Polysaccharide-Coated Thermosets for Orthopedic Applications: From

Apr 17, 2012 - Rescoll, Société de Recherche, 8 Allée Geoffroy Saint-Hilaire, 33615 Pessac, France. §. Orthopaedic Research Unit, Department of ...
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Polysaccharide-Coated Thermosets for Orthopedic Applications: From Material Characterization to In Vivo Tests Andrea Travan,*,† Eleonora Marsich,† Ivan Donati,† Marie-Pierre Foulc,‡ Niko Moritz,§ Hannu T. Aro,§ and Sergio Paoletti† †

Department of Life Sciences, University of Trieste, Via Giorgieri 1, Trieste I-34127, Italy Rescoll, Société de Recherche, 8 Allée Geoffroy Saint-Hilaire, 33615 Pessac, France § Orthopaedic Research Unit, Department of Orthopaedic Surgery and Traumatology, University of Turku, Kiinamyllynkatu 10, 20520, Turku, Finland ‡

S Supporting Information *

ABSTRACT: The long-term stability and success of orthopedic implants depend on the osseointegration process, which is strongly influenced by the biomaterial surface. A promising approach to enhance implant integration involves the modification of the surface of the implant by means of polymers that mimic the natural components of the extracellular matrix, for example, polysaccharides. In this study, methacrylate thermosets (bisphenol A glycidylmethacrylate/triethyleneglycol dimethacrylate), a widely used composition for orthopedic and dental applications, have been coated by electrostatic deposition of a bioactive chitosan-derivative. This polysaccharide was shown to induce osteoblasts aggregation in vitro, to stimulate cell proliferation and to enhance alkaline phosphatase activity. The coating deposition was studied by analyzing the effect of pH and ionic strength on the grafting of the polysaccharide. Contact angle studies show that the functionalized material displays a higher hydrophilic character owing to the increase of surface polar groups. The mechanical properties of the coating were evaluated by nanoindentation studies which point to higher values of indentation hardness and modulus (E) of the polysaccharide surface layer, while the influence of cyclic stress on the construct was assessed by fatigue tests. Finally, in vivo tests in minipigs showed that the polysaccharide-based implant showed a good biocompatibility and an ability for osseointegration at least similar to that of the titanium Ti6Al4V alloy with roughened surface.

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INTRODUCTION The development of tissue engineering and regenerative medicine techniques has introduced the new concept of “bioactive” biomaterial, able to induce specific biological responses by engaging interactions with the surrounding living tissue at the molecular level.1,2 In particular, for applications in orthopedics and dentistry, the ideal biomaterial surface should be able to induce osteogenesis and ensure a stable biological/ chemical bond between implant and bone. Upon implantation, the primary interactions of the biomaterial with the host tissue occur on a molecular scale and involve adsorption and reactions of water, inorganic ions, and biomolecules from the biological fluids on the biomaterial surface; the implanted material is thus colonized by proteins and eventually by cells. Because protein adsorption is an energy-driven process that involves electrostatic, hydrophobic, hydrogen bonding, and van der Waals interactions,3,4 it is important to investigate the wettability and surface energy of the biomaterial. From a mechanical point of view, materials for implant applications should possess load-bearing properties in order to withstand cyclic stress conditions within the body; in the absence of a biomechanical interlock, wear phenomena can © 2012 American Chemical Society

occur at the material−tissue interface, pointing out the importance of tribological aspects, which are closely related to the surface hardness and stiffness of the material. Various materials are clinically used for orthopedic implants, in particular, titanium, CoCrMo, and stainless steel alloys.5 In spite of their good biocompatibility, such metallic materials have some drawbacks, mostly related to high stiffness, possible wear of surface oxide, and lack of bioactivity. Alternative materials are being sought to overcome such limitations. Polymers like polyesters and methacrylic cements are good candidates because they can be tailored or used in composite formulations to match the mechanical properties of bone.6 In particular, methacrylic thermosets based on bisphenol A glycidylmethacrylate/triethyleneglycol dimethacrylate (BisGMA-TEGDMA) are employed for dentistry applications and are also being studied for orthopedic purposes.7 Although such resins are considered biocompatible, active research is aiming at improving the material−tissue interface.8,9 The long-term stability of the biomaterial may be promoted by deposition of Received: February 20, 2012 Revised: April 16, 2012 Published: April 17, 2012 1564

