Gelatine modified monetite as a bone substitute material: An in vitro assessment of bone biocompatibility

Accepted Manuscript Full length article Gelatine modified monetite as a bone substitute material: An in vitro assessment of bone biocompatibility Benjamin Kruppke, Jana Farack, Alena-Svenja Wagner, Sarah Beckmann, Christiane Heinemann, Kristina Glenske, Sina Rößler, Hans-Peter Wiesmann, Sabine Wenisch, Thomas Hanke PII: DOI: Reference:

S1742-7061(15)30272-5 http://dx.doi.org/10.1016/j.actbio.2015.12.035 ACTBIO 4049

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Acta Biomaterialia

Received Date: Revised Date: Accepted Date:

17 September 2015 25 November 2015 24 December 2015

Please cite this article as: Kruppke, B., Farack, J., Wagner, A-S., Beckmann, S., Heinemann, C., Glenske, K., Rö ßler, S., Wiesmann, H-P., Wenisch, S., Hanke, T., Gelatine modified monetite as a bone substitute material: An in vitro assessment of bone biocompatibility, Acta Biomaterialia (2015), doi: http://dx.doi.org/10.1016/j.actbio. 2015.12.035

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Benjamin Kruppke Revision_Acta_Biomaterialia_PPGC-Manuscript_Kruppke_Dresden_Text Only

26.12.2015

Gelatine modified monetite as a bone substitute material: An in vitro assessment of bone biocompatibility

Benjamin Kruppke1*, Jana Farack1, Alena-Svenja Wagner2, Sarah Beckmann1, Christiane Heinemann1, Kristina Glenske2, Sina Rößler1, Hans-Peter Wiesmann1, Sabine Wenisch2 and Thomas Hanke1

1

Max Bergmann Center of Biomaterials and Institute of Materials Science, Technische

Universität Dresden, 01069 Dresden, Germany 2

Department of Veterinary Clinical Sciences, Small Animal Clinic c/o Institute of Veterinary

Anatomy, Histology and Embryology, Justus-Liebig-University Giessen, 35392 Giessen, Germany

*

Corresponding author. Tel.: +49 351 463 42762; fax: +49 351 463 39401. Max-Bergmann-

Zentrum für Biomaterialien, Budapester Strasse 27, 01069 Dresden, Germany, E-mail address: [email protected] (B. Kruppke).

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Abstract Calcium phosphate phases are increasingly used for bone tissue substitution, and the load bearing properties of these inherently brittle biomaterials are increased by inclusion of organic components. Monetite prepared using mineralization of gelatine pre-structured through phosphate leads to a significantly increased biaxial strength and indirect tensile strength compared to gelatine-free monetite. Besides the mechanical properties, degradation in physiological solutions and osteoblast and osteoclast cell response were investigated. Human bone marrow stromal cells (hBMSC) showed considerably higher proliferation rates on the gelatine modified monetite than on polystyrene reference material in calcium-free as well as standard cell culture medium (α-MEM). Osteogenic differentiation on the material was comparable to polystyrene in both medium types. Osteoclast-like cells derived from monocytes were able to actively resorb the biomaterial. Osteoblastic differentiation and perhaps even more important the cellular resorption of the biomaterial indicate that it can be actively involved in the bone remodelling process. Thus the behaviour of osteoblasts and osteoclasts as well as the adequate degradation and mechanical properties are strong indicators for bone biocompatibility, although in vivo studies are still required to prove this. Keywords: Bone; Biocompatibility; Gelatine; Bone Marrow Stromal Cells; Osteoclast; Biomimetic Material

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Introduction

Biocompatibility is a material property that is closely linked to the target tissue in question, and both biomaterial function and host tissue response have to be taken into account here. For this reason the term “bone biocompatibility” was chosen to emphasize the importance of material and target tissue interaction [1,2]. In an in vitro study of bone biocompatibility, mechanical properties and unspecific degradation of the material have to be investigated as well as osteoblast and osteoclast cell response due to the involvement of the material in the bone remodelling process. From a functional point of view, the best suited bone substitute materials and their degradation products should thus be similar to naturally occurring ones; from a practical point of view, on the other hand, the material should be easy to use, meaning the components have to be stable over time, sterilisable and variable in shape. The mineral phase used is calcium anhydrous hydrogen phosphate also known as monetite (CaHPO4). Monetite was modified by incorporation of gelatine resulting in a material named PPGC (phosphate pre-structured gelatine mineralized with calcium). Monetite was chosen because it is easy to use and has the added benefit of a high solubility, thus guaranteeing an adequate calcium ion release during degradation [3]. Furthermore, the degradation products of gelatine, a partially denatured hydrolysis product of collagen, as well as calcium and phosphate ions all naturally occur in bone. Another aspect bone replacement materials may take into account is the fact that bone is a hierarchically organized material. Although gelatine modified calcium anhydrous hydrogen phosphate cannot completely mimic these hierarchical levels [4], the pre-structuring of gelatine by phosphate ions is reported to be an attempt to control the mineral nucleation and growth by proteins [5,6]. Monetite is one of the mildly acidic calcium phosphate phases (CaP) and soluble at physiological pH [7]. It is mainly used as a powder component of self-hardening CaP pastes or cements used for skeletal repair [8,9]. Monetite is commonly synthesized by thermal conversion of preset brushite cement or by crystallization from aqueous solutions at temperatures above 100°C, making it impossible to incorporate temperature sensitive components like peptides or proteins [3,10–14]. Nevertheless, monetite granules in rabbit calvaria bone defects showed a significantly enhanced bone regeneration compared to unfilled controls [15]. In alveolar bone defects in human patients, monetite granules also showed a faster bone regeneration as well as a better degradation rate compared to bovine hydroxyapatite [11]. Page 3 of 26

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In the present paper we show monetite preparation, combining gelatine mineralization with subsequent dehydration through simple lyophilisation. This was investigated to produce monolithic bone substitutes, by incorporation of gelatine, which are mechanically stable in case of uniaxial compression and cell culture, as well as degradable under physiological conditions.

