Nano forsterite biocomposites for medical applications: Mechanical properties and bioactivity Gabriel Furtos,1 Marieta-Adriana Naghiu,2 Heidi Declercq,3 Maria Gorea,2 Cristina Prejmerean,1 Ovidiu Pana,4 Maria Tomoaia-Cotisel2 1

Department of Dental Materials, Raluca Ripan Institute of Research in Chemistry, Babes-Bolyai University, Cluj-Napoca, Romania 2 Department of Chemical Engineering, Faculty of Chemistry and Chemical Engineering, Babes-Bolyai University, Cluj-Napoca, Romania 3 Department of Basic Medical Sciences, Tissue Engineering Group, Ghent University, Ghent, Belgium 4 Physics of Nanostructured Systems Department, National Institute for R&D of Isotopic and Molecular Technology, ClujNapoca, Romania Received 21 August 2014; revised 20 January 2015; accepted 8 February 2015 Published online 00 Month 2015 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.b.33396 Abstract: The aim of the present study was to obtain and to investigate nano forsterite and nano forsterite biocomposites for biomedical application. New self-curing forsterite biocomposites were obtained by mixing nano forsterite powder (5, 15, 30, 50, 70 wt %) with 2,2-bis[4-(2-hydroxy-3-methacryloyloxypropoxy)-phenyl]propane (bis-GMA) and triethyleneglycol dimethacrylate (TEGDMA) monomers. The new nano forsterite biocomposites were investigated for mechanical properties: compressive strength (CS) (143–147.12 MPa), compressive modulus (CM) (1.67–2.75 GPa), diametral tensile strength (DTS) (27.33–31.55 MPa), flexural strength (FS) (59.47–83.20 MPa) and flexural modulus (FM) (2.05–8.60 GPa). Increases of CS, DTS, FS with increasing amount of forsterite were observed up to 50 wt %. The highest CM and FM values were registered for 70 wt % and a direct correlation between

the forsterite volume fraction (%) was observed. SEM micrographs revealed the morphology of surface of fractured biocomposites after CS test. XPS indicated that these biocomposites promoted the hydroxyapatite formation on their surface immersed in simulated body fluid (SBF). AFM images showed that the growth of the hydroxyapatite layer occurs with a preferred orientation on the surface of forsterite biocomposites after immersion in SBF. Incorporation of nano forsterite in the polymer matrix (bis-GMA/TEGDMA) did show osteoblast adhesion and proliferation was improved on C 2015 Wiley Periodicals, Inc. J nano forsterite biocomposites. V Biomed Mater Res Part B: Appl Biomater 00B: 000–000, 2015.

Key Words: nano forsterite, biocomposites mechanical properties, biomaterials, polymers

materials,

How to cite this article: Furtos G, Naghiu M-A, Declercq H, Gorea M, Prejmerean C, Pana O, Tomoaia-Cotisel M. 2015. Nano forsterite biocomposites for medical applications: Mechanical properties and bioactivity. J Biomed Mater Res Part B 2015:00B:000–000.

INTRODUCTION

The considerable progress in nanotechnology provides different nanostructured biomaterials for various applications in bone tissue engineering and bone regeneration. Among them, biocomposites obtained by the incorporation of inorganic particles as fillers, like nano forsterite, within a polymer matrix based on methacrylates might have many applications in orthopaedic or dentistry field. The first bone cement used was based on poly(methyl methacrylate) (PMMA) in the 1960s by John Charnley.1 Poly(methyl methacrylate) (PMMA) based bone cement has been widely used for fixation of total joint replacement prostheses to periprosthetic bone tissue for over 40 years. The major drawbacks for the acrylic products

are the exothermic reaction, poor adhesion to bone surfaces and fillers, the rapid resorption, poor mechanical properties2 and lack of bioactivity. Discovery by Raphael Bowen3 of the molecule 2,2-bis[p-(2-hydroxy-3-methacryloxypropoxy)phenyl] propane (bis-GMA) led to the development of many types of materials for restorative dentistry. Advantages of bis-GMAbased composite over traditional PMMA include: possibility to be used as injectable cement, precise placement,4 polymerization at lower temperatures, superior mechanical properties as well as providing a direct bone surface bonding to the host bone compared to PMMA.5,6 Addition of bioactive particles to the polymer matrix could improve mechanical, biocompatibility and bioactivity

Additional Supporting Information may be found in the online version of this article. This article was published online on 24 June 2015. An error was subsequently identified. This notice is included in the online and print versions to indicate that both have been corrected 17 December 2015 Correspondence to: G. Furtos; e-mail: [email protected] Contract grant sponsor: National Research Plan; contract grant numbers: 171/2012, 189/2012

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TABLE I. Composition of the Nano Forsterite Biocomposites Code

Composition of Filler (P, wt %)

Organic Phase Composition (L, wt %)

P C5F C15F C30F C50F C70F

Forsterite Forsterite Forsterite Forsterite Forsterite Forsterite

Monomers Monomers Monomers Monomers Monomers Monomers

a

(0%) (5%) (15%) (30%) (50%) (70%)

0/100 5/95 15/85 30/70 50/50 70/30

P: forsterite nano powder; L: liquid monomers.

properties of biocomposites, may promote new bone growth at the interface with the implant, and in time can lead to the increase of in vivo longevity of the prosthesis.7 There are different types of fillers: hydroxyapatite, bioactive ceramics,8,9 beta-tricalcium phosphate,10 bioglass,11–13 glassceramic,9 or glass fibers14 with different sizes and shapes, mechanical properties and biological properties used for bone replacement. Hydroxyapatite (Hap, Ca10(PO4)6(OH)2) is a major mineral constituent of the bone matrix and is one of the most widely used bioceramic material in the field of biomaterials and tissue engineering, because Hap shows good biocompatibility, bioactivity, has high osteoconductive, and/or osteoinductive properties.8 Unfortunately, the low fracture toughness of Hap (0.6–1.0 Mpa m21/2) had the disavantage to be at the limits of the scope of clinical applications.15 To improve mechanical properties our article proposes to use forsterite as fillers for obtaining biocomposites. Forsterite (Mg2SiO4) is a mineral member of the Olivine group based on the magnesia–silica system and showed in vitro good bioactivity and biocompatibility, providing the opportunity to be a material that may be used as bioactive bone repair materials.16–19 Forsterite ceramics had higher fracture toughness (KIC 5 2.4 Mpa m21/2)17 compared with hydroxyapatite ceramics (KIC 5 0.6–1.0 Mpa m21/2) and higher than lower limit reported for bone implants.15 However, to the best of our knowledge, there are no reports on the use of nanostructured forsterite as reinforcing filler in the polymer matrix. Thus, biocomposites based on nano forsterite incorporated into polymer matrix might have potential applications in dental and orthopedic fields. The aim of this study was to obtain and characterize nano forsterite powder and to study the influence of its incorporation in polymer matrix for obtaining new nano forsterite biocomposites. The nano forsterite morphology, sizes and structure were determined by using transmition electron microscopy (TEM), granulometry and X-ray diffraction. A series of new nano biocomposites were obtained through the addition of different quantities of nano forsterite powder (5, 15, 30, 50, and 70 wt %) into a self-curing monomer mixture based on bis-GMA/TEGDMA The influence of nano forsterite quantity in new biocomposites was evaluated by the compressive strength (CS), compressive modulus (CM), diametral tensile strength (DTS), flexural strength (FS), and flexural modulus (FM) tests. The surface of forsterite biocomposites was investigated by scanning electron microscopy (SEM) before and after the CS tests. The topography of samples surfaces,