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Thermoset Preparation. The preparation of the thermoset samples (TS) and their coating with the polysaccharide have been previously reported.15 Briefly, BisGMA (70% w/w) and TEGDMA (30% w/w) were mixed with CQ (0.7% w/w) and DMAEMA (0.7% w/w) and UV-reticulated in Teflon mold Ø = 14 mm, h = 2.5 mm) with a Photopol IR/UV Plus oven (Dentalfarm, Italy) equipped with eight lamps and two spots operating in the wavelength range 320−550 nm. The thermosets were then sandpaper polished (granulometry: 1200). In the case of fiber-containing composites, thermoset bars (BarTS; 20 mm length × 2 mm width × 2 mm thickness) were reinforced by adding 50% (w/w) of longitudinally oriented E-glass fibers to the mold. Deposition of the Polysaccharide Coating on the Thermoset (Chitlac-TS). To introduce negative charges on the thermoset surface, an acidic hydrolysis of the methacrylic groups with hydrochloric acid (HCl) was carried out according to an activation procedure previously reported by the authors.15 Activated thermosets (HCl-TS) were immersed in 1 mL of Chitlac dissolved in water (4 mg/mL). Then the samples were gently stirred for 24 h, rinsed extensively with water, and air-dried for 18 h. For biological tests, all samples were UV-sterilized. Polysaccharide Adsorption as a Function of pH. To evaluate the influence of pH on Chitlac adsorption, the activated thermosets were incubated in 1 mL of fluorescein-labeled Chitlac18 (“Chitlac-fluo” 4 g/L) at different values of pH for 24 h. Initial pH was 5.5; it was modified by HCl or NaOH to obtain different values from pH = 2 to pH = 9 (see table S4 in the Supporting Information for the list of pH values). The fluorescence intensity of the supernatant solution was measured by a spectrofluorimeter Chameleon Hidex at 25 °C (excitation λ = 485 nm, emission λ = 535 nm). As a reference, the adsorption of Chitlac-fluo on both nonactivated samples and on the multiwell was measured. A calibration curve was performed at each chosen pH prior to quantitative analyses (R2 > 0.98). For each data point, the results from three disks were averaged. Polysaccharide Adsorption as a Function of Ionic Strength. To evaluate the influence of ionic strength (I), the activated thermosets were incubated in 1 mL of fluorescein-labeled Chitlac (“Chitlac-fluo” 4 g/L) for 24 h at four different NaCl concentrations (0, 0.015, 0.15, and 1.5 M). Then the supernatant solution was removed and its fluorescence intensity was measured by a spectrofluorimeter Chameleon Hidex at 25 °C (excitation λ = 485 nm, emission λ = 535 nm). As a reference, the adsorption of Chitlacfluo on the multiwell was measured. The fluorescence of nonlabeled Chitlac in contact with thermosets was found to be negligible. The measurements were performed at pH = 7.4 (in buffer solution HEPES 3 mM). A calibration curve was performed at each chosen ionic strength prior to quantitative analyses (R2 > 0.98). For each data point, the results from three disks were averaged. Nanoindentation Studies. Mechanical properties of the different surfaces were evaluated at 23 °C by depth-sensing indentation, using a Nanohardness Tester (CSM Instruments, Switzerland) at maximum load of 1 mN. The instrument used a Berkovich diamond indenter to produce triangular-shaped indentation marks on the coating surface. Depth-sensing indentation measurements were used to determine indentation hardness (H) and modulus (E) according to the procedure described by Oliver and Pharr.18 To have representative average values for the mechanical properties, 10 tests were performed for each sample. In the test, the applied load was increased at the speed of 2 mN/min until a nominal maximum load of 1 mN was reached. A creep period of 60 s was set before unloading at the speed of 20 mN/ min in order to limit viscoelastic effects of the material. A Poisson’s ratio of 0.4 was used for calculations. Fatigue Tests (Fixed Number of Cycles). The tests were performed on fiber-reinforced thermoset bars (BarTS) prior to and after the surface activation treatment. All the specimens were tested in three-point loading on a universal materials testing system (Galdabini Sun 500, Galdabini, Varese, Italy). The loading rate for the static test was 5 mm/min. The cyclic fatigue testing was done at the same loading rate of static conditions (which corresponds to 0.09 Hz) using a staircase approach of either 1000 cycles or specimen fracture as outlined by Drummond.19 Briefly, the specimen was loaded between

a bioactive coating capable of bonding to bone tissues. Surface functionalization by means of bioactive coatings have been proposed to provide the implant with osseointegrative properties. Various techniques may be employed according to the substrate material and the type of coating; typically these methods include the deposition of bioactive polymers and inorganic compounds like hydroxyapatite and calcium phosphates. Among different biopolymers, polysaccharides are employed in surface functionalization for their biocompatibility and ability to mimic the extracellular matrix.10,11 Recently, various polysaccharides have been used to coat implant materials with different purposes. Hu et al.12 coated titanium with carboxymethyl chitosan and hyaluronic acid modified with vascular endothelial growth factor (VEGF) to promote osteoblast functions; Cai et al.13 exploited the layerby-layer self-assembly technique for the adsorption of chitosan and gelatin on titanium to promote the viability of CAL-72 osteoblasts, while Pasqui et al.14 functionalized titanium with a modified carboxymethyl cellulose to increase adhesion and proliferation of MG-63 cells. In general, a convenient approach to graft polysaccharides onto biomaterials surfaces is to exploit cooperative electrostatic interactions between the substrate and the macromolecules, thus, avoiding the use of potentially toxic coupling agents. Our group recently devised a strategy to functionalize BisGMA/TEGDMA thermosets with a bioactive coating based on a lactose-modified chitosan (Chitlac) to promote bone cell proliferation15 or to prevent bacterial adhesion in vitro.16 In the present study, Chitlac was immobilized on BisGMA/TEGDMA thermosets by dip-coating; the effect of the polysaccharide on osteoblasts aggregation, proliferation, and activity was investigated in vitro, to provide a better biochemical and biological understanding of the previously reported behavior15 as well as to provide a sound basis to perform in vivo tests. The very encouraging in vitro results prompted us to expand our chemical and material investigations: the influence of pH and ionic strength was analyzed to optimize the deposition technique. The functionalized surface was characterized in terms of indentation hardness and modulus, wettability, and surface energy, while the influence of cyclic stress on the functionalized biomaterial was studied to predict the stability of the implant when subjected to fatigue conditions. Finally, the coated thermosets were implanted into minipig femurs to evaluate in vivo the biocompatibility of this third generation biomaterial.

