Integrin expression in human dental pulp cells and their role in cell attachment on extracellular matrix proteins[END_REF] which does not happen when silica is dispersed on the material surface due to conformational changes of fibronectin after adsorption on silica, reducing the availability of cell-binding domains [START_REF] Lord | The effect of silica nanoparticle coatings on serum protein adsorption and cellular response[END_REF].

Conclusions

Polymer/silica nanocomposites were prepared following a sol-gel process that leads to a material morphology of interpenetrated phases with controlled known concentration of silica both in the bulk and at the surface of the hybrid system. The incorporation of silica improves the mechanical properties of the polymer system and may influence the cellular response. Chondrocytes and dental pulp cells are both viable on the hybrid surfaces, but major differences are found between both cell types: chondrocytes are able to adhere and proliferate independently of the silica content of the surface, whereas dental pulp cells adhere only weakly and form clusters of cells that remain stable and viable up to 11 days.

Captions to Figures

Figure 1 .

 1 Figure 1. FTIR spectra of the different samples, in the 650-1850 cm -1 region.

  Figure 1. FTIR spectra of the different samples, in the 650-1850 cm -1 region.

Figure 2 .

 2 Figure 2. Phase light contrast microscopy images of human articular chondrocytes on the different substrates after 1 day of culture. (a) PEMA, (b) PHEA, (c) P(EMA-co-HEA) 70/30 %wt copolymer. Nanohybrids with increasing amount of SiO 2 (d) 5%, (e) 10%, (f) 15 %, (h) 20%, (i) 30%. Last image (j) represents control TCPS.

Figure 3 .

 3 Figure 3. Phase light contrast microscopy images of dental pulp cells on the different substrates after 2 day of culture. (a) PEMA, (b) PHEA, (c) P(EMA-co-HEA) 70/30 %wt copolymer. Nanohybrids with increasing amount of SiO 2 (d) 5%, (e) 10%, (f) 15 %, (h) 20%, (i) 30%. Last image (j) represents control TCPS.

Figure 4 .

 4 Figure 4. Phase light contrast microscopy images of pulp cells on the different substrates after 5 day of culture. (a) PEMA, (b) PHEA, (c) P(EMA-co-HEA) 70/30 %wt copolymer. Nanohybrids with increasing amount of SiO 2 (d) 5%, (e) 10%, (f) 15 %, (h) 20%, (i) 30%. Last image (j) represents control TCPS.

Figure 5 .

 5 Figure 5. Phase light contrast microscopy images of human dental pulp cells on the nanocomposites. (left) 10 culture days on TCPS after culture for 4 days on the substrate with 15 %wt SiO 2 . (right) Neutral red staining shows cell viability on H05 after 8 days.

Figure 2 100 µm 22 Figure 3 100 µm 23 Figure 4 100 µm 24 Figure 5

 2223234245 Figure 2

Table 1 .

 1 Average SiO 2 contents obtained by EDS, SiO 2 EDS , on the surface and in the bulk, water contact angles, θ w , and compressive elastic moduli, E.

	Sample	SiO 2 EDS (%) surface bulk	θ θ θ θ w (º)	E (GPa)
	PEMA	-	-	82.6±1.6	4.2±0.1
	PHEA	-	-	96.1±3.9	0.14±0.03
	H00	-	-	91.8±2.3	2.6±0.2
	H05	5.18±1.06	5.60±4.00	92.0±1.6	2.7±0.3
	H10	9.06±0.21	8.42±0.61	80.1±2.9	2.8±0.3
	H15	15.09±0.27 15.33±3.12	76.5±1.4	2.9±0.1
	H20	21.21±0.61 22.51±0.95	76.8±4.1	3.3±0.2
	H30	28.45±0.58 24.87±3.60	90.5±2.2	3.9±0.4

Acknowledgements

Financial support of the Spanish government through project MAT2009-14440-C02-01 is kindly acknowledged. CIBER-BBN is an initiative funded by the VI National R&D&i Plan 2008-2011, Iniciativa Ingenio 2010, Consolider Program, CIBER Actions and financed by the Instituto de Salud Carlos III with assistance from the European Regional Development Fund. This work was supported by funds for research in the field of Regenerative Medicine through the collaboration agreement from the Conselleria de Sanidad (Generalitat Valenciana), and the Instituto de Salud Carlos III.


M Lebourg

J Suay Antón

J L Gomez

J L Gomez Ribelles

Characterisation of calcium phosphate layers grown on polycaprolactone for tissue engineering purposes

Keywords: Polymer(A), Ceramics(A), Surface Treatment(B), X-ray diffraction(D), Fourier Transform Infrared Spectroscopy(D)

published or not. The documents may come

Characterisation of calcium phosphate layers grown on polycaprolactone for tissue engineering purposes

