Biomimetically Ornamented Rapid Prototyping Fabrication of an

Nov 9, 2015 - Here, an apatite–collagen–polycaprolactone (Ap-Col-PCL) composite construct was developed with unique nano–micro–macro hierarchi...
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Biomimetically Ornamented Rapid Prototyping Fabrication of an Apatite−Collagen−Polycaprolactone Composite Construct with Nano−Micro−Macro Hierarchical Structure for Large Bone Defect Treatment Jinbing Wang,† Dingyu Wu,‡,§ Zhanzhao Zhang,‡,§ Jun Li,† Yi Shen,† Zhenxing Wang,‡,§ Yu Li,‡,§ Zhi-Yong Zhang,*,‡,§ and Jian Sun*,† †

Department of Oral and Maxillofacial-Head and Neck Oncology, Shanghai Key Laboratory of Stomatology and ‡Department of Plastic and Reconstructive Surgery, Shanghai Key Laboratory of Tissue Engineering, Shanghai Ninth People’s Hospital, Shanghai Jiao Tong University School of Medicine, 639 Zhizaoju Road, Shanghai 200011, PR China § National Tissue Engineering Center of China, 68 Jiang Chuan East Road, Shanghai 200241, PR China ABSTRACT: Biomaterial-based bone graft substitute with favorable mechanical and biological properties could be used as an alternative to autograft for large defect treatment. Here, an apatite−collagen−polycaprolactone (Ap-Col-PCL) composite construct was developed with unique nano−micro−macro hierarchical architectures by combining rapid prototyping (RP) fabrication technology and a 3D functionalization strategy. Macroporous PCL framework was fabricated using RP technology, then functionalized by collagen incorporation and biomimetic deposition. Ap-Col-PCL composite construct was characterized with hierarchical architectures of a nanoscale (∼100 nm thickness and ∼1 μm length) platelike apatite coating on the microporous (126 ± 18 μm) collagen networks, which homogeneously filled the macroporous (∼1000 μm) PCL frameworks and possessed a favorable hydrophilic property and compressive modulus (68.75 ± 3.39 MPa) similar to that of cancellous bone. Moreover, in vitro cell culture assay and in vivo critical-sized bone defect implantation demonstrated that the Ap-Col-PCL construct could not only significantly increase the cell adhesion capability (2.0fold) and promote faster cell proliferation but also successfully bridge the segmental long bone defect within 12 weeks with much more bone regeneration (5.2-fold), better osteointegration (7.2-fold), and a faster new bone deposition rate (2.9-fold). Our study demonstrated that biomimetically ornamented Ap-Col-PCL constructs exhibit a favorable mechanical property, more bone tissue ingrowth, and better osteointegration capability as an effective bone graft substitute for critical-sized bone defect treatment; meanwhile, it can also harness the advantages of RP technology, in particular, facilitating the customization of the shape and size of implants according to medical images during clinical application. KEYWORDS: rapid prototyping technology, biomimetically functionalization, hierarchical structure, bone graft substitute, bone defect treatment, osteoconduction, osteointegration, bone regeneration brittleness, rigidity, and slow degradation rate.3−5 Natural polymers are advantageous for their biodegradability and favorable bioactivity but are hindered by their weak mechanical property for load-bearing applications.5−7 Synthetic polymers own good mechanical property, biodegradability, and easy processability but are limited by their inherent poor bioactivity, leading to poor cellular attachment and fibrosis encapsulation.4,8−11 To circumvent the inherent drawbacks of these materials and make full use of their advantages, a hybrid strategy was utilized to develop a composite material, which was made of ceramics and natural and synthetic polymers.12

1. INTRODUCTION Bone is one of the most frequently injured organs in humans, often requiring effective bone grafts to promote the healing of bone defects caused by congenital disorders, traumatic injury, or surgery for bone tumors.1 Currently, the treatment of large bone defects remains a major challenge. As the gold standard of the clinical practice, autologous bone tissue transplantation has been widely used since the early 1900s. Nevertheless, the use of autograft is often limited by graft supply, size, and donor-site morbidity.1,2 Therefore, biomaterial-based bone graft substitutes overcoming these limitations have been extensively investigated as alternatives for bone defect treatment.3 Currently, the most frequently used materials to construct bone graft substitute include ceramics and natural or synthetic polymers. Ceramic materials display excellent osteoconductivity and biocompatibility; however, ceramics are limited by their © XXXX American Chemical Society

Received: September 10, 2015 Accepted: November 9, 2015

A

DOI: 10.1021/acsami.5b08534 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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mechanical and biological characteristics, having the potential to be an effective bone graft substitute for large bone defect treatment.

Besides the material selection, the architecture of the scaffold is also very important. Previous studies indicated that scaffold with high porosity and interconnected pore networks could benefit cell migration, differentiation, and bone regeneration.13 As a fully computer-aided procedure, rapid prototyping (RP) technology enables the precise design and fabrication of implants that are highly reproducible and fully customizable with shape, size, and desired internal microstructure according to the medical images, such as CT or MRI imaging. Therefore, RP technology has quickly emerged as a promising technique to fabricate bone graft substitute with desired mechanical function, custom-tailored architectures, and anatomical shapes.14 RP-fabricated porous 3D PCL scaffolds have been used as bone graft substitutes in preclinical and clinical applications.8,9,15−18 However, the performance of PCL is generally limited by a suboptimal biological interaction with cells in vitro and fibrosis encapsulation with in vivo implantation. Efforts have been made to improve the bioactivity by combining them with specific bone extracellular matrix (ECM) components, such as calcium phosphate based particles. Nonetheless, use of this strategy is complicated by the difficulty of identifying optimal amounts and combinations of defined factors.17 Inspired by the components of bone (a hierarchical structure of mineralized collagen composite), several studies have investigated biomimetic strategy to fabricate hierarchical mineralized collagen scaffolds for bone tissue engineering.19−22 These scaffolds with a hierarchical architectures have been shown enhance cell viability and greater osteogenesis. In addition, the architectures and mechanical strength of the collagen−apatite scaffolds could be easily tailored by freezing conditions, collagen concentrations, and cross-linking methods.3,23,24 However, the compressive strength is commonly ∼200−300 kPa, less than ideal for surgical handling, fixation, and bearing osteogenic loads during healing.21,25 Previously, Zhou et al.26 fabricated a three-level hierarchical calcium phosphate/collagen/hydroxyapatite composite by combining microwave sintering with biomimetic synthesis; this composite showed favorable osteoinductivity and mechanical properties (with a compressive modulus of 352 ± 37 MPa). However, the traditional sintering method cannot precisely control the architectures, leading to poor interconnectivity and necrosis core structures, hindering cell ingrowth and flow transport of nutrients and metabolic waste.13,27 To construct a scaffold that matches function with structure, the combination of RP technology and biomimetic functionalization fabrication strategy to construct a hierarchical composite scaffold consisting of natural and synthetic polymers as well as biomimetic ceramics holds great promise. Specifically, we fabricated a 3D PCL construct using RP technology; the construct possesses at the first level macroporous frameworks with favorable mechanics to support nascent bone tissue ingrowth. Then, we meticulously functionalized this PCL framework through collagen evacuation to formulate the second level of microporous networks to mimic the ECM of bone. Finally, a third level of bonelike nanoapatite was introduced by biomimetic deposition. This apatite−collagen− polycaprolactone (Ap-Col-PCL) composite construct with unique nano−micro−macro hierarchical structure can not only harness the manufacturing advantages of RP technology, especially the capability for custom-made shape and size according to medical images, but can resemble natural bone’s hierarchical structures and compositions with favorable