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MATERIALS AND METHODS

Materials. Chitlac (lactose-modified chitosan, CAS registry number 85941-43-1) was prepared according to the procedure reported elsewhere,17 starting from a highly deacetylated chitosan (residual acetylation degree approximately 18%). The (viscosity average) relative molar mass of chitosan was estimated to be approximately 7 × 105. The composition of Chitlac determined by means of 1H NMR is glucosamine residue ∼20%, N-acetylglucosamine ∼18%, and 2-(lactit-1-yl)-glucosamine ∼62%. The relative molar mass of Chitlac was estimated to be approximately 1.5 × 106. Bisphenol A glycidylmethacrylate (Bis-GMA), triethyleneglycol dimethacrylate (TEGDMA), 2-[N-morpholino]ethanesulfonic acid (MES), 1-ethyl3-[3-(dimethylamino) propyl] carbodiimide (EDC), N-hydroxysuccinimide (NHS), 7-amino-1,3-naphtalene disulfonic acid (7ANA), and fluorescein isothiocianate (FITC) were purchased from Aldrich. Camphorquinone (CQ) and 2-dimethylamino ethylmethacrylate (DMAEMA) were purchased from Fluka. E-glass fibers were from Ahlström (Karhula, Finland). 1565

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Alkaline Phosphatase Assay. MG63 cells in adhesion and MG63 nodules were collected by centrifugation at 400 × g and lysed in a TritonX-100 solution (0.2% w/w TritonX-100 in 100 mM Tris/HCl buffer, pH = 9.8). Enzymatic activity was measured in a solution of 6 mM para-nitro-phenyl-phosphate and 1 mM MgCl2 in 100 mM TrisHCl at pH 9.8, after 60 min of incubation at 37 °C. Absorbance was measured at 410 nm. The results were normalized for the amount of protein content in the cellular extract calculated by means of BCA method according to the manufacturer’s protocol (Sigma). In Vivo Tests in Minipig. The in vivo testing was performed within the framework of the EU project “NewBone” (NMP3-CT2007−026279). The animal study protocol was approved by the Finnish National Animal Experiment Board, ELLA (permit No. ESLH2007-06829/Ym-23). The animal experiments were carried out in the Central Animal Laboratory of the University of Turku and the institutional guidelines for the analgesia and anesthesia were followed. The animal model included surgical placement of conical implants of a standard size (truncated cone, Ø 3.6−5.0 mm, height 8 mm), including one microroughened implant made of titanium alloy (Ti6Al4V) as a control material. This experimental setup allowed intra-animal comparisons of Chitlac-coated thermoset (Chitlac-TS) implants with the control implant. Three Chitlac-TS implants were implanted in the femur of three adult male Göttingen minipigs (Ellegaard Göttingen Minipigs A/S, Denmark). Surface roughness of the titanium alloy control implants, measured by a profilometer (Mitutoyo Surftest-301, Mitutoyo Corp., Japan), was Ra = 3.9 μm, while Chitlac-coated implants had Ra = 1.3 μm. Under general anesthesia and standard surgical sterility, the shaft of the left femur was exposed through an anteromedial approach. Unicortical holes were drilled in the anterior cortex of the femur using a high-speed dental drill under saline irrigation. The drill bit had the same conical geometric shape as the implants. The implants were tightly pressed-fit into the holes. After surgery, the functional activity of the animals was not limited and they received postoperative pain medication. After eight weeks, the animals were euthanized and the femurs with implants were harvested and analyzed using hard tissue histology. Histological Examination. The bone specimens with implants were dehydrated in a graded series of ethanol, cleared in xylene, and embedded in a light-curing methacrylate-based embedding medium (Technovit 7200 VLC, Heraeus Kulzer, Germany). Sections of 20 μm thickness were prepared with a cutting and grinding technique (Exakt Apparatebau, Hamburg, Germany) and stained by the van Gieson method. From each implant, a section was cut in the middle of the implant along the axis of the implant. The histological slices were imaged with a virtual microscope (Olympus BX51TF, Olympus Corp., Tokyo, Japan) controlled by a special imaging and analysis software (dotSlide 1.2, Olympus Soft Imaging Solutions GmbH, Münster, Germany). The quality of reactions at bone−implant interface was evaluated and scored using a standardized grading scale.23 In the histomorphometric analysis, a quantitative assessment of the direct bone−implant contact (BIC) was performed. In the cortical region, peri-implant bone segments in direct contact with the implant were measured. These measurements were summed and divided by the length of the implant surface available for bone apposition. The resulting ratio was expressed as a percentage. Due to the nature of the histological sections, the measurement was done twice for each implant (Figure 6C). Then the average of the two ratios was calculated. In addition, there were two distinct types of peri-implant bone observed; the mature lamellar bone similar to the original cortical bone and the newly formed woven bone. BIC was measured for both types of bone. No statistical analysis was feasible due to the small number of implants (n = 3). Statistical Analyses. To study differences among the behavior of different group samples, statistical analyses were performed using the Student’s t test. The level of statistical significance was considered to be either 0.05 or 0.01, as specified for the single experiments.

