Anomalous in Vitro and in Vivo Degradation of Magnesium

Article Cite This: ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX

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Anomalous in Vitro and in Vivo Degradation of Magnesium Phosphate Bioceramics: Role of Zinc Addition Kaushik Sarkar,†,§ Vinod Kumar,‡,§ K. Bavya Devi,† Debaki Ghosh,‡ Samit Kumar Nandi,*,‡ and Mangal Roy*,† †

Department of Metallurgical and Materials Engineering, Indian Institute of Technology-Kharagpur, Kharagpur 721302, India Department of Veterinary Surgery & Radiology, West Bengal University of Animal & Fishery Sciences, Kolkata 700037, India

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ABSTRACT: In vitro and in vivo degradation behavior and biocompatibility of magnesium phosphate (MgP) bioceramics and the potential role of zinc (Zn) on degradation were compared. Samples were prepared by conventional solid-state sintering at 1200 °C for 2h. Zn-doped MgP (0.5 wt %) showed 50% less degradation than that of pure MgP after immersion into simulated body fluid (SBF) for 8 weeks. Osteoblast-like cell (MG-63) proliferation was evident in MgP ceramics, which was significantly enhanced upon Zn addition. Both Alamar Blue assay and Live/Dead imaging showed the highest cell attachment and proliferation for 0.5 wt % Zn-doped MgP. In vivo biocompatibility of these MgP ceramics were studied after implantation in the rabbit femur. The micro computed tomography (μ-CT) analysis showed that in vivo degradability increased with the increase in the Zn content which is in contradiction to in vitro degradability. Histological evaluation showed large influx of osteoclast cells to the implantation site for Zn-doped MgP samples compared to that of undoped MgP, which is the primary reason of increased degradability of these samples. After 90 days of implantation, large sections of 0.5 wt % Zn-doped MgP samples were replaced by newly formed bones. Fluorochrome labeling showed 78 ± 3% new bone formation for 0.5 wt % Zn-doped MgP ceramics compared to 56 ± 3% for pure MgP samples. Our findings suggest that the addition of Zn in MgP ceramics alters their sintering and degradation kinetics that leads to decreased in vitro degradation, however, when Zn-doped MgP ceramics were implanted in rabbits, higher degradability was observed due to lower Mg2+ ion concentration in the degradation media. KEYWORDS: magnesium phosphate, zinc doping, degradation, μ-CT

1. INTRODUCTION Degradable biocermics are increasingly used for treating bone defects and craniofacial and maxilofacial reconstructions. The commonly used materials include calcium phosphate family [hydroxyapatite (HA) and tricalcium phosphate (TCP)] and magnesium-based compounds, such as magnesium silicate (MgS) and magnesium phosphate (MgP). The constant evolution in this active area is through control over degradation kinetics that results in tuned ion release and in vitro cytocompatibility and in vivo biocompatibility.1,2 For a specific composition, the degradability can be controlled through the change in grain size, density of the product, and/or metal ion doping. Metal ion doping in general increases the sinterability of the material and also decreases the mean grain size. The presence of metal ions also gives control over in vitro and in vivo biocomaptibility of the bioceramic. Magnesium phosphate (MgP) has drawn much attention for its higher degradability and both in vitro and in vivo biocompatibility. Mg is the key element of bones which influence bone mineralization and the metabolism process.3 Moreover, being present in the same group in periodic table, Mg possess similar chemical properties and can substitute calcium in © XXXX American Chemical Society

bones. The presence of Mg also increases the bone density and ductility and reduces the chance of bone fracture. Mg2+ ions also play an important role in bone tissue maturation via osteoblast proliferation and differentiation.4 The specific advantages of MgP ceramics are (i) biocompatiblity, similar to that of CaP ceramics, (ii) higher degradability, and (iii) suppression of the formation of insoluble apatite crystals in vivo and thereby leading to complete degradation of the bioceramic.5,6 The in vitro and in vivo biocompatibility of MgP is well studied. Tamimi et al.7 showed that mouse bone marrow cells can survive and actively differentiate on the surface of MgP ceramics. They also found that the expression of osteoblast differentiation markers, collagen (COL1A1), osteopontin, alkaline phosphatase, and osteocalcin was increased over the culture time. Ewald et al.8 also reported that MG-63, human osteoblast cells have the highest survival rate and can differentiate on MgP ceramics. MgP is also degradable in vivo and stimulate the new bone formation.5,6,9 It was reported that bone regeneration was Received: March 26, 2019 Accepted: August 13, 2019

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DOI: 10.1021/acsbiomaterials.9b00422 ACS Biomater. Sci. Eng. XXXX, XXX, XXX−XXX