Introduction

Polymer-ceramic composites are of interest in the field of bone tissue engineering [START_REF] Wang | Developing bioactive composite materials for tissue replacement[END_REF]: traditional polyesters used in biomedical field, as polylactide or polycaprolactone, are easily reabsorbed, show ductile properties, and can be easily processed to porous bodies as those used in tissue engineering, but they lack intrinsic bioactivity, thus not favouring any repair response from the body. On the other hand, ceramics from the calcium phosphate family, for example tricalcium phosphate and hydroxyapatites, have shown to induce a good response from bony cells. In fact they are currently in use in orthopaedics, though always with the drawback of fragility, scarce remodelling [START_REF] Van Landuyt | Influence of high sintering temperature on the mechanical properties of Hydroxyapatite[END_REF], and a difficulty in cutting to fill the defect shape. A combination of both material types thus reduces their drawbacks while benefiting from their respective advantages [START_REF] Hasegawa | In vivo evaluation of a porous hydroxyapatite/polyDL-lactide composite for bone tissue engineering[END_REF][START_REF] Shor | Fabrication of three-dimensional polycaprolactone/ hydroxyapatite tissue scaffolds and osteoblast scaffold interactions in vitro[END_REF][START_REF] Kenny | Bone cements and fillers: a review J of Mat[END_REF].

In recent years, much attention has been paid to the in vitro mineralisation of materials in protophysiological solutions (called simulated body fluid, hereafter mentioned as SBF) as an important step for demonstrating corresponding in vivo bone bonding.

The deposited calcium phosphate layer is described in literature mostly as "biomimetic apatite" for its composition is similar to that of bone mineral, that is to say a carbonated, calcium deficient and poorly crystalline apatite. Such nanostructured calcium phosphates, due to their high specific surface and soft preparation methods, are thought to be useful in tissue engineering as in vitro cell culture substrates for the study of cellmaterial interactions, or as carriers for proteins [START_REF] Azevedo | Incorporation of proteins and enzymes at different stages of the preparation of calcium phosphate coatings on a degradable substrate by a biomimetic methodology[END_REF] or drugs (like biphosphonates or strontium containing drugs) while being easily degraded in the body if implanted. Based on the assumption that such a "bone-like mineral" would show good bioactive properties for the culture of cells, many studies have been carried out to prepare composites by precipitating mineral on different substrates to study the response of cells from bone lineage (osteoblasts, SAOS cells, mesenchymal cells, etc) to such mineral, with variable results [START_REF] Chen | PLLA scaffolds with biomimetic apatite coating and biomimetic apatite/collagen coating to enhance osteoblast-like cells attachment and activity[END_REF][START_REF] Ciapetti | Osteoblast growth and function in porous poly( -caprolactone) matrices for bone repair: a preliminary study[END_REF][START_REF] Murphy | Effect of a bone-like mineral film on phenotype of adult human mesenchymal stem cells in vitro[END_REF][START_REF] Chim | A comparative analysis of scaffold material modifications for load bearing applications in bone tissue engineering[END_REF][START_REF] Lickorish | Collagen hydroxyapatite composite prepared by biomimetic process[END_REF][START_REF] Kim | Stimulation of osteoblast responses to biomimetic nanocomposites of gelatin-hydroxyapatite for tissue engineering scaffolds[END_REF][START_REF] Kim | Accelerated bonelike apatite growth on porous Polymer/ceramic composite scaffolds in vitro[END_REF]. Nevertheless, many papers do not characterise properly the mineral layer deposited, and as has been described among others by Chou et al. [START_REF] Chou | Effect of Hydroxylapatite coating crystallinity on biosolubility, cell attachment efficiency and proliferation in vitro[END_REF], the composition, Ca/P ratio, present phases, crystallinity [START_REF] Nagano | Differences of bone bonding ability and degradation behaviour in vivo between amorphous calcium phosphate and highly crystalline hydroxyapatite coating[END_REF] and biosolubility of the mineral layer is of outermost importance with respect to the cell response, since cells show exquisite sensitivity to the smallest changes in roughness, chemistry, surface tension, crystal structure or culture medium composition [START_REF] Chou | The effect of biomimetic apatite structure on osteoblast viability, proliferation, and gene expresión[END_REF]. In this paper, we describe how we obtained two differentiated calcium phosphate layers and their characterisation: we also present initial results of osteoblast-like cell culture on them.

