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Biomineralization of Recombinant Peptide Scaffolds: Interplay between Chemistry, Architecture and Mechanics Kendell M Pawelec, and Sebastiaan G.J.M. Kluijtmans ACS Biomater. Sci. Eng., Just Accepted Manuscript • Publication Date (Web): 03 May 2017 Downloaded from http://pubs.acs.org on May 9, 2017

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ACS Biomaterials Science & Engineering

Biomineralization of Recombinant Peptide Scaffolds: Interplay between Chemistry, Architecture and Mechanics Kendell M. Pawelec*†, Sebastiaan G.J.M. Kluijtmans Fujifilm Manufacturing Europe B.V., Oudenstaart 1, Tilburg, the Netherlands

ACS Paragon Plus Environment

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Abstract Biomineralized scaffolds are an attractive option for bone tissue engineering, being similar to native bone. However, optimization is difficult, due to complex interplay between architecture, chemistry and mechanics. Utilizing biomimetic nucleation, linear mineralized scaffolds were created from a collagen type I based recombinant peptide (RCP). Osteoblast mineralization was assessed, in response to changes in scaffold architecture, hydroxyapatite (HA) content and mechanics. Changes in scaffold pore size (150 - 450 µm) had little effect on mRNA levels, but influenced cell proliferation, achieving a balance between nutrient diffusion and surface area for cell attachment at 300 µm. Increasing the scaffold mechanical strength, from 2.9-5.2 kPa, enhanced the expression of osteocalcin, a late marker of mineralization. Further addition of HA, up to 20wt%, increased osteoblast mineralization, without altering the compressive modulus. Thus, it was shown that architectural cues influence cellular proliferation while the scaffold chemistry and mechanics independently contribute to gene expression.

Keywords mineralization, osteogenesis, scaffold, tissue engineering, hydroxyapatite

1 Introduction One of the key tools in regenerative medicine are porous scaffolds, which direct biological response through chemical, architectural and mechanical cues. Nowhere is the complex interplay between these three cues more apparent than in bone tissue engineering. Small changes in any aspect of the scaffold can elicit significant biological consequences, such as poor osteoconductivity or lack of integration with the host tissue. Only by considering all three types


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of signals can scaffolds achieve their final goal: well integrated and mineralized bone, which can withstand the large mechanical loads to which the skeletal system is subjected. As part of this biomedical challenge, many different materials have been utilized for bone repair, including metals, calcium phosphates, polymers, and composites structures.1 The material choice dictates many of the scaffold's features. Natural polymers have been used widely for their excellent biocompatibility, incorporating motifs for cell adhesion and drug delivery simultaneously.2 However, for bone applications, the addition of hydroxyapatite (HA), the calcium phosphate mineral predominant in bone, has been shown enhance osteogenic differentiation.3 Thus, biomineralization, or the process of introducing ceramic particles into polymer structures, is gaining attention. The end result is an environment similar chemically to the mineralized type I collagen extra-cellular matrix (ECM) of bone.4 Of the biomineralization routes explored in literature, biomimetic nucleation of calcium phosphates onto polymer chains has consistently demonstrated homogeneous scaffolds with crystals in the nanometer range, mimicking native bone.5 Both collagen type I and gelatin are compatible with the biomimetic nucleation process.6-8 However, with the batch to batch variation and questions of immunogenicity present with polymers from natural sources, it was hypothesized that recombinant peptides could offer an attractive alternative. Therefore, biomineralization, via a biomimetic route, was explored in this study, utilizing a recombinant peptide based on collagen type I (RCP), enriched in the RGD cell adhesion ligand. A range of scaffold compositions was defined, by varying the amount of calcium phosphate added into the structure. With the presence of cell adhesion ligands and hydroxyapatite, biomineralized scaffolds have an optimized chemistry for bone tissue engineering. However, scaffold architecture also plays a role in the final biological outcome. Scaffold structures must match certain key requirements to


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encourage cellular infiltration. These include pore sizes of 200 - 400 µm, with pore interconnections above 100 µm.9-11 In addition, high scaffold permeability to fluid flow has been shown to increase bone integration in vivo.12, 13 To capitalize on this, aligned lamellar structures were produced, which not only mimic the natural ECM organization of bone, but have been shown to have higher permeability than isotropic scaffolds.1, 14 Naturally, mechanical cues cannot be neglected, as they also direct osteogenesis and mineralization.1, 15 Stem cells can be pushed towards osteoblast differentiation due to stiffness alone, a process which is driven by the interplay of adhesion ligands in the matrix.16 Many chemical processes, such as cross-linking or HA addition, can alter not only the chemistry of scaffolds, but also their mechanical properties.17-19 It can therefore be difficult to separate which signals have the most impact on cellular behavior: chemical or mechanical. The optimization of tissue engineering scaffolds depends strongly on the interplay between all three signals: chemistry, architecture and mechanics. After producing a range of biomineralized scaffolds from RCP, it was further hypothesized that both the architecture and mechanics could be altered to tune the mineralization of osteoblast-like cells. Architecture was assessed by varying the pore size between 150 - 450 µm, keeping the chemistry and cross-linking constant. Maintaining a single architecture, the mechanics and effect of HA addition into the RCP scaffolds were investigated. The ability to systematically vary all three scaffold cues in this system provides the basis for a broad understanding of how they interact, and allows for the effective optimization of the scaffolds for bone tissue engineering. 2 Experimental 2.1 Biomimetic Mineralization


