Silicate bioceramic/PMMA composite bone cement with distinctive physicochemical and bioactive properties

Abstract

Polymethyl methacrylate (PMMA) bone cement has been widely used for orthopedic application due to its high mechanical strength and proper setting time. However, the major shortcomings of PMMA bone cement are its relatively low osseointegration and strong exothermic reactions. Silicate-based bioceramics, like akermanite (Ca₂MgSi₂O₇, AKT), have been demonstrated to possess excellent osteostimulation ability and controlled biodegradability. The purpose of this study is to harness the advantages of both PMMA and AKT in order to prepare a new kind of composite bone cement (AKT/PMMA) with superior mechanical strength, improved exothermic effect and osteogenic activity. AKT particles were uniformly incorporated into the matrix of PMMA cement. The effect of AKT on the in vitro setting behaviors, and mechanical and biological properties of resultant composite cements was systematically investigated. The results showed that the prepared AKT/PMMA composite bone cements revealed significantly decreased polymerization temperature as compared with pure PMMA, but maintained ideal setting times (12–14 min) and high mechanical strength (∼100 MPa for compressive strength). Most interestingly, the incorporation of AKT into PMMA improved its osteogenic activity, as indicated by the significantly enhanced apatite-mineralization ability and stimulatory effect on the proliferation and alkaline phosphate (ALP) activity of osteoblasts. The results suggest that AKT/PMMA composite bone cements possess distinctive physicochemical and bioactive properties, and are a promising injectable biomaterial for orthopedic applications.

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1. Introduction

Polymethyl methacrylate (PMMA) bone cement has been widely used for prosthetic fixation or as a bone substitute in orthopedics owing to its desirable mechanical strength, relatively low toxicity and good handleability in the surgical room.¹ Despite these advantages, it is found that PMMA usually fails to bond directly to the surrounding bone tissue, and instead an intervening soft tissue layer is formed between the bone and the cement.² Such cement–bone interface is considered as a weak-link zone, which would lead to aseptic loosening of implants from surrounding bone over a certain period.³ In addition, PMMA cement initiates high exothermic temperature during setting reaction and has potentially monomer-derived toxicity. The temperature during the polymerization process of MMA monomers is up to 100 °C, which is compromised to the surrounding cells and tissues.⁴ A large number of studies have been conducted to improve the clinical performance of PMMA bone cement. An effective approach is to incorporate bioactive inorganic fillers into PMMA matrix.⁵ Previously, incorporation of bioactive fillers such as bioactive glasses,⁶ bioactive glass-ceramics,^7,8 titania,^9,10 and hydroxyapatite (HA)^11–13 has been conducted. Although the method can improve their bioactivity, it is far from optimal for the prepared PMMA-inorganic composite bone cements. Among various bioactive fillers, titania filler is nonbiodegradable and lacks bioactivity. Apatite-wollastonite (A-W) glass-ceramic and hydroxyapatite lack sufficient biodegradability and osteostimulation activity. Although bioglasses possess osteostimulation activity, their mechanical strength is low, which limits its application.¹⁴ These composite cements can hardly combine good physicochemical and bioactive properties for better clinical applications. It remains a significant challenge to develop high-performance bone cement in combination of distinct mechanical strength, self-setting property and bioactivity.

In recent years, silicate-based bioceramics, like akermanite, has been demonstrated to possess excellent mechanical property and osteostimulation activity.^15–17 It has been found that AKT bioceramics could significantly stimulate the in vitro osteogenic and angiogenic differentiation of several kinds of stem cells, including bone marrow stromal cells,^18–20 adipose-derived stem cells,^21,22 periodontal ligament cells,^23,24 and human aortic endothelial cells.^20,25 Further in vivo studies have shown that AKT bioceramics possess ability to support bone regeneration and angiogenesis as compared to conventional β-TCP bioceramics.^19,25 All these studies reveal that AKT possesses improved mechanical property, controlled biodegradability and excellent osteostimulation activity, suggesting that AKT may be a suitable bioactive filler to improve the physicochemical and bioactive properties of PMMA. Therefore, the aim of this study is to prepare AKT/PMMA bioactive composite bone cements. The effect of AKT on the self-setting behavior and mechanical strength of the PMMA matrix was investigated. Furthermore, the ability of the composite cement to induce apatite formation and support the adhesion, proliferation and differentiation of osteoblast cells was systematically studied.