2. MATERIALS AND METHODS 2.1. Materials. PCL was obtained from Shenzhen Esun Industrial Co., Ltd., China. Collagen was obtained from Sichuan Mingrang Tech., China. Alizarin Red S, calcein, and collagen cross-linking agent N-(3-(dimethylamino)propyl)-N-ethylcarbodiimide hydrochloride crystalline (EDC) were purchased from Sigma-Aldrich, China. Simulated body fluid (SBF) solution was prepared in accordance with Kokubo’s methods,28 and all reagents for the preparation of SBF were purchased from Sinopharm Chemical Reagent Co., Ltd., China. 2.2. PCL Constructs Fabrication and Functionalization. PCL constructs were fabricated with a filament diameter of 500 μm and channel size of 1000 μm, with a 0−60−120° lay-down pattern as previously described.29 The surface-activated PCL (A-PCL) prepared by NaOH (5 M, 37 °C) treatment for 24 h, was used for the subsequent experiments.30,31 For functionalization, A-PCL constructs were immersed in the collagen solutions (2 wt %) in 0.1 mol/L acetic acid and placed under a vacuum for 10 min, frozen at −80 °C, freezedried at −50 °C, subsequently cross-linked by 50 mM EDC solutions for 24 h, thoroughly washed with 5 wt % glycine solution and distilled water three times, and freeze-dried a second time at −50 °C to obtain collagen−polycaprolactone (Col-PCL) composite constructs.3 2.3. Biomimetic Apatite Deposition. To accelerate biomimetic apatite deposition, the Col-PCL constructs were alternately soaked (AS) in CaCl2 solution and K2HPO4 solution to obtain a nucleation site for the next biomimetic coating as previously described.32,33 In brief, the Col-PCL constructs were soaked in 20 mL of 0.2 M CaCl2 aqueous solution for 3 min and then dipped in deionized water for 5 s, followed by air drying for 3 min. The samples were subsequently dipped in 20 mL of 0.2 M K2HPO4 aqueous solution for 3 min and then dipped in deionized water for 5 s, followed by air drying for 3 min. This process was repeated three times. The AS-treated Col-PCL constructs were immersed in SBF at 36.5 °C for biomimetic apatite deposition; after a 5 min vacuum treatment, the samples were placed at 37 °C for 24 h (AS-SBF 24 h), 72 h (AS-SBF 72 h), and 1 week (AS-SBF 1w). The SBF solution weas renewed every 12 h. 2.4. Field-Emission Scanning Electron Microscopy. The surface morphology of the scaffold was observed by field-emission scanning electron microscopy (FESEM, Hitachi, S-4800) at a beam intensity of 10 keV. The samples were gold-coated for 20 s at 10 mA before observation. Energy-dispersive X-ray spectroscopy (EDS) was carried out to quantitatively measure the calcium and phosphate ratio of the biomimetically deposited constructs (n = 5). 2.5. Micro-CT, TGA, XRD, and FTIR Analysis. Micro-CT (GE eXplore Locus SP Micro-CT, USA) was used to evaluate apatite deposition after biomimetic mineralization. The Ap-Col-PCL constructs were scanned at a resolution of 10 μm, a voltage of 45 kV, and a current of 80 mA at a grayscale threshold of 150. 3D images of the constructs were reconstructed from the scans by the micro-CT system software package. A collagen disk (with a diameter of 10 mm and height of 3 mm) was prepared and mineralized the same way, then compressed to a film (about 0.5 mm thickness), and used for the thermogravimetric analysis (TGA), X-ray diffraction (XRD), and Fourier transform infrared spectroscopy (FTIR) because it was difficult to perform these experiments directly on 3D Ap-Col-PCL constructs. TGA was carried out to determine the amount of apatite in the biomimetic collagen film using a TGA2050 system (Thermal plus EVO II series TG-DTA). The samples were heated from 30 to 800 °C at a heating rate of 10 °C min−1 in air. XRD of the biomimetic apatite was carried out at a step size of 0.02° and a scan rate of 1° min−1 with CuKα radiation (k = 1.54056 nm). FTIR (Thermo Fisher Scientific, Nicolet 6700) was carried out in absorption mode in the range 4000− 800 cm−1 at a resolution of 4 cm−1 with an average of 32 scans. 2.6. Pore Size and Porosity. The pore sizes of the PCL, A-PCL, Col-PCL, and Ap-Col-PCL constructs were measured under FESEM; B