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Materials and Methods

2.1

Mineralization of phosphate pre-structured gelatine

The experimental setup for the precipitation reactions [Figure 1] was inspired by basic research on apatite crystal formation under influence of gelatine reported by Kollmann et al. [5]. In particular with respect to the production of bone substitute materials, starting materials, pH control and processing parameters were extensively adjusted. First, porcine gelatine (300 bloom, 20 mesh, Gelita) was suspended in deionized water. After swelling the suspension was liquefied by heating in a water bath to 50°C. In parallel a phosphate solution was prepared by adding 12.8 g Na2 HPO4 (Roth) to 850 mL water. These two liquids were mixed and the pH adjusted to 7.0 ± 0.1 by adding 1 M HCl. During at least 8 h stirring the phosphate pre-structuring of gelatine occurred. After this process 150 mL of the mineralization solution (1 M CaCl2·2H2O-solution (Roth)) was added with 2.5 mL/min and allowed to ripen for 3 h, during which time the pH was either uncontrolled or kept constant by addition of 0.05 M NaOH with a TitraLab 90 (Radiometer). The PPGC (phosphate pre-structured gelatine mineralized with calcium) was then separated from the supernatant by centrifugation, and the pellet first lyophilized using a freeze dryer Epsilon 2-4 LSC (Christ), then pulverised by a mixer mill MM 400 (Retsch) at 30 Hz for 15 seconds. Changes in the material composition compared to the above described standard process of PPGC production are indicated by numbers right after that abbreviation. The numbers show the gelatine percentage of 0.0% (gelatine-free, PPGC0.0), 0.9% (PPGC0.9) and 4.5% (PPGC4.5) in the pre-structuring suspensions. 2.2

Mineral characterization by scanning electron microscopy and x-ray diffraction

Morphology of PPGC was characterized by scanning electron microscopy (SEM). To this end mineral powder was spread on a sample holder with a carbon tab and coated with carbon in a Balzers SCD 050 coater. Secondary electron images were taken using a Philips ESEM XL 30 scanning electron microscope working in high-vacuum mode with 3 kV acceleration voltage. X-ray diffraction (XRD) was used to characterize the crystal structure of PPGC mineralized and ripened at different pH values. PPGC was analysed using Cu-Kα radiation between 20° ≤ 2Θ ≤ 45° with a step width of 0.04° (in 2 Θ) in a Bruker D8 Discover using an area detector. Page 5 of 26

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Loss on ignition

About 200 mg lyophilized mineral without gelatine, with 0.9% and with 4.5% gelatine in the phosphate pre-structuring solution were weighed into melting pots. At least three samples each were pyrolyzed in a muffle furnace (TC 405/20, Padelttherm) in air atmosphere at a temperature of 1000°C for 1 h [16]. The remaining materials were weighed again and the amount of organic content was calculated using PPGC0.0 (gelatine-free PPGC) as a reference. 2.4

Sample preparation for cell culture and mechanical characterization

The mineral powder was pressed to both compact disc-shaped (13 mm in diameter, 400 mg) and cylindrical (5 mm in diameter, 100 mg) samples. The compaction was performed with an Instron 5566 uniaxial testing instrument (maximum load of 10 kN) or a hydraulic lever-press (Perkin Elmer, maximum load of 67.6 kN). For discs of 13 mm in diameter an amount of 400 mg was compressed with 75 MPa (10 kN, Instron 5566) and 510 MPa (67.6 kN, hydraulic lever-press), respectively, resulting in a sample thickness of 1.2 mm to 1.5 mm. Cylindrical samples with 5 mm diameter were compressed with 510 MPa (10 kN, Instron 5566) and 1000 MPa (20 kN, hydraulic lever-press), resulting in a sample height of ca. 2.6 mm. Increasing the amount of gelatine in PPGC4.5 caused difficulties during compression. It is assumed that a surplus of gelatine leads to a deposit of gelatine on the mineral instead of a thorough mixing. For this reason investigation of PPGC4.5 was not continued. 2.5

Mechanical investigation with ball on three balls test and Brazilian test

For mechanical characterization the ball on three balls test according to Börger et al. was used [17]. For this purpose, samples with 13 mm diameter were positioned on three balls arranged as a triangle. Each ball was 10 mm in diameter and bordered on the other two. The load ball above the sample also was 10 mm in diameter. During the measurement it was lowered with 0.5 mm/min, thus axisymmetrically applying a load until fracture. Due to the test geometry a biaxial stress state was generated within the sample. The maximum force leading to fracture was measured and used for calculation of the biaxial strength. The calculation according to the numerical solution of the stress field from a finite element analysis required the assumed Poisson’s ratio of 0.3. For determination of the tensile strength the indirect method of the Brazilian test was used to analyze twelve cylindrical samples with 5 mm diameter and ca. 2 mm height. The material was subjected to two opposing normal strip loads at the cylinder’s periphery. The load applied Page 6 of 26

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with 0.5 mm/min and the maximum force F leading to fracture was used to calculate the indirect tensile strength βSZ based on the following equation with cylinder diameter d and cylinder height h: βSZ = 2F/ (π*d*h). 2.6

Degradation and bioactivity

For degradation investigations cylindrical 100 mg/5 mm samples were stored in Dulbecco’s phosphate buffered saline without Mg2+ and Ca2+ (PBS, Biochrom) and m-simulated body fluid (SBF) according to Oyane et al. [18,19]. Buffer capacity of PBS is based only on 9.6 mM hydrogen phosphate in aqueous solution. In contrast, the buffer capacity of SBF is based on 10 mM bicarbonate, 1 mM hydrogen phosphate and 75 mM of the Good buffer HEPES [20], with the latter accounting for the majority of the buffering effect. PPGC samples were produced gelatine-free (PPGC0.0) and with gelatine (PPGC0.9) at the two different compressive stresses of 510 MPa and 1000 MPa. Part of the samples was gamma sterilized, another was not sterilized prior to further analysis. For each variation six samples were stored in 4 mL fluid at 37°C. After 1, 4 and 8 hours as well as after 1, 2, 4, 7, 14 and 21 days the pH, concentration of calcium and concentration of phosphate in the surrounding fluid was measured. To this end 70 µL of the fluid was removed without replacement at each time point. Ion concentration measurements were done in triplicate with 10 µL analysed with a colorimetric Fluitest® CA CPC test kit (Analyticon) and Fluitest ® PHOS (Analyticon). For all colorimetric measurements an Infinite 200 Pro microplate reader (Tecan) was used. The difference in mass was determined after 21 days of incubation after drying to mass constancy at 37°C and 15% relative humidity. The characterization of both the change in surface morphology and the shape of the precipitates was done by SEM during and after incubation. 2.7