2. Materials and methods

2.1 Preparation of bone cements

Akermanite (Ca₂MgSi₂O₇, AKT) powders were synthesized by the sol–gel process using tetraethyl orthosilicate ((C₂H₅O)₄Si, TEOS), magnesium nitrate hexahydrate (Mg(NO₃)₂·6H₂O) and calcium nitrate tetrahydrate (Ca(NO₃)₂·4H₂O) as raw materials according to our previously study.²⁶ A commercial PMMA cement (Acrylic bone cement III) with BaSO₄ as radio-opaque filler (Tianjin Synthetic Material Research Institute®) was used for the study. The size distribution of the original PMMA and AKT powders was examined with a laser particle size distribution instrument (Dandong Bettersize Instruments Ltd, BT-9300Z). The AKT/PMMA bone cement was prepared according to our previous patent.²⁷ To prepare AKT/PMMA mixtures, the AKT powder was uniformly mixed with the PMMA powders with different weight percentage. To prepare the cement paste, the mixture powders were manually added with the monomer liquid at a powder-to-liquid ratio of 2 : 1 (g mL⁻¹) until homogeneous pastes were obtained. Then the mixture was poured into a Telflon mold and stored for 3 hours. Afterwards, the hardened cement paste was demolded and stored at room temperature for further characterization. Details about the compositions of the cement powders were shown in Table 1. The composition and morphology of AKT powders, PMMA bone cement and AKT/PMMA bone cements after setting were characterized by X-ray diffraction (XRD; Geigerflex, Rigaku, Japan) and scanning electron microscopy (SEM; JSM-6700F, JEOL, Japan), respectively.

Table 1 Composition of AKT/PMMA bioactive cements

Cement	Filler loading (wt%)	Powder parts		Liquid parts (mL)	L/P P = PMMA + AKT
		PMMA powder (g)	AKT (g)
a Commercial bone cement (Tianjin Synthetic Material Research Institute®).
Control (PMMA^a)	0	2	0	1	0.5
10AKT/PMMA	10	2	0.2	1	0.45
30AKT/PMMA	30	2	0.6	1	0.38
50AKT/PMMA	50	2	1.0	1	0.33

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2.2 Self-setting properties and mechanical strength of bone cements

The setting properties of the bone cements were investigated according to ISO standard.²⁸ The temperature changes during the setting reaction were measured using a K-type thermocouple connected to data acquisition equipment (FLUKE HYDRA SERIES II). Immediately after bone cement was poured into a Teflon mold under ambient conditions (23 °C), the inner temperatures of the paste were measured and recorded once every 5 s intervals for total time length of 25 min. The profile of temperature (T) versus time (t) is applied to determine the peak temperature and setting time. The maximum temperature within the profile is assigned to the peak temperature during polymerization of the composite cement. The setting time is the time point when the exothermic temperature rises to the midpoint between the ambient and peak temperature.^4,7 The setting temperature was determined using eqn (1),

T_set = (T_max + T_amb)/2 (1)

where T_set was the setting temperature, T_max was the peak temperature and T_amb was the ambient temperature. The measurements were repeated three times for reproducibility.

To test the mechanical strength of the prepared composite cements, cylindrical specimens with 12 mm in height and 6 mm in diameter were prepared for compression testing according to ISO standard.²⁸ The compressive strength of the bone cements was measured by using a universal mechanical testing machine (Shimadza, AG-5KN, Japan) at cross-head speed of 20 mm min⁻¹. Six samples per group were tested to obtain an average value of the mechanical strength.

2.3 Apatite-mineralization ability of composite cements

Simulated body fluid (SBF) were prepared according to the method described by Kokubo.²⁹ PMMA and AKT/PMMA bone cement discs with 2.5 mm in height and 6 mm in diameter were immersed in 10 mL SBF solution and kept in a 37 °C water bath with a shaking rate of 90 rpm. After 1, 3, 7, 10 and 14 days of soaking, the pH values of the solution were measured using an electrolyte type pH meter (PHS-2C, Jingke Leici Co., Shanghai, China) without refreshment of the immersion medium. Three samples from each group were tested to obtain an average value.

To assess the apatite-mineralization ability and changes in ion concentrations of SBF solutions, the immersion solutions were updated and collected at every time period. The specimens were removed from the SBF after immersion for 7 days, gently rinsed with deionized water and then dried under vacuum at 37 °C for 24 h. The surfaces of the specimen were characterized by grazing incidence X-ray diffraction (GIXRD; Geigerflex, Rigaku, Japan) and SEM equipped with energy-dispersive spectrometry (EDS; INCA Energy, Oxford Instruments, UK). The concentration of Ca, Si, Mg, P and Ba ions of SBF was measured by using inductively coupled plasma optical emission spectrometer (ICP, Varian 715-ES). Three samples from each group were tested to obtain an average value.