10 and 175 N until 1000 cycles or the failure of the specimen; in the case of fracture, the initial load was lowered by 15 N, otherwise it was increased by 15 N. The flexure strength was calculated according to the equation σ = 3Fl/(2bh2), where F is the fracture load in Newtons, l is the span between the supports (20 mm), b is the width (2 mm), and h is the thickness (2 mm) of the specimen. The mean fatigue limit and standard deviation for each sample group composed of 15 specimens were determined using the procedure described by Draughn.20 Fatigue Tests (Variable Number of Cycles). Flexural tests were conducted on thermoset bars (BarTS) using a three-point bending configuration. A span ratio (length/cross section) of 20 was chosen for static and fatigue tests to minimize the effect of the shear stress and to maximize the effect of the flexural stress. Fixed displacement, sinusoidal, flexural fatigue tests were performed at frequency 1 Hz using a load frame (MTS 858 Table Top System) with a 100 kN load cell. The displacement ratio, which is the ratio of minimum to maximum displacement, was 0.1 for all tests. Tests were conducted at four cyclic displacement levels, which corresponded to initial applied stress levels ranging from 66.7 to 80% of the flexural strength. Contact Angle Studies. Contact angles of the surfaces were measured with a goniometer (Krauss DSA 100 Drop Shape Analyzer) using the sessile drop method.21 Both polar (ultrapure water and ethylene glycol) and apolar (ultrapure di-iodomethane) liquids were used in order to allow surface energy calculations. A droplet of liquid (2 μL for polar liquids, 1.5 μL for di-iodomethane) was placed on the surface with a 500 μL syringe (Hamilton, Switzerland; needle Φ = 0.75 mm) at a speed of 30 μL/min. The profile of the water drop on the surface was recorded after 30 s to avoid time-dependent angle variations among samples. The software used the tangent method to calculate the θL (left) and θR (right) angles from the digitalized image, and the mean value was considered only when they differed less than 3°. For statistical analysis, 15 measurements for each surface type were averaged, and the significance level was checked by Student’s t test. The surface energy parameters were calculated from the contact angle values of the probe liquids according to the acid−base method proposed by Van Oss.22 Briefly, the values of the contact angles of the three liquids are used in the Young−Dupré equation, which enables us to obtain the values of the Lifshitz−van der Waals contribution of surface tension γLW and the acid−base (AB) components γ+ and γ− of the material. The surface polarity is calculated as the ratio between the AB contribution and the total surface tension. More details on the calculations can be found in the Supporting Information. Cell Culture. Human osteosarcoma cell lines (MG63, ATCC no. CRL-1427) were used for the in vitro experiments. MG63 cells were cultured in Dulbecco’s modified Eagle’s medium (Sigma), 10% heat inactivated fetal bovine serum (Sigma), 100 U mL−1 penicillin, 100 μg mL−1 streptomycin, and 2 mM L-glutamine in a humidified atmosphere of 5% CO2 at 37 °C. MTT Assay. Cell proliferation was evaluated in terms of mitochondrial oxidative capacity by the MTT(3,4,5-dimethylthiazol2yl-2,5-diphenyltetrazolium bromide) assay. The experiments were performed by using MG-63 cells. The cells (5000 cells/well) were plated in a 96-well microtiter plate and incubated for 24, 48, and 72 h at 37 °C in the presence of different concentrations of Chitlac. After incubation, the MTT solution was added at a final concentration of 0.5 mg/mL and the plates were further incubated for 4 h at 37 °C in 5% CO2 atmosphere. The MTT-containing medium was then removed and 100 μL of dimethyl sulfoxide (DMSO) per well were added to solubilize violet formazan crystals. Absorbance was measured in a microplate reader (Tecan Trading AG, Switzerland) at 620 nm. MG63 Culture on Chitlac-Coated Wells. Volumes of 100 μL of 1% w/v polysaccharide water solutions (Chitlac or Chitlac-fluo) were homogeneously distributed to cover the bottom of each well of a 24multiwell plate and then allowed to air-dry. Coated wells were sterilized by a 12 h UV exposure. Thereafter, the MG63 cells were seeded at a density of 4 × 104 cells/well and maintained in culture for 8 days. Cells cultured on wells without polymer coating were used as adhesion control cultures. Cell morphology was investigated by means of light microscopy (Leitz Labovert) or by fluorescence microscopy (Leica DC 300F). 1566

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RESULTS AND DISCUSSION Chitlac Effects On MG63 Osteoblast-Like Cell Phenotype. Chitlac is a lactose derivative of a highly deacetylated chitosan, which has shown interesting biological effects toward primary cultures of pig chondrocytes. Previous studies have demonstrated that this glycopolymer exhibits the specific ability to induce chondrocyte aggregation, to stimulate extracellular matrix biosynthesis, and to alter cell cycle progression.17 Experimental evidence were identified in the S-type galactosebinding lectin Gal-1, with the protein acting as a bridge between membrane integrins and the polysaccharide.24 Similarly, experiments performed by seeding osteoblasts-like MG63 cells on Chitlac-coated culture wells have evidenced that Chitlac stimulates the aggregation of cells, inducing the formation of living cells nodules. As reported in Figure 1,