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Unless noted, all reagents came from Sigma Aldrich. Scaffolds were made from a recombinant collagen-like peptide (RCP, CellnestTM, Fujifilm).20 The range of mineral content in the final scaffolds were: 0, 20, 40, and 60 wt%. The percentage of RCP (7.5 wt%) was kept constant for each amount of mineral added. The desired calcium phosphate phase was hydroxyapatite, so the calculation of reactants was based on a Ca/P molar ratio of 1.67. All steps were done at room temperature. A solution of 20 wt% RCP was prepared in distilled water at 50°C, homogenized and degassed. The calcium hydroxide (Ca(OH)2) was hydrated in water; the amount of Ca(OH)2 was determined by the desired weight percentage of HA. Phosphoric acid (H3PO4, 15 M) was added to the RCP solution, dropwise, over 20 minutes with stirring. Immediately after, the RCPphosphate mixture was added dropwise into the Ca(OH)2 over 30 minutes with constant stirring. The pH was adjusted to 7 using 1 M HCl, and the solution was left stirring for 2 additional hours to complete mineralization. For 0 wt% HA samples, a solution of 7.5 wt% RCP was prepared. 2.2 Scaffold Production Linear scaffolds were produced via a directional ice-templating technique, relying on a controlled bath temperature to form the ice.14, 21 Prior to freezing, 1 wt% ethanol was added to the mineralized solutions and the mixture was transferred to an aluminum mold, 50 mm in diameter, with a 2 mm Teflon insert around the sides. The solution (20.4 g in total) was gelled for 20 minutes at 10°C. The bath temperature was adjusted to -30°C and held for 10 seconds to nucleate ice. The temperature was immediately adjusted to -4°C, and thereafter cooled with a constant slope until ice growth was complete. The final slope was varied to determine the transverse pore size: 0.5, 0.1, 0.03°C min-1. Unless specified, the slope was 0.1 °C min-1. Scaffolds were immediately lyophilized in a Zirbus freeze drier for 24 hours at -15°C, below 0.8


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Pa. Secondary drying was completed for 10 hours at 25 °C and scaffolds were stored at room temperature prior to use. Final scaffold size was a cylinder 45 mm in diameter, 11 mm high. Scaffolds were cross-linked using a dehydrothermal treatment (DHT). Prior to cross-linking, linear scaffolds were cut to the desired size. Scaffolds were dried at 60°C overnight, under vacuum. Cross-linking was done at 160°C, with a vacuum of less than 1x10-2 mbar. The crosslinking time was either 24 or 96 hours. 2.3 Mechanical Properties The mechanical properties of the cross-linked scaffolds were tested in uniaxial compression (n = 3), so that the direction of force was parallel to the long axis of the pores. Samples were cut to 10 mm diameter, 10 mm high and wetted in PBS for 30 minutes prior to testing, after vacuuming out the air. A constant displacement test (0.16 mm min-1) was performed on a RheoPlus MCR 301 (Anton Paar) and the compressive modulus and yield stress were calculated over the initial linear regime using the 0.2% off-set method. 2.4 Hydration The ability of the scaffolds to interact with water was tested via the percent hydration (n = 5). Scaffolds, 5 mm diameter, and 2 mm high, were weighed in the dry state, after cross-linking. Scaffolds were hydrated in PBS, removing the air with a vacuum. When fully hydrated, the scaffolds were removed from the PBS and weighed. This measure correlated to the total amount of water which could be associated with the scaffold structure, and was reported as the percent of mass change from the dry scaffolds. 2.5 Mineral Characterization