2.4 Attachment and proliferation of osteoblast-like cells

Fourth passage MC3T3 (MC3T3-E1 Subclone 14) cells were purchased from the cell bank of Chinese Academy of Sciences (Shanghai, China). To evaluate the attachment of cells on the surface of the materials, the cells were cultured on the PMMA and AKT/PMMA bone cements (Ø12 × 2 mm) placed in a 24-well culture plate at an initial density of 6 × 10³ cells per specimen, and incubated in α-MEM culture medium supplemented with 10% FBS maintained at 37 °C in a humidified atmosphere of 95% air and 5% CO₂. After 1 and 7 days of incubation, the specimens were removed from the culture wells and rinsed with phosphate buffered saline (PBS). For SEM observation, the cells on the specimens were fixed with 2.5% glutaraldehyde for 20 min, dehydrated in a grade ethanol series (30, 50, 70, 90, and 100% (v/v)) for 10 min, respectively, with final dehydration in absolute ethanol twice followed by drying in hexamethyldisilazane (HMDS) ethanol solution series. The morphological characteristics of the attached cells on the specimen were observed by using SEM.

The proliferation of MC3T3 cells on PMMA and AKT/PMMA bone cements were estimated using MTT assay. Briefly, MC3T3 cells were cultured on the bone cements (Ø12 × 2 mm) placed in a 24-well culture plate at an initial density of 6 × 10³ cells per specimen. The cells were cultured at 37 °C in a humidified atmosphere of 95% air and 5% CO₂. After 1, 3 and 7 days of incubation, the culture medium was replaced with 500 μL of MTT solution (freshly made 5 mg MTT powder in 1 mL PBS). Following incubation for 2 h at 37 °C in 5% CO₂, the MTT solution was removed, and 300 μL DMSO was added into each well to completely dissolve the precipitated formazan crystals. The formazan solution (100 μL) of each sample was subsequently added into individual wells of a 96-well plate and the absorbance at a wavelength of 590 nm was measured by a microplate reader (Bio-Tek Instruments Inc., USA). All experiments were performed in triplicate, and the results were shown as units of optical density (OD) absorbance value.

2.5 ALP activity assay of osteoblast-like cells

The ALP activity assay was performed on days 7 and 14 to assess the effect of AKT filler on the early osteogenic differentiation of MC3T3 cells. Cells were seeded at a concentration of 6 × 10³ cells per disc (Ø12 × 2 mm) placed individually in a 24-well plate. The cells were left to grow for 7 and 14 days at 37 °C in a humidified atmosphere of 5% CO₂. ALP activity was assayed by the widely used PNPP method. Aliquots of cell lysates were incubated with reaction solution (containing 2-amino-2-methyl-1-propanol, MgCl₂ and p-nitrophenylphosphate) at 37 °C for 30 min. The conversion of p-nitrophenylphosphate to p-nitrophenol was stopped by adding NaOH, and the absorbance at 405 nm was measured with a spectrophotometer (Epoch™ microplate spectrophotometer, Bio-Tek Instruments, USA). The ALP activity was normalized to the total intracellular protein content. ALP activity of the cells cultured on PMMA bone cement served as a control.

2.6 Statistical analysis

All date were expressed as means ± standard deviation (SD) and were analyzed using one-way ANOVA with a post hoc test. A p-value <0.05 was considered statistically significant.

3. Results

3.1 Characterization of bone cements

The SEM images and size distribution profile of PMMA and AKT powders were shown in Fig. 1. PMMA beads were partially covered with barium sulfate (BaSO₄) powders (Fig. 1a), and the size of PMMA beads fell in the range of 10–80 μm (Fig. 1b). The AKT powders showed an irregular shape with particle sizes ranging from 5 to 40 μm (Fig. 1c and d).

Fig. 1 The SEM images of the PMMA powders (a), AKT powders (c); the size distribution graphs of the PMMA powders (b), AKT powders (d).

The XRD patterns of the AKT, PMMA and 30AKT/PMMA were presented in Fig. 2. It was found that akermanite crystal phase were distinct in the patterns (Standard Card no: JCPD 83-1815). Barium sulfate as radio-opaque filler existed in PMMA bone cement (Standard Card no: JCPD 24-1035). The characteristic peaks for Ca₂MgSi₂O₇ and BaSO₄ crystals were obviously observed in the pattern of 30AKT/PMMA bone cements.