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(100%) (95%) (85%) (70%) (50%) (30%)

P/La (wt %)

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after storage into simulated body fluid (SBF), was determined by SEM and atomic force microscopy (AFM). The X-ray photoelectron spectroscopy (XPS) was used for qualitative and quantitative surface sample analysis by means of a SPECS custom build system. All the materials were also investigated for cell viability, adhesion, and proliferation. MATERIALS AND METHODS

Materials Reagent grade chemicals of magnesium nitrate hexahydrated Mg(NO3)2 6H2O, (99.5% purity Merck), and tetraethyl orthosilicate (TEOS (C2H5O)4Si, (Merck, Germany) were used for the preparation of forsterite. Nitric acid was purchased from Merck, Germany. The 2,2-bis[4-(2-hydroxy-3methacryloxypropoxy)phenyl]propane (bis-GMA) (Aldrich Chemical, Milwaukee, WI), triethyleneglycol dimethacrylate (TEGDMA) (Sigma Chemical, St. Louis, MO), N,N-dihydroxyethl-p-toluidine (DHEPT) (activator of polymerization) and benzoyl peroxide (BPO) (initiator of polymerization) (Merck–Schuchardt) were used as received without further purification. Deionized water was used in all experiments. Synthesis of forsterite nano powder Forsterite (Mg2SiO4) nano powders were prepared by sol– gel method using magnesium nitrate and TEOS. Magnesium nitrate (50 g) was disolved in distilate water (150 mL) for obtaining a transparent solution. To this solution 22 mL TEOS was added for obtaining a MgO/SiO2 molar ratio of 2. The mixture was vigorous stirred for 2 h at room temperature and pH of the mixture was adjusted to 3.5–4 using nitric acid solution. This wet gel was kept in the oven and aged at 120  C for 7 h to form a highly viscous gel. This highly viscous gel was calcined at 900  C for 2 h in air employing a heating rate of 5  C min21. Preparation of nano forsterite biocomposites The forsterite biocomposites were prepared as bicomponent systems (self-curing system), each consisting of two pastes (named paste A and paste B, respectively). The composite pastes were obtained by dispersing the forsterite filler in an organic phase (monomer mixture) in which was previously added the proper component of the initiation system. The experimental monomer mixture was obtained by mixing bisGMA (60% by weight) with TEGDMA (40% by weight). For obtaining paste A and paste B, respectively, the monomer mixture was divided in two liquids (liquid A and liquid B).

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BPO (initiator) was disolved in liquid A (1.5% by weight) and DHEPT (activator) was introduced in liquid B (1.5% by weight) . Forsterite filler was hand mixed with the two liquids (liquid A and liquid B) in different ratios (Table I) until when was obtained a homogeneous paste (Paste A and Paste B). The final composition of biocomposites after mixing the pastes A with the corresponding pastes B is shown in Table I. Forsterite characterization X-ray diffraction investigations. The diffraction data were collected in the 2u 5 15–85 angular domain with a Bruker D8 Advance diffractometer, using Cu Ka1 radiation (k 5 1.5406 Å) (40 kV; 40 mA). To increase the resolution, a Ge 111 monochromator was used to eliminate the Ka2 radiation. The crystallite size was estimated from the X-ray diffractograms using the Scherrer formula.20 Particle analysis The mean particle size and size distribution, of the forsterite nanoparticles was assessed by dynamic light scattering using a particle size analyzer (SALD-7101, Shimadzu Corporation, Tokyo, Japan) at 25  C. A transmission electron microscope (TEM: JEM 1010, JEOL, Tokyo, Japan) was employed to measure the particle size, to evaluate shape and the degree of agglomeration of the forsterite nanoparticles. TEM images have been recorded with JEOL standard software. Methods Sample preparation for mechanical test. Nano biocomposites were prepared by hand-mixing of paste composite A with the corresponding paste composite B in a ratio 1:1, at room temperature, until a homogeneus system is obtained. The self-curing pastes were inserted in a Teflon mold for compressive strength (cylindrical with internal diameter of 4.0 6 0.01 and length of 8.0 6 0.01 mm); diametral tensile strength (cylindrical with internal diameter of 6.0 6 0.01 mm in diameter and 3.0 6 0.01 mm in depth); flexural strengths (paralelipiped with 2 mm 6 0.01 3 2 mm 6 0.01 3 25 mm 6 0.01). In all cases the filled molds with forsterite biocomposites were covered with a transparent sheet of polyester and a glass slip, which were pressed in order to remove any excess of material. By mixing paste composite A with paste composite B, free-radicals are generated and polymerization reaction starts when methacrylate groups are able to make new ACACA covalent bonds. After 7–10 min the biocomposites became solids. Cured samples of forsterite biocomposites were measured with a micrometer and the samples with a thickness higher than the chosen value were sanded, using # 800 and 1200 SiC abrasive papers until their thickness was reduced to the selected values (60.01). All forsterite biocomposites samples used in this study for mechanical test (n 5 8) were immersed in 15 mL of simulated body fluid (SBF) and stored at 37  C for 24 h before the mechanical tests, exception the samples for FS test when was test also after 28 days. SBF solution were prepared according to Kokubo’s SBF solution21 buffered at the physiological pH of 7.40 at 37  C, with tris(hydroxymethyl)amino methane and hydrochloric acid.

Characterization of nano forsterite biocomposites. All mechanical tests were carried out in a universal testing machine (LR5K Plus, Lloyd instruments. Ltd., England) at a loading rate of 0.75 mm min21 until fracture. The load deflection curves were recorded with computer software (Nexygen; Lloyd Instruments). The compressive strength value in MPa was calculated by the following Eq. (1). Compressive modulus (CM) was determined from the slope in the elastic portion of the stress-strain curve. CS ¼

F p  r2

(1)

where F is the applied load (N), r—radius of the cylindrical sample measured before testing (2 mm). The flexural strength (FS) expressed in MPa was measured using Eq. (2) and according with ISO 4049 requirements22 for dental materials testing using a 20 mm span, and the cross-sectional diameter of loading tip 2 mm. Flexural modulus (FM) expressed in GPa was determined from the slope in the elastic portion of the stress-strain curve. FS ¼

3Fmax l 2bh2

(2)

where Fmax is the applied load (N), l is the span between the supports (mm), b is the width (mm), h the thickness (mm). DTS expressed in MPa was calculated using Eq. (3): DTS ¼ 2 F=ðp d t Þ

(3)

where F is the applied load (N), d is the diameter of the cylindrical sample measured before testing (6 mm), and t is the thickness of the cylindrical sample measured before testing (3 mm). Statistical analyses Data were statistically analyzed by one-way analysis of variance (ANOVA) SPSS (Version 11.5, SPSS, Chicago, IL) software package, with Tukey’s test with the level of significance set at 0.05 in order to determine the significant differences between the mean values of the tested materials. Forsterite volume fraction from nano forsterite biocomposites The forsterite content of the biocomposites (n 5 10) was measured by combustion analysis in a furnace (HT 04 - HT 450 Furnace, Nabertherm GmbH, Germany). The test biocomposites were dried in a desiccator at room temperature for 1 day before and after combustion, and weighed on an analytical balance with an accuracy of 0.0001 g (AW 220, Shimadzu, Tokyo, Japan). The combustion of the samples was performed at 700  C for 1 h, after which the samples were removed from the oven and kept in a desiccator at room temperature for 24 h. The weight percentage (wt %) of forsterite was determined by calculating the difference between the weight of the crucible before and after ashing in air. The forsterite weight