DOI: 10.1021/acsami.5b08534 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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ACS Applied Materials & Interfaces the porosity of Ap-Col-PCL was determined by analyzing the mass and volumes of the constructs according to eq 1:3

medium, or 5% DMSO. After 3 days, the cytotoxicity of the constructs was assessed using CCK-8. 2.10. In Vivo Study of the Restoration of Critical-Sized Bone Defect of Rabbit Radius. A total of 18 male New Zealand white rabbits with an average weight of 2.5 ± 0.2 kg and an age of 5 months were randomly divided into three groups (n = 6 for each group): Group A, nothing implanted (blank); Group B, implanted with APCL; and Group C, implanted with Ap-Col-PCL. The animals were anesthetized with an ear marginal veins injection of 1.5% sodium pentobarbital (30 mg/kg). Under sterile conditions, the midshaft of the left radius was exposed, and a section of the diaphysis (15 mm) was cut off using a dental drill, rinsed with 0.9% saline, and filled with the prepared constructs, respectively. At each time point, 4, 8, and 12 weeks after surgery, the animals were anesthetized and underwent Xray examination. The sequential fluorochrome markers Alizarin Red S (30 mg/kg, Sigma-Aldrich) and calcein (30 mg/kg, Sigma-Aldrich) were administered via intraperitoneal injection at 4 and 8 weeks after surgery. All animals were sacrificed by an overdose of pentobarbital at 12 weeks to obtain the defect sites. After retrieval, the samples were fixed in 10% formalin for at least 48 h; they were then detected and imaged using micro-CT to determine the newly formed bone in the defects. The scanning parameters were set at 45 kV and 80 mA with an exposure time of 3000 ms and a resolution of 20 μm. 3D images of the specimens were reconstructed from the scans by the micro-CT system software package. (The animal experiments were carried out with ethical approval by the Animal Care and Use Committee of Shanghai Ninth People’s Hospital.) 2.11. Histological Observation. After detection by micro-CT, three samples were decalcified in 10% ethylene diamine tetraacetic acid (EDTA) and subsequently embedded in paraffin. Sections near the central areas of the implants were used for hematoxylin and eosin (H&E) and Masson’s trichrome staining and visualized using an optical microscope. The other samples were dehydrated with ethanol and finally embedded in poly(methyl methacrylate) (PMMA). The embedded specimens were sectioned into 150 μm thick sections using a Leica SP1600 saw microtome (Leica, Hamburg, Germany) along the long axis of the radius at the central region. These sections were subsequently ground and polished to a final thickness of approximately 40 μm for fluorescence labeling observation under a confocal laser scanning microscope. The excitation/emission wavelengths of the chelating fluorochromes at 543/617 nm and 488/517 nm were used for Alizarin Red S (red) and Calcein (green), respectively.37 Finally, the undecalcified sections were stained with Van Gieson’s picrofuchsin staining. 2.12. Histomorphometric Analysis of Bone Tissue Ingrowth and Osteointegration. Image Pro Plus (Media Cybernetics, Silver Springs, MD, USA) was used to calculate the area of new bone. The new bone formation area was quantified from the pixels representing bone tissues. The total area was regarded as the implanted bone site (φ = 4 mm × 15 mm). The new bone formation capability was determined by the percentage of the new bone area over the total implant area (new bone area/total area ×100%). The bone mineral deposition rate was calculated by the distance between two fluorescence (green and red) labeling distances (μm/day). Additionally, the perimeters of new bone and scaffold were measured by Image Pro Plus. The osteointegration capacity of the scaffold was evaluated by calculating the length of the bone−material contact line/scaffold perimeter (CL/Ps) ratio, whereas the CL was calculated according to the following eq 3:18

Porosity (Ap−Col−PCL) ρap =1− ρm ⎛ ωcollagen ωapatite ⎞ ω ⎟/V = 1 − ⎜⎜ PCL + + scaffold ρcollagen ρapatite ⎟⎠ ⎝ ρPCL

(1)

where ρm is the density of the Ap-Col-PCL construct, and ρap is the apparent density of the construct. The densities of PCL, collagen, and apatite are 1.145 (ρPCL), 1.32 (ρcollagen), and 3.16 g cm−3 (ρapatite).3 The weight percent ratios of PCL (ωPCL), collagen (ωcollagen), and apatite (ωapatite) were measured by electronic microbalance before and after the inclusion of collagen and biomimetic apatite deposition. 2.7. Hydrophilicity and Water Absorption Ratio. The surface hydrophilicities of the constructs were measured by water contact angle (WCA) measurements (Automatic Contact Angle Meter model SL200B, Solon, China). At room temperature, a sessile ultrapure water droplet with a volume of 2 μL was dropped onto the constructs (n = 3). The WCAs at each time point (1 and 180 s) were averaged and presented as the means ± standard deviations. The water absorption was measured by weighing the constructs before and after soaking in distilled water for 2 h as previously described.15 The percent increase in water absorption was calculated according to eq 2: P = (W2h − W0)/W0 × 100%

(2)

W2h and W0 are the weights of the wet and dry constructs, respectively. 2.8. Mechanical Property Test. Compression tests were carried out to evaluate the mechanical properties using an Instron 5542 universal tester (Instron Corp., Norwood, MA) with a 500-N load cell. Constructs (n = 5) with a size of diameter of 10 mm and height of 3 mm were compressed at a loading rate of 1 mm/min to a strain level of 80% . The stress−strain (σ−ε) curves were obtained, and the compressive modulus was calculated from the stress−strain curves as the slope of the initial linear portion of the curves.34 2.9. In Vitro Cellular Cytotoxicity, Proliferation, and Viability Evaluation. Rabbit bone marrow stromal cells (r-BMSCs) were obtained from fetal rabbits (age: 28 days) according to a previous method;35 5 × 105 (passage 3) r-BMSCs in 100 μL of DMEM lowglucose containing 10% FBS and 1% penicillin/streptomycin (D10 medium) were seeded into the constructs in 6-well plates. For the quantification of cell adhesion capability, the cell-seeded constructs were transferred to a new plate after cell seeding for 4 h, and the cells that were attached to the plates were digested and counted (n = 6). The cell adhesion capability was quantitatively assessed as a percentage of the cell fraction into the constructs relative to the total amount of cells loaded. For cell proliferation evaluation, 5 × 105 r-BMSCs in 100 μL of D10 medium were seeded into the constructs in a 24-well plate; after 4 h of incubation, 1 mL of D10 medium was added and incubated at 37 °C. At 1, 3, and 7 days after cell seeding, the constructs were transferred to another new plate; and the cell proliferation was assessed using the Cell Counting KIT-8 (CCK-8, Dojindo, Kumamoto, Japan) at 450 nm. The cell number was correlated with the optical density (OD). Another group of cell-seeded constructs were used for the analysis of cell viability and morphology by LIVE/ DEAD Cell Imaging Kit staining. Briefly, the cell-seeded constructs were washed with phosphate-buffered saline (PBS) three times, incubated in 2 μM calcein (staining live cells) and 4 μM PI (staining dead cells) in PBS for 30 min at 37 °C, and washed again with PBS. Samples were imaged using a confocal laser scanning microscope (Leica, Germany) at 1 and 7 days after cell seeding. For cytotoxicity evaluation, the extracted fluid of the constructs was prepared as previously described.36 The constructs (n = 6) were immersed in 200 μL of D10 medium for 24 h in a 48-well plate. r-BMSCs were seeded onto a 96-well plate at a density of 103 cells/well. After 24 h, the medium was replaced with 100 μL of extracted fluid, fresh D10