Fig. 2 The XRD patterns of AKT powders, PMMA bone cement and 30AKT/PMMA bone cement.

Fig. 3 revealed the cross-sectional microstructure and compositions of PMMA and 30AKT/PMMA bone cements. It was observed that AKT powders were uniformly dispersed and interacted with the PMMA matrix (Fig. 3b), and there was no discernible defect within the composite bone cements. The BSEM image showed that two different contrast particles existed in the 30AKT/PMMA bone cements (Fig. 3c). EDS analysis revealed that sulfur (S) and barium (Ba) existed in the bright BaSO₄ particles (Fig. 3d), and calcium (Ca), silicon (Si) and magnesium (Mg) existed in the dark Ca₂MgSi₂O₇ particles (Fig. 3e).

Fig. 3 The cross-section microstructure of PMMA bone cement (a) and 30AKT/PMMA bone cement (b). The BSEM images of 30AKT/PMMA bone cement (c). The EDS analysis of two different particles (BaSO₄ and Ca₂MgSi₂O₇) in 30AKT/PMMA bone cement (d and e).

3.2 Self-setting and mechanical properties of bone cements

Fig. 4 illustrated the change in temperature of the cement paste during the setting process. In all cases, the inner temperature of the pastes increased as the reaction preceded, reaching a maximum temperature and then decreasing with time. It was observed that the peak temperature of polymerization of 10AKT/PMMA, 30AKT/PMMA and 50AKT/PMMA decreased with the increasing AKT content (85.4 ± 1.85, 80.2 ± 2.14 and 71.5 ± 2.85 °C, respectively), which was lower than that of pure PMMA bone cement (88.7 °C). It can be seen that the setting times for the composite cements slightly increased with the increase of AKT contents from 12.3 ± 0.32 min for pure PMMA to 14.08 ± 0.27 min for 50AKT/PMMA (Table 2).

Fig. 4 Graph of temperature versus time for exothermic temperature changes during the setting process of cement sample: PMMA, 10AKT/PMMA, 30AKT/PMMA, 50AKT/PMMA.

Table 2 Mechanical and setting properties of PMMA-based bone cements

	UCS (MPa)	Tmax (°C)	Tsetting (min)
a Commercial bone cement (Tianjin Synthetic Material Research Institute®).b ISO 5833 for determining tset, Tmax and UCS.^28
PMMA^a	96.70 ± 3.91	88.7 ± 3.05	12.3 ± 0.32
10AKT/PMMA	99.61 ± 7.55	85.4 ± 1.85	12.78 ± 0.28
30AKT/PMMA	97.63 ± 5.98	80.2 ± 2.14	13.18 ± 0.36
50AKT/PMMA	101.45 ± 4.05	71.5 ± 2.85	14.08 ± 0.27
Standard^b	>70	<90	5–15

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The results of the mechanical test were presented in Fig. 5. Fig. 5a showed the stress–strain curves of bone cements for the compressive tests. The compressive strength of the specimens was 96.70 ± 3.91, 99.61 ± 7.55, 97.63 ± 5.98 and 101.45 ± 4.05 MPa for pure PMMA, 10AKT/PMMA, 30AKT/PMMA and 50AKT/PMMA bone cements, respectively (Fig. 5b). There was no significant difference in compressive strength among the four cements (p > 0.05). The elastic modulus of 10AKT/PMMA and 30AKT/PMMA specimens were comparable to that of pure PMMA, whereas 50AKT/PMMA bone cements showed a significantly higher elastic modulus (p < 0.05) than other groups (Fig. 5c).

Fig. 5 Mechanical properties of AKT/PMMA bioactive cements. (a) Stress–strain curves for compressive tests, (b) compressive strength (MPa), (c) elastic modulus (MPa). *: Significant difference (P < 0.05).

3.3 Apatite-mineralization ability of composite cements

Fig. 6 showed GIXRD patterns of the cements after soaking in SBF for 7 days. The diffraction peaks for low crystallinity of hydroxyapatite phase (Standard Card No: JCPD 24-0033) were detected at 2θ 31.738° for 50AKT/PMMA bone cement after soaking in SBF.

Fig. 6 The GIXRD patterns of the bone cements after soaking in simulated body fluid (SBF) for 7 days.