Materials and Methods

Sample preparation

PCL (Polysciences, Mw=43000-50000 Da) and 1-4 dioxane (Scharlab 98% pure) were used without further purification. PCL bulk samples were prepared by solvent casting from a PCL solution in dioxane (15% by weight) in Petri dishes. PCL with hydroxyapatite nanoparticles used in cell culture (Hap particles, d<100nm, Sigma Aldrich) were prepared by solvent casting of PCL solution (15% by weight in dioxane) with 20% of Hap (dispersed with ultrasound); these samples are referred thereafter as PCL-HAp-P samples. After solvent evaporation, films were rinsed in ethanol and vacuum dried to constant weight. Then 12 mm-diameter disks were stamped out of the films and treated for apatite nucleation as described in [START_REF] Oyane | Simple surface modification of poly( -caprolactone) to induce its apatite-forming ability[END_REF]. Shortly after this, the samples were first treated with air plasma treatment, 300W power, 90 s on each side in an Electronic Piccolo plasma chamber (Plasma Electronic GmbH, Germany). PCL and PCL HAp P samples were submitted to plasma treatment just before cell culture as effect of plasma on surface tension decreases with time. Then, samples were immersed for ten seconds in a CaCl 2 alcoholic solution (0.1mM, water/alcohol, w/a, 50/50), washed for 1 second in w/a 50/50 mixture, then immersed for 10 s in a K 2 HPO 4 alcoholic solution (0.1mM, w/a=50/50), and washed in water-alcohol mixture. This process was repeated 5 times. Subsequently, samples were freely suspended in modified SBFs, prepared following the methodology of Müller et al [START_REF] Müller | Preparation of SBF with different HCO 3 content and its influence on the composition of biomimetic apatites[END_REF] out of concentrated solutions of KCl, NaCl, NaHCO 1. The main difference with Müller protocol was the concentration of sodium azide used as an antibacterial agent (when not prepared in totally aseptic conditions, SBF is likely to be colonised by phosphate eater bacteria [START_REF] Planell | Growth of bioactive surfaces on titanium and its alloys for orthoepaedic and dental implants[END_REF]) , this concentration was lowered from 1g/l to 10mg/l as sodium azide is very toxic and 10mg/l is sufficient for antibacterial activity: as a result, the sodium content in the prepared SBF was slightly lower than in traditional SBF, but it is hypothesized that since sodium is not a constituent of hydroxyapatite it does not modify its solubility product and thus does not influence its precipitation from the solution. Previous experiments (results not shown) showed that these SBFs are able to produce an apatite layer as do original SBF.

The SBF-A, (for amorphous), is a carbonate rich SBF prepared following the method by Müller and Müller and contains 15mM HCO 3 2-per litre. It was buffered at pH=7,4 at 37ºC. It is expected to produce small-crystallite, low crystalline apatite deposit.

The SBF-C, for crystalline, is mainly a SBFx2 (with concentration of electrolytes doubled) containing neither magnesium nor carbonate, since these ions are known to act as crystal growth inhibitors, and buffered at pH=7, (so as to lower the probability of spontaneous nucleation in the solution). It is expected to produce large-crystallite, highly crystalline apatite deposit.

SBF was changed every 4 days and SBF immersion lasted 14 days. After SBF treatment, the samples were washed with distilled water and dried (except samples for cell culture). During washing, leaching of some mineral particles was observed.

SEM observation

Samples were cut and stuck to a metallic support using graphite adhesive tape, so that a piece of each face of the sample could be observed. Samples were gold sputtered and observed in a Hitachi S3000 electron microscope with acceleration tension of 20 kV and work distance of 15cm. EDX analysis was performed and Ca/P values were calculated from at least three large field analysis measurements.

FTIR characterization

The surface of the samples was scraped and the resulting powder was mixed with KBr in relation 1:1000 and compacted to a disk. FTIR analysis was performed in a FT-IR Bruker IFS60v analyzer provided with MCT detector in transmission mode from 550 to 4000 cm-1 with a sensitivity of 4cm -1 . For evaluation of peaks, data was analyzed using the review on HAp characterization by Koutsopoulos [START_REF] Koutsopoulos | Synthesis and characterization of hydroxyapatite crystals: a review study on the analytical methods[END_REF].

X-Ray Diffraction

XRD analysis was carried out using a Siemens D-5000 diffractometer with /2 geometry equipped with parallel beam device for grazing incidence measurements and secondary monochromator. Measurements were performed in grazing incidence mode using monochromatic Cu (K ) radiation, with incidence angle 1º in order to observe mainly the mineral deposit, and scanning 2 between 10 and 100 degrees with increment steps of 0,04. For evaluation of peaks, data were compared with data from respective ICDD card files 026-1056(OCP) and 009-0432 (HAp).

Cell culture

Materials were sterilized with cold ethanol during two hours before cell culture. Before seeding, samples of pure PCL, PCL Hap-P, PCL-HAp-A and PCL-HAp-C (and glass coverslips as a control) were washed with PBS and coated with fibronectin solution in PBS (fibronectin from human plasma, Sigma, 20 g/ml) overnight. MT3C3 osteoblastic cells (Riken Cell Bank, Japan) at passage 20 were seeded at a density of ~9x10 3 cells/cm2 in 400 l standard growth medium (DMEM-LG supplemented with 1%

Penicillin and Streptomycin, 1% glutamine), and allowed to adhere for 3h. At the end of the culture time, samples were washed twice with PBS, blocked with 3.7% paraformaldehyde (Sigma-Aldrich) during 60 min at 4ºC and then stored in PBS at 4ºC until biological characterization was performed.

Biological characterization

Fluorescence microscopy: The actin cytoskeleton was visualized using the green-fluorescent Bodipy FL Phalloidin (Invitrogen). After staining according to standard protocols, samples were glued on microscope slides with mounting medium containing DAPI (Vectashield, ATOM) for visualization of nuclei and stored in the dark at 4 ºC until fluorescence microscopy was performed. Samples were observed in an inverted vertical microscope Axiovert200 (Zeiss) coupled to a CCD monochrome and color camera.

SEM: Samples were dehydrated in graded alcoholic solutions and air dried, then stuck to a metallic support using graphite adhesive tape. Samples were gold sputtered and observed in a JEOL JSM6300 electron microscope with acceleration tension of 15 kV and work distance of 15cm.