CL =

Ps + Pb − Pbs 2

(3)

where Ps is the scaffold perimeter, Pb is the bone perimeter, and Pbs is the bone and scaffold perimeter. 2.13. Statistical Analysis. All the presented data are expressed as the mean ± standard deviation. Statistical analysis was carried out using SPSS 15.0 for Windows (IBM/SPSS, Inc., Chicago, IL). A group t test and one-way ANOVA single-factor analyses of variance (ANOVA) were used to compare values among groups. The significance level was set at p < 0.05. C

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3. RESULTS 3.1. FESEM Evaluation. The RP-fabricated PCL constructs display complete, interconnected architectures and a honeycomblike pattern with a vertical pore size of ∼1000 um and a horizontal pore size of 346 × 546 μm2 (Figure 1a1−a3). The

Figure 1. FESEM images of the constructs. (a1−a3) The PCL frameworks have macroporous (∼1000 μm) architectures and complete interconnectivity. (b1−b3) The A-PCL (surface-activatedPCL) displayed porous architectures similar to those of PCL but had a rougher surface appearance under the higher magnification view (b3). (c1−c3) In the collagen-modified PCL scaffold (Col-PCL), microporous(126.2 ± 18.42 μm) collagen networks homogeneously occupied in the voids of the A-PCL frameworks (c1 and c2), with a thin layer (2.99 ± 0.52 μm, as indicated by the blue arrows) of collagen membrane coating the surface of A-PCL (c3).

Figure 2. FESEM and EDS analysis of the Ap-Col-PCL composite constructs after biomimetic apatite functionalization through CaCl2/ K2HPO4 alternative soaking (AS) and SBF treatment. (a1−d1) Left panel FESEM images show that the microporous collagen structures of composite constructs were maintained throughout the process of biomimetic apatite deposition. (a2−d2) Middle panel high-magnification FESEM images demonstrated that the nanotextured apatite was deposited onto the collagen fibers and that the scale of the nanoscaled apatite grew larger and more uniform as the immersion time increased (from 24 h to 1 week). (a3−d3) Right column shows the EDS analysis of the Ca/P ratios. One week of SBF treatment (d1−d3) led to the uniform formation of the composite constructs with the Ca/P ratio closest to that of natural bone tissue (1.67). (e1−e3) EDX mapping of biomimetic apatite structure of the composite constructs with 1 week of treatment (AS-SBF(1w)) further confirmed a homogeneous distribution of calcium and phosphate (e1: calcium; e2: phosphate; and e3: merger of e1 and e2).

surface activation did not change the porous architectures or dimension but rather increased the surface roughness (Figure 1b1−b3), which is in accordance with previous study.15 After A-PCL was functionalized through collagen incorporation, the collagen networks were cross-embedded among the voids and coated on PCL through physical integration and vacuum absorption to form a second level of microporous collagen networks (126.2 ± 18.42 μm) and a layer of collagen film coating (2.99 ± 0.52 μm, Figure 1c1−c3). After AS treatment, nanoscaled (∼50 nm diameter) round amorphous apatite particles were deposited on the surface of collagen fibers with a Ca/P ratio of 1.02 ± 0.06 (Figure 2a1− a3). When the AS-treated scaffolds were further immersed in SBF for different durations, the microporous structures of the collagen remained intact and open (Figure 2a1−d1), but the amorphous apatite particles on the collagen fibers were substituted by a nanosized flower- (Figure 2b2) or platelike (Figure 2c2,d2) apatite. The biomimetic apatite displayed a nanosized (∼50 nm thickness and ∼1 μm length) flowerlike appearance and a Ca/P ratio of 1.51 ± 0.24 after 24 h of SBF immersion (Figure 2b1−b3), then grew bigger (∼100 nm thickness and ∼1 μm length) and displayed a platelike appearance with a Ca/P ratio of 2.27 ± 0.49 as the immersion time increased from 24 to 72 h (Figure 2c2,c3). When the immersion time increased to 1 week, the platelike appearance become much more uniform, and the Ca/P ratio decreased to 1.65 ± 0.24 without a significant size change (Figure 2d2,d3). The Ca/P ratio of the 1 week SBF treatment sample is closer to the ratio of the natural bone of 1.67.38 EDX mapping of calcium

(Figure 2e1) and phosphate (Figure 2e2) distributions and the emergence picture (Figure 2e3) in the AS-SBF 1w group confirm that the calcium phosphate mineral phase was distributed uniformly throughout the constructs. 3.2. Micro-CT, TGA, FTIR, and XRD Analyses. Micro-CT was used to visualize and analyze the biomimetic apatite deposition and 3D distributions among the constructs, showing a great quantity of apatite anchoring within the Col-PCL constructs after AS treatment. However, the quantity of apatite deposition decreased within the first 24 h and gradually increased with the SBF immersion time, and maximum apatite precipitation was achieved at 1 week (Figure 3b). The apatite mass change was further confirmed by TGA analysis with apatite contents of 52.1, 46.6, 55.8, and 59.7% after 0 h to 1 week of SBF biomimetic deposition (Figure 3c). The decreased apatite in the first 24 h in SBF occurred because the amorphous apatite that formed during AS treatment was unstable, and when the AS treated scaffolds were immersed into the SBF, most of the amorphous material dissolved. From 24 to 72 h, the Ca2+ and PO43− in the SBF solution reprecipitated onto the D

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Figure 3. Characterization of the Ap-Col-PCL composite constructs: (a) optical images, (b) micro-CT 3D images, (c)TGA analysis, (d) FTIR, and (e) XRD analysis of the Ap-Col-PCL constructs. The Ap-Col-PCL composite constructs of the 1 week SBF (AS-SBF (1w)) treatment achieved the best biomimetic apatite functionalization, as indicated by the higher apatite deposition by micro-CT and TGA assay, and more typical apatite characteristic peaks by XRD analysis. Therefore, 1 week SBF treatment was utilized to fabricate Ap-Col-PCL composite constructs for the following in vitro and in vivo assays.