Cell culture experiments

HBMSC isolated from bone marrow aspirates of a 19 year old female donor were kindly provided by Professor Bornhäuser and co-workers, Medical Clinic I, Dresden University Hospital Carl Gustav Carus. Cells were expanded as described elsewhere and seeded in passage 5 [21]. Cultivation was performed either on gamma-sterilized (25 kGy) 100 mg/5 mm/510 MPa samples of PPGC or polystyrene (PS) as reference material in both calcium-free and calcium containing media. Unfortunately samples without gelatine disintegrated after 2 days of incubation and thus were not available as reference material for Page 7 of 26

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cell culture. The cells were seeded at a density of 10.000 cells by drop-seeding either directly on the scaffolds or on PS as reference. Prior to seeding the scaffolds were preincubated for 24 h in α-minimal essential medium (α-MEM, Biochrom) with 10% fetal calf serum (FCS, Biochrom), 100 U/mL penicillin, 100 µg/mL streptomycin (1% penicillin/streptomycin , Biochrom), and 2 mM L-glutamine. The seeded scaffolds and the PS controls were incubated in the same medium at 37°C, 5% CO2 in a humidified atmosphere. Differentiation was induced after three days by adding 10 mM β-glycerophosphate (β-GP, Sigma), 10 nM dexamethasone (Dex, Sigma), and 50 µM ascorbic-2-phosphate (ASC, Sigma) to the medium. Cell culture medium was replaced twice a week. Adhesion and proliferation was determined measuring lactate dehydrogenase (LDH) and DNA after 1, 3, 7, 14, 21, and 28 days. All measurements were performed with cell lysates prepared with 1% Triton X-100 (Sigma) in PBS. The determination of LDH activity in cell lysates was performed with a LDHCytotoxicity Detection Kit (Takara). The LDH activity was correlated with the cell number by a calibration curve using lysates of defined cell numbers. DNA amount was measured using QuantiT™ PicoGreen® dsDNA Reagent (Invitrogen). The alkaline phosphatase level (ALP) in relation to the cell number was determined as an indicator for osteogenic differentiation of hBMSCs. For this an aliquot of cell lysate was added to ALP substrate buffer containing 2 mg/mL p-nitrophenyl phosphate (Sigma), 0.1 M diethanolamine, 1 mM MgCl2, 0.1% Triton X-100 (pH 9.8), and the mixture was incubated at 37°C for 30 min. The enzymatic reaction was stopped by the addition of 0.5 M NaOH, and the absorbance was read at 405 nm. A calibration curve was constructed from different concentrations of p-nitrophenol (PNP). For cultivation of monocytes, PPGC is preincubated for 24 h in α-MEM with 10% FCS and 1% penicillin/streptomycin. Monocytes were isolated from human Buffy Coat by Ficoll density gradient centrifugation. Cultivation of monocytes on preincubated PPGC was performed over 12 days in α-MEM with 10% FCS, 1% penicillin/streptomycin (PAA Laboratories) and 25 ng/mL macrophage colony-stimulating factor (M-CSF; PeproTech). After two days of cultivation, 50 ng/mL receptor activator of NF-κB ligand (RANKL) was added to induce osteoclast formation. This point was regarded as day zero of osteoclast cultivation and medium was changed every two to three days henceforth. After 10 days of cultivation samples were rinsed with PBS and fixed with 2% paraformaldehyde (Carl Roth) in 0.1 M Na-cacodylate buffer (Cacodylic acid sodium salt trihydrate, Carl Roth) at pH 7.2 and 2% glutaraldehyde (Plano) and 0.02% picric acid (Riedel de Haen). Fixation was finalized Page 8 of 26

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with 1% osmic acid (osmium tetroxide, Carl Roth) in 0.1 M Na-cacodylate buffer at pH 7.2 and room temperature for 1 h. Samples were rinsed again, then dehydration in a graded ethanol series and Epon embedding was performed for 20 h at 60°C. For this samples were incubated in a mixture of glycidether 100 (Serva), methylnadic anhydride (Serva), 2dodecenylsuccinicacid anhydride (Serva) and 2,4,6-tris (dimethylaminomethyl) phenol (Serva). Semi-thin sections of 0.5 to 1 µm were prepared on an Ultracut (Leica Microsystems) and briefly stained with 1% toluidine blue (Carl Roth), 1% borax (di-sodium tetra borate, Merck), and 1% Safranin O (Merck). Cuts were examined on an Axiophot photomicroscope (Zeiss). The preparation of ultra-thin sections was also performed on an Ultracut (Leica Microsystems). Sections were applied onto collodion coated (Science Services) copper grids (Plano). Subsequent staining was performed with a Reichert Ultrostainer (Leica Microsystems) with 0.5% uranyl acetate (Leica Microsystems) for 30 min at 40°C and 3% lead citrate (Leica Microsystems) for 1 min and 20 sec at 20°C. The sample sections were analysed with a transmission electron microscopy (TEM) EM 109 (Zeiss) at 80 kV. Pictures were taken with a 2K CCD camera TRS (Tröndle image intensifier systems, Moorenweis). 2.8

Statistics

All measurements were done at least in triplicate and are expressed as mean ± standard deviation. One-way analysis of variance (ANOVA) with Bonferroni correction was applied for statistical analysis and P values <0.05 were considered significant and indicated by one asterisk, while ** indicate P < 0.01, and *** P < 0.001.