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ACS Biomaterials Science & Engineering

2.5. Mechanical Testing. A universal testing machine (KS50Tinius Olsen, UK) was used to measure the compressive strength, before and after immersion, for all the samples with a cross-head speed of 0.2 mm/min. Five samples for each composition was used to determine the compressive strength, and results are reported as mean ± standard deviation (SD). 2.6. Cytocompatibility. 2.6.1. Live/Dead Assay and Cell Proliferation Study. In vitro cytocompatibility of undoped and Zndoped MgP samples were studied by culturing MG-63 [National Centre for CELL Science (NCSS), Pune, India] cells on them. Details of the cell culture is given in our previous works.12 Briefly, media comprising of minimum essential medium (MEM Himedia, INDIA) and supplemented with 10% FBS, 50 mg/mL Gentamicin and antibiotics (100 U/mL penicillin, 100 mg/mL streptomycin, Himedia, INDIA) was used for cell culture experiments. A cell seed density of 1 × 105 cells/mL was used in all throughout the experiments. Samples were exposed to steam at 121 °C for 20 min to make it sterile, prior to the cell seeding. Alamar Blue (Himedia, India) assay was used for the determination of cell proliferation at 1 and 3 days after seeding of the cells and culture plate was considered as a control. Alamar Blue powder was dissolved in sterile distilled water for the preparation of Alamar Blue stock solution (5 mg/mL). The working Alamar Blue solution was prepared after diluting to 60 μg/mL solution using the cell culture media. Then, this 500 μL of working Alamar Blue solution was added to each well of a 48well plate and incubated for 2 h in dark. In a 96-well plate, this supernatant solution was transferred in a triplicate manner, and the fluorescence measurement was carried out at 530/25 nm excitation wavelength and 590/35 nm emission wavelength by using a multimode plate reader (BioTek, USA). Cell viability was checked after 1 and 3 days of culture by using a Live/Dead staining kit (Live/Dead Viability/Cytotoxicity Kit, catalog number: L3224, Invitrogen, USA). The images of cell were recorded by using a fluorescence microscope (Leica DM IL LED Fluo, Germany). The number of live cells was quantified from the fluorescence images using ImageJ software. 2.7. In Vivo Study. The animal study was approved by Institutional Animal Ethical Committee (IAEC), West Bengal University of Animal and Fishery Sciences (WBUAFS), West Bengal, India (approval no. IAEC/67(V) dated 28.07.2017). Sixteen New Zealand White rabbits (1.4−1.5 kg) were divided into four groups: group-1, the control (only defect); group-2, pure MgP; group-3, MgP−0.25Zn; and group-4, MgP−0.5Zn. All of the surgical procedures were performed under general anaesthesia, which is administrated by intramuscular injection of xylazine hydrochloride @6 mg/kg body weight (Xylaxin, Indian Immunologicals, India) and ketamine hydrochloride (Ketalar, ParkeDavis, Hyderabad, India) @33 mg/kg body weight.16 Both the femurs were shaved and cleaned with povidone-iodine and 75% ethanol solution. About 5 cm linear skin incision was made in the lateral aspect of femur, and the muscles and periosteum of the exposed part was incised. Then, a circular bone defect in the femur was created by a dental micro motor drill with continuous cold saline irrigation. The operated area was flushed with gentamicin solution, and the ceramics samples were implanted and the surgical area was closed immediately. Antibiotics cefotaxime sodium (Inj. Taxim-125 mg; Alkem, India) injection was administrated intramuscularly twice a day at 12 h interval for 5 days. Anti-inflammatory medicine meloxicam (Inj. Melonex 4mg/ mL, Intas Pharmaceuticals, India) was given 0.2 mL/animal I/M postoperatively followed by 0.1 mL I/M once daily for 2 days.16 2.7.1. Micro-CT and SEM Analysis. The retrieved femur bone samples after 30 and 90 days were studied by μ-CT (Phoenix V|tome|xs, GE, Germany). The machine was operated at a voltage of 100 kV, current of 90 μA, and voxel size of 10 μm. For visualization of the boneimplant interface, SEM (ZEISS, EVO 60) was carried out. Before testing, the formalin-preserved samples were completely dried. 2.7.2. OTC Labeling and Histology. Fluorochrome (OTC dehydrate; Pfizer India, India) was injected 25 days before sacrificing the rabbits, that is, on the days 5, 6, and 13, 14 for 30 day time period and for 90 day animals, on day 65, 66 and 73, 74 (2-6-2 manner) for double toning of newly formed bones.14 The undecalcified bone

