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Modification of Poly(propylene fumarate)−Bioglass Composites with Peptide Conjugates to Enhance Bioactivity Yanyi Xu,†,‡,⊥ Derek Luong,†,⊥ Jason M. Walker,§ David Dean,§ and Matthew L. Becker*,† †

Department of Polymer Science, The University of Akron, Akron, Ohio 44325, United States Department of Environmental Health, School of Public Health, Fudan University, Shanghai 200032, China § Department of Plastic Surgery, The Ohio State University, Columbus, Ohio 43210, United States ‡

S Supporting Information *

ABSTRACT: Poly(propylene fumarate) (PPF) has been highlighted as one of the most promising materials for bone regeneration. Despite the promising advantages of using polymer scaffolds for biomedical applications, their inherent lack of bioactivity has limited their clinical application. In this study, PPF was successfully functionalized with Bioglass and a novel catechol-bearing peptide bioconjugate containing bioactive short peptide sequences of basic fibroblast growth factor, bone morphogenetic protein 2, and osteogenic growth peptide. The binding affinity was assessed to be around 110 nmol/cm2 with the Bioglass content at 10 wt %. Fluorescence imaging studies show that the catechol-bearing modular peptide binds preferentially to the Bioglass. A 4 week in vitro cell study using human mesenchymal stem cells showed that cell adhesion, spreading, proliferation, and osteogenic differentiation at both gene and protein levels were all improved by the introduction of peptides, demonstrating the potential approach of dually functionalized polymers for bone regeneration.

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INTRODUCTION Bone repair and regeneration continues to be a challenging issue in orthopedics. Although bone autografts are considered the gold-standard clinical treatment in bone repair and regeneration, it remains a fact that there is no standard-ofcare “bone substitute”. Many bone tissue engineering approaches are viewed as more promising, such as the utilization of a synthetic, degradable material infused with stem cells and/or growth factors, which regenerate bone at the site while the material degrades away. Recent developments in degradable materials have concentrated on bioactive polymeric scaffolds with tunable degradation rates and mechanical properties.1 One such polymer, poly(propylene fumarate) (PPF), has been studied extensively as it is resorbable2 and nontoxic in vivo with an additional benefit of an unsaturated double bond for 3D printing purposes.2−7 Although PPF meets the stringent requirements of a polymeric biomaterial, its wide application in the bone regeneration field is hindered by its lack of osteoinductivity and osteoconductivity that most synthetic polymers face.8 Composite materials that incorporate ceramics such as hydroxyapaptite, 9 calcium phosphate, 10,11 and bioactive glasses12 have been utilized to impart osteoconductivity to naturally “bioinert” biomaterials such as polymeric scaffolds. In particular, bioactive glasses of certain compositions showed significant bony integration when implanted inside the body. 45S5 Bioglass is one such composition of bioactive glasses composed of 45 mol % SiO2, 24.5 mol % CaO, 24.5 mol % © XXXX American Chemical Society

NaO, and 6.0 mol % P2O5 that was developed by Larry Hench and co-workers in the 1960s.13 In depth mechanistic studies have shown that the SiO2 surface of Bioglass forms a bioactive carbonate-substituted hydroxyapaptite-like (HCA) layer. This layer is structurally and chemically similar to the mineral constituent of bone, resulting in a tight bond between bone and the implant through interfacial bonding. The formation of the HCA on the surface of the Bioglass with the release of ions such as Si, Ca, P, and Na eventually leads to recruitment/ activation of osteoprogenitor cells and bone formation.14−16 In addition to providing cues for osteogenic differentiation, studies have also shown that Bioglass and other bioactive glasses support vascularization through the stimulation of fibroblasts to secrete angiogenic growth factors.17,18 The osteogenic and angiogenic potential of these materials has spurred research into polymer/Bioglass composites and their potential orthopedic applications.19−21 The strategy of incorporating Bioglass additives into polymers has many advantages in addition to enhanced bioactivity, such as increasing the range of accessible moduli for tissue engineering22,23 and buffering of the decrease in local pH as a result of the acidic degradation byproducts that are generated from polyester polymers.21,24 The use of bioactive peptides to impart osteoconductivity and osteoinductivity in polymeric scaffolds has drawn a lot of Received: June 13, 2017 Revised: August 14, 2017

A

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Biomacromolecules Scheme 1. Molecular Structures of OGP-PEG-(Cat)4, bFGF-PEG-(Cat)4, and BMP2-PEG-(Cat)4

on Bone Morphogentic Protein 2 (73−92) and Basic Fibroblast Growth Factor (105−111), respectively (Scheme 1). We show that the incorporation of these bioconjugates increases osteogenic markers of MSCs at bone gene and protein levels, providing a new avenue of functionalization using the filler material as a binding source.