to cell cultures led to a slight but significant enhancement of cell proliferation with respect to the nontreated cells. As a control, Chitosan solutions at the same concentrations of Chitlac ones were used; in this case no significant effects on cell growth were observed (Figure S2 in the Supporting Information). Overall, the biological effects of Chitlac on osteoblasts could well be ascribed to a specific interaction of the polymer side groups with a receptor for galactose on osteoblast plasma membranes, resembling that found on the chondrocyte ones (in that case, Galectin 1). Work is in progress for the identification and characterization of this interactor. These promising results, together with the previous ones,15 provide a sound rationale for the use of Chitlac as a bioactive polymer for coating of BisGMA-based biomaterials in orthopedic applications, when stimulation of bone formation and development is highly desirable. Chitlac Adsorption on Activated Thermosets As a Function of pH and Ionic Strength. Our group has recently demonstrated that surface hydrolysis of BisGMA/TEGDMA thermosets can be successfully exploited to introduce carboxylate moieties, thus, obtaining an “activated” surface with a high density of negative charges. The activated thermoset can then be coated by dipping the material into the polysaccharide solution to exploit electrostatic interactions between the positively charged polysaccharide Chitlac and the negatively charged thermoset surface.15 In this work we aimed at studying the influence of charge density on polysaccharide adsorption by varying both pH and ionic strength of the Chitlac solution. To this end, the thermosets were immersed in Chitlac-fluo solutions with pH ranging from 2 to 9 and the amount of polysaccharide adsorbed on the thermoset was estimated as a function of pH by means of fluorescence measurements (Figure 3A). Figure 3A shows that the polysaccharide adsorption is a nonmonotonic function of pH, and it displays a maximum of approximately between pH = 5 and 6. This was attributed to the chemical nature of Chitlac amino groups and of carboxyl groups of the thermoset surface. In fact at this pH value all amino groups of Chitlac are protonated, while it is expected that about one-half of the carboxylate groups are deprotonated; this electrostatic interaction accounts for the maximum of Chitlac deposition. At low pH (∼2), both amino and carboxylate groups are protonated, thus, hampering electrostatic adsorption. Hence, at low pH, the small Chitlac adsorption could be ascribed only to weak interactions (ion− dipole) between amino groups and polar groups of the thermoset. At pH of between 6 and 9 the polysaccharide adsorption is also lowered due to partial deprotonation of amino groups because these values are close to the pKa of Chitlac (pKa = 5.97).17,25 The effect of the ionic strength on Chitlac adsorption on the thermoset was explored. The activated thermosets were immersed in Chitlac-fluo aqueous solutions at various ionic strengths at physiological pH (7.4). After 24 h of immersion, the amount of polysaccharide adsorbed was calculated by measuring the fluorescence of the supernatants. The results are shown in Figure 3, which shows that the adsorption stays at its maximum up to about I−1/2 = 2.25, corresponding to a NaCl concentration of 0.15 M. The variable I−1/2 was used, instead of just I, because the Debye (screening) length (κ−1) is directly proportional to I−1/2. The Debye length is a key parameter in polyelectrolyte physics, providing the length scale of significant electrostatic interactions. It appears that Chitlac adsorption

Figure 1. Morphology of MG63 cultured in adhesion (A) and in wells coated with Chitlac, 1 day (B), 5 days (C), and 8 days (D) from seeding. Magnification is 20×. Distribution of fluoresceine-labeled Chitlac (E) on single cells (magnification 100×) and (F) on nodules (magnification 20×).

control cells adhered to the well and displayed regular morphology (Figure 1A), whereas those grown on Chitlaccoated wells led to the formation of nodules (Figure 1B−D). The aggregation process started a few hours after cell inoculation, leading to nodules of various sizes and shapes (average dimensions 0.1−0.5 mm). Fluoresceine-labeled Chitlac enabled us to ascertain the presence of the polysaccharide both inside the aggregates and on the plasma membrane of isolated cells (Figure 1E,F). The level of alkaline phosphatase (ALP) activity in MG63 cells, a partially undifferentiated cell line, was determined after 8 days, with the aim of evaluating osteoblast functionality and/or differentiation processes. As reported in Figure 2A, ALP activity detected in the cell aggregates was higher in comparison with that assayed in adherent cells. The interaction of the polymer with cells also induced a clear and significant effect on growth kinetics, as reported in Figure 2B. Three different concentrations of Chitlac solutions added 1567

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Figure 2. (A) ALP activity in adhesion cells and in Chitlac-induced aggregates after 8 days of culture. Data are expressed as mean optical density ± standard deviation (n = 5). **p < 0.01 compared with cells growth in adhesion. (B) MTT assay performed on adherent MG63 cells in absence (0 gr/L) and in the presence of different concentrations (0.1, 0.5, 1, 2 gr/L) of Chitlac after 1, 2, and 3 days of culture. Data are expressed as mean optical density ± SD (n = 5) compared with cell growth in absence of Chitlac: *p < 0.05; **p < 0.01, n = 5.

Figure 3. (A) Chitlac adsorption on activated BisGMA/TEGDMA thermosets as a function of the pH of polysaccharide solution. Data points are the means of three independent experiments. The dotted line was drawn to guide the eye. (B) Chitlac adsorption on the surface of activated BisGMA/ TEGDMA thermosets as a function of the inverse of the square-root of the ionic strength (I) at physiological pH (7.4). The dotted line is drawn to guide the eye.

methane were measured (see Table S1 in the Supporting Information). The highest contact angles were obtained with water; the mean value for the uncoated material (TS) was 75.9 ± 2.7°, and it did not significantly change after the treatment with HCl (77.0 ± 1.9°). The presence of the polysaccharide coating induces a slight decrease in the value (Chitlac-TS = 70.9 ± 2.7°), ascribable to a higher hydrophilicity of the Chitlac macromolecules. In the literature, contact angle data for similar coatings based on the polysaccharide chitosan display a large range of values. In the case of water, Carneiro-da-Cunha et al.26 and Xu et al.27 reported angles of around 50° or lower for alginate−chitosan coatings, Wanichapichart et al.28 found values in the range of 65−78° for chitosan-based membranes, while other authors showed angles of 76.4°29 and even higher than 90°.30,31 This wide range of numbers points out the sensitivity of this technique to specific coating properties, like chemical composition of the coating, surface roughness, physical state of the macromolecules (e.g., coiled, stretched, cross-linked), pH, and temperature conditions. In our case, Chitlac-TS samples are associated with an angle decrease for polar liquids (water and ethylene glycol), while for the apolar di-iodomethane, the angle values increase, in agreement with the case of chitosan coatings for poly(hydroxyethyl meth-