stimulated after implantation of MgP scaffolds into 4 and 6 mm defect site of rabbit calvarial and about 80% new bone was formed compared to that of 37.7% for without scaffold after 8 weeks.10 Although in vitro degradation of MgP is well studied, details on in vivo degradability are not well known. Moreover, in vitro chemical degradation may not directly translate to overall in vivo degradability of MgP ceramics. Therefore, a detailed in vivo degradation and correlation with in vitro is of paramount importance. In the present work, Zn-doped MgP was taken as a model material to understand the in vitro and in vivo biocompatibility and degradability. Selection of Zn is based on the fact that Zn is an essential trace element which can be found in all human tissues and plays major biological roles. It is the cofactors of more than 300 enzymes and helps to stabilize the structure of proteins.11 Undoped and Zn-doped MgP ceramics were synthesized by a solid-state sintering process. In vitro degradation behavior and cytocompatibility, utilizing MG-63 cell line, was evaluated to understand the specific role of Zn doping. Porosity distribution within the sample was analyzed by μ-CT. The bone-forming ability of undoped and Zn-doped MgP was studied after implantation in the rabbit femur. Dopantinduced in vivo degradation and osteogenesis was analyzed by μCT and histology. Bone regeneration was quantified by oxytetracycline (OTC) labeling.

2. EXPERIMENTAL DETAILS 2.1. Materials Preparation. Magnesium phosphate (MgP) powders were synthesized in house and the detailed preparation technique is mentioned elsewhere.12 Briefly, magnesium hydroxide (Mg(OH) 2 ) and magnesium hydrogen phosphate trihydrate (MgHPO4·3H2O) were used in a molar ratio of 1:2 (12 g MgHPO4· 3H2O, 2 g Mg(OH)2) to prepare MgP where Zn doping was introduced using ZnO at 0.25 and 0.5 wt %. All of the precursor powder was mixed in a planetary ball mill for 2 h and thereafter calcined at 1150 °C for 5 h with a heating rate of 5 °C/min. Uniaxially pressed disc samples were sintered at 1200 °C for 2 h with a heating rate of 5 °C/min. In the similar sintering condition, Φ 3.5 mm × 4.5 mm cylindrical samples were prepared for in vivo experimentation. Compression test samples were prepared according to ASTM C778-33.13 2.2. Material Characterization. The phase composition of the samples was determined by X-ray diffraction (XRD, Bruker-D8 ADVANCE) equipped with Cu Kα radiation (λ = 0.154 nm at 40 kV and 40 mA). Data were collected from 10° to 50° for 2θ at a scan rate of 0.02°/s. Surface morphologies were studied using scanning electron microscopy (SEM) (MERLIN, CARL ZESIS, Germany) equipped with an energy-dispersive spectroscope (Oxford INCA Penta EETX3). The fracture surface of samples was analyzed using SEM (ZESIS, EVO 60). 2.3. In Vitro Dissolution Studies. To understand the degradation behavior, all of the samples were immersed in simulated body fluid (SBF) at 37 °C with (S/L = 150 g/L).14 The media was replenished two times per week. At predetermined time points (1, 2, 4, 6, and 8 weeks), samples were taken out from SBF, washed with deionization water, and dried before using for weight loss calculations. The metal ion release profile was determined by measuring the ion concentrations (using inductively coupled plasma-mass spectrometry, Agilent/7500ce, USA) in SBF as a function of the immersion time. Archimedes’ principle was used for relative density and porosity calculations. 2.4. Micro-Computed Tomography. μ-CT (Phoenix V|tome|xs, GE, Germany) scan of the samples was performed to study the pore distribution throughout the internal structure.15 The scanning was carried out at 100 kV, 90 μA, without any filter and exposure set to 500 ms. The voxel size was 10 μm, and 1000 slices were captured in one complete rotation. These digitally recorded projections were reconstructed into 3D and 2D images using VG studio MAX 2.2. B


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Figure 1. (A) XRD spectra of the synthesized Mg3(PO4)2 (a) MgP, (b) MgP−0.25Zn, (c) MgP−0.5Zn, (B) XRD spectrum of pure and Zn-doped MgP before and after immersion in SBF for 8 weeks.