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Thermal gravimetric analysis was used to quantify the amount of HA within the mineralized scaffolds (n = 3). Scans were performed on a Mettler Toledo STARe System, using an alumina cup, on 5-10 mg of scaffold, cut into fine pieces. Samples were heated from 25 to 800°C, with a rate of 10°C min-1, and 50 mm/ml air flow. The phase of the calcium phosphate was determined by x-ray diffraction (XRD). Scaffold was cut into fine pieces and dried overnight at 60 °C before milling (Retsch) using a 250 µm grid. The X-ray diffraction (XRD) patterns of the samples were recorded with a D8 Advance Diffractometer (Bruker, Karlsruhe, Germany) equipped with a Lynx-eye position sensitive detector using Cu Kα radiation (λ = 1.54178 Å) generated at 40 kV and 40 mA. XRD spectra were recorded in the 2θ range from 10 to 60° with a step size (2θ) of 0.02° and a counting time of 0.5 s. To investigate the chemistry, samples were prepared by slicing thin cross-sections of the scaffolds. The cross-sections were scanned, in the middle region, via Fourier transform infrared spectroscopy, using a PerkinElmer Frontier FT-IR Spectrometer + Spotlight 200. The spectra were acquired in the range of 4000 to 650 cm-1. 2.6 Scaffold Stability The stability of the scaffolds was tested by calculating the mass loss under physiological conditions. Scaffolds, with an HA content of 0 – 60 wt% were cut to dimensions of 5 x 5 x 5 mm, and cross-linked with DHT for 24 and 96 hours (n = 3). After weighing the scaffolds in the dry state, they were hydrated in phosphate buffer saline (PBS) and incubated at 37°C. After 7 days, scaffolds were washed in water for 1 hour, and dried overnight at 60°C before weighing. Stability is reported as the percentage of the initial mass which remained.


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2.7 Scaffold Porosity Pore size analysis was conducted using secondary electron micrographs, Section 2.12. Micrographs from three independent scaffolds were used for analysis. Individual pores were outlined and measured using Image J software; the pore size reported is the average Feret's diameter of at least 40 pores. Micro-computed tomography scans, Section 2.12, were used to compute the percent porosity of the scaffolds, using Image J software. 2.8 Cell Culture Scaffolds, 5 mm diameter and 2 mm thick, were sterilized via autoclaving in PBS at 121°C for 20 minutes, and incubated in complete media overnight (αMEM without ascorbic acid (Gibco #A104 90-01), 10% fetal bovine serum (FBS), 1% penicillin-streptomycin). MC3T3-E1 cells, a mouse osteoblast line, were seeded onto scaffolds via dynamic seeding. Scaffolds were placed in 50 ml Falcon tubes with a cell suspension of 5x105 cells/scaffold (2.5x106 cells/ml). The cells and scaffolds were rotated for four hours at 37°C, at less than 500 rpm. After seeding, the scaffolds were removed and placed in individual wells of a tissue culture plate. Scaffolds were further cultured in static conditions in mineralization media (αMEM without ascorbic acid (Gibco #A104 90-01), 10% FBS, 1% penicillin-streptomycin, 25 mM β-glycerophosphate, and 100 G/ml ascorbic acid). Scaffolds were cultured for a total of four weeks, replacing the media every 3-4 days. The effect of scaffold architecture on mineralization was measured by comparing MC3T3 response in 40wt% HA scaffolds, with a linear pore geometry. The pore size was varied over 150, 300 and 450 µm. To investigate the effects of mineral and weight percentage of HA, four


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sample types were used: 0wt% HA - 24 hr DHT, 0wt% HA - 96 hr DHT, 20wt% HA - 96 hr DHT, and 40wt% HA - 96 hr DHT. For all biological testing, n = 3. 2.9 Proliferation DNA quantification was done using Quant-iTTM PicoGreen ds DNA Reagent (Invitrogen). Scaffolds were washed in PBS, before freezing at -80°C for at least 6 hours. Scaffolds without any cells were used as controls. Once thawed, scaffolds were incubated overnight in 300 µl papain buffer (5 U/ml) at 60°C. A cell standard curve was created at the initial cell seeding: 0 5x105 cells. The assay was performed according to the manufacturer's instructions. Briefly, samples were added to a black 96 well plate with standard curves and controls. After adding the pico-green dye (1:200), the fluorescence was read at excitation 480 nm, emission 520 nm. 2.10 Polymerase Chain Reaction For RNA isolation, Direct-zol RNA Mini Prep columns were used (cat no. R2050). Briefly, scaffolds were homogenized in TRIreagent before proceeding with the on-column isolation according to the manufacturer's instructions, including an on column DNase step. Quantification of RNA was done using Quant-iTTM RiboGreen RNA Reagent and Kit (Invitrogen) and cDNA was obtained from QuantiTect Reverse Transcription Kit (Qiagen). Both were performed according to the manufacturer's instructions. Polymerase Chain Reaction (PCR) was conducted on an illumina EcoTM Real-Time PCR System (# EC-900-1001), using Sso Fast EvaGreen (BioRad). In the study on scaffold architecture, mRNA levels were normalized to the scaffold with 150 µm pores at day 7. For the study on mechanics and HA addition, the mRNA level of 0 wt% HA-24 hr DHT scaffolds at day 3 was used as the control.