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Results pH dependence of mineral formation

PPGC displayed a platelet shaped morphology comparable to brushite. Single platelets had dimensions of 5 to 100 µm in width and 2 to 50 µm in height, while the thickness was constantly less than 1 µm [Figure 2]. The structure of the precipitated mineral was closely connected to the pH value of the mineralization solution. At high pH (12, 10, 7.4) hydroxyapatite with small crystallites formed, identifiable by broadened peaks in the XRDspectra [Figure 3]. The broadened peaks also indicate a certain amount of amorphous mineral or nanoscale crystals, which might be caused by the incorporation of gelatine. The mineral precipitated at pH = 4 was identified as monetite. 3.2

Loss on ignition

Pyrolyzing gelatine-free reference mineral resulted in a loss on ignition of 15 w%. Mineral precipitated with 0.9% or 4.5% gelatine in the pre-structuring suspension revealed loss on ignition of 20 w% and 37 w%, respectively. The organic content of PPGC0.9 was thus 5%, that of PPGC4.5 was 22%. That means a 5-fold increase of gelatine in the pre-structuring suspension led to a 4.4-fold increase of gelatine in the precipitated mineral. 3.3

Biaxial strength and indirect tensile strength

The biaxial strength of PPGC was investigated with regard to the influence of gelatine amount and compaction pressure. In case of 400 mg/13 mm samples, an increase of the compaction pressure from 75 MPa to 510 MPa (which is equivalent to the compaction pressure for 100 mg/5 mm/510 MPa samples), led to an increase of biaxial strength by a factor of 3.5 [Figure 4]. Incorporation of gelatine significantly increased biaxial strength by a factor of 1.5. The indirect tensile strength of gelatine containing samples (PPGC0.9: 2.7 MPa) was significantly higher than tensile strength of gelatine-free samples (PPGC0.0: 1.3 MPa) [Figure 5]. In addition, the indirect tensile strength was significantly increased by 15% after gamma sterilization. The increase of compaction pressure was another possibility to increase the tensile strength, as a tensile strength of 7.5 MPa was achieved by doubling the applied load during compaction.

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Degradation and bioactivity

Samples without gelatine were not stable during incubation as they underwent a very rapid disintegration and degradation. Therefore, only gelatine containing PPGC samples were investigated regarding the influence of compaction pressure and gamma sterilization on degradation. The overall mass reduction of samples during incubation contains mass reduction due to degradation as well as the mass increase due to calcium phosphate reprecipitation (shown later). However, overall mass reduction was always higher for samples in PBS than in SBF [Figure 6], namely by factor 1.2 to 2.3. The highest mass reduction of 8.3% was measured for non-sterilized PPGC0.9 with 510 MPa compaction pressure in PBS (7.2% in SBF). Sterilization decreased mass reduction during incubation to 5.7% in PBS or 3.8% in SBF, respectively. Again, a doubled compaction pressure reduced PPGC0.9 degradation by factor of 1.6 in PBS and 2.5 in SBF. During incubation of PPGC samples the pH decreased. Samples without gelatine disintegrated, as previously mentioned, and a drop of pH from 7.4 to 5.8 in PBS and from 7.6 to 7.0 in SBF was detected. Samples with gelatine (PPGC0.9 510 MPa) showed a slight decrease of pH from 7.6 to 7.4 over 21 days in SBF. Similar to the gelatine-free samples, the decrease of pH was significantly higher in PBS than in SBF, as it dropped from 7.4 to 5.9 (PPGC0.9 510 MPa) or 6.3 (PPGC0.9 1000 MPa), respectively. Additional buffering effect of pure gelatine in relevant concentrations in deionized water, SBF, and PBS was verified by titration experiments (data not shown). Incubation in both SBF and PBS caused a release of calcium from PPGC0.9 [Figure 7]. After the first hour of incubation, the initial calcium concentration of around 2.8 mM in SBF was increased to up to 5.3 mM for PPGC0.9 (510 MPa) and 4.9 mM for PPGC0.9 (1000 MPa). The decrease was faster in case of the low-compacted PPGC0.9. For both PPGC0.9 variants an equilibrium was reached after 21 days in SBF at a calcium concentration of 0.6 mM. In PBS an initial calcium release up to 0.8 mM was measured. Afterwards a decrease was visible, followed by a second release ending after 21 days at 1.0 mM calcium in PBS for lowcompacted PPGC0.9 and at 0.5 mM for high-compacted PPGC0.9. The initial inorganic phosphate concentration in SBF was 1.0 mM on the average [Figure 8]. After 1 day incubation of both PPGC0.9 samples, the phosphate concentration was reduced to a minimum of about 0.3 mM. This was followed by an increase of phosphate that was faster Page 11 of 26

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for high-compacted PPGC0.9. The terminal concentration was 4.0 mM for PPGC0.9 1000 MPa and 3.4 mM in case of PPGC0.9 510 MPa. Incubation in SBF and PBS led to a surface coating on all samples. The resulting layers differed considerably in their morphologies [Figure 9]. In SBF closely packed needle-shaped crystals with a length of less than 1 µm covered the surface. In contrast to this, incubation in PBS caused a flake-like mineral precipitation. The flakes had dimensions in the low µm range and a thickness below 0.5 µm. They also appeared as a closely packed layer on the surface. 3.5

hBMSC and monocyte culture

Cultivation of hBMSC on PPGC0.9 revealed a comparatively low adhesion of less than 20%. Therefore, the proliferation rate (cell count at time x in comparison to day one) was used as an indicator for cells behaviour. Without osteogenic induction the hBMSC proliferated and doubled their number, irrespective of the calcium content in the medium [Figure 10]. After 28 days, the cell number on PPGC0.9 was increased by factor 10 for medium without calcium. For a standard calcium concentration the cell number on PPGC0.9 increased by a factor of 17. On PS the cell number increased by a factor of 3.5 for medium with and without calcium [Figure 11]. Cells on PPGC0.9 without calcium showed the lowest proliferation rate of 1.5 25 days after osteogenic induction. In contrast to this, cells on PPGC0.9 with calcium in the medium increased their cell number by a factor of 30. The proliferation was also determined by DNA measurements, which showed trends comparable to those measured by LDH. However, in own investigations a significant binding of DNA to calcium phosphates was observed (non-published results), making the LDH data considerably more reliable. The osteogenic differentiation was analysed by ALP measurement. Without osteogenic additives the ALP level was below 1 nmol PNP/min /10.000 cells (data not shown). For cells cultivated on PS, osteogenic additives induced faster ALP expression in comparison to cells on PPGC0.9 [Figure 12]. The ALP maximum might be reached between 21 and 28 days or after 28 days. There was no significant difference in the ALP level at day 28 between PPGC0.9 and PS for neither the medium with nor without calcium. Cultivation of monocytes over 12 days and induction of osteoclastogenesis was performed as a preliminary test. TEM investigations revealed multinucleated cells on PPGC0.9. Due to contrasting with heavy metals organic components are visible as dark parts and mineral parts are white in the TEM image [Figure 13]. The multinucleated cell had an osteoclast-like Page 12 of 26

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morphology as the basolateral microvilli prove, but missed the ruffled border to be considered a full mature osteoclast. The active resorption (endocytosis) or exocytosis of PPGC0.9 is visible at the right side of the cell, where a deep indentation of the basolateral membrane into the material can be seen.