Table 1. Relative Density, Grain Size, and Open Porosity of Pure and Zn-Doped MgP Samples sample

composition

relative density (%)

grain size(μm) (fracture surface)

open porosity (%) before/after immersion

MgP MgP−0.25Zn MgP−0.5Zn

MgP + 0 wt % Zn MgP + 0.25 wt % Zn MgP + 0.5 wt % Zn

90.15 ± 0.04 93.47 ± 0.92 96.80 ± 1.13

39.37 ± 13.25 14.37 ± 1.82 10.55 ± 2.13

6.46 ± 1.65/20.86 ± 6.24 3.33 ± 1.50/11.08 ± 1.75 1.66 ± 0.42/9.99 ± 1.62

Figure 2. SEM micrographs of surface morphology of pure and Zn-doped MgP before immersion (a−c) and after immersion (d−f) in SBF for 8 weeks (scale bar represents 20 μm). Insets images (d−f) representing apatite deposition on the surfaces. sections (15−20 μm) were prepared using a differently graded sand paper. The prepared specimens were observed under UV light using Leica DM 2000 Bright (Leica Microsystems, Wetzlar, Germany), where golden yellow fluorescence was observed from regenerated bone and sea green from the old bone. Using ImageJ (National Institute of Health, USA) software, the area was quantified from the pixels of golden yellow and sea green color, and average values were calculated and statistically analyzed. The retrieved bone samples after 30 and 90 days were washed with saline and fixed in 4% paraformaldehyde. Decalcification was carried

out using Gooding and Stewart’s fluid, containing formic acid (15 mL), formalin (5 mL), and distilled water (80 mL), and the fluid was changed once in 3 days. Using a microtome, vertical serial slices (5 mm) were prepared from the paraffin-embedded decalcified samples and stained with haematoxylin and eosin (H&E) and studied under an optical microscope. 2.8. Statistical Analysis. Statistical analysis for weight loss, pH, and Alamar Blue assay was performed using student’s t-test and presented as mean ± SD, where P value 0.05, n = 5). using SPSS software was carried out on fluorochrome labeling results. Mean values and SDs were calculated from five sample (n = 5).

2θ = 21.53°, 23.04°, 25.85° were characteristic of the farringtonite phase (JCPDS #33-0876). No peak shift or secondary phase formation was noticed for doped samples. 3.2. Densification and Microstructures. The relative densities of MgP, MgP−0.25Zn, and MgP−0.5Zn were 90.15 ± 0.04, 93.47 ± 0.92, and 96.80 ± 1.13%, respectively (Table 1).

3. RESULTS 3.1. Phase Determination Using XRD. The XRD patterns of the as sintered samples are shown in Figure 1A. The peaks at D


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the concentration of Mg2+ ions in the degradation medium as a function of the soaking time. It was found that Mg2+ ion concentrations in the degradation media increased during the soaking period for all the samples. The lowest rate of release was found for MgP−0.5Zn samples which directly correlated with the weight loss behavior of the samples. 3.3.2. Strength Degradation in SBF. The change in compressive strength of pure and doped MgP samples, due to degradation in SBF, is shown in Figure 5d. High initial compressive strength was noticed for Zn-doped MgP samples primarily due to higher sintered density. The initial compressive strength was 53.73 ± 18.22, 63.54 ± 7.09, and 64.32 ± 4.13 MPa for MgP, MgP−0.25Zn, and MgP−0.5Zn, respectively, that decreased by 27.82, 6.68, and 7.14% after 8 weeks of immersion. As MgP materials shows considerably higher degradation than TCP or other widely used CaP materials, a sharp decrease in compressive strength after few months is expected and also in line with recent publications.17 3.3.3. Micro-CT Analysis. 3D distributions of pores within the samples, before and after immersion in SBF, were analyzed using μ-CT and are shown in Figure 6. Corroborating to densification