interest. The utilization of peptides has risen compared to native, circulating cytokines as a consequence of the ability to control negative pleiotropic effects, their ease of synthesis, and cost effectiveness. The ability to synthesize these peptides on the benchtop also allows for facile functionalization with molecules specific for binding onto surfaces. Integrin-derived proteins such as the RGD motif25,26 and bone-related proteins such as bone morphogenetic protein-2 (BMP-2)27 and osteogenic growth peptide (OGP)28 have been successfully incorporated into polymeric scaffolds. In addition, basic fibroblast growth factor (bFGF) has been used to enhance angiogenesis and osteogenesis.29 Delivering these potent biotherapeutics in a controlled manner is important, as much of its downstream signaling is concentration dependent. For example, it has been shown that in situ administration of BMP2 may result in negative effects such as inflammation, hematoma formation, neoplasm, and osteoclastogenesis.30,31 Methods to direct the activity of these peptides include microsphere drug carriers32 and surface tethering33 of the whole cytokine or only its active site to the polymer surface. We have recently synthesized bioconjugate-containing peptides with distinct bioactive and surface-binding domains.34 The bioactive domain, consisting of peptide OGP (10−14), remained bioactive when tethered to TiO2 surfaces, showing the upregulation of osteogenic markers in mesenchymal stem cells (MSCs) compared to control samples. The surface binding domain, consisting of dendritic, 4-armed catechol groups, binds to a variety of metallic-oxide surfaces, including TiO2, CeO2, SiO2, Fe3O4, and ZrO2. Inspired by the surface binding variety of this binding domain, we were pleased to find that the catechol molecule binds to Bioglass. In this study, poly(propylene fumarate) (PPF) was mixed with Bioglass to form a composite. This composite was further functionalized with OGP-PEG-(Cat)4 from our previous studies as well as new bioconjugates BMP2-PEG-(Cat)4 and bFGF-PEG-(Cat)4 based

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MATERIALS AND METHODS

Fmoc-protected amino acids were purchased from Novabiochem (San Diego, CA). All of the solvents were purchased from Sigma-Aldrich (St. Louis, MO), and all were reagent grade and used as received unless otherwise stated. 45S5 Bioglass was graciously donated by the Hench lab. Catechol-Bearing Modular Peptides Synthesis. Synthesis of Short Peptide Sequences. The synthesis of Fmoc-OGP-Resin, FmocbFGF-Resin, and Fmoc-BMP2-Resin was conducted in the solid phase using a Liberty 1 peptide synthesizer (CEM Cooperation, Matthews, NC). Wang resin (0.25 mmol) containing the appropriate C-terminal amino acid of the sequence was loaded into the synthesizer and subjected to various deprotection and coupling steps under microwave irradiation to yield the target peptide. The resin was then transferred to a peptide synthesis reaction vessel for the following steps. The amino acid sequences used were bFGF 105−111 (YKRSRYT), BMP-2 73−92 (KIPKASSVPTELSAISTLYL), and OGP 10−14 (YGFGG). PEG6 Coupling to Short Peptide Sequences. The synthesized peptide sequence was swelled in N,N′-dimethylformamide (DMF) for 15 min under nitrogen bubbling; then, the terminal Fmoc group was deprotected using a 25 mL cocktail of 20% Piperidine in DMF solution for 1 h. Afterward, the deprotection cocktail was removed, and the resin was washed with 3× DMF, 3× MeOH, and 3× DCM. The resin was resuspended in DMF, and a solution of Fmoc-NHPEG6‑Propionic acid (4 equiv, 1 mmol) (AAPPTEC, Louisville, KY) in DMF was added along with a solution of hydroxybenzotriazole (HOBt, 4 equiv, 1 mmol) (AAPPTEC) in DMF. To start the coupling reaction, N,N′-diisopropylcarbodiimide (DIC, 4 equiv, 1 mmol) was added. The coupling was run for 3 h to yield the PEGylated peptide sequence. B