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percent was then converted to a volume percent (vol %) using the following formula from Eq. (4). Fraction of forsterite vol % ¼

wf =df 3100 wf =df 1wr =dr

(4)

where wf and wr are the weight fractions of forsterite and resin, respectively, and df and dr are the density of the forsterite and resin, respectively. The density value of forsterite from our study was 3.25 g cm23 and for the resin mixture 1.238 g cm23.23 Bioactivity testing In vitro bioactivity of forsterite biocomposites was studied by soaking samples in SBF solution.21 The composition and procedure for preparation of SBF solution was the same with solution used for soaking of the samples for mechanical test. The bioactivities of the biocomposites were assessed by their apatite forming ability in SBF solution. A disk-shaped biocomposites (Ø 5 8 mm, h 5 1 mm) were prepared to evaluate in vitro apatite-forming on the surface of forsterite biocomposites by atomic force microscopy (AFM) and X-ray photoelectron spectroscopy (XPS). For AFM investigation were prepared two samples of P, C5, C15, C30, C50, and C70 forsterite biocomposites of 8 mm in height and 4 mm in diameter were prepared. The samples for AFM investigation were PMMA embedded in the same mould and then, they were sectioned transversely in slices of 1.5 mm thickness using a diamond saw (Isomet 1000, Buehler, USA). All samples used for AFM/XPS investigation were sanded using 800, 1200, 2400, and 4000 SiC abrasive article. The samples for AFM investigation were soaked in 15 mL SBF solution at 37  C for 14 and 28 days. In case of XPS investigation samples were soaked in SBF solution 28 days. After this time the samples were removed from the SBF solution, washed with distilled water, and stored in a desiccator prior to AFM/XPS analysis.

a scan rate of about 1 Hz. The AFM images were obtained on at least five macroscopically separated areas, randomly selected, on each sample of specimen surface. All images were processed using the standard procedures for AFM. AFM images consist of multiple scans displaced laterally from each other in y direction with 512 3 512 pixels. The surface roughness of each specimen surface, as root mean square: RMS, was also determined by using AFM software. X-ray photoelectron spectroscopy (XPS) The surface of the C50F biocomposite after 28 days of SBF storage was analyzed by XPS for evidence of precipitated apatite formation at the surface of C50F forsterite biocomposites by means of a SPECS custom build system. The excitation was made using the Mg anode as X-ray source (hm 5 1253.6 eV).

Scanning electron microscopy (SEM) The sample morphology was investigated by scanning electron microscopy (SEM Quanta 3D FEG D9399, FEI Company, the Netherlands). After the CS test, the structure of fractured surfaces of samples P and C50F forsterite biocomposite were gold sputter-coated (Bio-Rad Polaron Division SEM Coating System, Polaron Instruments, Agawan, MN).

Cell culture Cytotoxicity. Human foreskin fibroblasts (HFF) (ATCC) were cultured (37  C, 5% CO2) in DMEM High Glucose GlutaMAXmedium supplemented with 10% fetal bovine serum, 1% sodium pyruvate and 0.5% penicillin–streptomycin (all from Gibco, Life Technologies, Ghent, Belgium). Medium was changed every other day. Discs (Ø 5 4 mm, h 5 1 mm) of polymer and polymer/ forsterite in different ratios were sterilized by ethylene oxide (Maria Middelares Hospital, Ghent, Belgium). Culture medium (1.13 mL) was added to a tube containing 6 discs corresponding to a surface area/volume ratio of 2 cm2 mL21 and incubated for 7 days at 37  C on a gyratory shaker (Gerhardt, Laboshake) at a stirring speed of 50 rpm. The (diluted) extraction medium was added to a monolayer of HFF cells, seeded 24 h before at a concentration of 20,000 cells/well of a 96-well culture plate. After 48 h, the culture medium was discarded and replaced by 200 mL of a 0.5 mg mL21 MTT containing medium, followed by an incubation for 4 h at 37  C (dark). The MTT-solution was discarded and replaced by 200 mL of lysisbuffer [0.1% of TritonTM-X-100 (Fluka, Sigma–Aldrich, Bornem, Belgium) in isopropanol/0.04M HCl (UCB, Brussels, Belgium)] and incubated for 30 min at 37  C (dark) on a gyratory shaker. Absorbance is measured at 570 nm with KCjunior software on a Universal Microplate Reader EL800 (BIO-TEK Instruments, Bad Friedrichshall, Germany). The viability was calculated as percentage of the control.

Atomic force microscopy (AFM) The structural features of various samples with polished surfaces (day 0) and the same samples after storage within simulated body fluid (SBF) at 14 and 28 days were visualized by tapping mode AFM on JEOL 4210 equipment. All measurements were performed in air at room temperature from large scan area of 20 lm 3 20 lm to 1 lm 3 1 lm area. Standard cantilevers, non-contact conical shaped of silicon nitride (NSC 11, purchased from MikroMasch) were used. The sharpened tips were on cantilevers with a resonant frequency in the range of 200–300 kHz and with a spring constant of 48 N m21. AFM images were collected at

Cell viability, adhesion, and proliferation. Saos-2 cells (human osteosarcoma cells, ATCC) were cultured in a-MEM L-glutaMAX (Gibco Invitrogen) supplemented with 10% fetal bovine serum (Gibco Invitrogen) and 0.5 vol % penicillinstreptomycin (10,000 U mL21 to 10,000 mg mL21, Gibco Invitrogen). Cells were cultured at 37  C in a humidified atmosphere of 5% CO2. The polymer and polymer/forsterite biocomposite discs (Ø 5 4 mm, h 5 1 mm) were sterilized using ethylene oxide-cold cycle (Maria Middelares, Ghent, Belgium). The discs were placed into 96-well tissue culture dishes (for suspension culture). Cells were seeded at a density of 40,000 respectively 5000 cells/200 mL/disc for cell

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crystalline forsterite with an orthorhombic structure (JCPDS 71–1081). Three diffraction peaks (130), (112), and (222), which have the advantage of being well separated and which have high intensities, were chosen for the measurement. The mean crystallite size of forsterite powder, evaluated from XRD patterns using the Scherrer formula [Eq. (1)], was found to be in the range of 17–25 nm.

FIGURE 1. X-ray diffraction pattern for the nano forsterite synthesized at 900  C.

adhesion, respectively proliferation and cultured for 14 days (5% CO2/95% air, 37  C). Cell viability, adhesion and proliferation were evaluated at different time points by fluorescence microcopy (1, 7, and 14 days) and MTT-assay (1 day). MTT assay. The colorimetric MTT assay, using a 3-(4, 5-dimethyldiazol-2-yl)-2, 5-diphenyltetrazolium bromide (MTT, Merck Promega) was performed to quantify cell adhesion after 1 day on the polymer or biocomposite discs. The tetrazolium component is reduced in living cells by mitochondrial dehydrogenase enzymes into a water-insoluble purple formazan product, which can be solubilized by addition of lysis buffer and measured using spectrophotometry. The cell culture medium was replaced by 0.2 mL (0.5 mg mL21) MTT reagent and cells were incubated for 4 h at 37  C. The MTT reagent was removed and replaced by 0.2 mL lysis buffer (0.1% Triton X-100 in isopropanol/0.04 N HCl) for 30 min. A 180 mL of the dissolved formazan solution was transferred into a 96-well plate and measured spectrophotometrically at 580 nm (Universal microplate reader EL 800, Biotek Instruments). Triplicate measurements were performed. The adhesion was calculated as percentage of the control.