Quantitative cell number analysis: In order to gain quantitative information on initial cell adhesion and seeding efficiency on the different surfaces, image analysis was performed using cell counting software [START_REF] Selinummi | Software for quantification of labeled bacteria from digital microscope images by automated image analysis[END_REF]: for each sample type at least five pictures of DAPI stained nuclei at magnification 5x were recorded from two different samples and analyzed with CellC.

Results

Deposition of mineral layer

Preliminary results showed that subjecting the samples to an air plasma treatment followed by nucleation treatment by immersion in CaCl 2 and K 2 HPO 4 solutions was effective in inducing the formation of an apatite-like layer. It is hypothesized that plasma treatment provokes the scission of ester bonds, thus multiplying the density of carboxylate moieties on the surface; those are known to be involved in the nucleation of calcium phosphates in simulated body fluid. Nevertheless, the effect of plasma alone is not decisive, and only slightly accelerates calcium phosphate nucleation on the surface of PCL. Alternative soaking in calcium and phosphate containing solutions after activation of the surface by plasma treatment or etching with sodium hydroxide has been shown to be effective in inducing apatite deposition. As described in [START_REF] Vallés Lluch | Biomimetic apatite coating on P(EMA-co-HEA)/SiO2 hybrid nanocomposites[END_REF] on the thermodynamical point of view, nucleation is the limiting step in apatite formation rate.

Once nucleated, the apatite grows without induction from simulated body fluid.

Morphology of mineral layer

The morphology of the mineral layers deposited can be observed in Figure 1.

The surface of the apatite layer deposited by immersion in SBF-A medium, that we will call hereafter HAp-A, is formed by globular structure close to the typical cauliflower usually described for hydroxyapatite, and relatively smooth. At the contrary the surface of HAp-C (deposited in SBF-C) is formed by large plate-like crystals (width ~2µm) that resemble the morphology of gypsum flowers and show considerable roughness at small scale; this kind of plate-like crystals has been described for OCP [START_REF] Mandel | Brushite (CaHPO4•2H2O) to octacalcium phosphate (Ca8(HPO4)2(PO4)4•5H2O) transformation in DMEM solutions at[END_REF] in the literature, and also in biomimetic apatite samples where it was inherited from crystallization of OCP as a primary phase that later on evolved towards hydroxyapatite, (observed by Müller et al. in OCP samples exposed to physiological conditions [START_REF] Horváthová | In vitro transformation of OCP into carbonated HA under physiological conditions[END_REF]). Layer thickness was determined on zones of the sample where the layer was peeled off from PCL surface and the profile appeared (Figure 2): it shown to be thicker for HAp-C (~12µm) than for HAp-A (~6µm) as expected from the respective ionic activities of the SBF solutions. As can be seen on profile for Hap-C, crystals grow vertically normal to the sample surface which should be related with preferred crystal orientation and was described by Müller in OCP samples [START_REF] Horváthová | In vitro transformation of OCP into carbonated HA under physiological conditions[END_REF]. Both sides of the samples had similar crystal microscopic morphology (see insets in Figure 1), whereas macroscopic morphology was a little different in roughness and surface profile which is likely to be due to different morphologies of the naked surfaces of PCL.

EDX analysis reveals that Ca/P ratio of HAp-A is higher [START_REF] Wang | Developing bioactive composite materials for tissue replacement[END_REF]65) than that of HAp-C

(1,35), likely due to carbonate substitution for phosphate (B-type substituted apatite) that lowers the proportional phosphate content in the mineral layer; the ratio of Hap-C is very close to Ca/P ratio of octacalciumphosphate [START_REF] Wang | Developing bioactive composite materials for tissue replacement[END_REF][START_REF] Taguchi | Apatite coating on hydrophilic polymer-grafted poly(ethylene) films using an alternate soaking process[END_REF], being likely OCP the main phase observed in HAp-C. Some substitution elements appear in Hap-A samples (Mg, Na to a lower extent) as usually seen when calcium phosphate is deposited from physiological solutions.

Vibrational characterisation of layers

The FTIR spectrum of the deposited layers is shown in Figure 3 in the 550-2000 cm -1 wavenumber interval. This is the most representative zone, since the 4000-2000 cm -1 zone only shows the large peak associated with hydrogen bonding and eventually water adsorbed on the samples and it does not differ significantly between the two samples.