Addition of Zn increased the density in a compositiondependent manner. Porosity in the sintered samples was measured using Archimedes’ principle and is tabulated in Table 1. Almost 50% of the pores in Zn-doped MgP samples were open pores compared to that of 66% for pure MgP. The surface morphology of sintered samples is shown in Figure 2a−c and corresponding EDX spectra in Figure 3a−c. No significant difference in morphology can be seen among the samples; however, when the fracture surface (Figure 4) was studied, some interesting features were noticed. The grains were much finer for both MgP−0.25Zn (14.37 ± 1.82) and MgP−0.5Zn (10.55 ± 2.13) samples compared to pure MgP (39.37 ± 13.25). The fractured grain surfaces were sharp for Zn-doped MgP samples, indicating the brittle mode of fracture. Very small closed pores were found within the grain of Zn-doped MgP samples supporting the density calculation which showed almost 15% higher close porosity for the Zn-doped samples. In comparison, large open pores at the grain triple junction were found for pure MgP, indicating poor sinterability of this material. Mostly, a transgranular fracture was seen where few of them got arrested by the intergrain porosity. In comparison, MgP−0.25Zn and MgP−0.5Zn showed small-sized residual pores near the grain boundaries or triple grain junctions. Both of the transgranular and intergranular fractures were present in MgP−0.25Zn samples. For MgP−0.5Zn samples, the fracture was predominantly intergranular in nature. 3.3. In Vitro Dissolution Studies. 3.3.1. Weight Loss and pH Analysis Mg2+ Concentration Analysis. The in vitro degradation behavior of pure and Zn-doped MgP was studied by immersing the samples in SBF at 37 °C for upto 8 weeks. The weight change during immersion in SBF is shown in Figure 5a. For comparative purposes, similarly sintered TCP samples (with a density of 80%) were also immersed in SBF and their degradation behavior is represented in Figure 5a. TCP samples, even with lower density, did show only about 0.25% weight loss compared to that of 3.35% for pure MgP (sintered density 90%). For Zn-doped MgP samples, the weight loss was 2.18 and 1.64%. The surface morphology after immersion for 8 weeks were studied in SEM and shown in Figure 2d−f. For all the samples, surface porosity increased considerably after immersion that is primarily due to surface dissolution. A considerable increase in open porosity, from 6.46 ± 1.65 to 20.86 ± 6.24%, was noticed for pure MgP. Open porosity increase for MgP−0.25Zn and MgP−0.5Zn was about 7.55 and 8.33%, respectively. After immersion in SBF for 8 weeks, the degraded surface of the undoped and Zn-doped MgP samples were studied by XRD, SEM, and EDX for the degradation and deposition on the surface. The XRD analysis confirms the presence of HA (2θ = 31.7°) in all of the samples (Figure 1B). The SEM images of the degraded surfaces are shown in Figure 2d−f. The inset of Figure 2d−f shows the high-magnification images of the deposited material after immersion in SBF. Although it is very difficult to get any specific feature of HA deposition, the EDX analysis (Figure 3d−f) shows the presence of Ca in the deposit. Therefore, correlating the XRD analysis and EDX results, we can say that the deposits are primarily HA. Figure 5b presents the change in pH of SBF solutions of undoped and Zn-doped MgP samples for different time periods. The pH values for MgP, MgP−0.25Zn, and MgP−0.5Zn were found to be 7.77 ± 0.02, 7.76 ± 0.02, and 7.78 ± 0.01, respectively, after 8 weeks of immersion. These results indicated a slight increase in the alkaline activity in the solution during the degradation of the samples. Figure 5c represents the change in

Figure 6. 3D (a−c) (d−f) and 2D (a1−c1) (d1−f1) μ-CT visualizations of pore distribution of pure and doped MgP before and after immersion in SBF (defect volume mm3).

and microstructural analysis, higher porosity was noticed in μCT images of pure MgP samples which further decreased with the increase in Zn doping. After 8 weeks of immersion, large change in the size of surface porosity was noticed for the pure MgP (Figure 6d) sample that can be attributed to the higher rate of degradation. In contrast, not much degradation (in terms of pore size) was noticed for MgP−0.25Zn and MgP−0.5Zn samples. 2D radiographs also support the fact that there is no preferential surface degradation for MgP−0.5Zn (Figure 6f1) samples after 8 weeks of immersion. For the pure MgP sample, porosities started to appear on the sample surface and just beneath it, indicating faster degradation. 3.4. In Vitro Cytocompatibility. In order to understand the cell−material interactions, MG-63 osteosarcoma fibroblast cells were cultured on pure and Zn-doped MgP samples. For comparative purposes, we have used TCP samples and culture plate as controls. Cell proliferation was studied using the Alamar Blue assay after culturing for 1 day and 3 days and is shown in Figure 7. At day 1, the number of attached cells was significantly lower for pure MgP compared to both the control samples. E


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Figure 7. MG-63 cell proliferation after day1 and day 3 (*p < 0.05, **p > 0.05, n = 9).

Figure 9. Number of live cells quantified from fluorescence images.