Particles sizes of nano forsterite Using particle size analyzer the size distribution of forsterite particles at different volumetric particle size distribution (Dv25, Dv50, and Dv90) was 17, 21, and 25 nm, respectively. TEM picture (Figure 2) of forsterite powder showed a size distribution between 10.41 and 25.88 nm with an average of about 18.68 nm SD 5 3.77 nm. Mechanical test of nano forsterite biocomposite At the 0.05 level, the values of mechanical tests for forsterite biocomposite were not significantly different (p > 0.05). The CS values (Table II) were between 143.72 and 167.49 MPa. Addition of forsterite until 50 wt % to polymer matrix led to an increase in CS values, in the following order: P < C5F < C70F < C15F < C30F < C50F. Using Tukey test showed a statistically significant difference between CS value of P samples and value of C30F and respectively C50F samples (p < 0.05). CM (Table II) increase with addition of forsterite from 1.67 to 2.75 GPa. Statistically significant differences were found when FM values of P, C5F, C15F, and C30F materials were compared with C50F and C70F (p < 0.05). DTS test showed values between 25.45 and 31.55 MPa (Table II) and a statistically significant difference was registered between C50F and C70F (p < 0.05). FS (Table II) ranged from 59.47 to 83.20 MPa after 24 h and decrease after 28 days between 50.04 and 77.73 MPa. The highest FS value of 83.20 MPa was registered for C50F biocomposites. FM (Table II) was in the range from 2.47 to 8.60 GPa after 1 day in SBF and 2.22–7.37 GPa after 28 days in SBF for forsterite biocomposite. The trends in CM values were similar to those seen in the FM, and they may all be related to the increases of filler content within composite materials.

Fluorescence microscopy. To visualize cell viability, adhesion and proliferation on the discs, cell/disc constructs were evaluated using fluorescence microscopy after performing live/ dead staining. After rinsing with PBS, the supernatant was replaced by 1 mL PBS solution supplemented with 2 mL (1 mg mL21) calcein AM (Anaspec, USA) and 2 mL propidium iodide (1 mg mL21) (Sigma–Aldrich). Cultures were incubated for 10 min at room temperature, washed twice with PBS solution and evaluated by fluorescence microscopy (Olympus inverted Research System Microscope, type U-RFLT, CellM software, Olympus, Belgium). Evaluations were done post-seeding at days 1, 7, and 14. RESULTS

X-ray diffraction of nano forsterite Figure 1 illustrates the XRD pattern of forsterite powder calcined at 900  C for 2 h. The X-ray diffraction revealed a

FIGURE 2. TEM image of nano forsterite powder.

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TABLE II. The Results of Percentage of Filler by Volume and Mechanical Properties for Nano Forsterite Biocomposites

Nr

Material

vol %

1

P

0

2

C5F

2.04

3

C15F

6.01

4

C30F

14.03

5

C50F

26.95

6

C70F

45.11

CS (MPa) SD a

128.70 (12.90) 143.72 (24.26) 149.18 (17.98) 162.33a (23.34) 167.49a (10.15) 147.12 (20.84)

CM (MPa) SD a

1.49 (0.20) 1.67b (0.13) 1.69c (0.11) 1.75d (0.16) 2.34a,b,c,d,e (0.18) 2.75a,b,c,d,e (0.23)

DTS (MPa) SD 29.72 (2.70) 27.33 (2.58) 28.45 (4.54) 30.25 (3.54) 31.55a (2.75) 25.45a (2.54)

FS (MPa) SD After 24 h a

80.55 (12.51) 70.54c (6.78) 70.92e (4.44) 76.56g (5.61) 83.20c,f,h,i (6.55) 59.47a,b,g,i (9.81)

FM (GPa) SD

After 28 Days b

77.73 (20.36) 69.94d (9.35) 68.18f (10.13) 64.22a.h (6.92) 70.70j (9.14) 50.04a,b,c,d,e,f,g,i,j (6.73)

After 24 h a

2.05 (0.41) 2.47c (0.24) 2.68e (0.30) 3.75a,b,c,d,g (0.55) 5.67a,b,c,d,e,f,g,h,i (0.75) 8.60a,b,c,d,e,f,g,h,i,j (1.39)

After 28 Days 1.94b (0.60) 2.22d (0.31) 2.56f (0.40) 3.58a,b,d,h (0.56) 4.88a,b,c,d,e,f,h,j (1.20) 7.37a,b,c,d,e,f,g,h,i,j (1.85)

Note: vol %: percentage of filler by volume; CS: compressive strength; CM: compressive modulus; DTS: diametral tensile strength; FS: flexural strength; FM: flexural modulus; SD: standard deviation; a–i: superscript letters within columns indicate mean values not statistically significant different from each other, when compared using the Tukey test, p > 0.05.

CS, CM, and FM values were correlated with forsterite volume fraction (%) from composites. Figure 3 shows that there could be established a linear dependence between the percentage of fillers and CS, CM, and FM of forsterite resin composites (r2CS 5 0.8176, r2CM 5 0.9659, r2FM 5 0.9944, r2FM 5 0.9967). SEM Figure 4(a–c) show micrographs of the surface for fractured C5F, C50F, and C70F biocomposites after compression test. Increasing of forsterite content change topography of fractured surface from smooth in C5F to smaller fractured fragments in C50F and C70F. Forsterite powder had the effect of reinforcing composites and will stop propagation of fracture inside of composite. AFM Topographic AFM images are given in Figure 5 for all samples: P (Figure 5.1), C5F (Figure 5.2), C15F (Figure 5.3), C30F (Figure 5.4), C50F (Figure 5.5) and C70F (Figure 5.6) taking into consideration the amount of forsterite in the

polymer matrix within the biocomposites and the storage time in SBF: initial polished samples [Figure 5(a)], after 14 days [Figure 5(b)] and 28 days [Figure 5(c)] immersion in SBF solution. Surface roughness values (RMS) derived from AFM measurements are presented in Table III. X-ray photoelectron spectroscopy (XPS) XPS was used for qualitative and quantitative analysis of sample surface. The spectra of Ca 2p, Cl 2p, P 2s and P 2p core-level lines were recorded and quantified. For instance in Figure 6(a) it is presented the core-level XPS spectrum of Ca 2p and its corresponding deconvolution. In Figure 6(b) the Cl 2p and P 2s core-levels spectra are shown together with their deconvolutions while Figure 6(c) shows the deconvolution of P 2p core-level line. Two chlorine positions namely Cl (A) and Cl (B) were detected. Regarding the XPS spectra of both phosphorous 2s and 2p core-levels two positions, labeled P (A) and P (B), are also evidenced. One can observe that the absolute intensities of (A) and (B) features from P 2s and P 2p spectra, respectively, are consistent to each other. The doublet separations used to fit the

FIGURE 3. (a) Correlation of compressive strength (CS) or compressive modulus (CM) with filler volume fraction of forsterite biocomposites; (b) Correlation of flexural modulus (FM) and filler volume fraction of forsterite in biocomposites. (Different letters denote statistically significant differences from each other by Tukey test p < 0.05).