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Biomacromolecules Dendron Coupling to PEGylated Short Peptide Sequences. The Fmoc-protected, PEGylated peptide sequence was deprotected using a 25 mL cocktail of 20% piperidine in DMF solution for 1 h. The resin was washed with 3× DMF, 3× MeOH, and 3× DCM. The resin was then resuspended in DMF, and a solution of Fmoc-Lys(Fmoc)−OH (4 equiv, 1 mmol) (AAPPTEC) in DMF was added along with a solution of HOBt (4 equiv, 1 mmol) (AAPPTEC) in DMF. To start the coupling reaction, DIC (4 equiv, 1 mmol) was added and run for 3 h to yield the bioconjugate-containing four-armed lysine groups. Catechol Coupling to Dendron. The conjugate containing four Fmoc protecting groups was deprotected using a 50 mL cocktail of 20% piperidine in DMF solution for 1 h. The resin was washed with 3× DMF, 3× MeOH, and 3× DCM. The resin was then resuspended in DMF, and a solution of acetal-protected 2,3-dihydroxyphenylpropionic acid (4 equiv, 1 mmol) (AAPPTEC) in DMF was added along with a solution of HOBt (4 equiv, 1 mmol) (AAPPTEC) in DMF. To start the coupling reaction, DIC (4 equiv, 1 mmol) was added and run for 3 h. Cleavage of Bioconjugate from Resin. Cleavage was performed by first washing the resin with 3× DMF, 3× MeOH, and 3× DCM. Next, a cleavage cocktail containing 15 mL of trifluoroacetic acid (TFA), 0.75 mL of triisopropylsilane (TIPS), and 0.75 mL of H2O was added. After 1 h, a second equivalent aliquot was added. After 2 total h of reaction, the solution was separated from the resin, and the TFA was removed under reduced pressure. The obtained peptide was redissolved in TFA and precipitated in ether three times. The precipitate was then dried and dialyzed against 70% ethanol for 1 day. The dialyzed product was dissolved in DMF and purified by a Waters 1525 HPLC system using an XBridge Peptide BEH C18 Prep Column (300 Å, 5 μm, 10 mm × 250 mm). A gradient was used from 20 to 80% methanol with 0.1% TFA with a flow rate of 3.3 mL/min. The purified products were freeze-dried and analyzed by MALDI-TOF mass spectrometry. Lys(Mca)-Labeled OGP-PEG-(Cat)4. Fluorescent analogues of the bioconjugates were prepared by adding Fmoc-Lys(Mca)−OH (4 equiv, 1 mmol) (AnaSpec, Fremont, CA) to the N-terminus of the peptide sequence. The following synthesis steps were identical to the regular bioconjugates. Fabrication of Bioglass-Encapsulated PPF Thin Films. Resin Creation. Resin was mixed as described previously with some modifications.35 Briefly, PPF (42 g) was weighed and mixed with DEF (42 g) in a 1:1 w/w ratio. While stirring at 60 °C, BAPO (2.52 g, 3% by wt of PPF/DEF) was slowly added followed by HMB (0.59 g, 0.7% by wt of PPF/DEF) and then Irgacure 784 (0.34 g, 0.4% by wt of PPF/DEF) to create the working resin solution. Bioglass (8.74 g, 10% by wt of the working solution) was added with the dispersant Darvan 811 (4.39 g, 5% by wt of the working solution). Resin was used while stirring constantly at 60 °C. Thin Film Preparation. To create Bioglass-encapsulated PPF thin films, approximately 1 mL of resin was pipetted onto each glass slide. A second glass slide was placed on top of the resin and was pressed to force the resin into a thin film covering the area of the glass slide. Slides containing resin pressed between them were then cured under UV light (365 nm) for 30 min. The slides were split apart using a razor blade, leaving the film of PPF resin on one of the two slides. A razor was used to strip the film from the slide, and the films were then cured under UV light for an additional 7.5 h to complete cross-linking under a thick glass cover to keep the films from curling. Conjugation of Catechol-Bearing Modular Peptides onto PPF/Bioglass Thin Films. Catechol-bearing modular peptides OGPPEG-(Cat)4, bFGF-PEG-(Cat)4, and BMP2-PEG-(Cat)4 were separately dissolved in DMSO at a concentration of 20 μM. The Bioglassencapsulated PPF thin films were then immersed into the corresponding solution (0.5 mL for one circular sample with a diameter of 6 mm) and incubated at ambient temperature overnight. After the reaction, thin films were thoroughly rinsed with DMSO three times to remove unbonded molecules and then rinsed with ethanol and stored dry for further tests. All solutions used here were filtered using Nylon filters (0.2 μm pore size) for sterilization, and all the processes were conducted in a sterile biohood.