Fig. 7 revealed the SEM and EDS analysis for the surface of the bone cements after soaking in SBF for 7 days. No particle deposition was observed on the surface of the PMMA bone cements (Fig. 7a and b). EDS analysis demonstrated that there were no Ca and P elements, but only S and Ba elements existed in pure PMMA bone cement (Fig. 7c), which was attributed to the presence of BaSO₄. In contrast, newly formed clusters of particles were observed on the surface of the AKT/PMMA bone cements after exposure to SBF (Fig. 7d, g and j). It can be seen that the coverage of deposition layer on the surface increased with the increasing AKT contents. Only a small number of granular crystals appeared on the surface of 10AKT/PMMA bone cement (Fig. 7d). As the AKT content increased up to 50%, the surface of the sample was completely covered by newly formed apatite layer (Fig. 7j). Higher magnification images showed that the newly formed apatite layers on the composite cements were composed of aggregates of nanocrystals with worm-like morphology (Fig. 7e, h and k). The EDS analysis confirmed that the Ca/P ratio of the formed apatite on the 10AKT/PMMA, 30AKT/PMMA and 50AKT/PMMA bone cements was 1.35, 1.53 and 1.64, respectively (Fig. 7f, i and l).

Fig. 7 SEM morphologies and EDS results of AKT/PMMA bone cements, after soaking in simulated body fluid (SBF) for 7 days (a, b and c: PMMA bone cement; d, e and f: 10AKT/PMMA bone cement; g, h, i: 30AKT/PMMA bone cement; j, k and l: 50AKT/PMMA bone cement. a, d, g and j: Lower magnification; b, e, h and k: higher magnification, size bars represent 100 nm).

The changes in ion concentrations and pH value of SBF solutions during the immersion test were shown in Fig. 8. It can be seen that the Ca, P, Mg and Si concentrations of SBF solutions remained nearly constant for PMMA bone cement. However, the concentration of Ca and P in SBF decreased distinctively after 3 days of soaking with AKT/PMMA bone cements (Fig. 8a and b), and the Mg and Si ions release of AKT/PMMA bone cements increased with the increasing AKT contents (Fig. 8c and d). It was obvious that at the first 7 days, the release of SiO₄⁴⁻ ions from AKT/PMMA composites led to a significant increase of Si concentration in SBF. After 7 days of soaking, the apatite layer was formed on the surface of composite cements, in which parts of Si ions might be incorporated into the newly formed apatite layer, and the apatite layer might further inhibit the release of SiO₄⁴⁻ ions from cements, leading to the decrease of Si concentration in SBF at the late stage of soaking. In addition, the Ba ion release of all the groups showed similar variation trends, and decreased with the increasing AKT contents (Fig. 8e). Fig. 8f revealed the change of pH value of the SBF solution after soaking PMMA and AKT/PMMA bone cements for various times. The pH value of AKT/PMMA-soaked solution increased with the increase of AKT contents, which was higher than that of PMMA-soaked solution.

Fig. 8 Changes of Ca (a), P (b), Mg (c), Si (d), Ba (e) ion concentrations and pH value (f) of the SBF solution after soaking PMMA and AKT/PMMA bone cements for various times. (a–e): the SBF solution was replaced and collected at every time period. (f) The SBF solution was not replaced during the whole periods of soaking test.

3.4 Attachment, proliferation and ALP activity of MC3T3-E1 cells on composite bone cements

The cell morphology on different bone cements was characterized by SEM (Fig. 9). It can be seen that, after 1 and 7 days of culture, all of the bone cements supported MC3T3 cell attachment. On the surfaces of PMMA bone cement, the cells displayed a spindle shape with discernible filopodia (Fig. 9a and b). However, for AKT/PMMA bone cements, the cells exhibited flattened morphology. Furthermore, high contents of AKT addition lead to the better cell spreading on the surface of the composite bone cements (Fig. 9c–h).

Fig. 9 The SEM images of MC3T3-E1 cells cultured on the PMMA bone cements and AKT/PMMA bone cements for 1 day (a, c, e and g) and 7 days (b, d, f and h). (a and b) PMMA bone cements; (c and d) 10AKT/PMMA bone cements; (e and f) 30AKT/PMMA bone cements; (g and h) 50AKT/PMMA bone cements.