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FIGURE 4. SEM micrographs of fractured surface of biocomposites after CS test. (a) C5F biocomposite; (b) C50F biocomposite; arrows fracture propagated inside of biocomposite, (c) C70F biocomposite; arrows forsterite–polymer interface.

peaks were 3.5, 1.7, and 0.8 eV for Ca 2p, Cl 2p, and P 2p lines, respectively. The lines intensities were normalized by dividing with real sensitivity, transmission and mean free path factors as given in the CASA software database. The calibration was made by using C 1s from CAC, CAH bindings positioned at 284.6 eV. The results are summarized in Table IV presented in the Supporting Information. Biocompatibility Cytotoxicity. The viability of HFF cells in the presence of extraction media of the discs was determined by MTT-assay. Discs were extracted for 7 days. At the highest surface area/volume ratio (1 cm2 mL21), the viability of the cells was minimal 70% for all the discs, independent of the forsterite ratio (data not shown). Cell viability, adhesion, and proliferation. The amount of viable cells attached on the polymer and polymer/forsterite biocomposite discs after 1 day, as determined by the MTT assay, are shown in Figure 7. After 1 day, the amount of cells on the polymer is 57.8% 6 24.4%. The amount of viable cells on the discs increased after incorporation of forsterite toward minimum 58.0% 6 6.7% or maximum 82.6% 6 6.6%. Adherent cells cultured for up to 14 days on the discs were examined by fluorescence microscopy after live/dead staining (Figure 8). One day after seeding, less live (green) cells were detected on the polymer disc. A majority of the cells were round, not well spread and dead (red) [Figure 8(a,a’)]. In contrast, cells on the polymer/forsterite biocomposite discs, independent of the forsterite ratio, were well spread and showed a polygonal morphology [Figure 8(b–f,b’–f’)]. Details of the proliferating cells on the discs after 7 days are shown in Figure 8(g’–l’). After 14 days, the complete surface of the polymer/forsterite biocomposite discs was covered with viable, well attached cells [Figure 8(h–l)] in contrast to the polymer, which only showed a limited amount of cells [Figure 8(g)]. FIGURE 5. AFM images: 2D-topographies on polished surfaces for polymer (P, noted 1) and forsterite biocomposites [sample code is given in Table I]; C5F, (2); C15F, (3); C30F, (4); C50F, (5); and C70F, (6); before (a) 0 day and after their immersion in SBF for different time: (b) 14 days and (c) 28 days; scanned area 2.5 lm 3 2.5 lm.

DISCUSSION

Forsterite fillers is known to have bioactivity properties in the literature16–19,24 and was synthesized in our study in order to obtain new biocomposites. Because the topic of

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FIGURE 6. XPS spectra of (a) Ca 2p; (b) Cl 2p together with P 2s and (c) P 2p core-levels together with their corresponding deconvolutions and assignments in case of C50F forsterite biocomposite after 28 days of immersion in SBF.

composites based forsterite filler was not published we tried to compare with Hap composites or other materials from the biomedical market in order to present a image for the comparative properties of new biocomposites obtained in this study. Increasing the quantity of Hap filler in the composites showed an increase of biocompatibility.25 Our materials were designed by addition of nano forsterite powder (5, 15, 30, 50, and 70 wt %) in order to see improving biological properties and mechanical properties. In vivo, forces applied to a bone replacement material could generate crack initiation, propagation, and the fracture initiation in materials may be a critical factor in failure of the implant.26 The influence of the quantity of Hap fillers on the mechanical properties of bone cements is not clear regarding to different publications. The change of CM was not significant by the addition of 14% Hap to the cement.27 Kwon et al.7 showed that addition of 30% Hap to bone cement increases the push-out strength after implanting into the distal end of rabbit femora. In our study forsterite filler is mechanicaly immobilized into the polymer matrix and has a positive effect in stopping crack propagation inside the polymer matrix. This is why the CS increases with increasing forsterite content up to 50 wt % Above 70 wt % the mechanical properties of the composite decrease. Because forsterite acts as rigid filler, the use of much higher filling ratios may produce phase segregation thus determining more stress fractures into the polymer matrix. This observation is in agreement with other studies.27,28 Kurtz et al.29

R P bone cement showed CS values of 97 MPa for SimplexV (containing 10% BaSO4 radiopacifier), (Stryker Orthopedics, Mahwah, NJ) and 111 MPa KyphX HV-R bone cement (containing 30% BaSO4 radiopacifier), (KyphX HV-R, Kyphon Inc.). A slight increase in the mean CS after 30 days than after 1 day of storage in SBF were observed from 99 MPa R P and 100 to 113 MPa for (1 day) to 113 MPa for SimplexV 30 R R (OrthoSimplexV. The compressive strength of CortossV vita) showed a significant decrease with maturation (179– 91 MPa, from 1 to 30 days).30 CM of forsterite biocomposites increased with increasing addition of nano forsterite fillers from 5 to 70 wt % The best evidence was obtained for C50F and C70F when CM was improved with 0.84 GPa and respectively 1.25 GPa more than P samples. As seen in Figure 3(a) strong correlation was found between CS and nano forsterite volume fraction (r2 5 0.8176) respectively, as well as between CM and nano forsterite volume fraction (r2 5 0.9659). Figure 3 showed a strong correlation between FM after 24 h and 28 days with filler volume-fraction (r2 was between 0.9944, respectively 0.9967). It increase progressively and almost linearly from the unfilled to less filled (2.04 vol % forsterite) and to the highest filled polymer biocomposites (45.11 vol % forsterite). This behavior is in accordance with another publication.31 Values registered from CS test (Table II) are also in

TABLE III. Surface Roughness (Root Mean Square: RMS) for Polymer and Forsterite Biocomposites, Before, and After Storage in SBF for Different Time Roughness (RMS) of Surface (nm 6 SD nm) Code

0 day

14 days

28 days

P C5F C15F C30F C50F C70F

18 6 2 17 6 3 21 6 2 20 6 3 15 6 2 18 6 4

19 6 3 47 69 37 6 5 47 6 8 42 6 8 46 6 3

39 6 8 49 6 9 38 6 5 51 6 9 47 6 6 54 6 7

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FIGURE 7. Percentage of viable Saos-2 cells cultured on polymer/forsterite biocomposite discs after 1 day relative to the control (tissue culture polystyrene). The amount of viable cells was quantified with the MTT assay. Mean and SD, n 5 3.