Characterization of Peptide Functionalization onto PPF/ Bioglass Thin Films. Fluorescent conjugates prepared above were dissolved in DMSO to make a 20 μM solution. Ten PPF/Bioglass thin films (diameter of 6 mm) were immersed in 5 mL of solution for 12 h. Samples were washed with 2 mL of DMSO. The combined solution and washes were assayed by excitation at 340 nm, and the emission was measured at 405 nm. A standard curve was created by serially diluting the 20 μM solution to 0.625 μM. The amount of bioconjugate absorbed to PPF/Bioglass thin films was measured by taking the concentration of the solution before coating and subtracting the concentration of the solution after coating. Human Mesenchymal Stem Cell (hMSC) Culture and Seeding onto PPF/Bioglass Films. Male hMSCs (Lonza, Wakersville, MD) were cultured following manufacturer’s protocol using Lonza MSC growth medium (supplied with 10 vol % of FBS, 10 mL of L-glutamine, 30 μg/mL of Gentamicin, and 15 ng/mL of Amphotericin) without any osteogenic additives to avoid MSC differentiation. hMSCs within passages 3−5 were used for all the tests here, and earlier works have shown that cells within these passages maintained their multipotency.36 Cells were seeded onto PPF/Bioglass films at the density of 200 cells/mm2, and the growth media was changed every 2 days. hMSCs Viability and Adhesion on Peptide-Functionalized PPF/Bioglass Films in Vitro. Cell adhesion capability on films was detected and quantified using a CyQUANT cell proliferation assay kit (Invitrogen) following the manufacturer’s protocol 48 h after cell seeding onto OGP-PEG-(Cat)4, bFGF-PEG-(Cat)4, or BMP2-PEG(Cat)4 functionalized thin films (nonfunctionalized ones were used as the control). In brief, cell growth medium was aspirated, and samples were frozen and thawed at rt before adding CyQUANT GR dye/celllysis buffer mixture (1 mL/sample). After vortexing and incubation in the mixture for 10 min protected from light, sample supernatant fluorescence was measured using a fluorescence microplate reader with excitation at ∼480 nm and emission at ∼520 nm. A calibration curve was obtained using standard DNA at concentrations of 0, 10, 50, 100, 200, 400, 600, 800, and 1000 ng/mL. N = 3 replicates were studied for each group, and each sample was tested three times to eliminate pipeting error. hMSC Spreading on Peptide-Functionalized PPF/Bioglass Films in Vitro. hMSC spreading on thin films was observed by staining of cytoskeletal actin 48 h after cell seeding. All samples were collected and first prefixed in 3.7% paraformaldehyde in CS buffer for 5 min. After aspiration, they were fixed in 3.7% PFA solution for another 5 min and then thoroughly washed with 1× PBS. Triton X100 in CS buffer (0.5% v/v) was then added to permeabilize the cells for 10 min. The samples were washed with 1× TBS three times, and freshly prepared 0.1 wt % NaBH4 in 1× PBS was then added for 10 min at rt to quench the aldehyde fluorescence. After aspiration, the samples were incubated in blocking buffer (10 vol % donkey serum) for 1 h at rt, stained in vinculin primary antibody at 4 °C overnight, washed with 1× TBS three times, and stained with rhodamine phalloidin (v/v 1:40). Cell nuclei were then stained with DAPI (300 nM) for 15 min at ambient temperature in the dark. After washing with 1× TBS three times, samples were mounted and viewed under an IX81 Microscope (Olympus, Center Valley, PA) with mercury bulb excitation and filters of DAPI and TRITC. N = 12 areas were randomly picked and imaged for each group. Cell spreading area was calculated by drawing an area around the actin cytoskeleton in ImageJ and calculating the area for each cell. hMSC Proliferation on Peptide-Functionalized PPF/Bioglass Films in Vitro. Cell proliferation was measured using the CyQUANT cell proliferation assay following the manufacturer’s protocol. In brief, dsDNA was isolated and quantified as described above. All the data were normalized to day 1 to assess the cell proliferation rate on functionalized and nonfunctionalized PPF/Bioglass films using N = 3 replicates for each time point. hMSC Osteogenic Differentiation on Peptide-Functionalized PPF/Bioglass Films in Vitro. Reverse Transcription Polymerase Chain Reaction (RT-PCR). RNA extraction and isolation from samples was conducted using an RNeasy Mini kit following manufacturer C

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Biomacromolecules instructions (Qiagen, Valencia, CA). In brief, the cell-seeded thin films were first homogenized in 600 μL of lysis buffer, mixed with ethanol by pipetting, and added to the RNeasy Mini Column for total RNA isolation. DNase digestion was performed during the RNA isolation process using a Qiagen RNase-free DNase set (Qiagen, Valencia, CA). RNA quantity and purity were detected using a Take3 Multi-Volume Plate and a Synergy Mx Microplate Reader (BioTek, Winooski VT) at 260 nm. RNA was then reverse transcribed into cDNA using the Taqman Reverse Transcription Reagent kit (Life Technologies, Grand Island, NY) following the manufacturer’s protocol. The synthesized cDNA was stored at −20 °C for further tests. Real time RT-PCR was performed with a 7500 Real-time PCR System (Applied Biosystems) using SYBR Green Master Mix and designed primers. Ten nanograms of cDNA, 1× SYBR Green Master Mix, forward and reverse primers (209.4 nM for each), and a corresponding amount of DNase/RNasefree water were included in a 100 μL reaction mixture. Glyceraldehyde-3-phosphate dehydrogenase (GAPDH) was used as the housekeeping gene, and osteogenic markers Runx2, BSP, and OCN as well as endothelial marker CD31 were quantified. All of the primer information is listed in Table 1. Gene expression levels for hMSCs cultured in culture flasks were used as the control, and a standard ΔΔCt method was applied to calculate the single fold difference.

post hoc Tukey method using JMP software. Significant difference is defined as p < 0.05. Experiments were performed in triplicate unless otherwise noted.