Figure 4. Hydrophilic property and water uptake ability analysis. (a and b) Water contact angles (WAC) were measured at 1 and 180 s for PCL, APCL (surface-activated-PCL), Col-PCL, and Ap-Col-PCL (AS-SBF (1w)) and revealed the highest hydrophilicity of Ap-Col-PCL. (c) Water uptake analysis showed the best water absorption ability of Col-PCL and Ap-Col-PCL constructs compared to those of PCL and A-PCL. (***, p < 0.001; NS, not significant).

amorphous apatite and formed more stable apatite. The ASSBF (1w) scaffold displayed higher apatite content (59%)

closer to that of bone (67%). FTIR was used to analyze the chemical component of the constructs (Figure 3d). For the E

DOI: 10.1021/acsami.5b08534 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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Figure 5. Mechanical property evaluation. (a) Stress−strain curves of the scaffolds. (b) Compressive modulus analysis demonstrated the best mechanical properties of Col-PCL and Ap-Col-PCL composite constructs compared to collagen sponges (Col), PCL constructs, and A-PCL constructs. (c) Stiffness of Col-PCL and Ap-Col-PCL fell into the stiffness range of native cancellous bone tissues. (**, p < 0.01; ***, p < 0.001; NS, not significant).

typical collagen bands, the peak at approximately 1600−1700 cm−1 for the CO stretch of amide I, an N−H deformation of approximately 1500−1550 cm−1 for amide II, and an N−H deformation of approximately 1200−1300 cm−1 cm for amide III were observed in collagen. In the mineralized collagen, the main absorption bands (1028 cm−1) refer to those vibrations of the PO43− groups of apatite.3,39 Figure 3e shows the XRD patterns of the native and the mineralized collagen films; the broad diffuse peak at approximately 20° is attributed to collagen.40 In biomimetic mineralized collagen, the peaks at 26.48, 28.1, 33.52, 39.81, 46.73, 49.46, and 53.52° were detected in all scaffolds except for the AS scaffold; these peaks correspond to the (002), (102), (211), (310), (222), (213), and (004) diffraction peaks of apatite, respectively.1,26 All scaffolds had broad diffraction peaks, indicating that the biomimetic apatite was poorly crystallized. Given this character, AS-SBF 1w referred to as Ap-Col-PCL was used for the following in vitro and in vivo experiments. 3.3. Pore Size and Porosity. SEM evaluation indicated that the A-PCL constructs display a pore size of 1000 × 546 × 346 μm3. After functionalization through collagen incorporation, the collagen networks possessed a pore size of 126.2 ± 18.42 μm, which slightly decreased to 116.7 ± 15.82 μm after biomimetic deposition by AS treatment and subsequent immersion in SBF for 1 week. The porosity analysis demonstrated that the A-PCL, Col-PCL, and Ap-Col-PCL constructs had porosities of 59.05 ± 1.67, 57.5 ± 0.1, and 55.0 ± 0.1%, respectively. 3.4. Hydrophilic Property and Water Uptake Ability Analysis. The hydrophilicity of the constructs was evaluated by water contact angles (WCA) analysis. Compared to the pristine PCL (with a WCA of 83° at 1 s and 78° at 180 s), APCL displayed significantly lower WCA values (with a WCA of 33° at 1 s and 30° at 180 s). The WCA differences between PCL and A-PCL indicate that alkaline surface treatment

increases the hydrophilicity of PCL because the weak chemical bonds of PCL are replaced by highly reactive carboxyl (−COOH) and hydroxyl (−OH) groups.31 The WCA of Col-PCL was 85° at 1 s and decreased to 0°, and the WCA of the Ap-Col-PCL was 0° initially (Figure 4a,b). The results indicate that functionalization with collagen and biomimetic deposition increases hydrophilicity dramatically because of the hydrophilic property of collagen and nano apatite. The water uptake can be defined as the ability of the scaffolds to maintain water and water permeation, which can influence the transfer of nutrients and cell proliferation.15 Compared to PCL, A-PCL showed higher (1.44-fold) water uptake ability, and collagen functionalization and biomimetic deposition could increase the water uptake ratio twofold compared to that of PCL (Figure 4c). 3.5. Mechanical Properties of the Constructs. Compression testing was carried out to evaluate whether there was any detrimental effect of the biomimetic 3D functionalization on the mechanical properties of the constructs. The results suggest that the pristine PCL had a compressive modulus of 41.40 ± 3.93 MPa, which slightly (p > 0.05) decreased to 33.62 ± 4.74 MPa after surface activation (A-PCL). However, the EDC cross-linked collagen displayed a compressed modulus of 0.53 ± 0.12 MPa, and collagen incorporation of A-PCL (ColPCL) and subsequent biomimetic apatite deposition by SBF treatment (Ap-Col-PCL) significantly (p < 0.01) increased the compressive modulus of the A-PCL constructs. The reinforced Col-PCL and Ap-Col-PCL constructs displayed compressive moduli of 56.48 ± 4.21 and 68.75 ± 3.39 MPa, respectively, falling into the stiffness range of native spongy bone (Figure 5a−c).13,21 3.6. In Vitro Cellular Evaluation. Ap-Col-PCL showed a significantly (p < 0.001) higher cellular adhesion capability (2.0-fold) compared to that of A-PCL (Figure 6a). A CCK-8 assay was used to measure the cytotoxicity of the extracted fluid F

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Figure 6. In vitro cellular evaluation of the Ap-Col-PCL composite constructs. (a) Cell adhesion capability assay. (b) Cytotoxicity evaluation of the extracting liquor of the scaffolds. (c) Cell proliferation on the scaffolds. (d) Confocal microscopy imaging of r-BMSCs that were cultured on the constructs with LIVE/DEAD staining at days 1 and 7 (green color: live cells, red color: dead cells). Composite constructs possess great biocompatibility and more favorable biological properties, achieving better cellular adhesion and proliferation capacities. (***, p < 0.001; NS, not significant; scale bar: 200 μm).