However, the lower cell attachment was compensated by the addition of Zn. With increasing culture time, significantly higher cell proliferation was noticed for MgP−0.5Zn samples. A live/ Dead cell viability kit was used to visually verify the cell viability on the samples after 1 and 3 days of culture. Figure 8 shows the fluorescence images of cells and followed the similar trends of Alamar Blue results. High numbers of live cells were noticed on 0.5 wt % Zn-doped MgP samples after 3 days of culture. Figure 9 shows the number of live cells after being quantified from the fluorescence images. The quantitative data also show that higher number of live cells on day 3 for the MgP−0.5Zn sample compared to the MgP and MgP−0.25Zn sample. 3.5. In Vivo Osteogenesis. 3.5.1. SEM and Micro-CT Analysis. In vivo bone-implant material interaction was studied on retrieved bone samples using μ-CT, and the representative images are shown in Figure 10. Figure 10a1−d1 shows 3D images of the bone along with the implants after 30 days of implantation, whereas Figure 10a2−d2 shows the 2D images of the same samples. The control sample showed the formation of fibrotic tissues within the defect site. On the other hand, the formation of new trabeculae was observed for undoped and Zndoped MgP samples. Significant implant degradation was also noticed for implants which was highest for the MgP−0.5Zn sample. Along with high degradation, new bone formation was higher for MgP−0.5Zn samples (Figure 10d2). Ninety days post-surgery 3D and 2D μ-CT images of the bone with implants are shown in Figure 10e1−h1. The defect for the control sample was not completely filled with the bone tissue even after 90 days.

It was observed that the trabeculae density had increased for pure MgP and Zn-doped MgP samples. The ceramic implant had gradually degraded and engulfed by the new bone tissue that can also be seen from Figure 10f1−h1. The MgP−0.5Zn sample showed a large portion of the irregular degraded surface with the newly formed trabecular bone. The thick bone strut formation and maturation was greater in extent for MgP−0.5Zn among all of the compositions. Pure MgP and MgP−0.25Zn samples also showed a prominent irregular bone strut formation adjacent to the ceramic implant which is the indication of osteoconductive behavior of MgP ceramics in vivo. SEM images of the samples at different time points were taken to visualize the nature of gap filling between the bone and implant. After 30 days of operation, images (Figure 11a−c) of the samples showed irregular gap bridging between the implant and host tissue. After 90 days (Figure 11d−f) of implantation, the gap between the bone and implant was completely covered up by the newly formed bone tissue. 3.5.2. Fluorochrome Labeling and Histological Analysis. Histological H&E staining was carried out at the interface of the bone and implant material, which is shown in Figure 12A. At day 30 [Figure 12A(a−d)], the control sample without the ceramic implant showed less cellular infiltration, and the medullary cavity contains few R.B.C and fat cells. The defect with undoped MgP ceramic showed a well-formed matrix with canaliculi and haversian canal organization containing mucinous deposits and infiltration of mononuclear cells. The cortical portion of the bony plate showed mild regeneration of the bony tissue.

Figure 8. Live/Dead images of MG-63 cells culturing for 1 day (a−e) and 3 days (f−j) (scale bar = 200 μm). F


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Figure 10. Representative 3D and 2D (front view) tomographs of bone-implant interface of control, pure, and Zn-doped MgP. (Red dotted circle showing defect site, yellow arrowceramic implant, green arrownew bone formation around the implant).

Figure 11. SEM micrographs of bone-implant interface of pure and Zn-doped MgP after 30 days and 90 days of operation. Red arrowshowing the bone and implant interface. (Scale bar = 100 μm).

Figure 12. (A) Representative micrographs of histological H&E stained sections of the implanted samples after 30 (a−d) and 90 (e−h) days (scale bar200 μm) (yellow arrowosteoblast, blue arrowosteoclast). (B) Fluorescence images of bone-implant interface of pure and Zn-doped MgP after 30 (a−d) and 90 (e−h) days (bright yellow area representing new bone formation; and green area representing the old bone) (scale bar = 500 μm).