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RESULTS Catechol-Bearing Modular Peptides Were Successfully Synthesized. The synthesis of catechol-bearing modular peptides bFGF-PEG-(Cat)4, BMP2-PEG-(Cat)4, and OGPPEG-(Cat)4 were successfully carried out via Fmoc-based solid phase synthesis method (Scheme S1). Products were purified by reverse-phase HPLC and confirmed by MALDI mass spectrometry (Figures S2−S4). Bioglass Was Successfully Encapsulated into PPF Thin Films and Functionalized with Synthesized CatecholBearing Modular Peptides and Proved Noncytotoxic. Bioglass was successfully and homogeneously encapsulated into the PPF thin films (Figure 1B). In addition, fluorescent

Table 1. Primers Used in Real-Time RT-PCR to Detect hMSC Differentiation primer

sequence

hRunx2

forward, GGACGAGGCAAGAGTTTCAC reverse, CAAGCTTCTGTCTGTGCCTTC forward, CCTGGCACAGGGTATACAGG reverse, CTGCTTCGCTTTCTTCGTTT forward, CATGAGAGCCCTCAC reverse, AGAGCGACACCCTAGAC forward, TCTATGACCTCGCCCTCCACAAA reverse, GAACGGTGTCTTCAGGTTGGTATTTCA

hBSP hOCN hCD31

Figure 1. (A) Scheme of functionalization of Bioglass-encapsulated PPF surfaces via catechol-bearing modular peptides OGP-PEG-(Cat)4, bFGF-PEG-(Cat)4, or BMP2-PEG-(Cat)4; (B) fluorescence image of PPF/Bioglass films before functionalization; (C) bright field image of PPF/Bioglass films before functionalization; and (D) fluorescence image of PPF/Bioglass films after functionalization with the Lys(Mca)functionalized conjugate.

Alkaline Phosphatase (ALP) Activity Test. ALP activity was measured at 1, 7, and 14 days by a SensoLyte pNPP ALP Assay Kit (AnaSpec Inc., San Jose, CA) following the vendor’s protocol. Growth media was aspirated from the samples, and 1× Assay Buffer from the kit was added to lyse the cells. The samples were stored at −80 °C until use. A standard curve was measured with an ALP solution at concentrations of 0, 3.1, 6.2, 12.5, 25, 50, 100, and 200 ng/mL. Fifty microliters of sample/standard solution and 50 μL of pNPP solution were added to each well in a 96-well plate. The solution was mixed by gently shaking for 30 s. After incubation for 1 h, the 96-well plate was shaken for 1 min before measuring the absorbance at 405 nm. Three replicates were measured for each sample (N = 3 for each group at each time point). The standard curve was fitted with a linear relationship by plotting absorbance vs ALP concentration. The ALP activity result was normalized by the total DNA amount of the corresponding samples, which was quantified with a CyQUANT assay following the previously described method. Histology Analysis. Calcium deposition conditions of the samples at 7 and 14 days were assessed with Alizarin Red S staining. Samples were fixed in a 3.7% paraformaldehyde (PFA) buffer for 1 h and washed three times with 1× PBS. Freshly made Alizarin Red S solution (0.8 g in 40 mL of ddH2O at pH 4.2) was then added to the substrates, and the samples were incubated at rt for 15 min. Next, the samples were dehydrated using increasing concentrations of ethanol (50, 75, 95, and 100% EtOH in DI water) followed by increasing concentrations of xylenes (50:50 EtOH/xylene and 100% xylene). The dehydrated samples were then mounted between glass slides using DPX mounting medium and observed under a bright field microscope. Statistics. All quantitative data are presented as average ± standard deviation. Statistical comparisons were performed by ANOVA with

analogues of the catechol-bearing modular peptides were synthesized by adding a coumarin-functionalized Lysine amino acid to the N-terminus of each molecule to obtain the binding affinity of the conjugate to PPF/Bioglass films. Binding studies using fluorescence microscopy revealed that the peptide conjugates were successfully bound to the PPF/Bioglass thin films. Further analyses showed that the peptide functionalization was concentrated at the Bioglass crystals on the surface of the PPF thin film (Figure 1B and D). Quantitative assessment based on a plate reader (excitation at 340 nm and emission at 405 nm) showed that 110.354 nmol/cm2 of peptide was chemically bonded to the Bioglass-encapsulated PPF thin films. hMSC Survival and Improved Adhesion on the Peptide-Functionalized PPF/Bioglass Films in Vitro. Surfaces functionalized with the catechol-bearing modular peptides showed similar cell survival ratios (Figure 2A and B). For all the experimental groups, above 85% of the cells survived successfully, indicating the nontoxicity of the PPF/ Bioglass films. In addition, significantly increased cell adhesion 48 h after cell seeding was also observed for peptidefunctionalized PPF/Bioglass films compared to that of nonfunctionalized controls (Figure 2C). hMSC Proliferation and Increased Spreading on the Peptide-Functionalized PPF/Bioglass Films in Vitro. hMSCs on PPF/Bioglass films functionalized with each of the three peptides showed an obviously larger spreading morpholD

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Figure 2. (A) Live/dead images of cells; (B) cell survival ratio based on live/dead assay result; (C) cell adhesion assessed by the cell DNA amount by CyQUANT assay on PPF/Bioglass thin films functionalized with or without bFGF-OGP-(Cat)4, OGP-PEG-(Cat)4, or BMP2-PEG-(Cat)4 48 h after cell seeding. *p < 0.01, one-way ANOVA.