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FIGURE 8. Cell viability and proliferation on polymer and polymer/forsterite biocomposite discs. Fluorescence microscopy (CaAM/PI staining) of Saos-2 cells cultured on the discs for 1 day (a–f, a’–f’), 7 days (g’–l’) and 14 days (g–l). P (a, a’, g, g’), C5F (b, b’, h, h’), C15F (c, i, i’), C30F (d, d’, j, j’), C50F (e, e’, k, k’), C70F (f, f’, l, l’).

accordance with other studies for direct core buildups used in dentistry: 166.6–331.7 MPa32 and 61.1–250 MPa33 respectively. In our study FS (Table II) decreases at 11.95% for C5F-C15F samples but slightly increases for C50F. C70F

composite was brittle, having a decreased FS with 28.53% (after 24 h in SBF) and 29.23% (after 28 days in SBF) as compared to C50F composites. This could be explained by high filler load and decreasing the degree of cross-linking monomers in polymer matrix and around nano forsterite particles. In opposite to this, investigation of FM (Table II) for C70F composites showed an increase of 34.04% as compared to C50F composites and 76.12% with respect to P sample. After 28 days of SBF storage for samples, the water sorption combined with the solubility of the material will increase the plasticity of the polymer matrix and degradation of the interface filler-polymer matrix all will decrease FS and FM. These results are in accordance with other studies.9,32,34–36 Addition of silica nanoparticles (20, 30, 40, 50, and 60 wt %) to a polymer matrix bis-GMA/TEGDMA showed an increase of FS between 103.41 and 149.74 MPa.37 The highest FS value was obtained when composites were loaded it with 40 wt % silica nanoparticles beyond that FS decreases sharply when the mass fraction was further increased to 50 wt %37 The lower FS values in our study were lower than the study of Hosseinalipour et al.37 and could be explained by the higher mechanical properties of silica and silane A-174 treatment at the surface of silica. The FS results were in agreement to other results: direct core buildups 15.8–165.8 MPa,32 flowable composites 62.8– 133.0 MPa.38 Flowable composites from the market showed FM between 1.6 and 6.4 GPa.38 DTS values showed a slower increase for C30F and C50F than P sample (Table II). The increasing nano forsterite filler at 70%wt. showed a high decrease of DTS values of 19.33% for C70F than C50F (p < 0.05). The DTS results were in agreement with other results for direct core buildups 9.6–52.2 MPa,32 18.3–55.1 MPa,33 flowable composites 28.4–53.6 MPa.38 One explanation for the higher mechanical properties of dental composites than our materials could be explained by using glass fillers and silane A-174 coating able to bond by the hydroxyl group to the filler and by the methacrylate group to polymer matrix. If Hap filler is not silane-functionalized to be chemically bound to the polymer matrix, it will act as rigid filler inside of composite and slightly improve the fracture resistance, flexural modulus and yield strain, up to a certain content.39 This behavior is the same as in our materials because forsterite does not have chemical bonds with the polymer matrix. Figure 4(a) shows micrographs of the surface for fractured C5F biocomposite after compression test. The surfaces of C5F biocomposite after fracture show less roughness and smaller fractured components of polymer at the surface than in C50F biocomposite [Figure 4(b)]. The rough surface of the C50F fractured sample shows some small components of composite. In Figure 4(b), the arrows showed some fractures propagated inside the biocomposite bulk. Forsterite powder had the effect of reinforcing in composites and will try to stop propagation of fracture inside of biocomposite. These could explain the higher CS for C50F than P sample. In Figure 4(c), arrows show particles of nano forsterite inside of cross-linking polymer matrix of biocomposites. Fracture was propagated also at the interface polymer-

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forsterite particle and that could be explained by the lack of a chemical bond between powder and polymer matrix. AFM images, given in Figure 5, reveal the surface features for all samples, namely for polymer (Figure 5.1) and forsterite biocomposites (Figures 5.226). The morphological changes depend on the amount of forsterite within each sample [Figures 5.2(a)–6(a)] and the storage time in SBF [Figure 5(b,c)]. From AFM images, it can be noticed that the polymer sample shows its initial morpholology [Figure 5.1(a)] even after 28 days of immersion [Figure 5.1(c)] in SBF solution. A small increase in surface roughness was determined after 14 days in SBF (Table III). A significant increase in RMS is observed after 28 days in SBF (Table II), particularly due to long term contact of P sample with the SBF. However, the P sample does not generate new features regarding the surface morphology confirming the assumption that forsterite particles might generate new structures in contact with SBF solution. Figures 5.2(a)–6(a) shows the surface characteristics of polymer/forsterite biocomposites before immersion into SBF. The morphology is similar for all biocomposites due to immobilized forsterite particles in polymer matrix. For instance, polymer fragments are observed on the surface of C5F biocomposite alternating with forsterite filler bonded in polymer matrix [Figure 5.2(a)]. With increasing amount of forsterite in biocomposites [Figure 5.3(a)–6(a)], a rather uniform granular aspect is still observed indicating a good homogeneity of samples and a low agglomeration tendency of nano forsterite particles. Even for C70F biocomposites [Figure 5.6(a)] a compact granular structure is formed because forsterite nanoparticles are well connected with the polymer matrix. After 14 days of storage in SBF solution, the polymer sample does not show any morphological changes [Figure 5.1(b)]. The forsterite biocomposites reveal new fibrous structures [Figure 5.2(b)–6(b)] made due to the interaction between forsterite nano particles embeded into polymer matrix with the SBF solution. Generally, these fibers have a diameter between 30 and 50 nm, and a length ranging between 200 and 500 nm. In case of C5F biocomposite, only few fibrous structures were identified [Figure 5.2(b)]. If forsterite content is progressively increased, the multiple fibrous structures are in consequence observed in the AFM images [Figure 5.3(b)–6(b)]. The appearance of fibrous structures on the surface of biocomposites causes a significant increase of the surface roughness as seen in Table III. After 28 days in SBF, the polymer sample shows no morphological changes, but a significant increase in surface roughness is determined. Most likely the SBF started to deepen some imperfections on the surface of the polymer and a slight erosion took place. In contrast to these findings, the forsterite biocomposites promoted multiple fibrous structures [Figure 5.2(c)–6(c)] even for samples with little amount of forsterite (C5F and C15F biocomposites, Figures 3(c)–5.2(c)]. Further, the C30F, C50F, and C70F biocomposites with forsterite content between 30-70%wt generated fibrous structures with preferred orientation [Figure 5.4(c)– 6(c)]. The average diameter of fibers is around 80 nm and their length is ranging from 500 to 1000 nm. As observed