of Ap-Col-PCL composite constructs achieved partial bridging of the defect at 4 weeks and full defect union at 12 weeks (Figure 7a). Additionally, micro-CT was used to evaluate the new bone formation after 12 weeks postimplantation. The blank group and the A-PCL group barely exhibited bone formation at the edge of the defects. In the Ap-Col-PCL group, successful defect bridging occurring with partially circumferential cortical regeneration, and new bone tissue was observed to surround and integrate with the scaffolds in the defect sites (Figure 7b). Moreover, quantification of the newly formed bone demonstrated that much more bone tissue was regenerated in the Ap-Col-PCL group (147.6 ± 27.9 mm3) compared to that in the A-PCL (47.7 ± 15.9 mm3) and blank groups (28.4 ± 12.2 mm3) (Figure 7c). 3.8. Histomorphometric Analysis of Bone Tissue Ingrowth and Osteointegration. Photomicrographs of

of the constructs, showing no apparent cytotoxicity of any constructs (Figure 6b). The cell proliferation was significantly greater on the Col-PCL and Ap-Col-PCL scaffolds (Figure 6c), possibly because of much more cell adhesion. The morphology, viability, and cell distribution of r-BMSCs cultured on the constructs were observed by confocal laser microscopy at 1 and 7 days after cell seeding. Many more live cells were observed distributed in the Col-PCL and Ap-Col-PCL composite constructs compared to those in A-PCL at 1 day; in addition, Ap-Col-PCL exhibited a much higher growth rate after the culture time was prolonged to 7 days (Figure 6d), although all the scaffolds in this study showed good cytocompatibility. 3.7. X-ray and Micro-CT Evaluation of Critical-Sized Bone Defect Treatment. The X-ray images results revealed that the bone defects remained nonunion in the blank group and the A-PCL group until 12 weeks, whereas the implantation G

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Figure 7. In vivo evaluation of the Ap-Col-PCL composite constructs for critical-sized long bone defect treatment in a rabbit radial defect model. (a) X-ray examination at different time points (4, 8, and 12 weeks) after implantation. (b) Micro-CT images at 12 weeks (row 1: 3D reconstructed images; row 2: longitudinal section views parallel to the long axis of bone; row 3: cross-sectional views at the center of the defect). (c) Quantitative micro-CT analysis of the new bone formation. Ap-Col-PCL constructs have successfully repaired the critical-sized radial bone defect with complete defect bridging and achieved a significantly higher volume of new bone regeneration compared to that of A-PCL. (**, p < 0.01; NS, not significant).

Figure 8. Histological evaluation. (a) H&E staining: In the A-PCL group, the majority of the defect area is occupied by fibrous connective tissue encapsulating the struts of A-PCL constructs (a1 and a2, black arrow), whereas in the Ap-Col-PCL group, the defect is almost filled with newly formed bone tissues that bridged the segmental defects and integrated well with the struts of Ap-Col-PCL constructs (a3 and a4, black arrows). (b) Masson’s trichrome staining further confirmed the previous findings. In the A-PCL group, the fibrous tissue filled most of the defect region (b1), and the A-PCL constructs did not integrate well with native bone tissue at the end of the defect, where the bone−material interface was separated by fibrous tissue (b2, black arrows). In contrast, the implantation of Ap-Col-PCL composite constructs achieved complete defect bridging with more bone tissue formation and better osteointegration between material and bone tissue at the defect end (DE in b3 and b4, black arrows); furthermore, the bone-marrow-like structure (b3, black stars) was found as well, indicating the generation of mature bone structure (S: scaffolds, NB: newly formed bone, FB: fibrosis; black bar = 5 mm; red bar = 1 mm).

as seen with Masson’s trichrome staining (Figure 8b3,b4). These newly formed bone tissues directly integrate with the struts of the Ap-Col-PCL constructs (Figure 8b4, black arrows) and bridge the scaffolds with the native bone; bone marrow formation was observed in the newly regenerated bone tissues in Ap-Col-PCL group (Figure 8b3, black stars). The undecalcified sections were further stained with Van Gieson’s picrofuchsin for quantitative evaluation of the bone tissue ingrowth and the osteointegration degree between bone tissue and implants. Newly formed bone tissue was stained red

longitudinal sections (parallel to the long axis of the radius) of the bone defect sites are presented (Figure 8). H&E staining of the A-PCL group (Figure 8a1,a2) demonstrated that the majority of defect sites were filled with loose connective fibrous tissues encapsulating the A-PCL, which is further confirmed by Masson’s trichrome staining (Figure 8b1), showing the dense fibrosis surrounding the struts of PCL constructs (Figure 8b2, black arrows). In contrast, the H&E staining of Ap-Col-PCL groups showed defects that were mainly filled with bone tissues (Figure 8a3,a4), comprising both immature and mature bone, H

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Figure 9. (a) Histomorphometric analysis of bone tissue ingrowth and osteointegration. Van Gieson’s picrofuchsin staining further confirmed that Ap-Col-PCL composite constructs have much more bone regeneration and better bone-scaffold osteointegration. (b) Quantitative analysis showed that new bone covered 49.27 ± 12.65% of the total area in the Ap-Col-PCL group, compared to only 5.48 ± 4.58% in the A-PCL group 12 week after implantation. The osteointegration was evaluated by calculating the length of the bone−material contact line/scaffold perimeter (CL/Ps) ratio. Pb, Ps, and Pbs were derived from the perimeter of white areas in Van Gieson’s picrofuchsin staining imaging (c), and CL was calculated according to the formula and schematic illustration in d. (e) The Ap-Col-PCL constructs showed a 7.2-fold higher osteointegration capability than A-PCL (55.95 ± 7.80% vs 6.99 ± 4.44%). (f) Sequential fluorescent labeling observations indicated new bone mineral deposition at 4 and 8 weeks, labeled by Alizarin Red S (red) and Calcein (green), respectively. (g) The distance between the red line and green line was measured to evaluate the mineral deposition rate. The Ap-Col-PCL group showed a faster new bone deposition rate, 2.9-fold higher than that of the A-PCL group (7.08 ± 0.66 μm/ day vs 2.44 ± 0.52 μm/day). (Ps: scaffold perimeter, Pb: bone perimeter, Pbs: bone and scaffold perimeter, ***, p < 0.001; NS, not significant).