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2.11 Mineral Content The kit OsteoImageTM Mineralization Assay (Lonza) was used to quantify the change in hydroxyapatite within the samples. The kit is specific for HA, rather than generic carbonate or phosphate salts, allowing it to be used semi-quantitatively. The manufacturer's instructions were adapted for use with scaffolds. Samples were washed once in PBS and transferred into wells of a black 96 well plate. The samples were then fixed in 4% paraformaldehyde for 15 minutes and washed once in Wash Buffer (provided by the kit). Staining reagent was incubated for 30 minutes with the scaffolds, and, after three washes, the fluorescence was read at excitation/emission of 492/520. Measurements were taken at day 14 and day 28 (n = 3), and the fluorescence intensity is reported, which is proportional to the amount of HA present. To correct for the HA content inherent in the scaffolds, that was not due to cellular processes, all measurements were normalized by controls without cells (n = 2). Controls were incubated in complete cell culture media at 37°C for the same amount of time as the samples. 2.12 Microscopy Scaffolds were characterized with both scanning electron microscopy (SEM) and microcomputed tomography (µCT). All samples with cells were imaged via SEM. Scaffolds were gently washed in PBS twice and fixed in 3.7% paraformaldehyde for 15 minutes. Scaffolds were then incubated for 45 minutes in graded ethanol dilutions, at room temperature: 70%, 80%, 90%, and 100% ethanol. Finally, the scaffolds were washed for 10 minutes in hexamethyldisilazane (HMDS) and allowed to dry overnight in a fume hood. After drying, the scaffolds were cut with a razor blade, mounted on SEM stubs, and sputter coated with platinum prior to imaging on a JEOL JSM-6335F.


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To image the scaffolds in 3D, µCT was utilized. Dry scaffolds, 3 mm diameter, 10 mm high, were imaged with a Skyskan 1072 (Bruker). All scans were taken at 35 keV, 150 µA and reconstructions were performed using Skyskan software. In all scans, the resolution was below 3µm/pixel. 2.12 Statistics All statistics were done with GraphPad Prism software using ANOVA and a post-hoc Tukey test. A confidence interval of 95% was reported in all cases. The standard deviation was marked on all graphs, and significance is indicated. 3 Results 3.1 Biomineralized Scaffolds Calcium phosphate was nucleated biomimetically onto the RCP before scaffold formation. Up to 60 wt% calcium phosphate could be incorporated into the scaffold. In all cases, an interconnected porous structure was produced, with predominantly linear pores. All scaffolds had a high percent porosity: 85.2, 87.2, and 85.4% for 0, 20, and 40wt% HA respectively. The amount of calcium phosphate, measured via TGA, was close to the calculated amount, Figure 1(a-b). All scaffolds had a calcium phosphate distribution from the scaffold top to base, of around 10% of the ceramic weight percent, which is probably caused by sedimentation of the HA-RCP composite. Thus, with a 60 wt% addition, a 6% difference in mineral content was observed. To determine whether the calcium phosphate was indeed hydroxyapatite, XRD spectra were taken as the percentage of ceramic varied. The prominent peaks all corresponded to the


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hydroxyapatite phase, and the broad size of the peaks indicated a small crystal size, Figure 1(c). In addition, the interaction between the RCP and HA was examined via FTIR. With the addition of hydroxyapatite to the RCP, characteristic peaks for the phosphate group become apparent, especially the peaks between 1050 and 1090 cm-1.22 Figure 1(f) shows a shift in the carboxyl peak at 1400 cm-1, confirming the close interaction of HA with the RCP, as reported during the biomineralization of other polymers.6 Further, 1 wt% ethanol, added during scaffold manufacture to expand the pore size, did not alter the interaction in any way. Stability under physiological conditions (aqueous environment, 37 °C) is an important prerequisite for tissue engineering constructs. It was found that the mineralized scaffolds underwent a rapid loss of RCP, within one week of initial wetting, and were stable thereafter. This suggests that the RCP lost did not form cross-links during DHT treatment, and was solubilized after immersion in PBS. The amount of RCP loss was dependent on mineral content and cross-linking time. The addition of HA decreased the stability of the scaffolds dramatically, leading to significant mass loss within one week, Figure 1(e). Scaffolds without HA (0 wt%) showed little mass loss at either 24 or 96 hour DHT cross-linking. However, with HA addition, the effectiveness of DHT cross-linking was reduced. Scaffolds with 40wt% HA, 24 hours DHT, experienced over 40% mass loss. Increasing the cross-linking time to 96 hours ensured that the mass loss remained below 10% for HA additions of up to 40wt%. At 60wt%, the scaffolds were not stable enough for use in cell culture. Only 16 ± 2% and 68 ± 2% of the original mass remained for 24 and 96 hours of cross-linking, respectively.