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by AFM investigations, by increasing the nano fosterite content in the polymer matrix the thickness of the fibrous layer also increases. We choose to investigate the surface of C50F kept for 28 days in SBF by XPS since, according to its determined mechanical properties and AFM results, it appears to be the best material for biomedical applications. By analyzing the absolute XPS line intensities, as shown in Table IV of Supporting Information, one can observe that, in case of P 2p corelevel lines, the peaks labeled (A) may be associated with the formation of hydroxyapatite Ca10(PO4)6(OH)2 with a Ca excess content. On the other hand, it is known that, in many cases, the resulting apatite compounds are calcium deficient with Ca/P ratios as: 1.06, 1.27, or 1.6.40–42 It appears that the total amount of calcium at the sample surface is divided between two compounds belonging to the apatite family. By analyzing Figure 6(c) it can be seen that, besides the most intense doublet peaks (A), an additional P 2p core-level doublet feature, marked as (B) also appears in the spectrum. It is positioned at higher binding energies namely 138.07 eV and 138.94 eV for (3/2) and (1/2) peaks respectively. It is an indication that phosphorus atoms corresponding to (B) doublet are positioned into a more electronegative neighborhood as compared to hydroxyapatite. Their intensity (Table IV of Suplementary material) multiplied by 3 mach the Cl 2p (A) doublet intensity as appearing in Figure 6(b) thus revealing, besides the formation of hydroxyapatite associated to P 2p core-level doublet (A), the formation of chlorapatite Ca10(PO4)6(Cl)2 as a second apatite type compound. The higher binding energies of phosphorus (B) are dues to the presence of more electronegative chlorine neighbors. Under this consideration the ratio between calcium absolute intensity and the sum absolute intensities of both (A) and (B) phosphorous features is 1.51 and corresponds to the values expected in case of Ca deficient apatite. The resulting molar ratio between hydroxyapatite and chlorapatite is 7.2. A small amount of chlorine, (B) labeled peaks in Figure 6(b), still remains unassigned thus indicating the presence of small amounts of some other insoluble Cl compounds. This was in agreement with the other studies regarding the formation of apatite layer to forsterite19,24,43,44 with possible application in bone tissue engineering.43 In particular, C50F and C70F biocomposites showed a preferred alignment of apatite which is generally found in bone and calcified tendon.45 Moreover, these results clearly suggest that the formation of the fibrous apatite might be a predictor of the bioactivity of forsterite biocomposites in vivo. To reach this goal, as a first step the viability of osteoblasts on these biocomposites is tested and further interpreted. Multiple fibrous structures and surface roughness results obtained from AFM images [Figure 5(b,c)] after SBF storage and based on XPS results regarding the apatite layer formation at the surface of C50F biocomposites, we could conclude that the topography changes observed in Figure 5(b,c) are due to modifications of the apatite layer. Bioactivity of biocomposite in SBF represents a criterion of bone bonding ability in body environment.21 This means that osteoblasts might preferentially proliferate and differentiate

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to produce apatite and collagen on the surface of the replacement material.21 As a consequence, in vivo we could obtain a strong chemical bonding through the apatite layer at interface between replacement material and bone. The apatite layer hypothesis could be explained by the contact of forsterite particles from the composite surface with SBF solution. Nano forsterite particles will be able to be dissolved and to send Mg21 ions into the SBF forming silanol groups at its surface. Further, the silanol groups from could act as a nucleation sites46,47 for the ions of SBF solution. Thus, negatively charged silanol groups will first chelate the Ca21 ions from the SBF solution. Then, by further adsorbing phosphates as well as hydroxyl, carbonate, and Ca21 ions a cluster of critical size will be formed. Moreover, the threedimensional cluster could act as the nucleus for the formation of a new apatite crystal. After the formation of the first apatite crystals, there will be a spontaneous growth of apatite layer by using ions from SBF solution.46,47 Therefore, by increasing the addition of nano forsterite into the biocomposite the bioactivity in vitro and in vivo will be increased. Lastly, the increase of the apatite layer thickness will create a stronger bond at the interface with the material-bone. The biocompatibility of forsterite biocomposites did not alter with increasing nano forsterite filler content (from 5 to 70 wt %). A viability of minimum 70 wt %, independent of the nano forsterite ratio, was obtained when cells were cultured in extraction medium (data not shown). After 1 day, cell adhesion was improved on the nano forsterite biocomposites. After 7 and 14 days, the nano forsterite biocomposites (5–70 wt %) were completely covered with viable, well attached cells. Our results are in agreement with Lopes et al.48 who revealed an increased osteoblast cytocompatibility on silicate glass/PMMA composites. However, their positive effect decreased with increase in the percentage of incorporated glass. Their results were attributed to the release of Si ions in the culture medium at high silicate glass/PMMA ratios. In our study, we did not observe a decrease in cell viability, suggesting that no forsterite is released in the culture medium and proven by the biocompatibility assay of cells in contact with a 7 days old extraction media. CONCLUSIONS

The results reveal that the mechanical properties CS, FS, DTS increase with addition of nano forsterite filler up to 50 wt % and CM and FM up to 70 wt % forsterite. From this work we can conclude that C50F biocomposite could be used as bone cement for orthopedic application or as dental cement, direct core buildups and flowable composites for dental restoration. AFM, and XPS investigations confirm that forsterite filler induces biomineralization of hydroxyapatite as well as small quantities of chlorapatite forming Ca deficient materials (Ca/P 5 1.51) at the surface of new forsterite biocomposites. Biological tests showed an improved cell adhesion after addition of forsterite powder, that confirm possible applications of forsterite biocomposites for biomedical fields.

ACKNOWLEDGMENTS

The authors also thank UEFISCDI for financial support through grant numbers: 171/2012 and 189/2012. The authors thank the COST Action MP1301 for COST meeting support. REFERENCES 1. Charnley J, McKee GK, Coltart WD, Scales JT. Arthroplasty of the hip by the low friction technique. J Bone Joint Surg 1961;43B: 601. ndez L, Gurruchaga M, Gon ~ i I. Injectable acrylic bone 2. Herna cements for vertebroplasty based on a radiopaque hydroxyapatite. Formulation and rheological behaviour. J Mater Sci Mater Med 2009;20:89–97. 3. Bowen RL. Dental filling material comprising vinyl silane treated fused silica and a binder consisting of the reaction product of bisphenol and glycidyl acrylate. US Patent 3066112, Nov. 27, 1962. 4. Szpalski M, Descamps PY, Hayez JP, Raad E, Gunzburg R, Keller TS, Kosmopoulos V. Prevention of hip lag screw cut-out by cement augmentation: Description of a new technique and preliminary clinical results. J Orthop Trauma 2004;18:34–40. 5. Saito M, Maruoka A, Mori T, Sugano N, Hino K. Experimental studies on a new bioactive bone cement: Hydroxyapatite composite resin. Biomaterials 1994;15:156–160. 6. Ikeda D, Saito M, Murakami A, Shibuya T, Hino K, Nakashima T. Mechanical evaluation of a bio-active bone cement for total hip arthroplasty. Med Biol Eng Comput 2000;38:401–405. 7. Kwon SY, Kim YS, Woo YK, Kim SS, Park JB.Hydroxyapatite impregnated bone cement: In vitro and in vivo studies. Biomed Mater Eng 1997;7:129–140. 8. Hench LL. Bioactive ceramics: Theory and clinical applications. In bioceramics, 7th ed. Anderson OH, Yli-Urpo A, editors. Butterworth-Heinemann Ltd; Oxford, 1994. pp 3–14. 9. Shinzato S, Kobayashi M, Mousa WF, Kamimura M, Neo M, Kitamura Y, Kokubo T, Nakamura T. Bioactive polymethyl methacrylate-based bone cement: Comparison of glass beads, apatite- and wollastonite-containing glass-ceramic, and hydroxyapatite fillers on mechanical and biological properties. J Biomed Mater Res 2000;51:258–272. 10. Huan Z, Chang J. Novel bioactive composite bone cements based on the beta-tricalcium phosphate-monocalcium phosphate monohydrate composite cement system. Acta Biomater 2009;5:1253– 1264. 11. Renno AC, van de Watering FC, Nejadnik MR, Crovace MC, Zanotto ED, Wolke JG, Jansen JA, van den Beucken JJ. Incorporation of bioactive glass in calcium phosphate cement: An evaluation. Acta Biomater 2013;9:5728–5739. 12. Lopes P, Corbellini M, Ferreira BL, Almeida N, Fredel M, Fernandes MH, Correia R. New PMMA-co-EHA glass-filled composites for biomedical applications: Mechanical properties and bioactivity. Acta Biomater 2009 5:356–362. 13. Mota J, Yu N, Caridade SG, Luz GM, Gomes ME, Reis RL, Jansen JA, Walboomers XF, Mano JF. Chitosan/bioactive glass nanoparticle composite membranes for periodontal regeneration. Acta Biomater 2012;8:4173–4180. 14. Furtos G, Tomoaia-Cotisel M, Baldea B, Prejmerean C. Development and characterization of new AR glass fiber reinforced cements with potential medical applications. J Appl Polym Sci 2013;15:1266–1273. 15. Suchanek W, Yoshimura M. Processing and properties of hydroxyapatite-based biomaterials for use as hard tissue replacement implants. J Mater Res 1998;13:94–117. 16. Kharaziha M, Fathi MH. Synthesis and characterization of bioactive forsterite nanopowder. Ceram Int 2009;35:2449–2454. 17. Ni S, Chou L, Chang J. Preparation and characterization of forsterite (Mg2SiO4) bioceramics. Ceram Int 2007;33:83–88. 18. Kharaziha M, Fathi MH. Improvement of mechanical properties and biocompatibility of forsterite bioceramic addressed to bone tissue engineering materials. J Mech Behav Biomed Mater 2010; 3:530–537.