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Discussion

The field of calcium phosphates for use as bone substitutes is a large one and has been under investigation for years, but the calcium phosphate phase monetite seems to be underrepresented except in its usage as a component of calcium phosphate cements. There, it shows a significant support of bone growth [15]. In the case of bone defects of a supercritical size, healthy bone growth and remodelling have to be guided by a resorbable biomaterial, and the properties bioactivity, hBMSC response and resorbability are crucial to the performance of such a material. It was the aim of this study to develop a biomaterial that stimulates in vivo osteoblasts and at the same time is resorbable by osteoclasts, and to analyse it with regard to the mentioned properties. Specifically, the present study was to evaluate the feasibility and performance of a monetite deposition on an organic template at low temperatures. Monetite is a highly soluble calcium phosphate phase, which makes it well suited regarding the intended release of calcium from the bone substitute. Commonly, it is prepared either at high temperatures, which precludes the use of an organic template. The synthesis of submicrometer monetite particles by room-temperature-reaction of calcium carbonate in an ethanol solution containing small aliquots of concentrated H3PO4 (orthophosphoric acid) by Tas et al. is also unsuitable for use with an organic template [7], and the same is true for monetite precipitation in a sodium meta silicate gel by Sivakumar et al. [22]. In our approach we used lyophilization to eliminate the mineral water by drying at low pressure (0.001 mbar, 30°C, 120 min). This is a particularly gentle method and thus suitable for a biomaterial based on a gelatine template. The importance of gelatine to the formation of stable monolithic samples was shown by the Brazilian test. Loss on ignition revealed an amount of 5 w% gelatine in PPGC0.9. This led to an increase of biaxial strength by a factor of 1.5 compared to samples without gelatine. Even more important was the fact that samples without gelatine disintegrated after 2 days of incubation in SBF or PBS, respectively, while samples with gelatine were stable for at least 28 days. Besides the improvement of the mechanical properties of calcium phosphates by gelatine, gelatine as template also influences mineral nucleation and crystal growth [5]. The ball on three balls test as well as the Brazilian test revealed a huge impact of the compacting force during sample production on its later mechanical stability. It is assumed that gelatine is liquefied during compaction and pressed as a sort of glue into small gaps and Page 14 of 26

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cracks, leading to an increase of indirect strength values by a factor of 2 compared to gelatinefree samples. Surprisingly, gamma-sterilization after sample production also caused a significant increase of indirect tensile strength, which might be due to cross-linking processes during irradiation. The cross-linking effect of reactive species produced from water by gamma sterilization was observed for hydrogels (10% gelatine dissolved in water) sterilized with doses between 10 to 100 kGy [23]. The dose-dependent decrease of swelling of these hydrogels as well as an increase of dry sample mass after incubation in water over 7 days at 37°C proved crosslinking. Regarding PPGC, it seems likely that moisture adsorbed during storage prior to sterilization leads to occasional cross-linking reactions. The increased stability of gammasterilized samples was also measurable during degradation tests, where samples without irradiation showed the highest mass reduction during incubation both in SBF and in PBS. The higher mass reduction in PBS compared to SBF is probably due to the fact that there are no calcium ions in PBS, and thus a smaller degree of reprecipitating apatite. Mineral precipitation in SBF is assumed to simulate the mineralization in living organisms. It is thermodynamically controlled by the solubility products of the used mineral salts in that particular solution, the corresponding ion concentrations and the pH of the solution. The precipitation is initiated by a nucleation that is governed by kinetic hindrances. The typical morphology of the mature precipitate is that of apatite, as can be seen in the SEM pictures. During degradation in SBF, the mass of this layer caused a partial compensation of the mass reduction of the whole sample. Surprisingly, a complete surface coverage of precipitated mineral was found on samples in PBS as well as in SBF. This effect is due to the fast calcium release. In combination with the high phosphate concentration of PBS, this leads to a reprecipitation showing a coarser structure compared to SBF. The pH drop in PBS during incubation was due to its low buffer capacity (c(HPO42-, H2PO4-) = 9.6 mM) in comparison to the HEPES-buffer containing SBF (c(HEPES) = 75 mM). Therefore reprecipitating calcium phosphates, as visible on the SEM pictures, caused an increase of hydronium ion proportion when phosphate ions and metal cations were withdrawn from solution. The both desired and effective release of calcium ions from the material into SBF is due to its high solubility during the first days. The high initial values of 5 mM calcium lead to a fast reprecipitation due to oversaturation, and thus to comparably low calcium concentrations of less than 1 mM after 28 days of storage. The decrease of calcium ion concentration in the surrounding fluid accompanied by an apatite layer formation indicates PPGC to be highly Page 15 of 26

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bioactive. This bioactivity is recently discussed to be critical for osteoblasts due to the unphysiological calcium ion decrease [24]. It is known that the behaviour of osteoblasts is not only influenced by the absolute extracellular calcium concentration but also by variations in calcium ion concentration, regardless of whether the concentration is increasing or decreasing [25]. It was shown that the intracellular calcium increased for a fast increase of extracellular calcium concentration as well as for a fast decrease. The ability of cells to sense the time derivative of extracellular calcium concentration might be a hint for their proliferation even on highly bioactive materials. In PBS there is an initial increase of calcium, which initiates a reprecipitation process, decreasing the measurable calcium ion concentration. This is finished after 4 days, after which calcium ion concentrations in PBS and SBF both seem to stabilize at an equilibrium value of about 0.5 mM. In contrast to the calcium concentration in SBF, the concentration of phosphate ions is increased after a short initial reduction. It is assumed that the initial reduction of phosphate is associated with a superficial apatite deposition, whereas the release of phosphate from day 1 indicates a transformation of the samples crystal structure. This transformation of monetite into hydroxyapatite can also be achieved by immersion into 0.1 M NaOH (60 °C, 48 h), which was described as a dissolution and re-deposition process and gave rise to a comparable surface coverage [26]. Monetite probably converts slower in hydroxyapatite during incubation in body fluids at pH 7.4 by phosphate release and calcium ion uptake. Surprisingly, this process is enhanced for the higher compacted PPGC. Regarding surface morphology it is important to take into account that yet no proteins interfered with the reprecipitation process. As they are present in both cell culture and in vivo, altered morphologies of calcium phosphate precipitates or even impeded crystallization are possible [27]. Regarding the bone biocompatibility of the material, the cellular response of hBMSC and monocytes is of special interest. As mentioned before the high bioactivity, which means a fast decrease of calcium concentration in the medium, represents a tough condition for the cells. As a result the adhesion (cell number by LDH after 24 h) of hBMSC is quite poor with about 20% of initially seeded cells. However, the low adhesion of hBMSC on PPGC was compensated by good cell proliferation, which is in agreement with the compensation shown for osteoblastic cell lines on powder printed monetite scaffolds [13]. The influence of calcium concentration on cells (osteoblast-like SAOS-2 and MG63 cells as well as mesenchymal stromal cells) was intensively investigated by the group of Engel et al. Page 16 of 26