(due to charge attraction) is rather insensitive to the Debye length until the NaCl concentration reaches the value of approximately 0.15 M. Above this I range, the increasing concentration of ionic species shields the electrostatic (attractive) interactions between amino and carboxylate groups, which lowers polymer adsorption. The adsorption study pointed out that for an optimal absorption of the polysaccharide the pH range should be between 5 and 6 with low ionic strength; thus, for the characterizations performed in this paper, Chitlac was dissolved in deionized water with pH = 5.7. Because the stability of the coating in saline solutions was previously assessed,15 culture medium and physiological solution should not detach the polysaccharide layer. Physical-Chemical Properties: Contact Angle Measurements. The sessile drop method was used to study wettability and surface energy of the functionalized thermosets; because the variation of wetting conditions depends on the properties of the material layer in direct contact with the liquid phase, this technique was used to spot the effect of the acidic treatment and of the polysaccharide deposition on the thermoset surface. The contact angle values of the investigated surfaces obtained from water, ethylene glycol, and di-iodo1568

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acrylate) described by Yakup and co-workers.32 Because, in the case of apolar liquids, the angles tend to increase when the content of polar functional groups is higher,33 the wetting behavior of di-iodomethane is in line with the presence of the hydrophilic polysaccharide on the material surface. Physical Chemical Properties: Surface Free Energy Parameters and Work of Adhesion. According to the Van Oss theory, it is possible to correlate contact angles with surface energy of a solid material by means of the Young-Duprè equation which enables the dispersive (LW) and acid−base (AB) components of a solid surface to be calculated.22 This approach was used to calculate surface free energy parameters and work of adhesion of the investigated surfaces (all values are shown in Table S2 in Supporting Information). The total surface energy γTOT is comparable between the three different surfaces and it ranges from 45.3 to 46 mJ/m2. For the uncoated thermoset (TS and HCl-TS), the acid−base γAB term gives a small contribution because the surfaces exhibit predominantly monopolar electron-donicity (γ− ≫ γ+ ∼ 0). At variance with this, in the case of Chitlac-TS, there is an increase in γ+ contribution, which is ascribable to the presence of surface hydroxyl sites; in the case of the polysaccharide coating, the higher value of γAB and the decrease of the dispersive component γLW jointly cause a 3-fold increase in the surface polarity (12.7%) with respect to HCl-TS; this result was expected, given the chemical structure of the macromolecules adsorbed on the surface. Similar behavior was observed in the literature in the case of chitosan coating for poly(hydroxyethyl methacrylate).32 The thermodynamic work of adhesion measures the degree of intermolecular interactions between surface and liquid; from the contact angle data for polar liquids (water, ethylene glycol), the acid−base work of adhesion WAB and its percentage with respect to the total work of adhesion (WAB %) were calculated for the different surfaces. The results of Table S2 show that the acidic treatment does not induce significant changes in the work of adhesion. On the contrary, the polysaccharide coating is associated with an 8% increase of acid−base interactions, which reflects the presence of polar functional groups of Chitlac; the adhesion work on the coatings is higher with both polar liquids (water and ethylene glycol). These results suggest that surface energy parameters and interfacial interactions of polar liquids provide a reasonable description of the acid−base character of the Chitlac-coated thermoset. When biomaterials are implanted into body, protein adsorption onto the foreign surface occurs within seconds from implantation, so that cells arriving at the biomaterial surface interact with the adsorbed protein layer rather than directly with the material itself. This initial protein adsorption onto a biomaterial plays a fundamental role in how the body responds to an implanted biomaterial. Material surface properties, including hydrophobicity, wettability, charge, roughness and softness, are key driving mechanisms for protein absorption processes, and thus, biological response.34 It is worth noting that coating chitosan on Ti surfaces decreased the wettability of the Ti, but increased protein adsorption and osteoblast-like cell attachment and proliferation.35 Additional experimental evidence will be necessary to clarify whether the enhanced proliferation rate of osteoblasts on Chitlac-coated thermoset previously described by some of the authors15 can be ascribed solely to the biological activity of the polysaccaride or also to the combined effect of the more hydrophilic and polar layer.

Material Properties: Nanoindentation Studies. Nanoindentation assays enabled us to calculate the values of indentation modulus (E) and hardness (H) of the three different surfaces, namely the unmodified thermosets (TS), the activated HCl-treated thermosets (HCl-TS) and the thermoset coated with Chitlac (Chitlac-TS). For Chitlac-TS the average maximum indentation depth was lower than 600 nm. The thickness of the coating layer on Chitlac-TS samples in the dry state was previously reported to be approximately 10 μm.15 Because the effect of the substrate on the mechanical properties is negligible when the indentation depth is lower than 10% of the film thickness,30 the methacrylic substrate is not supposed to influence the mechanical response of the coating in these measurements. The indentation values of E and H for the three different surfaces are reported in Figure 4.

Figure 4. Nanoindentation values of BisGMA/TEGDMA samples, that is, the unmodified thermoset (TS), the HCl-treated thermoset (HCl-TS), and the Chitlac-coated thermoset (Chitlac-TS). The left y axis reports the indentation modulus E (■ = E). The right y axis reports the indentation hardness (○ = H). Differences between all mean values are statistically significant according to Student’s t test (p < 0.01, n = 10).