The choice of measuring the signal associated to the mineral layer only (by measuring the crushed coating after peeling it off from the substrate, instead of a reflectance measurement of the whole sample) gives good results and produces a reduction in the signal from the subjacent PCL. The peaks in the zone 1350-1550 cm -1 are associated with carbonate moieties (bending modes 3 and 4, stretching mode 3). Peaks between 1000-1100 cm -1 are due to the triply degenerated asymmetric stretching mode of P-O bond in phosphate group. Peaks between 550 and 650cm-1 are due to the triply degenerated bending mode of O-P-O bonds of the phosphate group. As can be easily seen in the spectrum, the signal associated with carbonate is higher in HAp-A relative to HAp-C, whereas the signal associated with phosphate groups is higher in HAp-C. Thus not only the carbonate content is higher in HAp-A but the phosphate content is lower, an indication that carbonate is substituting for PO 4 3-in the lattice, as described for apatite deposited from SBF for carbonate concentrations lower than 20mM [START_REF] Müller | Biomimetic apatite coatings-carbonate substitution and preferred growth orientation[END_REF]. A further clue indicating that the B-substitution is predominant is the absence of a peak at 1550cm -1 that appears when CO 3 2-stands for OH (A-type substitution). OH typical signal is not observed at 630 and 3570 cm -1 ; this is frequently the case when analysing calcium phosphates deposited from SBF and also the case in apatite from bone that was not calcined [START_REF] Pasteris | Lack of OH in nanocrystalline apatite as a function of degree of atomic order: implications for bone and materials[END_REF]. Peak at 960 cm- It can also be observed that in general the peaks from HAp-A are smoother and wider than the peaks of HAp-C, which is indicative of its amorphous nature (or very thin crystallite structure).

The spectrum obtained by X-ray diffraction is presented in Figure 4. The spectrum of pure PCL is shown below as a comparison, since PCL is semi-crystalline, and PCLassociated signal appears and must be differentiated from the signal associated with the mineral layer. PCL background spectrum mainly consists of two intense peaks at 2 =21.6 and 2 =23.8 (inset in Figure 4), that account for diffraction on the [110] and

[200] planes respectively (PCL has a crystalline structure with polyethylene-like orthorhombic cell disposition, with lattice parameters a=0.748nm, b=0.498nm and c=1.727nm) [START_REF] Bittiger | Crystal structure of polycaprolactone[END_REF]. Here again, the signal of HAp-C speaks for a more crystalline layer than that of HAp-A: the peaks are more defined, and their shape is more acute (in both cases the small size of the crystallites prevents from obtaining high peaks as usual when considering high-temperature sintered apatite as a matter of fact c-dimension of HAp lattice is 0,688 nm whereas in OCP c= 0,6855 nm. The peak that is generally used to differentiate both phases, and corresponds to reflection on [010] plane of OCP, at 2 =4.72, is out of the measure range in this case.

Nevertheless other peaks may match the corresponding reflections in range from 25 to 50 degrees (see figure 4). As discussed later, in physiological solutions and biological environments HAp and OCP phases often coexist. The peaks of HAp-A cannot be precisely defined, due to generalized broadening related with the lattice distortion due to carbonate substitution but they correspond generally to those of HAp-C.

Cell observation

As can be seen on Figure 5 and 6, cells were able to adhere and spread on all supports tested. Actin cytosqueleton is more stretched on PCL and PCL-HAp-P surfaces, cells appear wider. On calcium phosphate surfaces, cell edges are sharper and irregular, and cell spreading is limited due to the roughness of the apatite surface, much higher than that of PCL bulk surface; as a result cell size is generally tinier. In terms of adhesion efficiency, PCL and PCL-HAp show low adhesion efficiency at 3h in the culture conditions (11% and 15% respectively when compared to initial seeding density) and cells were not evenly distributed on the samples. Calcium phosphate surfaces show a good cell adhesion, comparable to that of glass coverslip controls (Figure 7). There is a significative difference (p<0.05, Student t-test) between cell number observed on HAp-A and HAp-C, with more cells on HAp-C.

