Bilayered BMP2 Eluting Coatings on Graphene Foam by

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Bilayered BMP2 Eluting Coatings on Graphene Foam by Electrophoretic Deposition: Electro-responsive BMP2 Release and Enhancement of Osteogenic Differentiation Qingqing Yao, Jiajia Jing, Qingyan Zeng, TL Lu, Yu Liu, Xiao Zheng, and Qiang Chen ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.7b10180 • Publication Date (Web): 27 Oct 2017 Downloaded from http://pubs.acs.org on October 28, 2017

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ACS Applied Materials & Interfaces

Bilayered

BMP2

Eluting

Coatings

on

Graphene

Foam by

Electrophoretic Deposition: Electro-responsive BMP2 Release and Enhancement of Osteogenic Differentiation Qingqing Yao a, b, *, Jiajia Jing c, Qingyan Zeng c, TL Lu c, Yu Liu a, b, Xiao Zheng a, b, Qiang Chen c, * a

School of Ophthalmology and Optometry, Wenzhou Medical University, 270 Xueyuan Xi Road, Wenzhou,

Zhejiang 325027, China b

Institute of Advanced Materials for Nano-Bio Applications, Wenzhou Medical University, Wenzhou, Zhejiang

325027, China c Key Laboratory for Space Bioscience and Biotechnology, School of Life Sciences, Northwestern Polytechnical University, Xi’an, 710072, China * Corresponding authors: [email protected], [email protected]

ABSTRACT Recent development of three-dimensional graphene foam (GF) with conductive and interconnected macro-porous structure is attracting particular attention as platforms for tissue engineering. However, widespread application of GF as bone scaffolds is restricted due to its poor mechanical property and inert surface character. To overcome these drawbacks, in this study, a bilayered biopolymer coating was designed and successfully deposited covering the entire surface area of GF skeleton. A poly(lactic-co-glycolic acid) layer was firstly dip-coated to strengthen the GF substrate, followed by the electrophoretic co-deposition of a hybrid layer, consisting of chitosan and BMP2, to functionalize GF with the ability to recruit and induce osteogenic differentiation of hMSC. Our data indicated that the mechanical property of GF was significantly increased without compromising the macroporous structure. Importantly, the immobilized BMP2 exhibited sustained and electro-responsive release profiles with rapid response to the electric field exerted on GF, which is beneficial to balancing BMP2 dose in physiological environment. Moreover, the osteogenic differentiation of hMSC was significantly improved on the functionalized GF. Taking advantages of the unique macrostructure from GF as well as the superior mechanical properties and BMP2 release profile supported by the deposited coatings, it is therefore expected that the developed GF could be a promising alternative as innovative bone-forming favorable scaffolds. Keywords: Graphene foam, Coating, Electrophoretic deposition, Electro-responsive BMP2 release, Osteogenic differentiation 1

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1. INTRODUCTION The worldwide incidence of bone disorders and conditions remains a significant clinical challenge. Engineered bone tissue has been viewed as a promising alternative to autogenous and allogenous bone grafts, due to their donor site morbidity, disease transmission and immunoreactions.1-5 The combined use of stem cells, scaffolds and growth factors are three essential components for tissue engineering.6 Nowadays, 3D scaffolds are widely used in bone tissue engineering, especially, nanomaterials-based scaffolds are attracting more attention in medical diagnosis and treatment due to their excellent mechanical, optical, electrical, and magnetic properties.7 Bone morphogenic proteins (BMPs) were originally identified as unique proteins in demineralized bone matrix that induce ectopic bone formation upon implantation into muscular tissues.8,9 BMPs were later shown to regulate the differentiation and function of cells that are involved in bone and cartilage formation and degradation, including osteoblasts, chondrocytes, and osteoclasts.10 However, clinical application of BMPs on bone regeneration was impeded by several factors, such as: super-physiological dose requirements, high costs, short half-life, and other serious side effects, including the potential for ectopic bone formation, inflammation and even cancer.11-13 To address these challenges in bone tissue engineering, novel strategies should be developed to ensure a low dose use of BMPs with high stability and controllable release profile during bone-healing process. Graphene is a 2D layer of sp2 bonded carbon atoms, which has been widely investigated in nearly all fields of materials science due to its excellent electrical,14-17 thermal,14,18 and

mechanical

properties.19-21

Recently,

scaffolds

made

of

three-dimensional graphene foam (GF) is attracting particular attention as platforms for tissue engineering, for example, some studies indicated that application of electrical stimulation on GF could supported stem cell for neural and myoblasts differentiation.22,23 Though graphene foam was reported to slightly promote the differentiation of hMSC cells,24 it is not enough to promote new bone formation in vivo according to our preliminary data. In order to improve the osteogenic activity of GF, the surface of GF should be modified to promote osteogenic differentiation and 2