As to the methacrylic surface alone, the treatment of TS with HCl (surface activation) improves the surface properties. This fact could be explained by invoking the presence on the thermoset surface of a layer in which the reticulation process is partially inhibited by the contact with air oxygen,36,37 which affects the mechanical behavior of the surface at the nanoscale level; this “inhibition layer” could be subjected to degradation upon treatment with concentrated HCl, thus, exposing the underlying substrate with a higher degree of reticulation, which corresponds to higher values of E and H. Moreover, the treatment with HCl is associated with a decrease of residual monomers, as pointed out by ATR/FTIR analyses (see Figure S3 and Table S3 in the Supporting Information). When dealing with polymeric materials, nanoindentation data can be compared between materials with the same viscoelastic behavior. In fact, the mechanical properties of a polymeric material are dominated by the deformation, the applied stress, and also by the solicitation time; indeed, polymers show a time-dependent mechanical behavior. Thus, the comparison of indentation values between two different polymeric materials is not rigorous, and in this study, the mechanical values should be directly compared only between the two methacrylate surfaces (TS and HCl-TS); nevertheless, 1569

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the values of the polysaccharide-coated sample (HCl-TS) are reported in the same graph assuming a similar viscoelastic behavior on this observation scale. In this perspective, the deposition of the polysaccharide (Chitlac-TS) corresponds to an enhancement of E and H values. This behavior is in line with the results reported by several authors who described polysaccharide-based coatings on various substrates; for instance, Carneiro-da-Cunha26 found an increase of indentation hardness for PET coated with a mixed alginate−chitosan multilayer, and Martin et al.30 showed similar indentation hardness values (150−190 MPa) in the case of chitosan-based coating of titanium. Material Properties: Fatigue Tests. The acidic treatment employed for BisGMA/TEGDMA surface activation was previously shown to cause a decrease of 18% in the flexural strength of the construct in static conditions;15 in this work we analyzed the mechanical behavior in a condition more closely resembling the stress of implants in vivo, namely, fatigue stress. In fact, because these methacrylate-based materials are being developed for load-bearing applications such as orthopedic prosthesis, it is important to evaluate the effect of dynamic loading on the material resistance. As the fatigue stress is considered a property of the bulk, the presence of the coating is not expected to bring about a notable contribution; therefore, the comparison is made between nontreated and HCl-treated samples. The effect of fatigue stress was first evaluated by determining the flexure fatigue limit at a fixed number of loading cycles: thermoset samples reinforced with 50% (w/w) of uniaxial Eglass fibers (BarTS) were subjected to three-points bending for 1000 cycles at 0.09 Hz to determine the flexure fatigue limit using the staircase method proposed by Draughn.20 In the case of nontreated samples, the flexural strength decreased from 845 ± 26 MPa in static conditions to 698 ± 20 MPa when the bars were cyclically loaded, which means a performance reduction of 17% when considering the fatigue limit after 1000 cycles. In the case of HCl-treated samples, the flexural strength decreased from 650 ± 62 MPa in static conditions to 600 ± 56 MPa when the bars were cyclically loaded, with a performance reduction of 8%. In both cases, the decrease could be ascribed to crack propagation and weakening of the fiber−matrix interface due to fatigue stress. Dynamic fatigue tests have been performed at higher frequency (1 Hz) in a three-points bending setting using variable displacement levels; larger displacements produce larger stresses in the composite, increasing the probability of initiation and propagation of damages within the material, with a consequent decrease in the number of cycles to failure. Four different stresses were exerted on the materials, ranging from 66 to 80% of the static flexural strength. Experimental data were compared for both activated and nontreated composites and a relationship was found between the number of cycles to failure and applied stress level (Figure 5); this type of curve (Wöhler curve) is typically used by designers to predict fatigue performance of constructs. The figure shows that, for both HCl-treated and nontreated samples, the number of cycles to failure decreases as the stress increases in a log−linear fashion. The scatter in fatigue life among similar samples can be attributed to natural variability in the material properties and to the slight inhomogeneous and anisotropic nature of the handmade composites. This test confirms that surface activation has a non-negligible influence on the fatigue behavior of the composite thermoset, which

Figure 5. Stress level−number of cycles to failure graph for flexural fatigue behavior (1 Hz) of BisGMA/TEGDMA-based composites (BarTS). Y axis displays the stress percentage with respect to static flexural strength. Dotted bands are drawn to guide the eye.

should be considered when designing composite implants for cyclic load-bearing applications. In Vivo Tests in Minipigs. In vivo experiments were designed to perform basic screening of Chitlac-based implants for bone incorporation. The comparison was made with control implants made of titanium Ti6Al4V alloy with clinical surface microroughening, which represents the state of the art in orthopedic practice. The implants were harvested from minipig femurs after 8 weeks, and the histological examination of implants demonstrated no signs of inflammation or foreign body reactions around both test and control implants (Figure 6). Notably, Chitlac-TS implants (Figure 6A) showed direct bone-implant contact with minimal soft tissue interlayer indicating good biological compatibility of the material; the thickness of the reaction zone was narrower (2 μm) than titanium control (14 μm; see Table S5 in the Supporting Information) and the peri-implant bone was similar to the original cortical bone. At variance, the control Ti6Al4V implants exhibited a zone of peri-implant woven new bone formation with local areas of direct bone-implant contact, remodelling lacuna and localized fibrous tissue (Figure 6B). The bone−implant interfaces were compared with both qualitative and quantitative analysis according to a standardized grading scale23 and the numbers are shown in detail in Table S5 (Supporting Information). As to the bone reaction score, it is based on a scale where the maximum grade (4) indicates a tissue similar to the original cortical bone, while lower values indicate progressively increasing amounts of fibrous tissue; in the case of Chitlac-TS, the score was 3.8, while for the titanium implant it was 3.2. Similarly, the interface score is based on a scale where the maximum grade (4) indicates a bone-toimplant contact without soft tissue interlayer, while lower values indicate, progressively, the formation of a fibrous tissue capsule; in the case of Chitlac-TS, the score was 3.9, while for titanium implant it was 3.2. All in all, the three evaluation parameters considered indicate the good biological response of the polymeric implant. 1570

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newly formed fraction to 50%. This finding and the data on bone−implant interface (Table S5) indicate that, while ChitlacTS does have biocompatibility and does not cause adverse bone remodeling, while its ability to stimulate peri-implant new bone formation and osseointegration is as least as good as that of the control Ti6Al4V with roughened surface. The latter surface property is absent in the Chitlac-TS sample; therefore, a combination of surface microroughness with bioactive surface properties could be a possible direction of the future research to increase osseointegration.