MTT analysis showed that all bone cements supported cell proliferation with the increase in culture time (Fig. 10). There was no significant difference in cell number between the PMMA and AKT/PMMA bone cement groups (p > 0.05) after 1 and 3 days of culture. Whereas after further prolonged culture (at day 7), the cell numbers on 30AKT/PMMA and 50AKT/PMMA bone cements were significantly higher than those on the PMMA and 10AKT/PMMA groups (p < 0.01). 50AKT/PMMA bone cement showed highest cell proliferation rate among all groups (p < 0.05).

Fig. 10 The proliferation of MC3T3-E1 cells on the bone cements after different culture periods. (*) p < 0.01 compared with PMMA and 10AKT/PMMA at day 7. (#) P < 0.05 compared with 30AKT/PMMA at day 7.

The ALP activities of MC3T3 cells after culture on the PMMA and AKT/PMMA bone cements for 7 and 14 days were shown in Fig. 11. It was found that there were no significant differences in ALP activity among the different groups after 7 day culture (p > 0.05). With prolonged incubation from 7 to 14 days, ALP activities of MC3T3 cells on all groups significantly increased. Furthermore, the AKT/PMMA bone cements had higher ALP activities than the PMMA bone cement (p < 0.05), and there were no significant differences in ALP activity among AKT/PMMA composite cements with different AKT contents.

Fig. 11 Alkaline phosphatase activity of MC3T3-E1 cells cultured on the bone cements for 7 days and 14 days. *: Significant difference (P < 0.05).

4. Discussion

One of the major concerns over the use of PMMA bone cement is that the polymerization of MMA into PMMA initiates a tremendous heat release, which inevitably leads to a significant increase in local temperature up to 100 °C and thus results in cell death and damage to the surrounding tissue.³ Therefore, the self-setting properties of PMMA bone cement are considered one of the most important factors that influence its clinical application. In our study, the peak temperature of AKT/PMMA composite bone cements was significantly lower than that of pure PMMA cement (88.7 °C) and ISO standard (90 °C).²⁸ Moreover, the peak temperature decreased with the increase of the AKT contents in the composite cement. The main reason is that ceramic materials can act as heat insulators and absorb part of the heat produced by the exothermal polymerization reaction, and then leading to a reduction of peak temperature.^3,7,30,31 The relative low temperature may be beneficial to improved biocompatibility. In addition, it was found that the setting times of AKT/PMMA composite bone cements were comparable to that of pure PMMA (∼12–14 min), indicating that AKT/PMMA composite bone cements maintain a proper self-setting time.²⁸

Besides the decrease of exothermal effect of AKT/PMMA composite bone cements, they still maintained high mechanical strength as compared to pure PMMA cements. Previous studies have found that the addition of bioactive particles could improve the bone-bonding ability of PMMA cements, but an inappropriate preparation process would result in particle aggregation in the cement matrix, thus causing mechanical deterioration.^32,33 In our study, the AKT powders were firstly uniformly mixed with PMMA powders, and then were stirred with the liquid monomer. This process could make the AKT particles uniformly disperse in the PMMA cement matrix, even when the content of AKT was up to 50%. The results indicated that AKT particles did not result in any significant mechanical deterioration of the PMMA matrix. The uniform dispersion of AKT within the PMMA matrix is essential to retain the good mechanical properties of PMMA cement, as the presence of particle agglomerates is frequently related with the occurrence of mechanical failure.^34,35 The compressive strength of AKT/PMMA composite bone cements was around 100 MPa, which was comparable to that of commercial PMMA bone cement (73–120 MPa),^5,36 and higher than that of commercial CPC bone cement (12–55 MPa).³⁷ The strength of AKT/PMMA composite bone cements matched well the requirement of clinical applications according to international standard organization (ISO5833: 70 MPa).²⁸