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[28–30]. It was shown that the calcium sensing receptor of rat bone marrow mesenchymal stromal cells forced cell migration and proliferation in a concentration-dependent manner [31]. Additionally, an increase of extracellular calcium stimulated overexpression of osteogenic markers like ALP, collagen I and osteocalcin. This was similarly shown for calcium releasing hydroxyapatite, where ALP, OPN and BSP expression by MC3T3-E1 cells were enhanced [32]. Our experimental setup took these findings into account, and the influence of calcium concentration and of the material itself was accordingly investigated using both a calcium-free medium and a standard α-MEM cell culture medium. Owing to the addition of fetal calf serum to the medium, an additional amount of about 0.4 mM calcium has to be considered, so that the calcium-free α-MEM contains 0.4 mM calcium and the standard α-MEM contains 2.2 mM calcium. Interestingly, the proliferation rate of hBMSC with and without osteogenic supplements on PS was not influenced by the amount of calcium in medium. Cells on PPGC, on the other hand, showed higher proliferation rates when cultivated in standard α-MEM compared to calcium-free medium. This effect was even more pronounced when cultivated with osteogenic induction; where hBMSCs revealed a huge proliferation rate of about 30 on PPGC in calcium containing medium and almost no proliferation on PPGC in calcium-free medium. Based on these results it can be concluded that at least low levels of calcium are necessary to maintain hBMSC proliferation if there is no osteogenic stimulation, while for cells with stimulated osteogenic differentiation a higher amount of calcium is necessary. Only an assumption can be made on this. It seems as the calcium ion concentration, due to calcium release from PPGC0.9, is not sufficient to allow proliferation of hBMSCs and osteogenic differentiation in parallel. Thus it seems that the (few) cells in calcium-free medium exposed to osteogenic supplements are subject to a greater differentiation pressure than those cells in standard medium. The apparently contradictory fact that hBMSCs continue to proliferate on PPGC after osteogenic induction might be explained as follows: Due to the low adhesion, the initial number of cells on PPGC is quite low. Therefore, the cells had not reached confluence when the osteogenic induction was started. Osteogenic induction slightly decreased the proliferation rate but did not terminate proliferation. The resulting ongoing proliferation reached after 28 days the usually higher cell number of the PS reference. Additionally, the presence of dexamethasone as well as ascorbate, both supplements for osteogenic differentiation, are also known to promote cell proliferation [33,34]. Page 17 of 26

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The ALP level was measured as a marker for osteoblastic differentiation. After 14 days of cultivation with osteogenic supplements, the ALP level of cells on PS was higher compared to those on PPGC, an effect that was independent of the calcium concentration in the medium. On PS the increase of ALP was faster, but after 28 days osteoblasts on both PS and PPGC showed comparable values of ALP activity. This may indicate a delayed osteogenic differentiation of cells on PPGC material compared to PS, an effect possibly due to differences in surface morphology. The morphology of large platelet-shaped PPGC particles [Figure 2] was considered to be favourable for osteoclastic resorption. The platelet-shape of PPGC was still visible after compression to cell culture samples. Even though osteoclast-like cells cultivated on PPGC did not develop the ruffled border indicative of the full osteoclast phenotype, the active resorption of PPGC was confirmed by TEM images [Figure 13]. Due to the accumulation of the microfilaments along the inner aspect of the apical plasma membrane, there is evidence to suggest that adhesion of the cells is mediated by actin filaments which are a typical feature of the sealing zone. Of the two osteoclastic material transport mechanisms - the receptor-bound transcytosis and the vesicles bound exo- or endocytosis [35] - the latter can be seen here. Since osteoclasts derive from the haematopoietic stem cell within the monocyte-macrophage lineage, incomplete formation of the osteoclast phenotype might also indicate the formation of macrophages, something that could be prevented by the addition of RANKL. Resorbable biomaterials are subject to the requirement of degradation within an adequate time period, both through chemical dissolution and resorption by osteoclasts. Instead of holding osteoclastic resorpion responsible for weakening in the defect area and consequently trying to limit their activity, biomaterials should be designed to be resorbed in the same speed as osteoblasts lay new bone on its surface [35]. As recently reviewed by Detsch and Boccaccini there is still only limited knowledge about the effects of osteoclasts on biomaterials, and vice versa [36]. They reported that crystallinity, grain size and density of the surface seem to be a less significant influence for osteoclastic resorption than surface chemistry and dissolution rate. PPGC complies with the material requirement of Detsch and Boccaccini as it shows both a positive influence on proliferation and differentiation of osteoblasts but, even more important, a resorbability by osteoclast-like cells. The mutually stimulating effect of these two processes was shown by Sugimoto et al. to depend on the extracellular calcium concentration. They could increase osteoblastic differentiation and chemotaxis of murine pre-osteoblast cells (MC3T3-E1) by culturing them Page 18 of 26

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in a medium conditioned by human monocytes at elevated calcium concentrations [37,38]. Apart from this indirect – osteoclast mediated – way of stimulating of bone formation through increased calcium concentrations, the direct stimulation of osteoblasts by calcium release from PPGC could also be shown in the present study. Jung et al. determined that the calcium efflux from bone microdamages resulted in an intracellular calcium response in MC3T3-E1 through voltage-gated calcium channels [39]. Therefore, it can be assumed that the biomaterial related calcium efflux induces the same cascade of regulating cell functions for matrix damage repair and bone substitute replacement. However, it has to be considered that the activation of osteoblasts by direct calcium signalling is bound to determined temporal concentration changes, like repetitive spikes and waves [40]. Therefore, indirect osteoblastic stimulation via osteoclasts might be more advantageous. Resorbability and bioactivity promoting an “appropriate host response for a given application” represents the combination of biomaterials properties claimed by Hench and Polak to be the “Third-Generation Biomedical Materials” [41,42]. In their papers and in the present as well it is suggested that a controlled release of chemicals in the form of ionic dissolution products may act as a stimulus to specific cellular responses at the molecular level. Therefore, the resorbability of the biomaterial is a major key in tissue regeneration.