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Figure 1. Characterization of biomineralized-RCP scaffolds. (a-b) The percentage of HA incorporated was calculated via TGA at several points in the scaffolds, illustrated in (a). (c) The calcium phosphate was identified as amorphous hydroxyapatite, via XRD. * indicates HA peaks (d) FTIR spectra of the mineralized scaffolds, with or without 1wt% ethanol, demonstrate the addition of HA; the shift at 1400 cm-1 (f) suggests that HA interacts with a carboxyl group on the RCP chain. (e) Scaffold stability was evaluated at physiological conditions to determine the optimum cross-linking and percentage of HA. 3.2 Effect of Architecture on Mineralization To assess the impact of architecture on MC3T3-E1 cells, scaffold pore size was varied from 150 to 450 µm, Figure 2(a-c), while keeping the addition of HA constant, at 40 wt%. All scaffolds were able to support the initial attachment and spreading of the osteoblast cells, Figure 2(d-f), and proliferation through the linear structure, Figure 2(g-i). Over the course of 28 days, the


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scaffolds with an average pore size of 300 µm appeared to promote the greatest level of proliferation. On the other hand, the mineralization response, assessed from the mRNA levels of mineralization markers (ALP and osteocalcin) was not significantly affected by the architecture, Figure 2(k-l). In addition, ECM production, measured by the level of collagen mRNA, was unaltered by pore size. As the greatest proliferation was observed with a pore size of 300 µm, all further scaffolds were made with this architecture.

Figure 2. Architecture had the greatest effect on cellular proliferation in 40 wt% HA scaffolds with (a) 450 µm, (b) 300 µm, and (c) 150 µm pores. Cells were able to (d-f) attach and (g-i) proliferate on 40 wt% HA of varying pore size: (d, g) 450 µm pores, (e, h) 300 µm pores, (f, i) 150 µm pores. (j) Architecture of 40wt% HA scaffolds had the greatest effect on cell proliferation. mRNA levels were not significantly altered by architecture: (k) ALP, (l) osteocalcin, (m) collagen type I. Scale bar (a-c): 200 µm, (d-i): 50 µm.


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3.3 Mechanics and Hydroxyapatite Addition The balance between chemistry and mechanics was assessed by varying the HA addition and the DHT cross-linking time, in linear scaffolds, with a mean pore size around 300 µm. It was verified, using µCT, that changes to the amount of HA incorporated into the RCP scaffolds did not alter the architecture significantly, Figure 3(a-c). Cells were able to adhere to the scaffolds, regardless of cross-linking time or HA addition. However, the percent attachment of MC3T3-E1 cells was significantly decreased with the addition of HA, from around 17% to 1.5%, for 0wt% and 40wt% HA, respectively. Adhesion was most likely dictated by changing availability of RGD ligand between different sample types. Percent growth was highest in scaffolds with 20wt% HA, reaching 710 ± 150% after 28 days, significantly higher than in scaffolds without HA. The mechanical properties of the hydrated scaffolds varied in response to both the crosslinking time and the mineral addition, Figure 4(i). The compressive modulus increased significantly with the cross-linking time, going from 2.9 ± 0.1 to 5.2 ± 0.1 kPa, as the crosslinking time went from 24 to 96 hours, respectively. The addition of 20wt% HA did not significantly alter the modulus, but with 40 wt% HA, the modulus was significantly decreased to 4.0 ± 0.2 kPa. The yield stress also increased with cross-linking time in 0 wt% HA scaffolds, from 26 kPa to 39 kPa. The yield stress peaked at 47 kPa in 20 wt% HA scaffolds. Hydration properties of the scaffold were also changed significantly with mineral addition. Longer crosslinking times tended to increase the overall hydration of the scaffolds. With increasing HA content, however, the overall hydration decreased, dropping from 729% to 599% as HA increased from 20 to 40 wt%.


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Figure 3. Scaffold morphology remained open and linear regardless of HA content when visualized via µCT: (a) 0 wt% HA, (b) 20 wt% HA, (c) 40 wt% HA. Cell attachment and proliferation on scaffolds with varying mineral content and mechanics: (d) 0 wt% HA 24hr DHT, (e) 0 wt% HA 96hr DHT, (f) 20 wt% HA, (g) 40 wt% HA. (h) Percent attachment (striped bars) and percent growth (solid bars) followed opposite trends. *Significantly greater than mineralized samples (p < 0.05). **Significantly greater than 40wt% HA (p < 0.05). All µCT reconstructions are 2 x 2 x 2 mm. Scale bar (d-g): 50 µm.