Attachment and spreading of cells belong to the first phase of the cell/material interactions and influence their capacity to proliferate and differentiate in contact with the implants.³⁸ It is well known that material characteristics are able to influence cellular response. In our study, both PMMA and AKT/PMMA bone cements were able to support cells adhesion, while the cells on the surface of AKT/PMMA bone cements showed a better spreading behavior. Furthermore, significantly higher increase in cell number and ALP activity was observed for MC3T3 cells cultured on AKT/PMMA than PMMA. The improvement of cell proliferation and bone-related differentiation can be related to the formation of the apatite layer, the release of Ca, Mg and Si ions from AKT/PMMA bone cements and beneficial pH microenvironment. Firstly, the formation of an apatite layer on the surface of the biomaterial is known to favor cell proliferation and differentiation.^29,39–41 In our study, there was no discernable deposition on the surface of pure PMMA bone cement. In contrast, apatite deposition was distinct on the surface of the AKT/PMMA composite bone cements after immersion in SBF. As AKT possessed excellent apatite-formation ability in SBF,²¹ it was obvious that the formation of the apatite layer on the composite cements was attributed to the presence of AKT particles within the PMMA matrix. With the ability of inducing apatite formation on the surface, it is understandable that the AKT/PMMA bone cements have better ability to favor cell proliferation and differentiation as compared with PMMA cements. Secondly, the Ca²⁺, and SiO₄⁴⁻ containing ionic products released from AKT within the composite cements could contribute to the improved cell viability and ALP activity, in which the stimulatory effect of AKT bioceramics on bone cells was indicated in our previous studies.^18–22 Thirdly, previous studies have shown that pH microenvironment plays an important role to influence cell response.^42–44 In this study, it is obvious that the incorporation of AKT powders into PMMA leads to a more neutral pH value in simulated body fluids. Therefore, the improved viability and ALP activity of MC3T3 cells on the composite cements could be related to their beneficial pH microenvironment as compared with pure PMMA cements. Our results suggested that the incorporation of AKT into the PMMA matrix could offer a more beneficial microenvironment for cell attachment, proliferation and differentiation. Further in vivo study is necessary to investigate the osseointegration of AKT/PMMA cements.

5. Conclusions

New bioactive AKT/PMMA bone cements were successfully prepared. The incorporation of AKT into PMMA significantly improved the self-setting properties, including the lower polymerization temperature and proper setting time, and high mechanical strength. Furthermore, AKT/PMMA bone cements possessed excellent apatite-mineralization ability in SBF and showed better ability to support cells attachment, proliferation and ALP activity as compared with PMMA cements. Our results indicated that the new AKT/PMMA composite bone cement may be a promising candidate material for orthopedic applications due to its improved self-setting and biological properties.

Acknowledgements

This work was supported by grants from the Natural Science Foundation of China (grants no. 81190132), the Funds of the Clinical Research Center for Biomaterials, Shanghai Institute of Ceramics, CAS (grants no. BMCRC2010001), Key Research Program of Chinese Academy of Sciences (Grant KGZD-EW-T06), Innovative Project of SIC, CAS, and One Hundred Talent Project, SIC-CAS (Y36ZB1110G).