Discussion

Deposition of calcium phosphate phase on biomaterials

Deposition of apatite on materials when immersed in simulated body fluid has been described as a guarantee for their bioactivity in vivo. SBF is an aqueous solution that mimics the electrolyte content of blood plasma. The mechanism proposed for the deposition of a precursor phase of amorphous calcium phosphate that later evolves towards apatite involves the negative surface potential of materials [START_REF] Bodhak | Role of surface charge and wettability on early stage mineralization and bone cell-materials interactions on polarized Hydroxyapatite[END_REF], normally due to deprotonation of functional groups such as -PO 4 H 2 [START_REF] Oliveira | Nature-inspired calcium phosphate coatings: present status and novel advances in the science of mimicry[END_REF],-TiOH, -SiOH, -COOH [START_REF] Kawashita | Apatite forming ability of carboxyl containing polymer gels in a simulated body fluid[END_REF], -OH at physiological pH (~7,4) (this was demonstrated for hydroxyapatite [START_REF] Kim | Process and kinetics of bonelike apatite formation on sintered hydroxyapatite in a SBF[END_REF], and the process is the same for many biomaterials). Negative surface potential has been related both with HAp deposition from SBF -whereas positive surface potential leads to preferential deposition of sodium chloride [START_REF] Bodhak | Role of surface charge and wettability on early stage mineralization and bone cell-materials interactions on polarized Hydroxyapatite[END_REF]-and good osteoblast behaviour. This negative surface potential attracts calcium ions and a calcium-rich calcium phosphate layer is formed. Then due to an excess of calcium ions the surface potential turns positive, and phosphate ions are now attracted in majority. This process is repeated and a layer is formed: this amorphous calcium phosphate layer, highly hydrated, is then remodelled to hydroxyapatite. Here, the plasma treatment performed is meant to generate carboxyl and hydroxyl moieties by scission of the esters of the PCL chain. As a matter of fact a strong decrease of the water contact angle of the PCL surface after plasma treatment (from ~72 to ~34º, data not shown) show that polar groups have been generated on the surface. In previous experiments, the plasma treatment was seen to enhance the deposition rate of apatite, but to a lower amount than plasma treatment and nucleation. Nucleation by seeds of calcium phosphate was proposed first by Taguchi [START_REF] Taguchi | Apatite coating on hydrophilic polymer-grafted poly(ethylene) films using an alternate soaking process[END_REF], and later by Oyane [START_REF] Oyane | Simple surface modification of poly( -caprolactone) to induce its apatite-forming ability[END_REF] as a mean to effectively accelerate apatite deposition: SBF is saturated towards apatite, and calcium phosphate seeds provide sites for heterogeneous nucleation, consequently the formation of solid phases is much easier. This is also the result we obtained in previous experiments where we also observed dramatically accelerated apatite formation rates on PCL scaffolds (reinforced or not with synthetic hydroxyapatite nanoparticles) [START_REF] Lebourg | Hybrid structure in PCL-HAp scaffold resulting from biomimetic apatite growth[END_REF] in traditional SBF after nucleation, as well as in previous experiments with this modification of SBF prepared with the Müller method; nevertheless it is important to mention that the decrease in sodium content associated with our modification of the Müller method seemed to cause a slower mineralisation than traditional SBF, which is likely to be due to the lower ionic strength of the solution. However, the lower sodium concentration did not impede the proper precipitation of the apatite layer, it being an adequate solution for performing the Kokubo-test (if not comparing the results obtained with the results obtained with traditional SBF) or for covering the materials with a mineral layer. From a practical point of view, the method from Müller has unbeatable advantages such as simplicity, easiness of use and reproducibility. Matching the traditional SBF concentrations but using the methodology of the stock solutions is also possible (also mentioned by Lemaitre et al. in its useful critic of Kokubo test [START_REF] Bohner | Can bioactivity be tested in vitro with SBF solution?[END_REF]).

Nature of the mineral layer

It is known from the literature that the composition, concentration [START_REF] Kim | Composition and structure of the apatite formed on PET substrates in SBF modified with various ionic activity products[END_REF] and pH of SBF greatly influences the nature of the apatite layer deposited. A more basic pH, due to OH being one of the components of hydroxyapatite, favours the precipitation of apatite raising the supersaturation in the solution [START_REF] Planell | Growth of bioactive surfaces on titanium and its alloys for orthoepaedic and dental implants[END_REF]. Concentrated SBF, such as 1,5xSBF, 2xSBF, has been shown to result in a more crystalline, less carbonated structure [START_REF] Stoch | FTIR absorptionreflection study of biomimetic growth of phosphates on titanium implants[END_REF].

Müller et al. have shown that carbonate content in SBF not only influences the carbonate content in the deposited layer [START_REF] Müller | Preparation of SBF with different HCO 3 content and its influence on the composition of biomimetic apatites[END_REF], but also the preferred orientation growth of the crystals as well as the main crystallite dimension [START_REF] Müller | Biomimetic apatite coatings-carbonate substitution and preferred growth orientation[END_REF]. As a matter of fact, carbonate as well as magnesium has been shown to act as crystal growth inhibitor, in the case of the carbonate this may be due to the inactivation of crystal growth sites by the adsorption of these divalent cations in place of Ca 2+ [START_REF] Johnsson | The role of brushite and octacalcium phosphate in apatite formation[END_REF]. In this work, the influence of these ions was predominant for tuning the crystallinity and crystal size, as observed by electron microscopy, FTIR and XRD. The crystal size and crystallinity of the mineral deposited has a significant influence on biosolubility [START_REF] Rey | Physico-chemical properties of nanocrystalline apatites: Implications for biominerals and biomaterials[END_REF], as well as on the topography and likely protein adsorption on the surface and has thus a decisive influence on the cell behaviour [START_REF] Okada | Nucleation of biomimetic apatite in synthetic body fluids: dense and porous scaffold development[END_REF].

From FTIR and XRD results the deposited phase in HAp-A appears to be a carbonated hydroxyapatite with low crystallinity due to carbonate content and high carbonate substitution. FTIR results lead us to believe that the deposited apatite is B-Substituted with carbonate standing for phosphate. This is more likely when phosphate appears as hydrogen phosphate due to the equality of charges carried by both cations. As a result of the equilibrium between phosphate and hydrogen phosphate at physiological pH, a part of the phosphates is often replaced by hydrogen phosphates, broadening the triply degenerated band for asymmetric stretching mode of P-O at 1000-1100 cm -1 and lowering the Ca/P value [START_REF] Koutsopoulos | Synthesis and characterization of hydroxyapatite crystals: a review study on the analytical methods[END_REF]41]. In the case of HAp C, the strong signal associated with hydrogenphosphate and higher peak definition and crystallinity leads to think that hydrogenphosphate may not be present as a substitution but as a stoechiometric component of octacalciumphosphate, an acidic calcium phosphate, which has been described as an apatite precursor in vitro [START_REF] Dorozhkin | Calcium Orthophosphates in Nature, Biology and Medicine[END_REF]43]. Although XRD data is not totally conclusive (only peak at 4,72 is only seen for OCP), the Ca/P ratio observed, abundance of hydrogenophosphates and crystal morphology observed speak for OCP being the main phase deposited. In physiological solution, although apatite is the thermodynamically most stable phase, OCP is often formed due to faster kinetics of deposition [START_REF] Horváthová | In vitro transformation of OCP into carbonated HA under physiological conditions[END_REF]. Later transformation to HAp is driven by the metastable character of OCP , and is favoured by the very low interfacial tension between both components [START_REF] Liu | Intergrowth of calcium phosphates: an interfacial energy approach[END_REF].