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release profile as a function of immersion time. In this present study, it was found that pure MgP has the highest weight loss where MgP−0.5Zn has lowest. The results can be easily correlated to their density, grain size, and porosity. From the fracture surface morphology of the as-sintered samples, it is evident that addition of ZnO has the dual effect of decreased grain size and increased density of MgP. It is believed that ZnO substitution increases lattice diffusion and restricts grain boundary movement of MgP, leading to smaller grain size and increased sintered density. Higher density for MgP−0.5Zn samples results in lower degradability which in turn results in lowest Mg2+ ion concentration in the degradation media. Irrespective of composition, pH values were around 7.77 after 8 weeks of immersion which reveal the alkaline condition of the degradation media. This change in pH can be attributed to the fast ion exchange of cations from samples with H+/H3O+ ions of SBF.18 Research suggests that weak alkaline condition is favorable for cell proliferation and differentiation.19 Mechanical strength is also an important property for clinically adoptable implant materials. For degradable bioceramics, the strength starts to decrease as the material starts to be resorbed. In this study, a significant decrease in strength was due to the increasing porosity in all of the MgP samples after immersion. Extensive μCT study showed a decrease of the sample volume, and SEM images showed the degraded surface morphology of all of the samples after 8 weeks of immersion in SBF. The primary advantage of MgP-based ceramics is its fast degradation behavior along with cytocompatibility. As addition of Zn alters the degradation kinetics of MgP ceramics, MG-63 cells were cultured for 3 days to understand the cell material interaction. Live/Dead images and Alamar Blue assay shows higher cell attachment and proliferation on Zn-doped MgP samples compared to pure MgP samples which can be attributed to the beneficial effect of Zn on various biochemical processes. Zn can markedly enhanced the insulin-like growth factor-I (IGFI) and transforming growth factor-β l production.20,21 The IGF-I can stimulate cellular protein synthesis, protein concentration, and the DNA content of osteoblastic cells. Zn also modulates the leucyl-tRNA synthetase activity for the biosynthesis of bone proteins.22 Earlier studies have already shown that presence of Zn can effectively stimulate the cell proliferation and differentiation on bioceramics.23,24 The MG-63 cell viability and proliferation rate were increased for Zn-doped HA ceramics compared to pure HA samples.25 Mg2+ can help in maintaining the cell structural integrity and stimulates platelet-derived growth factor (PDGF). The PDGF can enhance DNA and collagen synthesis and promotes adhesion, proliferation, and migration of osteoblastic cells.26 The presence of Mg2+ ions can effectively increase the MG-63 cell proliferation and differentiation of the biomaterials.27,28 On a comparative note, widely studied TCP bioceramic showed higher cell proliferation compared to all of the MgP samples, that is due to more stable material surface and calcium release from TCP samples. For the clinically adoptable biomaterial, the in vivo boneimplant interaction study is an important parameter to understand the suitability of the material. To determine the in vivo bone forming ability, undoped and Zn-doped MgP ceramics were implanted in the rabbit femur. The 3D μ-CT, quantitative fluorochrome labeling studies and SEM analysis show that Zn-doped implants support enhanced osseointegration and bone formation than pure MgP. Newly formed bone quantification by fluorochrome labeling studies shows signifi-

However, Zn-doped MgP showed a well-organized bony structure containing a haversian canal and fibro-mucinous deposition. The proliferating osteocytes and sufficient amount of R.B.Cs, mononuclear cells, and fibrinous cells in the medullary cavity indicated that the healing process was started. After 90 days, the cellular level interaction was also studied and represented in Figure 12A(e−h). The control sample showed irregular angiogenesis with large number of osteocytes . The bony section was invaded by few osteoclastic and osteoblastic cells. The medullary cavity was filled with the adipose tissue, R.B.C., and few mononuclear cells. The pure MgP sample showed the formation of a bony matrix containing haversian canal, bony lamellae, and sinusoidal space. The medullary cavities are irregular and contained few lining fibroblastic cells. While the Zn-doped samples indicated a well-ordered bony structure comprising of abundant haversian canal and few sinusoidal space along with infiltrated RBCs, abundant mononuclear cells, and osteoblasts. Among all of the implanted samples, the MgP−0.5Zn sample showed the highest osteoclast formation among all of the implant samples (as indicated by a blue arrow in Figure 12Ah). OTC labeling was used for the quantitative analysis of new bone formation. The OTC-labeled samples under UV radiations emit bright golden yellow fluorescence from the newly generated bone and dark green fluorescence from the old bone. After 30 days [Figure 12B(a−d)], MgP−0.25Zn and MgP−0.5Zn emitted higher intense golden yellow color than MgP sample which is the indication of an enhanced bone regeneration. The percentage of the new bone formation was calculated after 30 days and found to be 47 ± 3, 53 ± 3, and 57 ± 3% for MgP, MgP−0.25Zn, and MgP−0.5Zn, respectively and are shown in Table 2. After 90 days [Figure 12B(e−h)] of Table 2. Percentage of New Bone Formation in the Operated Site of Pure and Zn-Doped MgP Ceramicsa composition

30 days

90 days

control MgP MgP−0.25Zn MgP−0.5Zn

36a ± 3 47b ± 3 53c ± 3 57d ± 3

44a ± 2 56b ± 3 69c ± 2 78d ± 3

a Means (n = 4) along with different letters (a−d) above each value differ significantly (p < 0.05) among the groups (one way ANOVA).

surgery, quantification of newly formed bone using fluorochrome labeling showed a stimulatory effect of Zn in bone formation. At this time point, 0.25 and 0.5 wt % Zn-doped MgP showed 69 ± 3% and 78 ± 3% new bone formation, respectively, whereas only 56 ± 3% bone was regenerated for undoped MgP ceramics.