Figure 3. Spreading of hMSCs on PPF/Bioglass films with or without functionalization of bFGF-OGP-(Cat)4, OGP-PEG-(Cat)4, or BMP2-PEG(Cat)4 3 days after cell seeding. (A) Immunohistochemical staining images of cells on films; red corresponds to F-actin in cytoskeleton and blue corresponds to cell nuclei; (B) Immobilized peptides increased cell spreading area; data was collected based on the IHC images (N = 12 for each group). *p < 0.01, one-way ANOVA.

ogy, whereas those on nonfunctionalized films were comparatively slim and smaller spreading circumference (Figure 3A). Quantitative assessment based on these F-actin staining results showed that cells on peptide-functionalized PPF/ Bioglass films had a significantly larger spreading area compared with those of the controls (Figure 3B). As expected, CyQUANT results showed that hMSCs seeded on PPF/ Bioglass composites proliferated for all experimental groups, with or without catechol-bound peptides, from 1 to 14 days. No significant differences on proliferation rate between the groups were found, indicating that the introduced peptides at a concentration of around 110 nmol/cm2 did not have a crucial influence on cell division (Figure 5A). Introduction of Peptides Enhances hMSC Osteogenic Differentiation Significantly at Both Gene and Protein Levels in Vitro. Gene expression of hMSCs seeded on PPF/ Bioglass composites with and without peptide functionalization was measured by quantitative reverse-transcriptase PCR (RTPCR). Key osteogenic markers, namely, RUNX-2, BSP, OCN,

and endothelial marker platelet endothelial cell adhesion molecule (CD31), were quantified and normalized to hMSCs cultured on tissue culture plates. As expected, hMSCs seeded on PPF/Bioglass composites showed comparable gene expression levels to those of the cell controls. In addition, cells on catechol-bearing modular peptide-functionalized PPF/ Bioglass films showed significantly improved expression of the middle stage osteogenic marker BSP and slightly increased expression of late stage osteogenic marker OCN, whereas the early stage marker RunX2 showed no significant difference when comparing functionalized PPF/Bioglass films (except for the bFGF-functionalized group) with the nonfunctionalized one. Surprisingly, CD31, as a crucial marker for endothelial differentiation, was significantly upregulated for all the peptidefunctionalized groups compared with the nonfunctionalized one (Figure 4). Consistent with the gene expression results of RunX2, ALP activity (ALP is also an early stage osteogenic marker) slightly decreased from 1 to 14 days with no significant differences E

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Figure 4. Gene expression of hMSCs on functionalized PPF/Bioglass thin films at 2 and 4 weeks after cell seeding. One star (*) indicates significance (p < 0.05) compared to cell control at the same time point. Two stars (**) indicates significance (p < 0.05) compared to the PPF/ Bioglass group at the same time point.

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DISCUSSION Previous work regarding Bioglass has shown sufficient promise to further explore composites of Bioglass with other polymers.15,37 In addition, our previous study has demonstrated the surface-binding capabilities of a catechol-bearing modular peptide bioconjugate to TiO2 surfaces while enhancing bioactivity for stem cells.34 On the basis of these studies, the work presented herein explores the dual functionalization of poly(propylene fumarate) (PPF) using bioceramics (i.e., Bioglass) and bioactive peptide bioconjugates (i.e., bFGF, BMP2, and OGP) to create a functional polymer scaffold for bone tissue engineering applications. In this study, the feasibility of PPF/Bioglass composites for bone tissue engineering applications were explored by culturing human mesenchymal stem cells (hMSCs) on the scaffods for up to 4 weeks. This was performed to assess the adhesion, proliferation, and differentiation of these multipotent cells into mature osteoblasts. In addition, these PPF/Bioglass composites were further functionalized with catechol-bearing modular peptides containing different bioactive peptide sequences. Basic fibroblast growth factor (bFGF) 105−111 (YKRSRYT), bone morphogenetic protein-2 (BMP-2) 73−92 (KIPKASSVPTELSAISTLYL), and osteogenic growth peptide (OGP) 10−14 (YGFGG) peptides were chosen because they have been extensively examined in previous studies of bone tissue engineering. bFGF 105−111 is derived from the heparin binding domain of basic fibroblast growth factor (FGF-2) and has been shown to increase bone marrow-derived hMSC cell attachment and osteoblastic differentiation in vitro and induce angiogenesis.38 BMP-2 73−92 peptide is derived from the “knuckle epitope” of the parent protein BMP2 and is one of the most widely studied peptides for bone regeneration purposes.27,39 OGP 10−14 is derived from the active domain of the parent protein OGP and has been shown to also enhance osteogenic proliferation, differentiation, and matrix mineralization in a variety of osteoblasts cell lines both in vitro and in vivo.28

Figure 5. (A) Cell proliferation and (B) ALP activity of hMSCs on PPF/Bioglass substrates with or without bioconjugate functionalization.

between the groups (Figure 5). This showed that these hMSCs may have already moved on to the next osteoblast differentiation stage (i.e., middle stage as demonstrated by BSP and late stage as demonstrated by OCN) and had begun to mature. In addition, the calcium deposition condition of the samples was also obtained using the Alizarin Red S staining approach. At 2 and 4 weeks, higher levels of calcium deposition were observed for all three ligand-bearing groups compared with those of the nonfunctionalized PPF/Bioglass group (Figure 6). This result is consistent with the improved gene expression of osteogenic markers BSP and OCN (Figure 4). F