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further bone regeneration. It is therefore expected that the immobilization of BMPs on the surface of GF could be a promising alternative to develop innovative bone-forming favorable scaffolds, taking advantages of the strong osteogenic ability from BMPs and the unique microstructure from GF. Various surface modification techniques have been applied to immobilize biomolecules on bioinert surfaces for biomedical applications,25-27 however, the poor mechanical properties and the 3D porous structure of GF make it rather difficult to deposit BMPs homogeneously on the entire surface area of GF skeleton while maintaining their reactivity. Considering the conductive character of GF, we think about using electrophoretic deposition (EPD) to functionalize GF with BMPs in this study. EPD is a simple, rapid, and room-temperature coating technique that involves the movement of charged particles/molecules under an appropriate electric field, leading to their consolidation on the oppositely charged electrode to form films and coatings with high microstructural homogeneity and tailored thickness.28 With the proper design of EPD suspension, e.g. the addition of polyelectrolytes (chitosan ‘+” or alginate ‘–’) which act as the charger in suspension and as the binder in the deposit, functional molecules dispersed in EPD suspension are able to migrate and robustly bonded to the oppositely electrode under an electric field.29-31 Furthermore, the presence of amino or carboxylic groups on polyelectrolyte chains facilitated the pH-responsive property of the deposited coatings, which could be used to design stimuli-responsive drug delivery systems. Combining the conductive charter of GF and the pH-sensitivity of EPD coatings, it is expected that the electric field exerted on GF may play a role on controlling the BMP2 release from the deposited coatings. In the present work, for the first time, we explored the use of EPD method to functionalize the GF with BMP2 eluting coatings to develop innovative GF scaffolds for bone tissue engineering. A thin PLGA layer was firstly dip-coated on GF substrates to support the mechanical properties, a second layer consisting of chitosan and BMP2 was then electrophoretically deposited aiming to enhance the osteogenic activity. EPD parameters were optimized to produce robust coatings without compromising the porous structure of raw GF. The mechanical properties, BMP2 3



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release profiles under the influence of electric field as well as the osteogenic properties of the coated GF were thoroughly evaluated by means of a combination of relevant characterizations. 2. EXPERIMENTAL 2.1 Materials. Chitosan powder (CTS, medium molecular weight, CAS: 9012-76-4) was purchased from Sigma Aldrich. BMP2 powder was purchased from Zhong Ke Wu Yuan Biotechnology Co., Ltd., Beijing, China). Poly(lactic-co-glycolic) powder (PLGA, mean MW of ~10kDa, LA:GA of 75:25 in mol %) was purchased from Daigang Biomaterial Co., Ltd (Jinan, China). All other chemicals were supplied by Sinopharm Chemical Reagent Co,Ltd (Xi’an, China). 2.2 Preparation of Graphene Foam. GF was synthesized via a modified CVD method as described elsewhere.6 Typically, nickel foam (250-450 g/m2, 20×20×2 mm3) were used as the templates for the CVD growth of GF scaffold. The nickel foams were placed at the center of a quartz tubular furnace and heated at 1000 °C under N2 (200 s.c.c.m.) and H2 (60 s.c.c.m.) to remove the thin surface oxide layer on their surface. A small amount of CH4 (50 s.c.c.m) was then introduced into the reaction tube at ambient pressure for 10 min, leading to the growth of an ultra-thin graphene layer on the Ni template. A second layer of poly(methyl methacrylate) (PMMA) was dip-coated on the surface of the graphene films as a support to prevent the graphene network from collapsing during the removal of the nickel template (4 wt% of PMMA in ethyl acetate). The Ni template was then removed by soaking the graphene coated Ni foam into the diluted HNO3/H2O2 solution for 72h. GF scaffolds were subsequently obtained after removing PMMA layer in a hot acetone solution. 2.3 Coating Process. 2.3.1 Dip-coating of PLGA Layer. 5 wt% of PLGA solution was prepared using a mixture of DCM and DMF (7:3 in vol%) as the solvent. The as-prepared GF was slowly dipped in PLGA solution for 30 s, withdrawn and immediately laid on the filter paper to remove the residual solution. 4