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CONCLUSIONS Surface functionalization of BisGMA/TEGDMA thermosets with the polysaccharide Chitlac was optimized with respect to the dip-coating procedure previously devised. In particular, the amount of polysaccharide adsorbed on the activated thermoset was shown to depend on the pH and ionic strength of the dipping solution; both parameters trigger the electrostatic interactions between the positively charged macromolecule and the oppositely charged substrate. The polysaccharide deposition causes an increase in indentation hardness and modulus of the material surface and improves wettability due to higher surface polarity. The effect of fatigue stress on the composite construct determines a decrease in flexural strength, which depends on cycles frequency and displacement level; this occurs for both nontreated and surface-activated samples, the latter displaying reduced stress resistance. In vitro biological tests on MG63 osteoblasts proved that Chitlac drives cell aggregation, stimulates alkaline phosphatase production, and enhances cell proliferation. Finally, in vivo tests in minipigs proved that BisGMA/TEGDMA thermosets functionalized with the polysaccharide Chitlac show a biological compatibility comparable with that of titanium, with higher scores as to bone reaction and interface parameters. This approach appears particularly appealing for orthopedic and dentistry applications, expanding the potential of Chitlac as a novel biomaterial with respect to what has already been shown for cartilage.

Figure 6. Light micrographs of the (A) Chitlac-TS implant and (B) microroughened control Ti6Al4V implant in cortical bone (red) after 8 weeks of implantation in a minipig model. Areas with direct bone− implant contact (arrowheads) alternate with areas with fibrous tissues (arrows). Asterisks indicate newly formed woven bone. (C) Schematic illustration of a cross section of an implant after healing in cortical bone. The growth of newly formed woven bone from periosteal and endosteal side is shown. Lines illustrate the bone−implant contact (BIC) measurements in the cortical area.

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For the control Ti6Al4V implants, the total bone−implant contact (BIC) was 46% (min 30%, max 60%). All the periimplant bone was the newly formed woven bone as the main component of the osseointegration process. For the Chitlac implants, the total BIC was 72% (min 59%, max 80%), with lamellar bone constituting 44% (min 18%, max 68%) and the newly formed woven bone constituting 28% (min 12%, max 41%). The reaction zone, where bone remodeling had taken place, was narrower for the Chitlac-TS implants. Histomorphometric analysis indicated that control Ti6Al4V with roughened surface implants induced an extensive periimplant bone remodeling. All peri-implant bone was newly formed, with 46% being attached to the surface of the implant, characterizing an expected osseointegration process of a titanium alloy implant. In the case of Chitlac-TS with nonroughened surface implant, 72% of the implant interface was in close contact with the cortical bone. However, the cortical bone observed in histological sections was mainly lamellar cortical bone (44%). This finding can either be attributed to the excellent bioactive properties of the material (then with an excellent 72% of total BIC) or be a sign of no new bone formation due to the tight press-fit of the implant in bone. Under the latter conservative hypothesis, the observed value of 28% of newly formed cortical bone contact should then be rescaled to exclude the pre-existing component, bringing the

ASSOCIATED CONTENT

S Supporting Information *

This section includes data from ATR-FTIR studies, biological tests (MTT assay) regarding the effect of Chitosan on MG63 osteoblast-like cells, and histological parameters from in vivo tests. It also includes tables with the list of contact angles, surface free energy parameters, work of adhesion values, and a table with values of Chitlac adsorption on activated thermosets as a function of pH (displayed in Figure 3A). This material is available free of charge via the Internet at http://pubs.acs.org.

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected] Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This study was supported by the EU-FP6 Project “NEWBONE” (Contract No. 026279-2) and by the COST action MP0701. The financial support of the University of Trieste to I.D. (through the InterDepartment Center for Molecular Medicine - CIMM) and to A.T. (through the Post-Doc 1571

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(31) Silva, S. S.; Goodfellow, B. J.; Benesch, J.; Rocha, J.; Mano, J. F.; Reis, R. L. Carbohydr. Polym. 2007, 70, 25−31. (32) Yakup, A., M.; Yilmaz, M.; Yalçin, E.; Bayramoğlu, G. J. Membr. Sci. 2004, 240, 167−178. (33) Rotta, J.; Ozorio, R. A.; Kehrwald, A. M.; de Oliveira Barra, G. M.; de Melo Castanho Amboni, R. D.; Barreto, P. L. M. Mater. Sci. Eng., C 2009, 29, 619−623. (34) García, A. J. Interfaces to Control Cell−Biomaterial Adhesive Interactions Polymers for Regenerative Medicine, 203rd ed.; Werner, C., Ed.; Springer: Berlin/Heidelberg, 2006; pp 171−190. (35) Bumgardner, J. D.; Wiser, R.; Elder, S. H.; Jouett, R.; Yang, Y.; Ong, J. L. J. Biomater. Sci., Polym. Ed 2003, 14, 1401−1409. (36) Vallittu, P. K. J. Oral Rehabil. 1999, 26, 208−212. (37) Yatabe, M.; Seki, H.; Shirasu, N.; Sone, M. J. Oral Rehabil. 2001, 28, 180−185.

Fellowship Programme) is gratefully acknowledged. The authors would like to thank Mrs. Suzanne Cann for proofreading.

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