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Conclusions

Combining inorganic materials with organic macromolecules not only in a structural way, but also regarding their functionality is one of the most promising approaches of bone substitute research. Taking into account mechanical requirements for handling during production, surgery and in its implanted state a load bearing material benefits from its composite structure on the sub-micrometre level. Mineralized gelatine provides furthermore degradability and resorbability by osteoclast-like cells. As it was moreover possible to produce calcium anhydrous hydrogen phosphate by a low temperature precipitation process, this is the first monolithic compact composite of monetite and gelatine. As gelatine and monetite both are resorbed to body familiar elements a reduced irritation of the surrounding tissue is supposed. Viability, proliferation and differentiation of human osteoblasts as well as resorption by multinuclear osteoclast-like cells are strong indicators for the materials bone biocompatibility. We have chosen the term ‘indicators’ because in vitrostudies at all do not produce evidence for the long term biocompatibility. Nevertheless, this approach is consistent with the opinion that biomaterials for orthopaedic applications should not fight pharmacologically an osteoclastic weakening of the bone [35]. Rather, they should be designed to “smooth assimilate” into the balance of resorption and regeneration of the bone. This material indicates the postulated shift by Jones and Hench from tissue replacement to tissue regeneration, where in situ repair is initiated in the host tissue [43]. Survival and phenotypic behaviour of both major bone cells qualifies this material for a role in the bone remodelling process, which is considered to be the key step for successful bone defect healing. Most notably, a more distinguished investigation of osteoclastic cell reaction is supposed to give a detailed view on in vitro bone biocompatibility after this first promising insight. Therefore, continuing the present study, the action of mature osteoclasts on the PPGC is being investigated in a current in vitro study.

Acknowledgements The authors are grateful to Prof. Hartmut Worch1 for giving the inspiring impulse for developing this material, to Axel Mensch1 for XRD investigations and to Anne Hild2 for her support in TEM sample preparation and for taking TEM images. We gratefully acknowledge Page 20 of 26

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the Deutsche Forschungsgemeinschaft (DFG Collaborative Research Centre TRR 79/SP M3) for financial support. We thank Prof. M. Bornhäuser and co-workers (Medical Clinic I, University Hospital Carl Gustav Carus, Dresden) for providing hBMSC.

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[10] Galea LG, Bohner M, Lemaître J, Kohler T, Müller R. Bone substitute: transforming beta-tricalcium phosphate porous scaffolds into monetite. Biomaterials 2008;29:3400– 7. [11] Tamimi F, Torres J, Bassett D, Barralet J, Cabarcos EL. Resorption of monetite granules in alveolar bone defects in human patients. Biomaterials 2010;31:2762–9. [12] Tamimi F, Torres J, Gbureck U, Lopez-Cabarcos E, Bassett DC, Alkhraisat MH, et al. Craniofacial vertical bone augmentation: a comparison between 3D printed monolithic monetite blocks and autologous onlay grafts in the rabbit. Biomaterials 2009;30:6318– 26. [13] Klammert U, Reuther T, Jahn C, Kraski B, Kübler a C, Gbureck U. Cytocompatibility of brushite and monetite cell culture scaffolds made by three-dimensional powder printing. Acta Biomater 2009;5:727–34.

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[14] Tamimi F, Sheikh Z, Barralet J. Dicalcium phosphate cements: brushite and monetite. Acta Biomater 2012;8:474–87. [15] Tamimi F, Torres J, Kathan C, Baca R, Clemente C, Blanco L, et al. Bone regeneration in rabbit calvaria with novel monetite granules. J Biomed Mater Res A 2008;87:980–5. [16] Hoyer B, Bernhardt A, Heinemann S, Stachel I, Meyer M, Gelinsky M. Biomimetically mineralized salmon collagen scaffolds for application in bone tissue engineering. Biomacromolecules 2012;13:1059–66. [17] Börger A, Supancic P, Danzer R. The ball on three balls test for strength testing of brittle discs: Stress distribution in the disc. J Eur Ceram Soc 2002;22:1425–36. [18] Dulbecco R, Vogt M. Plaque Formation and Isolation of Pure Lines with Poliomyelitis Viruses. J Exp Med 1954;99:167–82. [19] Oyane A, Kim H-M, Furuya T, Kokubo T, Miyazaki T, Nakamura T. Preparation and assessment of revised simulated body fluids. J Biomed Mater Res A 2003;65:188–95. [20] Good NE, Winget GD, Winter W, Connolly TN, Izawa S, Sing RMM. Hydrogen Ion Buffers for Bifological Research* 1966. [21] Heinemann C, Heinemann S, Worch H, Hanke T. Development of an osteoblast/osteoclast co-culture derived by human bone marrow stromal cells and human monocytes for biomaterials testing. Eur Cell Mater 2011. [22] Sivakumar GR, Girija EK, Narayana Kalkura S, Subramanian C. Crystallization and Characterization of Calcium Phosphates: Brushite and Monetite. Cryst Res Technol 1998;33:197–205. [23] Haema K, Oyama TG, Kimura A, Taguchi M. Radiation stability and modification of gelatin for biological and medical applications. Radiat Phys Chem 2014;103:126–30. [24] Malafaya P, Reis R. Bilayered chitosan-based scaffolds for osteochondral tissue engineering: influence of hydroxyapatite on in vitro cytotoxicity and dynamic bioactivity studies in a specific double-chamber bioreactor. Acta Biomater 2009;5:644-60. [25] Habel B, Glaser R. Human osteoblast-like cells respond not only to the extracellular calcium concentration but also to its changing rate. Eur Biophys J 1998;27:411–6. [26] Prado Da Silva MH, Lima JHC, Soares GA, Elias CN, De Andrade MC, Best SM, et al. Transformation of monetite to hydroxyapatite in bioactive coatings on titanium. Surf Coatings Technol 2001;137:270–6. [27] Schinke T, Amendt C, Trindl A, Pöschke O, Müller-Esterl W, Jahnen-Dechent W. The Serum Protein α 2 -HS Glycoprotein/Fetuin Inhibits Apatite Formation in Vitro and in Mineralizing Calvaria Cells. J Biol Chem 1996;271:20789–96.