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The coated samples were then vaccum-dried at 37 oC for 24 h to completely remove the organic solvent. 2.3.2 EPD of CTS-BMP2 Layer. Chitosan solution was prepared by dissolving chitosan powder in an aqueous solution (containing 1 vol% of acetic acid) under continuous magnetic stirring. BMP2 powder was then added into chitosan solution with magnetic stirring for 10 min. In order to suppress the electrolysis of water, ethanol was added dropwise into the blend suspension under vigorous magnetic stirring. The composition of the final solvent for EPD was water-acetic acid-ethanol mixture (49:1:50 in vol%), in which the concentrations of chitosan and BMP2 were 1 mg/ml and 300 ng/ml, respectively. In order to elucidate the possible interaction between BMP2 and chitosan molecules in EPD suspension, the zeta potential and size distribution of BMP2, chitosan, and their mixtures with different ratios (in wt%) dispersed in the same solvent for EPD, were measured based on dynamic light scattering (DLS) technique (Zetasizer nano ZS, Malvern, UK). The EPD cell included a cathodic substrate (PLGA coated GF) centered between two parallel 316L stainless steel counter electrodes, as indicated in Fig. S1. The distance between the cathode and the counter electrode was 10 mm and EPD was performed using constant current (CC) mode. A trial-and-error optimization process of both current density (1-20 mA/cm2) and deposition time (10-300 s), the deposition process was finally carried out using the current density of 4 mA/cm2 and deposition time of 90 s yielding a relative thick and homogeneous coating while maintaining the interconnected porous structure of GF. After deposition, the electrodes were slowly withdrawn from the EPD suspension and dried horizontally at room temperature. 2.4 Coating Characterizations. The layer structure and defect states of the as-prepared GF were analyzed using SEM (Zeiss Leica) and Raman spectrometer (Thermo Scientific Nicolet Almega XR) with an excitation wavelength of 532 nm. The cross-sectional morphology of GF at different coating stages (bending in liquid nitrogen) was observed using scanning electron microscopy (SEM, Auriga, Zeiss, Germany). The thickness of the coatings 5



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was measured from at least 15 different areas of the struts in the SEM images using Nano Measurer 1.2 software. The weight of the GF (mg/cm2) after each coating stage was recorded using an analytical balance (accuracy of 0.01 mg). Tensile strength of the bare and coated samples was characterized using an Instron 3400 tensile tester. The dimension of the specimen for tensile test is approximately 15×20×1 mm3with a crosshead speed of 0.5 mm/min. At least 7 specimens were tested for each sample condition. 2.5 Long-term and electro-response BMP2 Release. BMP2 release from the GF/BMP2 (control group) and GF/PLGA/CTS-BMP2 scaffolds was carried out in PBS buffer at 37 oC. The control group was prepared by soaking GF scaffold in BMP2 solution (300ng/ml) for 1h at room-temperature, followed by removal of residual liquids inside the scaffold. Both of the samples were immersed in 2 ml PBS at 37 ºC, and the supernatant was completely collected and refreshed with identical volume of PBS at 1, 4, 8, 24, 96, 168, 240 h after incubation. The possible presence and the amount of BMP2 was determined using human BMP2 ELISA kit according

to

the

manufacture’s

instruction.