Notes and references

1. J. Charnley, J. Bone Jt. Surg., Br. Vol., 1960, 42, 28–30.
2. M. A. R. Freeman, G. W. Bradley and P. A. Revell, J. Bone Jt. Surg., Br. Vol., 1982, 64, 489–493.
3. P. Liu-Snyder and T. J. Webster, Curr. Nanosci., 2008, 4, 111–118.
4. G. Lewis, J. Biomed. Mater. Res., 1997, 38, 155–182.
5. G. Lewis, J. Biomed. Mater. Res., Part B, 2008, 84, 301–319.
6. S. Shinzato, T. Nakamura, T. Kokuba and N. Kitamura, J. Biomed. Mater. Res., 2001, 54, 491–500.
7. H. Abd Samad, M. Jaafar, R. Othman, M. Kawashita and N. H. Abdul Razak, Bio-Med. Mater. Eng., 2011, 21, 247–258.
8. W. F. Mousa, M. Kobayashi, S. Shinzato, M. Kamimura, M. Neo, S. Yoshihara and T. Nakamura, Biomaterials, 2000, 21, 2137–2146.
9. C. Fukuda, K. Goto, M. Imamura, M. Neo and T. Nakamura, Acta Biomater., 2011, 7, 3595–3600.
10. C. Fukuda, K. Goto, M. Imamura and T. Nakamura, J. Biomed. Mater. Res., Part B, 2010, 95, 407–413.
11. M. J. Dalby, L. Di Silvio, E. J. Harper and W. Bonfield, Biomaterials, 2001, 22, 1739–1747.
12. S. Y. Kwon, Y. S. Kim, Y. K. Woo, S. S. Kim and J. B. Park, Bio-Med. Mater. Eng., 1997, 7, 129–140.
13. S. B. Kim, Y. J. Kim, T. R. Yoon, S. A. Park, I. H. Cho, E. J. Kim, I. A. Kim and J. W. Shin, Biomaterials, 2004, 25, 5715–5723.
14. L. L. Hench and I. Thompson, J. R. Soc., Interface, 2010, 7, S379–S391.
15. C. T. Wu, J. Chang, S. Y. Ni and J. Y. Wang, J. Biomed. Mater. Res., Part A, 2006, 76, 73–80.
16. C. T. Wu and J. Chang, J. Inorg. Mater., 2013, 28, 29–39.
17. M. Razavi, M. Fathi, O. Savabi, S. Mohammad Razavi, B. Hashemi Beni, D. Vashaee and L. Tayebi, Ceram. Int., 2014, 40, 3865–3872.
18. H. L. Sun, C. T. Wu, K. R. Dai, J. Chang and T. T. Tang, Biomaterials, 2006, 27, 5651–5657.
19. Y. Huang, X. Jin, X. Zhang, H. Sun, J. Tu, T. Tang, J. Chang and K. Dai, Biomaterials, 2009, 30, 5041–5048.
20. W. Y. Zhai, H. X. Lu, C. T. Wu, L. Chen, X. T. Lin, K. Naoki, G. P. Chen and J. Chang, Acta Biomater., 2013, 9, 8004–8014.
21. Q. H. Liu, L. Cen, S. Yin, L. Chen, G. P. Liu, J. Chang and L. Cui, Biomaterials, 2008, 29, 4792–4799.
22. H. Gu, F. Guo, X. Zhou, L. Gong, Y. Zhang, W. Zhai, L. Chen, L. Cen, S. Yin, J. Chang and L. Cui, Biomaterials, 2011, 32, 7023–7033.
23. L. G. Xia, Z. Y. Zhang, L. Chen, W. J. Zhang, D. L. Zeng, X. L. Zhang, J. Chang and X. Q. Jiang, Eur. Cells Mater., 2011, 22, 68–83.
24. Y. Zhou, C. Wu and Y. Xiao, J. Mater. Chem. B, 2014, 2, 3907.
25. W. Y. Zhai, H. X. Lu, L. Chen, X. T. Lin, Y. Huang, K. R. Dai, K. Naoki, G. P. Chen and J. Chang, Acta Biomater., 2012, 8, 341–349.
26. C. T. Wu and J. Chang, Mater. Lett., 2004, 58, 2415–2417.
27. J. Chang and L. Chen, Invention Patent, CN104307035A, 2015.
28. ISO-5833/1, International Standards Organization, 1987.
29. T. Kokubo and H. Takadama, Biomaterials, 2006, 27, 2907–2915.
30. K. Serbetci, F. Korkusuz and N. Hasirci, Polym. Test., 2004, 23, 145–155.
31. T. Endogan, K. Serbetci and N. Hasirci, J. Appl. Polym. Sci., 2009, 113, 4077–4084.
32. C. F. Jasso-Gastinel, S. G. Enriquez, J. Flores, I. Reyes-Gonzalez and E. M. Mijares, Macromol. Symp., 2009, 283–284, 159–166.
33. T. Miyazaki, C. Ohtsuki, M. Kyomoto, M. Tanihara, A. Mori and K.-I. Kuramoto, J. Biomed. Mater. Res., Part A, 2003, 67, 1417–1423.
34. Z. Hong, P. Zhang, C. He, X. Qiu, A. Liu, L. Chen, X. Chen and X. Jing, Biomaterials, 2005, 26, 6296–6304.
35. A. Liu, Z. Hong, X. Zhuang, X. Chen, Y. Cui, Y. Liu and X. Jing, Acta Biomater., 2008, 4, 1005–1015.
36. K.-D Kuhn, Bone Cements: Up-to-date Comparison of Physical and Chemical Properties of Commercial Materials, Springer-Verlag, Berlin, 2000.
37. L. L. Cunningham, J. Long-Term Eff. Med. Implants, 2005, 15, 609–615.
38. K. Anselme, Biomaterials, 2000, 21, 667–681.
39. A. El-Ghannam, P. Ducheyne and I. M. Shapiro, Biomaterials, 1997, 18, 295–303.
40. N. Olmo, A. I. Martín, A. J. Salinas, J. Turnay, M. Vallet-Regi and M. A. Lizarbe, Biomaterials, 2003, 24, 3383–3393.
41. Y.-F. Chou, W. Huang, J. C. Y. Dunn, T. A. Miller and B. M. Wu, Biomaterials, 2005, 26, 285–295.
42. I. A. Silver, J. Deas and M. Erecińska, Biomaterials, 2001, 22, 175–185.
43. K. K. Kaysinger and W. K. Ramp, J. Cell. Biochem., 1998, 68, 83–89.
44. J. E. Gough, J. R. Jones and L. L. Hench, Biomaterials, 2004, 25, 2039–2046.

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