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As the mechanical and chemical properties of the scaffold varied, the mineralization response of the osteoblast cells was also significantly affected, Figure 4. RUNX2, an early marker of mineralization, peaked earliest in scaffolds with 20 wt% HA. By day 14, RUNX2 mRNA levels had peaked in all scaffolds. However, scaffolds with 40 wt% HA had the lowest expression of RUNX2 mRNA overall. Collagen type I, on the other hand, was sensitive only to mineral addition, and was significantly up-regulated in the absence of HA. ALP expression peaked in all samples at day 14, with sustained expression in 0 wt% HA-96 hour DHT, and 40wt% HA scaffolds. In contrast, osteocalcin, a later marker of mineralization was significantly upregulated, both with increased cross-linking and with the addition of 20 wt% HA. Even though 40 wt% HA addition had a higher compressive modulus than 0wt% HA with 24 hr DHT, mRNA expression of RUNX2 and osteocalcin in 40 wt% HA was comparable or significantly lowered. Between days 14 and 28, cells deposited an increasing amount of HA in the scaffolds, confirming their mineralization potential, Figure 4(j). The amount of mineral deposited, at day 28, correlated to the mRNA levels of osteocalcin in the samples at day 21. The greatest mineralization was found in 20wt% HA scaffolds.


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Figure 4. Mineralization was affected by the mechanical strength and the HA content. Cells were attached and proliferated over 28 days: (a) 0 wt% HA, 24 hr DHT, (b) 0 wt% HA, 96 hr DHT, (c) 20 wt% HA, and (d) 40 wt% HA. Markers for osteoblast maturity and mineralization were followed over 21 days: (e) RUNX2, (f) collagen Type I, (g) ALP, (h) osteocalcin. The dashed line represents a fold-change of one; all results are normalized to 0 wt% HA 24hr DHT on day 3. (i) The mechanical properties and swelling were altered in the scaffolds. Mineral content (j) was significantly higher on scaffolds with 20 wt% HA. *Significantly greater than all


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other groups (p < 0.05). ^ Significantly less than all other groups (p < 0.05). # Significantly greater than mineral samples (p < 0.05). ** Significantly greater than 40wt% HA (p < 0.05). ^^ Significantly less than 0wt%, 96 hr DHT (p < 0.05). V Significantly less than 0wt% and 20wt% HA, 96 hr DHT (p < 0.05). Scale bar in (a-d): 50 µm.

4 Discussion Mimicking the native structure of bone requires attention to all possible biological signals: chemical, architectural, and mechanical. Incorporating all of these cues into porous tissue engineering scaffolds can be a challenge, given the considerable interplay between them. In this study, chemical cues were represented by exploring the range of HA addition possible in the scaffold. Architecture was controlled using ice-templating, to create a variety of scaffold pore sizes, while maintaining high permeability to fluid flow.14 At the same time, the mechanics of the scaffolds, related to cross-linking degree, were examined to determine if there were synergistic affects with the biomineralized chemistry. Effective strategies for bone regeneration rely on a deep understanding of all possible cellular signals incorporated into porous scaffolds, and the way in which they are interrelated. 4.1 Stability of Biomineralized Scaffolds It was demonstrated that a recombinant peptide based on collagen type I (RCP) could be successfully mineralized via biomimetic nucleation, to form scaffolds similar to the structure of natural bone. This approach has been used successfully with collagen and natural gelatin in literature, but unlike natural polymers, the RCP incorporated additional cell adhesion signals and


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had a defined structure.6-8 Biomimetic nucleation of HA, on the recombinant peptide chain, took place via an interaction with a carboxylate group, as observed by a shift in the FTIR peak, consistent with work on natural gelatin.7, 22 Previous studies have demonstrated that HA crystals, in the nanometer range, nucleate homogeneously on recombinant peptide chains.22 The fact that ethanol addition did not disturb this chemical bond suggests that the chemical state of the RCP chain was unaffected by the low amounts used (1wt%). The calcium phosphate phase was positively identified from the XRD spectra as hydroxyapatite. The broad XRD peaks supported the hypothesis that the HA was amorphous.23 Due to its higher solubility, amorphous HA can serve as a source of calcium and phosphate ions to drive osteoblast differentiation and mineralization.24 One of the greatest advantages to utilizing biomimetic nucleation, is that there was no aggregation or settling of HA particles in the solution during ice-templating. While a biomimetic HA-RCP scaffold could be produced, it was important to examine the range of HA addition. The key limiting factor for the amount of HA which could be incorporated into the RCP structure, was the scaffold stability. With greater HA addition, more carboxyl groups, needed for DHT cross-linking, interacted instead with HA crystals. Thus, at higher HA content, DHT cross-linking was ineffective, even at long time periods of 96 hours. Indeed, a large mass loss was observed almost immediately when 60 wt% HA–RCP scaffolds were immersed in an aqueous environment, due to a loss of uncross-linked RCP. For this reason, the practical limit of HA addition was 40 wt%, with a cross-linking time of 96 hours. In transitioning to environments of greater complexity, for example in native tissue, many additional cellular processes will be present. The bone niche contains several different cell types, including osteoclasts, which play a crucial role in bone resorption and can influence scaffold degradation. While no direct osteoclast studies were performed, the peptide is subject to