4. DISCUSSION In this present study, we have incorporated Zn in MgP ceramics with dual purpose of controlling its in vitro and in vivo degradability and biocompatibility. In order to know the effects of Zn concentration on MgP ceramics, a comparative in vitro and in vivo test was carried out with undoped and doped MgP ceramics. To begin with, we have confirmed, through XRD analysis, that the synthesized material is farringonite [Mg3(PO4)2] and Zn addition has no effect on phase purity or peak positions. For degradable bioceramics, the well-established technique for the degradation study is weight change, pH change, ions H


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to that of 56 ± 3% for pure MgP. The controlled in vivo degradation combined with increased ossointegration makes Zn-doped MgP ceramics an attractive alternative material for bone tissue engineering according to patient specific requirement.

cantly higher bone regeneration for Zn-doped samples, which can be attributed to the control degradation kinetics. The in vivo degradability of MgP−0.5Zn is significantly higher (as seen in Figure 10d2,h2) compared to pure and MgP−0.25Zn which is in complete contradiction to in vitro degradability. The results can be explained by in vivo degradation mechanisms of bioceramics which has two primary components: chemical degradation and cellular (osteoclast mediated) degradation. In the in vitro model, we have primarily determined the chemical degradability of MgP which showed higher degradation of MgP samples owing to their low density and high open porosity. The higher degradability of MgP results in higher Mg ion concentration in the degradation media. Recent studies have indicated that high concentrations of Mg can significantly reduce osteoclast activity.29 Therefore, MgP, which has higher chemical degradability, does have reduced osteoclast cell-mediated degradation. On the contrary, the MgP−0.5Zn sample, owing to their low in vitro degradation has lower Mg ion release and thereby has higher in vivo degradability. The results are also supported by the fact that more number of osteoclast cells were found in histological results of MgP−0.5Zn samples both at 30 and 90 days compared to undoped MgP. The higher osteoclast activity for MgP−0.5Zn sample also results in a higher bone remodeling rate around the implant−bone interface which is reflected in fluorochrome labeling results. After 90 days of implantation, abundant thicker bone strut formation and maturation were greater in the extent for MgP− 0.5Zn among all of the compositions that can be attributed to the higher bone remodeling and beneficial roles of Zn. The presence of Zn can enhance the osteogenesis by stimulating the transforming growth factor beta and bone morphogenetic protein.30,31 It is also reported that Zn can induce osteoblastogenesis, as it is a cofactor of more than 300 enzymes and plays a major role in stabilization of protein and signaling enzymes.32 Although we could not detect Zn in the in vitro dissolution study, we confirmed the presence of Zn through the EDX analysis and believe that the enhanced osseointegration of MgP−0.5Zn is influenced by Zn doping. The present result indicates that selection of a dopant for MgP bioceramics can significantly alter the in vitro and in vivo degradability and overall performance of degradable bioceramics.

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AUTHOR INFORMATION

Corresponding Authors

*E-mail: [email protected] (S.K.N.). *E-mail: [email protected] (M.R.). ORCID

K. Bavya Devi: 0000-0002-5513-6066 Samit Kumar Nandi: 0000-0002-6487-4546 Mangal Roy: 0000-0002-1490-9624 Author Contributions §

K.S. and V.K. have equally contributed and are “Joint First Author”.

Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS M.R. would like to acknowledge the financial assistance from Science and Engineering Research Board (SERBSB/FTP/ ETA-0114/2014), Department of Science and Technology (DST), India and technical personnel at the Central Research Facility, IIT-Kharagpur, for their support in μ-CT analysis. S.K.N. would like to acknowledge the support from the Honorable Vice Chancellor, West Bengal University of Animal and Fishery Sciences, Kolkata.

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REFERENCES

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5. CONCLUSIONS In the present work, we have successfully prepared Zn-doped MgP bioceramics from magnesium hydroxide and magnesium trihydrogen phosphate by the conventional solid-state sintering method. Addition of Zn in MgP effectively increased the sintered density from 90% for pure MgP to 96% for MgP−0.5Zn samples. Weight loss measurement indicated reduction in degradability of Zn-doped samples which was also supported by FESEM and μ-CT study. The in vitro cell material interaction using human osteoblast-like cells (MG-63) proved that all of the samples were cytocompatible where cell proliferation was significantly increased for Zn-doped samples in a concentration-dependent manner. In vivo degradability and bone regeneration capability were studied after implanting the samples in the critical size defect in the rabbit femur. Detailed μ-CT and histological analysis of samples showed that MgP− 0.5Zn samples had the highest degradation, osseointegration, and bone formation, which is primarily due to increased number of osteoclast cells and their activity. The quantitative analysis using OTC labeling after 90 days of post implantation has shown 78 ± 3% bone regeneration for MgP−0.5Zn samples compared I


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