To

evaluate

the

electro-responsive BMP2 release, the coated GF was used as the cathode and immersed in PBS solution (10 ml) while a platinum plate was used as the anode with a distance of 10 mm. The BMP2 release was measured by applying cathodic voltages of −1.5, 0 (control), +1 V, respectively. The applied potential values for adjusting drug release are in agreement with a relevant investigation,32 which is high enough to induce the change of local pH around GF scaffold while avoiding visible bubble formation by water electrolysis. At predetermined time intervals (0.5, 1, 2, 4, 8 h), 0.5 ml of the released medium was taken out and refilled with 0.5 ml of fresh PBS to balance the drug release. The concentration of BMP2 was measured in triplicate using ELISA to calculate the cumulative BPM-2 release under different electric field conditions. 2.6 Cell Culture. 2.6.1 Cell Seeding on GF and GF/PLGA/CTS-BMP2. GF and GF/PLGA/CTS-BMP2 were firstly cut into cubes (10 mm ×10 mm ×2 6


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mm thickness) with tissue punch, and then sterilized in 70 vol% ethanol for 30 min followed by rinsing with PBS for 3 times. Subsequently, these samples were incubated in minimum essential medium α (α-MEM, Gibco, Waltham, MA) for 30 min. The residual medium on a scaffold was removed with sterile gauzes before human mesenchymal stem cells (hMSCs) were seeded into the scaffold. All of the cells/scaffolds were cultured in 24-well plate with an orbital shaker (30 rpm) at 37 °C with 5% CO2. 2.6.2. Cell Viability. Cell viability was quantitatively analyzed using CellTiter 96 AQueous One Solution Cell Proliferation Assay (MTS, Promega, USA) according to the manufacture’s instruction. Briefly, after culturing for 1 and 3 days, the culture medium was removed; fresh medium with 10% MTS was then added, and incubated at 37 °C with 5% CO2 in dark for 1 h. The absorbance was measured at 490 nm using a microplate reader (Infinite M200, Tecan, USA). GF was chosen as the control group. The relative cell viability (%) was expressed as percentage relative to the control group. hMSCs morphologies on GF and GF/PLGA/CTS-BMP2 were visualized by staining with Texas red-X Phalloidin (Life technologies, OR, USA) and DAPI (Southern Biotech, Birmingham, AL), which could label F-actin and cell nuclear, respectively.33 Briefly, cells seeded on the scaffold were fixed in 3.7% paraformaldehyde for 10 min and the permeabilized with 0.1% TritonX-100 for another 5 min. Thereafter, the samples were blocked with 1% bovine serum albumin for 30 min before they were stained with Texas red and DAPI for 20 and 5 min, respectively. The stained samples were examined using a laser scanning microscope (FV1200, Olympus, Japan). 2.6.3. ALP Activity and Calcium Content. ALP activity was carried out using an EnzoLyte pNPP Alkaline Phosphatase Assay Kit (AnaSpec, San Jose, CA), as we previously described with minor modifications.33 Briefly, after 7 days of culture, cells/scaffolds (GF, GF/PLGA/CTS, GF/PLGA/CTS-BMP2) were rinsed with PBS solution and lysed with lysis buffer for 1−2 min at room temperature. The lysate was then transferred into a tube and 7



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centrifuged for 15 min at 2500 g at 4 °C.

The collected supernatant or standard

solution (50 µl) was mixed with p-nitrophenyl phosphate and incubated for 30 min at 37 °C. Following the incubation, the reaction was stopped by adding 100 µl terminated liquid. ALP activity was measured at 405 nm and normalized against total protein content. The total protein content was measured with a BCA kit (Thermo Scientific™, Waltham, MA) according to the manufacture’s instruction. Briefly, 25 µl of the collected supernatant (the same from ALP activity) or standard solution was mixed with 200 µl BCA working reagent and incubated for 30 min at 37 °C. Following the incubation, the protein content was measured at 562 nm. The cell-scaffold constructs were also examined for calcium deposition by using a total calcium LiquiColor® kit (Stanbio laboratory, TX). After 3 weeks of culture, cells/scaffolds were rinsed with PBS and cut into small pieces with a sharp blade. The calcium was extracted by using 1 ml 6 M hydrochloric acid. Thereafter, 10 µl extraction solution or 10 µl standard solution was added into 1 ml working solution prepared according to the manufacturer’s instruction. The absorbance was measured at 550 nm, and the calcium content was calculated from the following equation: Calcium ( mg / dL ) =

Au × 10 As

where Au and As are the absorbance values of sample and standard, respectively.