JOURNAL OF BIOMEDICAL MATERIALS RESEARCH B: APPLIED BIOMATERIALS | MONTH 2015 VOL 00B, ISSUE 00

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19. Tavangarian F, Emadi R. Improving degradation rate and apatite formation ability of nanostructure forsterite. Ceram Int 2011;37: 2275–2280. 20. Klug HP, Alexander LE. X-ray Diffraction Procedures. New York: Wiley; 1954. 21. Kokubo T, Takadama H. How useful is SBF in predicting in vivo bone bioactivity? Biomaterials 2006;27:2907–2915. 22. ISO standard 4049. Dentistry–polymer-based filling, restorative and luting materials. International Organization for Standardisation. Printed in Switzerland; 2000. pp 15. 23. Garoushi SK, Lassila LV, Vallittu PK. Short fiber reinforced composite: The effect of fiber length and volume fraction. J Contemp Dent Pract 2006;7:10–17. 24. Naghiu MA, Gorea M, Mutch E,.Kristaly F, Tomoaia-Cotisel M. Forsterite nanopowder: Structural characterisation and biocompatibility evaluation. J Mater Sci Technol 2013;29:628–632. 25. Di Silvio L, Dalby MJ, Bonfield W. Osteoblast behaviour on HA/PE composite surfaces with different HA volumes. Biomaterials. 2002;23:101–107. 26. Molino LN, Topoleski LD. Effect of BaSO4 on the fatigue crack propagation rate of PMMA bone cement. J Biomed Mater Res 1996;31:131–137. 27. Serbetc¸i K, Korkusuz F, Hasirci N. Mechanical and thermal properties of hydroxyapatite-impregnated bone cement. Turkish J Med Sci 2000;30:543–549.  pez JM, 28. Vallo CI, Montemartini PE, Fanovich MA, Porto Lo Cuadrado TR. Polymethylmethacrylate-based bone cement modified with hydroxyapatite. J Biomed Mater Res 1999;48:150–158. 29. Kurtz SM, Villarraga ML, Zhao K, Edidin AA. Static and fatigue mechanical behavior of bone cement with elevated barium sulfate content for treatment of vertebral compression fractures. Biomaterials 2005;26:3699–3712. 30. Boyd D, Towler MR, Wren A, Clarkin OM. Comparison of an experimental bone cement with surgical Simplex P, Spineplex and Cortoss. J Mater Sci Mater Med 2008;19:1745–1752. 31. Masouras K, Silikas N, Watts DC. Correlation of filler content and elastic properties of resin-composites. Dent Mater 2008;24:932– 939. 32. Saygili G, Mahmali SM. Comparative study of the physical properties of core materials. Int J Periodontics Restorative Dent 2002; 22:355–363. 33. Cho GC, Kaneko LM, Donovan TE, White SN. Diametral and compressive strength of dental core materials. J Prosthet Dent 1999; 82:272–276. 34. Harper EJ, Behiri JC, Bonfield W. Flexural and fatigue properties of a bone cement based upon polyethylmethacrylate and hydroxyapatite. J Mater Sci Mater Med 1995;6:799–803.

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FURTOS ET AL.

35. Hamizah AS, Mariatti M, Othman R, Kawashita M, Noor Hayati AR. Mechanical and thermal properties of polymethylmethacrylate bone cement composites incorporated with hydroxyapatite and glass-ceramic fillers. J Appl Polym Sci 2012;125:E661–E669  lu ND, Alkurt M. Evaluation of the flexural 36. Duymus ZY, Yanikog strength of dual-cure composite resin cements. J Biomed Mater Res B Appl Biomater 2013;101:878–881. 37. Hosseinalipour M, Javadpour J, Rezaie H, Dadras T, Hayati AN. Investigation of mechanical properties of experimental Bis-GMA/ TEGDMA dental composite resins containing various mass fractions of silica nanoparticles. J Prosthodont 2010;19:112–117. 38. Ilie N, Hickel R. Investigations on mechanical behaviour of dental composites. Clin Oral Investig 2009;13:427–438. 39. Montemartini P, Cuadrado T, Frontini P. Fracture evaluation of acrylic bone cements modified with hydroxyapatite: Influence of the storage conditions. J Mater Sci Mater Med 1999;10:309–317. 40. Vandiver J, Patel N, Bonfield W, Christine O. Nanoscale morphology of apatite precipitated onto synthetic hydroxyapatite from simulated body fluid. Key Eng Mater 2005;284/286:497–500. 41. Radin SR, Ducheyne P. The effect of calcium phosphate ceramic composition and structure on in vitro behavior. II. Precipitation. J Biomed Mater Res 1993;27:35–45. 42. Kanzaki N, Onuma K, Ito A, Teraoka K, Tateishi T, Tsutsumi S. Direct growth rate measurement of hydroxyapatite single crystal by moire phase shift. J Phys Chem B 1998;102:6471–6476. 43. Diba M, Kharaziha M, Fathi MH, Gholipourmalekabadi M, Samadikuchaksaraei A. Preparation and characterization of polycaprolactone/forsterite nanocomposite porous scaffolds designed for bone tissue regeneration. Compos Sci Technol 2012;72:716– 723 44. Tavangarian F, Emadi R. Nanostructure effects on the bioactivity of forsterite bioceramic. Mater Lett 2011;65:740–743 45. Wenk HR, Heidelbach F. Crystal alignment of carbonated apatite in bone and calcified tendon: Results from quantitative texture analysis. Bone 1999;24:361–369. 46. Rhee SH, Lee SJ, Tanaka J. Synergistic effect of silanol group and calcium ion in chitosan membrane on apatite forming ability in simulated body fluid. J Biomater Sci Polym Ed 2006;17:357–368. 47. Lee KH, Rhee SH. The mechanical properties and bioactivity of poly(methyl methacrylate)/SiO2–CaO nanocomposite. Biomaterials 2009;30:3444–3449. 48. Lopes P, Garcia M, Fernandes M, Fernandes M. Properties and osteoblast cytocompatibility of self-curing acrylic cements modified by glass fillers. J Biomater Appl 2012;22:1–16.

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