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Figure 6. Mineralization of hMSCs on PPF/Bioglass substrates by Alizarin Red S staining with or without functionalization of bFGF-OGP-(Cat)4, OGP-PEG-(Cat)4, or BMP2-PEG-(Cat)4 at 2 and 4 weeks after cell seeding.

shown that the ionic dissolution products from Bioglass upregulate proliferation and differentiation of osteoblasts.41 Upon further functionalization with bFGF-PEG-(Cat)4, BMP2PEG-(Cat)4, and OGP-PEG-(Cat)4, an additional amplification in gene expression was observed. Interestingly, amplification is not seen early at the chosen time points, suggesting that osteogenic differentiation occurred quickly on these scaffolds, likely during the first week. RunX-2 and ALP are key markers early in osteoblastic differentiation and did not show continued upregulation after the first week. Indeed, the decrease in ALP expression after the first week is consistent with the osteoblast maturation evidenced by OCN upregulation and mineralization of the hMSC-secreted extracellular matrix observed in weeks 24. Indeed, ALP is crucial enzyme for ossification and observed early in hMSC differentiation. RunX-2, also an early stage osteoblastic differentiation marker, is a necessary transcription factor at this stage and regulates expression of later osteoblastspecific genes. RunX-2 gene expression was either the same or significantly lower than cell controls at both 2 and 4 weeks, whereas ALP activity as measured by a pNPP assay showed no differences between the groups. The disparity may indicate that cell differentiation had already moved on to the next stages before week 2. ALP activity decreases when mineralization has progressed and this was supported by the presence of calcium deposition at 2 weeks.42 Enhanced early differentiation was further verified by the significant enhancement of expression of the middle and late stage markers BSP and OCN in Bioglass-bearing and Bioglass/ peptide-functionalized specimens compared to cell controls. BSP is a component in mineralized tissues and plays an essential role in cell adhesion and matrix mineralization.43,44 The increase in BSP gene expression at 2 weeks shows that the Bioglass itself is bioactive enough to induce the early differentiation of hMSCs. Further functionalization with bFGF-PEG-(Cat)4 amplifies this effect. The additional amplification was not seen in the BMP2 and OGP peptideloaded groups (PPF/Bioglass + BMP2 and PPF/Bioglass + OGP, respectively). This may be attributed to the additional proliferative or differentiative effects of the bFGF peptide. bFGF is well-known for recruiting and differentiating stem cells toward an endothelial fate, and recent evidence suggests that the protein plays a role in the growth and patterning of the limb as well as bone homeostasis.29,45−47 In addition to enhancements seen in osteogenic markers, there was an increase in gene expression of platelet endothelial

Through an established solid phase synthesis procedure developed previously,34 catechol-bearing modular peptides consisting of bFGF, BMP-2, and OGP (named bFGF-PEG(Cat)4, BMP2-PEG-(Cat)4, and OGP-PEG-(Cat)4, respectively) were synthesized successfully (Figures S2−S4). A fluorescent analogue of OGP-PEG-(Cat)4 was synthesized by addition of a lysine reside (Fmoc-Lys-(Mca)−OH) at the Nterminus of the peptide sequence, and this molecule was used to quantify the amount of peptide bonding on PPF/Bioglass surfaces. The fluorescence studies showed that the molecule was bound to the PPF/Bioglass surface at a concentration of 110 nmol/cm2 with a given Bioglass concentration of 10 wt %. In addition, fluorescence microscopy showed that peptide binding was preferentially onto the Bioglass itself. This is not surprising considering that Bioglass is made of oxides such as SiO2 and CaO, which are surfaces that form strong coordinative bonds with the catechol functional groups. The preferential binding of the catechol-bearing modular peptides to Bioglass is advantageous as peptide loading of scaffolds can be directly controlled by changing the Bioglass content of the PPF/ Bioglass composite. This current in vitro study demonstrates the potential performance of dually functionalized PPF scaffolds for tissue engineering applications. PPF/Bioglass composites that were functionalized with bFGF-PEG-(Cat)4, BMP2-PEG-(Cat)4, and OGP-PEG-(Cat)4 showed increased cell adhesion (Figure 2) and cell spreading area (Figure 3) compared to those of the nonfunctionalized controls. However, the increased adhesion and spreading did not seem to affect proliferation in any way, as the proliferation rates for the three peptide coating groups and control were similar up to 14 days (Figure 5). The proliferative effects of the peptides and/or the Bioglass may have been masked by the cell seeding density. Even though the cell seeding density is considered “low,” it may have not been low enough to see any differences in proliferation.40 Analysis of the gene and protein expression showed that the PPF/Bioglass with and without peptide coating enhanced osteogenic differentiation compared to cell controls. This suggests that both the Bioglass and PPF may be inherently bioactive. A follow up qRT-PCR study on hMSCs seeded on PPF samples for 4 weeks showed undetectable levels of the genes analyzed in this study (BSP, OCN, RunX2, and CD31; Figure S5), implying that the incorporated Bioglass is the sole source for osteogenic differentiation in the nonfunctionalized PPF/Bioglass samples. Many reports in the literature have G