2.6.4. Bone Related Gene Expression. Quantitative gene expression analysis was carried out as we previously described with minor modifications. 33 Briefly, total RNA was extracted using the GeneJET™ RNA Purification Kit (Thermo Scientific™, Waltham, MA) by following the manufacturer’s instruction. RNA concentration was measured with UV−vis spectroscopy (DU 730, Beckman coulter) at 260 nm. An equivalent amount of RNA was processed to generate cDNA by using the High Capacity cDNA Reverse Transcript kit purchased from Applied Biosystems (Forster City, CA). Quantitative PCR was performed with Taqman gene expression assays (Applied Biosystems, Forster City, CA) using the Applied Biosystems 7500 Fast Real-Time PCR System (Applied Biosystems, Carlsbad, CA). Triplicates were performed for each sample, and 8


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results were normalized to β-actin. Gene primers of β-actin (Hs01060665), ALP (Hs00758162) and Bone sialoprotein (BSP) (Hs00173720) were purchased from Applied Biosystems (Forster City, CA).

2.6.5. Statistic Analysis. To determine statistical significance of observed differences between the study groups, a two-tailed homoscedastic t-test was applied. A value of p < 0.05 was considered to be statistically significant. Values are reported as the mean ± standard deviation (SD).

3. RESULTS and DISCUSSION. 3.1. Physical-chemical Characterization. In this study, the structure of as-prepared GF was firstly analyzed using both Raman spectroscopy and SEM. As shown in Fig. 1(a), the D band which is related to the defect of graphene, was not observed on the Raman spectrum, suggesting that a high quality of graphene was prepared. The presented ratio between 2D and G band indicated that a few-layer of graphene was formed during the CVD process on Ni templates.34,35 In order to keep the 3D structure of GF and to bear the liquid capillary force caused during removal of PMMA layer, several layers of graphene are necessary, otherwise the GF should tend to collapse and break into small pieces. The SEM image of GF scaffold showed an interconnected structure (Fig. 1(b1)) and continuous growth of graphene on the surface of structs (Fig. 1(b2)). Moreover, wrinkles and ripples were observed on the as-synthesized graphene.

9



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Fig. 1. (a) Raman spectrum and (b1, b2) SEM images of as-prepared graphene foam. In order to enhance the mechanical property of raw GF, a thin PLGA layer was dip-coated on the GF at first. Chitosan-BMP2 layer was then deposited on the GF/PLGA from their blend suspension via EPD. The weight of GF at different coating stages was shown in Fig. 2. The weight of GF (5.04±0.6 mg/cm2) was increased after coating with PLGA layer (9.61±0.4 mg/cm2) and the further deposition with chitosan-BMP2 layer (10.29±0.1 mg/cm2). These results demonstrated that PLGA and chitosan were successfully deposited onto the GF.

Fig. 2. the deposit mass as a function of holding time at different coating stages. Cross-sectional morphologies of GF, GF/PLGA and GF/PLGA/CTS were shown in Fig. 3. Raw GF showed interconnected macroporous structure with the pore size of 309±86 µm (Fig. 3 (a1), and the GF could retain the macroporous structure after PLGA and chitosan decoration (Fig. 3 (b1, c1)). The GF struts showed a hollow 10


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interior tubular structure with multiple layers of graphene on the exterior (Fig. 3 (a3)). The polymer layers were also observed inside the GF struts after dip-coating with PLGA (Fig. 3(b3)) and EPD with CTS-BMP2 (Fig. 3(c3)). The thickness of PLGA and chitosan layers was measured to be 395±60 nm and 51±8 nm, respectively. The PLGA and CTS-BMP2 layers seemed to be homogeneously and tightly deposited on the surface of GF (Fig. 3(b2, c2)), which could provide enhanced mechanical strength and uniform distribution of BMP2 in the coating.