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enzymatic cleavage, and would most likely be recognized as a substrate by osteoclasts due to its close resemblance to native bone. Thus, with the addition of osteoclasts, in the in vivo environment, an increased degradation rate of the scaffolds is anticipated. However, given the lack of noticeable degradation during the 4 week in vitro study, the structures are expected to remain stable during wound healing and regeneration. The loss of stability with increased HA was supported by the mechanical data, which showed a decrease in compressive modulus as the amount of HA increased from 20 wt% to 40 wt%. In literature, the addition of HA often produces an increase in mechanical properties.3, 8, 18 One possible difference might be due to the differing crystal structure of the calcium phosphates used; in this study, amorphous nanoparticles were produced, which were associated with the peptide chain. Another possibility is that the recombinant peptide, with additional and regularly spaced cell signaling motifs, might encourage a more even mineral layer, encapsulating the entire chain. 4.2 Mineralization Response to Architecture Scaffold architecture has been observed to play a major role in the final clinical outcome of scaffolds in vivo.12, 13 The need for a highly permeable structure, to maximize nutrient and cell infiltration, dictated that scaffolds with anisotropic, linear pores were manufactured.25 There is a fine balance between an increased permeability, a result of larger pore sizes, and a reduction in the surface area available for cell attachment as the pore size increases.9, 14 This principal was illustrated in the current study, where a balance between diffusion and surface area was found at 300 µm, which is within the range shown necessary for promoting blood vessel infiltration and mineralization in bone.10 Above and below this pore size, the proliferation was decreased. In contrast, the initial cell attachment did not change significantly in the linear scaffolds with the


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same amount of HA. Studies using isotropic scaffolds, have shown a relationship between attachment and pore size, in part due to the improved permeability of scaffolds with a larger pore size.9 With permeability an order of magnitude greater than isotropic scaffolds, even at 150 µm pores, the effect of pore size was negated in this study. While proliferation was affected, mRNA levels of mineralization markers were not significantly altered by the scaffold architecture alone. Architecture can influence cell fate when it changes the availability of cell adhesion ligands or hampers the cell's access to nutrients by restricting diffusion.16 Clearly, the scale of the architecture, 150 - 450 µm, was too large to regulate cell fate, although it made a measurable difference in the proliferation. 4.3 Chemistry and Mechanics of Mineralized Scaffolds The interplay between chemical and mechanical signals was examined, after determining the architecture which encouraged optimal cell proliferation. Alterations to the chemistry impacted both the stability and mechanical properties of the scaffolds. Therefore, the cellular response to mechanics alone was assessed on scaffolds without HA addition, which have already been shown to support the osteogenic differentiation and mineralization of mesenchymal stem cells.26 Initial cellular attachment was significantly reduced with mineral incorporation. Decreased attachment may be due to disruption of the RGD adhesion ligands present in the recombinant peptide. Cell adhesion is sensitive to the ligand confirmation, which could be altered by the presence of HA in several ways. The crystals increased the surface roughness of the scaffolds up to 5-fold, possibly altering the presentation of RGDs to the cells.22 In addition, the interaction of HA with carboxyl groups in the RGD adhesion ligand may have reduced sites available for attachment. Lower availability of the carboxyl groups on the peptide chain also played a role in


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reducing scaffold stability, as the addition of HA increased above 40wt%. While adhesion was reduced, all adhered cells were well spread and viable. Osteoblast-like cells were able to proliferate on all materials over 28 days, reaching the highest proliferation rates in 20wt% HA scaffolds. A slight decrease in cell proliferation with higher HA content was observed in this study. This may possibly be due to an increased surface roughness, consistent with other studies on mineralized gelatin.8, 22 It must be noted that the percentage of proliferation may vary with cell type, with less proliferative cell types, such as primary osteoblasts, demonstrating a lower percentage proliferation than the osteoblast-like cells. However, the trend is expected to remain the same. Changing cross-linking and HA addition significantly impacted the mechanics of the scaffolds. Using dehydrothermal cross-linking, which does not introduce any potentially toxic chemical into the scaffold, the mechanical strength increased with increasing cross-linking time.27 The chemical linkages occur via condensation reactions between carboxyl groups and amines. As the amount of HA increased, fewer carboxyl groups were free to react, and the compressive modulus decreased. The variation in compressive modulus and yield stress suggests a balance between the decrease in cross-linking effectiveness, lowering the strength with higher HA content, and the increase in strength due to a greater solid content.1 Scaffold hydration was also decreased in the presence of HA, which is known to absorb less water than the recombinant peptide.8 Substrate stiffness modulates cell behavior, pushing stem cells towards an osteogenic lineage as the stiffness increases.3, 16, 28 In the linear scaffolds without HA, higher compressive moduli promoted the up-regulation of mineralization markers, such as osteocalcin, significantly. The ECM marker, collagen type I, also increased as the compressive moduli increased. The high