The mechanism proposed for OCP hydrolysis to HAp is a topoaxial conversion by ion diffusion within the crystal lattice, which would explain the conservation of macroscopic crystal structure upon formation of HAp from OCP [START_REF] Müller | Biomimetic apatite coatings-carbonate substitution and preferred growth orientation[END_REF][START_REF] Suzuki | Review: Octacalcium phosphate: Osteoconductivity and cristal chemistry[END_REF]. This hydrolysis process has been described both in vitro [START_REF] Horváthová | In vitro transformation of OCP into carbonated HA under physiological conditions[END_REF] and in vivo [START_REF] Suzuki | Surface Chemistry and Biological Responses to Synthetic Octacalcium Phosphate[END_REF] with synthetic OCP; some researchers relate the osteoconductive potential of OCP to its transient character, considering its conversion to HAp a key factor for bone regeneration, osteoblast and osteoclast activation [START_REF] Suzuki | Review: Octacalcium phosphate: Osteoconductivity and cristal chemistry[END_REF]. Nevertheless it is not clear that OCP appears as an endogen apatite precursor during the mineralization of human tissues [START_REF] Grynpas | Transient precursor strategy or very small biological apatite crystals?[END_REF]. Here the clues that point towards presence of both OCP and HAp phases, could be explained by the formation of OCP and partial hydrolysis of OCP phase to HAp: it would explain the low Ca/P ratio observed [START_REF] Wang | Developing bioactive composite materials for tissue replacement[END_REF][START_REF] Bohner | Can bioactivity be tested in vitro with SBF solution?[END_REF] for HAp-C: partially hydrolyzed OCP samples from Suzuki et al. had similar ratios ranging from 1,33 to 1,49 [START_REF] Suzuki | Review: Octacalcium phosphate: Osteoconductivity and cristal chemistry[END_REF][START_REF] Suzuki | Surface Chemistry and Biological Responses to Synthetic Octacalcium Phosphate[END_REF]. Although OCP has a stoechiometric Ca/P ratio of 1,33 in the aforementioned papers synthetic OCP had a ratio before hydrolysis of 1,26 and this ratio raised upon hydrolysis. High intensity ratios of the peaks [002]and [004] seen in the diffractogramm lead to think that crystal growth along c-axis is favoured.

Cell behaviour

Apatite morphology is decisive for cell behaviour, and whereas synthetic hydroxyapatite and calcium phosphates have been studied from the biological point of view, there is still a lack of information about cell response to different biomimetic apatites; the available studies show that behaviour is altered by very slight changes in morphology and chemistry and a thin analysis of the layer formed is thus of uttermost importance. Hydroxyapatite as a bulk sintered material has been shown be osteoinductive in vivo [START_REF] Klein | Osseous substance formation induced in porous calcium phosphate ceramics in soft tissues[END_REF][START_REF] Yamasaki | Osteogenic response to porous Hydroxyapatite ceramics under the skin of dogs[END_REF][START_REF] Ripamonti | Osteoinduction in porous hydroxyapatite implanted in heterotopic sites of different animal models[END_REF] and to induce stem cell differentiation towards the osteogenic pathway in vitro, and the response of osteoblasts to different mimetic apatites has also been described. A toxic effect of very amorphous calcium phosphate has been described [START_REF] Oreffo | In vitro and in vivo methods to determine the interactions of osteogenic cells with biomaterials[END_REF] likely due to dissolution of both calcium and phosphate that causes apoptosis [START_REF] Adams | Matrix Regulation of Skeletal Cell Apoptosis: Role of calcium and phosphate ions[END_REF], but there is no much data on the effect of different types of hydroxyapatite on bony cells. Our first and preliminary results about osteoblast-like cells behaviour on the engineered apatite layers with different topographies, chemistry and crystallinity show that both types of surfaces support cell adhesion and short term viability of osteoblast-like stem cells. In this first set of experiments, morphologic differences have been seen between cells adhered on PCL or PCL-HAp-P surfaces on one side, and HAp layers on the other. A low adhesion on PCL surfaces was seen, maybe due to poor protein adsorption (very smooth surface) or due to cytotoxic effect of plasma treatment; this should be further explored since in previous experiments without plasma treatments PCL and PCL-HAp P supports elicited cell adhesion and proliferation. Morphology on calcium phosphate surfaces was very different; cells were not wide spread, appeared tinier and grained actin was observed inside of the cells. This was also described in the work of Okada [START_REF] Okada | Nucleation of biomimetic apatite in synthetic body fluids: dense and porous scaffold development[END_REF]. Recent cell research shows that cell morphology and cell cytosqueleton organization in consequence of cell-material interaction and focal adhesion organization is closely related to cell fate [START_REF] Demali | Integrin signaling to the actin cytoskeleton Current Opinion in[END_REF], including phenotypic expression, intracellular metabolic and catabolic pathways signalization, migration behaviour and so on [START_REF] Lock | Cell-matrix adhesion complexes: Master control machinery of cell migration[END_REF]. Significant difference in cell density at 3h after seeding was observed between both calcium phosphates, so at longer times behaviour should arise many more significant differences. 1 We are looking forward to further characterize these coatings from the biological point of view in order to gain fundamental knowledge about cell-material interaction, which will help us design better composite biomaterials for efficient bone tissue engineering; in particular, study of protein conformation on substrate with different morphology and crystal chemistry should gain in understanding cell behaviour. Although this type of coating may not be used in load bearing applications due to its tendency to crack and peel off, it could be used for the entrapment of drugs and proteins or as a laboratory tool for studying cellmaterial interactions.