Fig. 3. Cross-sectional overview, outer and inner surface of the struts on (a1, 2, 3) GF, (b1, 2, 3) GF/PLGA and (c1, 2, 3) GF/PLGA/CTS samples, the arrows in figures show the presence of polymer layers. The tensile strength of the prepared scaffolds was evaluated by macroscale tensile testing. The characteristic stress-displacement curves of the samples at different coating stages were shown Fig. 4(a), in which the typical morphology of failure crack is also included. Raw GF samples failed typically at a displacement of 0.1 mm, while GF/PLGA reached a value of 0.18 mm and GF/PLGA/CTS reached a value of 0.22 mm, indicating a significant increase of the ductility after the coating treatment. In addition, the tensile strength of GF scaffolds, as listed in Table 1, increased drastically from 47.2±8 kPa of raw GF samples to 126.1±6 kPa of PLGA coated GF samples. 11



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The tensile strength was further improved was also detected after the chitosan coating (150.1±18 kPa), which is about three folds higher than that of raw GF. As shown in Fig. 4(b), qualitative evaluation was carried out to visualize the robustness of the GF/PLGA/CTS sample, which indicated that the coated GF scaffold possessed an excellent flexibility and high porosity.

Fig. 4. (a) Characteristic stress-displacement curve of the samples at different coating stages, the inset shows the characteristic signs of cracking on GF, (b) qualitative tests showing the flexibility and ductility of the GF/PLGA/CTS sample. Table 1 Tensile strength of the sample at different coating stages. Sample code Tensile strength/kPa

GF

GF/PLGA

GF/PLGA/CTS

47.2±8

126.1±6

150.1±18

3.2. BMP2 Release profiles It is expected that the immobilized BMP2 molecules are able to release in a controllable manner, in order to regulate the BMP2 induced osteogenic activities and meanwhile minimize the potential side effects. The long-term BMP2 release from GF/BMP2 and GF/PLGA/CTS-BMP2 scaffolds were shown in Fig. 5(a). Burst release was observed on both of scaffolds during the first 8 h, however, the GF/PLGA/CTS-BMP2 scaffold showed lower burst release compared to GF/BMP2 scaffolds. BMP2 on the GF/BMP2 scaffold was released completely in 24 h, while the GF/PLGA/CTS-BMP2 scaffold exhibited a sustained BMP2 release up to 10 days. 12


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The total amounts of BMP2 released from GF/BMP2 and GF/PLGA/CTS-BMP2 scaffolds were approximately 12.8 ng and 19.4 ng, respectively. These data indicated that the EPD of CTS layer is promising as the platform for high loading and controlled release of BMP2. In addition, the detectable release of BMP2 up to 10 days indicated that the CTS layer plays a role on protecting BMP2 from degradation. The possible effect of the electric field on BMP2 release profiles was studied under different applied potential (–1.5, 0, +1 V), as shown in Fig. 5(b). The cumulative release of BMP2 was normalized by the weight of CTS layer for a comparative analysis. Burst release of BMP2 during the first 2 h was observed when all the samples were immersed in PBS medium.

Compared to the control group (0

V), retarded BMP2 release was detected when applying a positive voltage (+1 V). In addition, application of a negative voltage (–1.5 V) led to substantially accelerated BMP2 release compared to the control group. Significant difference of BMP2 release profiles was observed starting from the first detection point (30 min), indicating a rapid response of BMP2 release with respect to the influence of electric field. It should be mentioned Fig. 5(b) presented the BMP2 release only in the first 8 h (rather than saturation stage) in order to show clearly their different behavior of BMP2 release influenced by electrical triggers. In addition, no detectable weight loss of the sample was found after the release experiment (using analytical balance with an accuracy of 0.1 mg). Considering the relative low degradation rate of chitosan in vitro,36 we consider that the chitosan layer in current study should retain its microstructure after the release test.

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Fig. 5. (a) BMP2 release profiles from GF/BMP2 and GF/PLGA/CTS-BMP2 scaffolds in PBS at 37 ºC for a period of 10 days (n=3), (b) Cumulative BMP2 release profiles on GF/PLGA/CTS samples under different applied voltages (n=3). 3.3. MTS and Live/Dead Staining. The cell viability of hMSCs on GF and GF/PLGA/CTS-BMP2 scaffolds were quantitatively measured by MTS assay after culture for 1 and 4 days. As shown in Fig. 6(a), GF/PLGA/CTS-BMP2 scaffold showed much higher cell viability than GF scaffold (P

Bilayered BMP2 Eluting Coatings on Graphene Foam by

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