rate, 2.9-fold higher than that of A-PCL group (7.08 ± 0.66 μm/day vs 2.44 ± 0.52 μm/day, Figure 9f,g).

with a woven, trabecular appearance (Figure 9a). In the A-PCL group, only a trivial amount of new bone was formed on the margin of the scaffold. In contrast, the Ap-Col-PCL implanted group showed a much larger new bone area (49.27 ± 12.65%) compared to that of the A-PCL group (5.48 ± 4.58%, Figure 9b). The osteointegration capacity of the constructs was evaluated by calculating the length of the bone−material contact line/scaffold perimeter (CL/Ps) ratio (Figure 9c,d), and the Ap-Col-PCL constructs showed a 7.2-fold higher osteointegration degree than did the A-PCL constructs (55.95 ± 7.80% vs 6.99 ± 4.44%), (Figure 9e). Red (Alizarin) and green (Calcein) fluorescent labeling were detected throughout the entire defect of Ap-Col-PCL groups, whereas merely fluorescent labeling was observed in the margin of the defect in the A-PCL group, suggesting that the Ap-Col-PCL group experienced a much faster new bone tissue deposition

4. DISCUSSION An excellent bone graft substitute for clinical use should possess not only mechanical function with highly porous architecture but also favorable biological performance to promote rapid bone tissue ingrowth. Generally, increasing the porosity, interconnected pore networks, and surface area of the scaffold should benefit cell migration and differentiation, bone ingrowth, and vascularization. However, excessive porosity compromises the scaffold’s mechanical performance. Therefore, a balance between a denser scaffold providing better mechanical function and a more porous structure providing better biological performance should be carefully considered during scaffold fabrication.13 In the past decades, the emergence of computer-aided RP I

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and structure and more readily promotes osteointegration and subsequent bone tissue formation.45 The in vitro cellular evaluation results indicate that hierarchical constructs demonstrated outstanding biological characteristics, as evidenced by the significantly higher cell adhesion capability (twofold compared to that of PCL). The improved cell attachment is greatly due to the microporous collagen networks, which increase in the scaffold surface area for cell adhesion and act like a net bag that prevents cell leakage from the PCL macropores during the process of cell culture.27 Additionally, the Ap-Col-PCL showed a higher cell proliferation and more uniform cell distribution, indicating that nanoscaled apatite surfaces might be more effective for cell proliferation and distribution because they simulate the hierarchical structure of typical bone (a nanoapatite-reinforced collagen composite).45 In addition, it should be noted that cells that are attached to and grew on the stiff Ap crystals experience different mechanical constraints compared to those experienced in other types of scaffolds because MSCs can sense or “feel” the mechanical properties of the matrix and transduce that information into morphological changes and lineage specification, which could favor MSCs osteogenic differentiation.46,47 Moreover, in the current study, our in vivo evaluation indicated that the hierarchically porous Ap-Col-PCL scaffolds induced much more (5.2-fold) new bone formation, better osteointegration (7.2-fold), and a faster new bone deposition rate (2.9-fold) than did pure PCL scaffolds, indicating that the hierarchically porous structure of the scaffold plays an important role in enhancing the in vivo osteogenesis and osteointegration. The improvement of in vivo osteogenesis is related to the enhanced cell attachment and proliferation, as indicated by in vitro cell experiments. Additionally, the biodegradable collagen networks that homogeneously distribute over and fill the voids among the PCL struts can also serve as a 3D microporous barrier or membrane to prevent fibrous tissue infiltration and preserve the space for new bone formation, making full use of the principle of guided bone regeneration.48,49 Moreover, a thin collagen film (approximately 3 μm) was found to coat directly upon the PCL struts (Figure 1c3), which can work together with the collagen networks to improve the osteoconductive properties of PCL constructs and lead to significantly enhanced osteointegration at the bone−PCL interfaces, as demonstrated by the histomorphometric analysis (Figure 9).50,51 Third, the ceramic biomaterials based on nanosized Ap exhibits enhanced resorbability and much higher bioactivity than microsized ceramics; in addition, previous study indicated that calcium ions released from nanosized HAp are similar to those from biological apatite and significantly faster than that from coarser crystals.45 In this study, the nanosized Ap, which has a grain size of ∼100 nm to1 μm in at least one direction, had high surface activity and an ultrafine structure, similar to that of the mineral found in hard tissues.45 All of the above served as synergistic factors promoting bone regeneration and bone implant osseointegration. Because bone is a load-bearing tissue, the graft substitutes for bone defect treatment must be able to provide sufficient temporary mechanical support to withstand physiological loading and protect the defect space for osteogenic cells and bone tissue ingrowth.21,41,42 However, if the Young’s modulus of the graft substitute is too high, then it is also detrimental to bone tissue regeneration, leading to the absorption of native bone tissue, which is known as stress shielding.18,52 Therefore, the mechanical properties of the bone graft substitute should be