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expression of collagen mRNA, compared to biomineralized scaffolds, suggested that cells were laying down their own matrix, which could then be mineralized.29 However, a higher compressive modulus alone did not result in a significant increase in deposited mineral. Literature has suggested that stiffnesses of between 25-40 kPa are necessary to ensure osteogenic differentiation.15 However, within three-dimensional scaffolds, cell mineralization can be sensitive to mechanical properties between 0.5 - 4 kPa for osteoblastic cell lines.28, 30 Although the linear RCP scaffolds fall within this range, the change in compressive modulus, from 2.9 to 5.2 kPa was apparently not enough to enhance the final mineralization. It was observed that both mechanics and chemistry acted synergistically to promote mineralization in the RCP scaffolds. A positive influence of HA addition on osteoblast mineralization has been reported in the past.3, 31, 32 However, this study demonstrated that HA content was only beneficial until it interfered with scaffold stability. Thus, cells in 20 wt% HA scaffolds appeared to differentiate earlier and produce more mineral than those in 40 wt% HA scaffolds, which had reduced stability and lowered mechanical properties. Both the 20 wt% HA scaffold and 0wt% HA scaffold, with 96 hours cross-linking, had similar mechanical and architectural properties. Of the two, 20 wt% HA scaffolds had a significant increase in mineralization, verifying an independent and positive affect of the mineral addition. For the first time, this study highlights the relationship between HA addition and mechanics, in the range of 2 - 6 kPa, using biomimetic nucleation of recombinant peptides. The conclusions extend the relationships found in earlier studies, which focused on higher stiffness ranges, from 10 - 25 kPa.3 Taken together, it has been demonstrated that pore architecture ensures cell attachment and proliferation, while the mechanics and chemistry influence osteoblast differentiation and


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mineralization. By knowing how the signals inherent in porous structures affect cells, a new class of biomineralized scaffolds can be defined for optimum bone tissue regeneration. 5 Conclusion Balancing architecture, chemistry and mechanics is important for creating successful scaffolds for bone tissue engineering. Homogeneous biomineralized linear scaffolds, consisting of amorphous hydroxyapatite and recombinant peptide, were developed. Up to 40wt% HA could be incorporated into the scaffolds, after which, the scaffold stability and mechanical properties were impaired. The architecture of the scaffold affected osteoblast proliferation, with a pore size of 300 µm providing the optimum surface area and nutrient diffusion. Scaffold chemistry, controlled by varying the amount of HA addition, affected initial cell attachment. Osteoblast mineralization in the scaffolds was dependent on both chemistry and mechanics. With increased mechanical properties, markers such as osteocalcin and collagen type I increased. HA addition, up to 20wt%, was shown to have an added beneficial effect, independent of the scaffold mechanics. Overall, this study demonstrates how all three scaffold signals: chemistry, mechanics and architecture, are important for directing cell response and promoting bone regeneration. AUTHOR INFORMATION Corresponding Author *University of Michigan, 2350 Hayward St. Ann Arbor, MI 48109, USA; e-mail: [email protected] Present Addresses † Present address: University of Michigan, 2350 Hayward, Ann Arbor, MI 48109, USA


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Author Contributions The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. Funding Sources The research leading to these results has received funding from the European Union Seventh Framework Programme FP7/2007-2013 under grant agreement n° 607051. Notes Conflicts of Interest The recombinant peptide based on collagen type I was provided by Fujifilm Manufacturing Europe; SGJM Kluijtmans was employed by Fujifilm.

ACKNOWLEDGMENT The research leading to these results has received funding from the European Union Seventh Framework Programme FP7/2007-2013 under grant agreement n° 607051. REFERENCES [1] Wu, S.; Liu, X.; Yeung, K.W.K.; Liu, C.; Yang, X. Biomimetic porous scaffolds for bone tissue engineering. Mater. Sci. Eng. R 2014, 80, 1-36, DOI: 10.1016/j.mser.2014.04.001. [2] Friess, W. Collagen - biomaterial for drug delivery. Eur. J. Pharm. Biopharm. 1998, 45(2), 113-136.


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Biomineralization of Recombinant Peptide Scaffolds: Interplay between Chemistry, Architecture and Mechanics Kendell M. Pawelec*, Sebastiaan G.J.M. Kluijtmans


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Biomineralization of Recombinant Peptide Scaffolds - American

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