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[28] Gustavsson J, Ginebra MP, Engel E, Planell J. Ion reactivity of calcium-deficient hydroxyapatite in standard cell culture media. Acta Biomater 2011;7:4242–52. [29] Gustavsson J, Ginebra MP, Planell J, Engel E. Osteoblast-like cellular response to dynamic changes in the ionic extracellular environment produced by calcium-deficient hydroxyapatite. J Mater Sci Mater Med 2012;23:2509–20. [30] Gustavsson J, Planell J, Engel E. Ion-selective electrodes to monitor osteoblast-like cellular influence on the extracellular concentration of calcium. J Tissue Eng Regen Med 2013;7:609–20. [31] González-Vázquez A, Planell JA, Engel E. Extracellular calcium and CaSR drive osteoinduction in mesenchymal stromal cells. Acta Biomater 2014;10:2824–33. [32] Jung GY, Park YJ, Han JS. Effects of HA released calcium ion on osteoblast differentiation. J Mater Sci Mater Med 2010;21:1649–54. [33] Jaiswal N, Haynesworth SE, Caplan AI, Bruder SP. Osteogenic Differentiation of Purified , Culture-Expanded Human Mesenchymal Stem Cells In Vitro. J Cell Biochem 1997;64:295–312. [34] Heinemann C, Heinemann S, Lode A, Bernhardt A, Worch H, Hanke T. In vitro evaluation of textile chitosan scaffolds for tissue engineering using human bone marrow stromal cells. Biomacromolecules 2009;10:1305–10. [35] Schilling A, Filke S, Brink S, Korbmacher H, Amling M, Rueger J. Osteoclasts and Biomaterials. Eur J Trauma 2006;32:107–13. [36] Detsch R, Boccaccini AR. The role of osteoclasts in bone tissue engineering. J Tissue Eng Regen Med 2014. [37] Sugimoto T, Kanatani M, Kano J, Kaji H, Tsukamoto T, Yamaguchi T, et al. Effects of high calcium concentration on the functions and interactions of osteoblastic cells and monocytes and on the formation of osteoclast-like cells. J Bone Miner Res 1993;8:1445–52 (Abstract). [38] Kanatani M, Sugimoto T, Fukase M, Fujita T. Effect of elevated extracellular calcium on the proliferation of osteoblastic MC3T3-E1 cells:its direct and indirect effects via monocytes. Biochem Biophys Res Commun 1991;181:1425–30. [39] Jung H, Best M, Akkus O. Microdamage induced calcium efflux from bone matrix activates intracellular calcium signaling in osteoblasts via L-type and T-type voltagegated calcium channels. Bone 2015;76:88–96. [40] Berridge MJ. Calcium signalling and cell proliferation. BioEssays 1995;17:491–500. [41] Hench LL, Polak JM. Third-generation biomedical materials. Science 2002;295:1014– 7.

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[42] Narayan RJ. The next generation of biomaterial development. Philos Trans R Soc 2010;368:1831–7. [43] Jones JR, Hench LL. Biomedical materials for new millennium: perspective on the future. Mater Sci Technol 2001.

Figure captions Figure 1

Process scheme for mineral preparation

Figure 2

SEM of PPGC0.9 after precipitation and lyophilization

Figure 3

Mineral structure analysed by XRD in dependence of pH

Figure 4

Biaxial strength of PPGC with/without gelatine and different compaction pressures determined by ball on three balls test

Figure 5

Indirect tensile strength (Brazilian test) of PPGC influenced by sterilization, incorporation of gelatine, and compaction pressure

Figure 6

Degradation of PPGC0.9 in SBF and PBS influenced by sterilization, and compaction pressure characterized by mass reduction

Figure 7

Calcium concentrations in SBF and PBS during incubation of PPGC0.9 as indicator for bioactivity

Figure 8

Phosphate concentrations in SBF during incubation of PPGC0.9

Figure 9a

SEM of PPGC0.9 after incubation in SBF for 21 days

Figure 9b

SEM of PPGC0.9 after incubation in PBS for 21 days

Figure 10

Change in cell number (LDH) relative to material specific number of adherent cells on day 1 during hBMSC cultivation without osteogenic stimulation on polystyrene (PS) and PPGC0.9, with and without calcium in α-MEM

Figure 11

Change in cell number (LDH) relative to material specific number of adherent cells on day 1 during hBMSC cultivation with osteogenic stimulation (Dex, ASC, β-GP) on PS and PPGC0.9, with and without calcium in α-MEM

Figure 12

ALP activity during hBMSC cultivation with osteogenic stimulation (Dex, ASC, βGP) on PS and PPGC0.9

Figure 13a

Multinucleated, osteoclast-like cells localized along the surface of PPGC0.9 after 10 d of cultivation, semi-thin section.

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Microfilaments accumulate along the apical domain of the plasma membrane (arrows) facing the PPGC0.9 surface. The cellular regions enriched with the filaments correspond spatially to the region of the sealing zone and ruffled border. However, the latter cannot be identified morphologically, but frequently finger-like protrusions arise from the apical membrane and extend into small spaces of the PPGC0.9. The protrusions are devoid of organelles but enriched with microfilaments (arrow heads). Figure 13b

Ultrastructure of a single osteoclast-like cell adhering apically at the surface of the PPGC0.9. The cell is multinucleated and polarized revealing morphological features of an osteoclast such as the basolateral microvilli (asterisks). Actin-enriched areas form long slender processes which range from the apical membrane of the cell into small spaces of the PPGC0.9 (arrowheads). The nuclei (N) are distributed throughout the cytoplasm which is enriched with mitochondria (M). A deep indentation of the basolateral membrane (arrow) containing a material of high electron density supports the assumption that they may be a site of endo-/exocytosis.

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