technology has made scaffolds with designed characteristics possible.2,13,41 The RP-fabricated PCL scaffold with a completely interconnected porous structure and favorable mechanical properties has been utilized as a bone substitute graft to promote bone regeneration.42 However, the poor biological performance of PCL scaffolds results in limited cell adhesion and elicits in vivo fibrous encapsulation, compromised osseointegration, insufficient bone tissue regeneration, and eventually the failure of bone defect treatment.9 A previous study indicated that alkaline treatment could improve the PCL scaffold activity by increasing surface roughness, increasing the hydrophilic property and decrease fibrous encapsulation.9,43 In our study, the alkaline-activated PCL scaffold (A-PCL) displayed a rough surface appearance and significantly better hydrophilicity property and water uptake ability but was still insufficient to improve cell entrapment and proliferation in vitro, resulting in fibrosis encapsulation and the failure of bone tissue regeneration in vivo. Therefore, we further propose a unique hierarchical nano− micro Ap-Col-functionalizing 3D macroporous PCL scaffolds. Specifically, the macroporous 3D PCL constructs were meticulously fabricated using the RP technique with a pore size of 1000 × 546 × 346 μm3 and 59% porosity. This design of a macropore (approximately 1000 μm) in the vertical provides much more space for subsequent collagen incorporation; meanwhile, horizontal pores with a size of 346−546 μm are optimal for bony ingrowth.9 The collagen networks display a microporous architecture (126.2 ± 18.42 μm) after evacuated incorporation and subsequent lyophilization at −80 °C in accordance with previous study.44 The collagen networks crossembed among the voids and coat the PCL through physical integration and vacuum absorption to mimic the ECM of the bone. This porous collagen not only increased the surface of the PCL scaffold for much more efficient cell entrapment and adhesion but also provided the carboxyl groups as the template for subsequent biomimetic deposition. When the Col-PCL composite was immersed in the AS solution, the carboxyl groups of the collagen served as nucleation sites for amorphous apatite precipitation by covalent bonding.33 The formation of amorphous apatite may be attributed to the following points. First, Ca2+ ions bind to the carboxyl sites of Col because of ionic attraction and the subsequent accumulation of PO43− ions at calcium complexes and grow to the critical apatite nuclei or precursors.26 When the AS-treated specimen is subsequently subjected to SBF treatment, the apatite nuclei or precursors spontaneously grow into the apatite layer in SBF by consuming calcium ions and phosphate ions from the solution.32,33 The AS treatment could provide apatite nuclei or precursors that greatly promote Ap precipitation effectively compared to that of the classic method requiring several weeks.32 Interestingly, the quantity of apatite decreased from 52.1 to 46.6% and gradually increased to 59.7% after 1 week of SBF treatment (Figure 3b,c). This change could be explained as dissolving and reprecipitation. The nanosized (∼50 nm) amorphous apatite in the AS scaffold showed a broad peak of apatite, indicating its low crystallinity. When it is transferred to SBF, most of it dissolved, and the residual amorphous apatite served as a precursor to promote the relatively high crystalline and stable Ap deposition. These Ap accumulate and deposit gradually to form the third level of hierarchical nanosized platelike Ap coating on the collagen matrix, which mimics the bone mineral in composition J

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carefully tailored, matching the natural bone tissue as closely as possible. Some studies have reported the use of composite scaffolds with hierarchical pore structure for bone tissue engineering.27,53 Louet et al. prepared hierarchical PLLA-βTCP composite scaffolds with first-level pores (50−300 μm) and second-level pores (0.5−10 μm) by combining thermalinduced phase separation and salt leaching techniques.53 However, the porous structure of the scaffolds prepared by this method were not uniform, and the mechanical strength was relatively low (less than 1 MPa). In the present study, the RPfabricated macroporous pristine 3D PCL and surface-activated A-PCL constructs shared a similar (p > 0.05) compressive modulus (41.40 ± 3.93 and 33.62 ± 4.74 MPa); 3D functionalization with collagen incorporation and biomimetic deposition further improved the mechanical property of the ApCol-PCL composite construct significantly to a compressive modulus of 68.75 ± 3.39 MPa, which is close to that of the cancellous bone tissues (Figure 5c) and falls within the suggested compressive modulus range for bone tissue regeneration(10−1500 MPa) by Hollister et al.13 Compared to the traditional collagen or collagen−ceramics composite constructs, which are frequently used in clinics,6,21 our Ap-ColPCL composite construct dramatically improved its mechanical strength with a compressive modulus that was at least 2 orders of magnitude greater than that of freeze-dried HA-Col scaffolds (5−300 kPa) and 3 orders of magnitude greater than that of absorbable collagen sponges. Furthermore, this mechanical property of the Ap-Col-PCL composite construct allows great feasibility for surgical handling, fixation, and load-bearing implantation during the clinical operation procedure. This mechanical enforcement can be explained by the classic “brick-and-mortar” reinforcement theory as stated by Zhou et al.26 According this theory, the good mechanical behavior of Ap-Col-PCL constructs could be due to the special three-level hierarchical structure, comprising an elastic PCL frameworks, elastic collagen matrix, and stiff Ap crystals. In the RP-fabricated 3D PCL porous framework, there are substantial free spaces; after collagen incorporation, the elastic collagen fibers filling the space could respond rapidly to external forces and absorb the forces simultaneously, resulting in a relatively high strain and compressive modulus. Furthermore, the stiff Ap crystals were embedded in the collagen forming a compact 3D network and reinforcing the compressive modulus.

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AUTHOR INFORMATION

Corresponding Authors

*Tel.: +86-21-34291002. Fax: +86-21-34292305. E-mail: [email protected]. *Tel.:+86-21-23271699 x5161. E-mail:[email protected]. Author Contributions

J.W., D.W., and Z.Z.Z. contributed equally. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This work was supported by National Natural Science Foundation of China (nos. 81371964 and 81572137), the Program for Professor of Special Appointment (Eastern Scholar) at Shanghai Institutions of Higher Learning (no. 1220000187), National Young Thousand-Talent Scheme, Shanghai Rising-Star Program (13QA1402400), and Shanghai Jiao Tong University Medicine-Engineering Integrated Research Grant (YG2012MS44) to Z.Y.Z., and the Shanghai Municipal Science and Technology Commission program (no. 124119b0102) to J.S.

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REFERENCES

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5. CONCLUSIONS In this study, we developed an Ap-Col-PCL composite construct by combining RP fabrication technology and a 3D functionalization strategy of collagen incorporation and biomimetic deposition. The composite construct shows favorable characteristics with the three nano−micro−macro levels of hierarchical architectures, mechanical properties similar to those of cancellous bone, biodegradability, and outstanding bioactivity to promote rapid bone regeneration in the segmental long bone defect model in rabbit radius. Moreover, it can also make full use of the advanced manufacturing features of RP technology, particularly, the feasibility to fabricate custom-made shape and structure according to medical images, favoring its clinical application. Therefore, the Ap-Col-PCL constructs demonstrated great potential as an effective bone graft substitute for large bone defect treatment. K

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DOI: 10.1021/acsami.5b08534 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX

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ACS Applied Materials & Interfaces (52) Liu, H.; Niinomi, M.; Nakai, M.; Cho, K. Beta-Type Titanium Alloys for Spinal Fixation Surgery with High Young’s Modulus Variability and Good Mechanical Properties. Acta Biomater. 2015, 24, 361−369. (53) Lou, T.; Wang, X.; Song, G.; Gu, Z.; Yang, Z. Fabrication of Plla/Beta-Tcp Nanocomposite Scaffolds with Hierarchical Porosity for Bone Tissue Engineering. Int. J. Biol. Macromol. 2014, 69, 464−470.

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DOI: 10.1021/acsami.5b08534 ACS Appl. Mater. Interfaces XXXX, XXX, XXX−XXX