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cell adhesion molecule (CD31). CD31 is considered one of the defining markers for endothelial cells, which play an important role in the vasculature of bone remodeling.29 Unlike the results gathered for BSP, CD31 was only upregulated for the PPF/ Bioglass groups that were functionalized with bFGF-PEG(Cat) 4, BMP2-PEG-(Cat)4, and OGP-PEG-(Cat)4 . The upregulation in CD31 expression may be a result of the peptides’ role in proliferation and/or differentiation toward an endothelial fate. Although previous studies have suggested that dissolution products of Bioglass are angiogenic,18 this effect was not observed as the PPF/Bioglass group showed less expression than the cell control. Further studies into the dissolution kinetics and products of the Bioglass from this composite would provide a clearer picture. The enhanced CD31 expression (Figure 4) seen in bFGF over BMP2 and OGP supports the idea that bFGF provides additional proliferative/ differentiative effects. In combination with the previous observation that proliferation remained generally the same between all three peptide-functionalized groups, the amplified effects of bFGF-PEG-(Cat)4 seem to be mostly attributed to enhanced differentiation. Although middle to late stage markers BSP and CD31 were upregulated at 2 and 4 weeks, another marker, osteocalcin (OCN), did not show the same enhancements. This could be attributed to the timelines chosen as osteocalcin shows a very abrupt increase at the end stages of differentiation followed by a sharp decrease.42 Taken together, the observation of similar proliferation rates between the four experimental groups suggest that any enhancements in gene expression can be mostly attributed to the differentiative effects of the peptides and/or Bioglass used in this study. Further experiments to determine whether the increase in gene expression translates to a difference in protein expression (i.e., quantification of calcium deposition) will provide a more robust story. In the context of bone tissue engineering, the increased adhesion, spreading, and osteogenic differentiation seen in these PPF/Bioglass composites demonstrate their translational potential. A bone tissue engineering implant should fulfill several criteria, including being nontoxic, degradable, bioactive, and patient specific. Additive manufacturing techniques such as 3D printing are attractive as they can be used to fabricate defect-specific devices, such as bone tissue engineering implants, that match a bone defect seen in a patient's image (e.g., a 3D CT image). Although the initial studies performed herein were on PPF thin films to facilitate quantitation of biochemistry and imaging, translation to a 3D scaffold is straightforward as PPF composites containing hydroxyapatite have recently been fabricated into 3D structures.48,49 Current efforts are now focused on 3D printing these composites into scaffolds and implantation into a relevant animal model.

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Article

ASSOCIATED CONTENT

S Supporting Information *

The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acs.biomac.7b00828. Synthetic scheme (Figure S1) and mass spectrum data of the bioconjugates bFGF-PEG-(Cat)4 (Figure S2), BMP2-PEG-(Cat)4 (Figure S3), and OGP (OGP-PEG(Cat)4 (Figure S4) (PDF)

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AUTHOR INFORMATION

Corresponding Author

*Phone: 330-972-2834. E-mail: [email protected]. ORCID

Matthew L. Becker: 0000-0003-4089-6916 Author Contributions ⊥

Y.X. and D.L. contributed equally to this work.

Notes

The authors declare the following competing financial interest(s): The methods for synthesizing PPF by ring opening polymerization have been licensed by 21MedTech. M.L.B. and D.D. have financial interests in 21MedTech.

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ACKNOWLEDGMENTS The authors acknowledge partial support from the Army, Navy, NIH, Air Force, VA, and Health Affairs to support the AFIRM II effort under Award No. W81XWH-14-2-0004. The U.S. Army Medical Research Acquisition Activity is the awarding and administering acquisition office for Award No. W81XWH14-2-0004. Opinions, interpretations, conclusions, and recommendations are those of the authors and are not necessarily endorsed by the U.S. Department of Defense. M.L.B. acknowledges financial support from 21st Century Medical Technologies and the W. Gerald Austen Endowed Chair in Polymer Science and Polymer Engineering from the Knight Foundation. The authors also acknowledge Hernan LaraPadilla, Kevin Martin, and Mary Beth Wade for the PPF/ Bioglass thin film fabrication and Alex Kleinfehn and Yuanyuan Luo for the PPF synthesis.

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ABBREVIATIONS PPF, poly(propylene fumarate); OGP, osteogenic growth peptide; BMP2, bone morphogenetic protein 2; bFGF, basic fibroblast growth factor; hMSC, human mesenchymal stem cell

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REFERENCES

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CONCLUSIONS

The results of our study demonstrate that the dual functionalization of poly(propylene fumarate) with Bioglass and bioactive peptides enhances hMSC osteogenic differentiation. The combination of ceramics and peptide bioconjugates will be crucial to improve the mechanical properties and biological activity of these polymer scaffolds to meet the stringent demands of a clinically translatable biomaterial. H

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