Conclusions

An apatite-like mineral layer can be deposited on the surface of polycaprolactone substrates by immersing them in simulated body fluids. SBF preparation from stock solutions proved to be more efficient, user-friendly and repetitive than direct preparation from salt weighing. Previous activation of the PCL surface by plasma treatment followed by nucleation by alternative immersion in CaCl

  2 and K 2 HPO 4 solutions proved itself to be very effective in accelerating apatite deposition. The composition and morphology of the apatite layer produced can be easily tailored by modifying the concentration of the different ions in SBF medium; the immersion of nucleated PCL sheets in a carbonate rich SBF medium produces a layer of low-crystallinity apatite with the classical cauliflower morphology. The absence of magnesium and carbonate ions increases crystallinity, and thickness of the apatite layer changing the crystal morphology and decreasing the Ca/P ratio to values of 1.35; it is likely octacalcium that has been partially hydrolyzed to HAp. Preliminary cell culture results show that cells can adhere to the HAp-coated materials independently of their composition; nevertheless the cells show different morphologies on the different supports and adhesion is influenced by topography and chemistry, being adhesion most efficient on HAp-C; effects related with these changes in morphology and further influence will give clues on how to produce the best environment for bone tissue engineering.

Table 1 .Figure 1 .Figure 2 .Figure 3 .

 1123 Figure 1. Morphology of the deposited calcium phosphate layers Figure 2. Profile of the layers: A-HAp, left (bar size 10 m) and C-HAp, right (bar size 30 m) Figure 3. FTIR spectrum of deposited layers: amorphous apatite (A-HAp) and crystalline apatite (C-HAp) in the range 550-2000 cm 1 . Peaks marked with stars are due to PCL residue. Figure 4. X-ray Diffraction spectrum of PCL and HAp covered PCL samples with corresponding Miller indices of HAp (black) and OCP (gray) Figure 5. MT3C3 cells adhered on PCL, PCL with HAp nanoparticles (PCL HAp P), and calcium phosphate surfaces (bar size up 30 m, bottom 60 m) Figure 6. Cell morphology on the different supports as observed by fluorescence microscopy (green actin filaments-blue nuclei, bar size 50 m) Figure 7. Cell density on the different surfaces as determined by image analysis

Figure 4 .Figure 5 .Figure 6 .Figure 7 .

 4567 Figure 3. FTIR spectrum of deposited layers: amorphous apatite (A-HAp) and crystalline apatite (C-HAp) in the range 550-2000 cm 1 . Peaks marked with stars are due to PCL residue. Figure 4. X-ray Diffraction spectrum of PCL and HAp covered PCL samples with corresponding Miller indices of HAp (black) and OCP (gray) Figure 5. MT3C3 cells adhered on PCL, PCL with HAp nanoparticles (PCL HAp P), and calcium phosphate surfaces (bar size up 30 m, bottom 60 m) Figure 6. Cell morphology on the different supports as observed by fluorescence microscopy (green actin filaments-blue nuclei, bar size 50 m) Figure 7. Cell density on the different surfaces as determined by image analysis

  3 , MgSO 4 -7H 2 O, CaCl 2 , Tris-HCl buffer, NaN 3 , KH 2 PO 4 . The final electrolyte concentrations in SBFs are presented in Table

Acknowledgements

Support of the Spanish Ministry of Science through project No. MAT2007-66759-C03-01, including the FEDER financial support, with complementary funding of Generalitat Valenciana with project ACOMP/2009/112 and Universidad Politécnica de Valencia with 2911-2008 project is acknowledged. JLGR and JSA acknowledge funding from the Centro de Investigación Principe Felipe in the field of Regenerative Medicine through the collaboration agreement of the Conselleria de Sanidad (Generalitat Valenciana), and the Instituto de Salud Carlos III (Ministry of Science and Innovation). Myriam Lebourg thanks Patricia Rico and Cristina Martinez for help and advice with culture experiments and CIBER-BBN for funding. CIBER-BBN is an initiative funded by the VI National R&D&i Plan 2008-2011, Iniciativa Ingenio 2010, Consolider Program, CIBER Actions and financed by the Instituto de Salud Carlos III with assistance from the European Regional Development Fund.