Impacts of a Nano-Laponite Ceramic on Surface Performance, Apatite Mineralization, Cell Response, and Osseointegration of a Polyimide-Based Biocomposite
Authors Zhang Y, Jiang W, Yuan S, Zhao Q, Liu Z, Yu W
Received 7 August 2020
Accepted for publication 21 October 2020
Published 24 November 2020 Volume 2020:15 Pages 9389—9405
Checked for plagiarism Yes
Review by Single anonymous peer review
Peer reviewer comments 3
Editor who approved publication: Dr Yan Shen
Yiqun Zhang,1,* Weibo Jiang,2,* Sheng Yuan,3 Qinghui Zhao,4 Zhongling Liu,5 Wei Yu1
1Department of Hand Surgery, China–Japan Union Hospital of Jilin University, Changchun 130033, People’s Republic of China; 2Department of Orthopedics, Second Hospital of Jilin University, Changchun 130022, People’s Republic of China; 3Department of Orthopedics, Peoples’ Hospital of Huolinguole City, Tongliao 029200, People’s Republic of China; 4Translational Medical Center for Stem Cell Therapy and Institute for Regenerative Medicine, Shanghai East Hospital, Tongji University School of Medicine, Shanghai 200123, People’s Republic of China; 5Department of Hospital Infection Control, China–Japan Union Hospital of Jilin University, Changchun, 130033, People’s Republic of China
*These authors contributed equally to this work
Correspondence: Wei Yu Email firstname.lastname@example.org
Introduction: Polyimide (PI) exhibits good biocompatibility and high mechanical strength, but biological inertness that does not stimulate bone regeneration, while laponite possesses excellent bioactivity.
Methods: In this study, to improve the bioactivity of PI, nano–laponite ceramic (LC)–PI composites (LPCs) were fabricated by melt processing as implantable materials for bone repair.
Results: The compressive strength, hydrophilicity, and surface roughness of LPCs with 40 w% LC content (LPC40s) were higher than LPC20s, and LPC20s higher than pure PI. In addition, no apatite mineralization occurred on PI, while apatite mineralized on LPCs in simulated body fluid. Compared with LPC20, more apatite deposited on LPC40, indicating good bioactivity. Moreover, the adhesion, proliferation, and alkaline phosphatase activity of rat bone mesenchymal stem cells on LPCs significantly increased with LC content increasing in vitro. Furthermore, the evaluations of animal experiments (micro-CT, histology, and pushout load) revealed that compared with LPC20 and PI, LPC40 significantly enhanced osteogenesis and osseointegration in vivo.
Discussion: Incorporation of LC into PI obviously improved not only surface physicochemical properties but also biological properties of LPCs. LPC40 with high LC content displayed good biocompatibility and bioactivity, which markedly promoted osteogenesis and osseointegration. Therefore, with its superior biocompatibility and bioactivity, LPC40 could be an alternative candidate as an implant for orthopedic applications.
Keywords: laponite, polyimide, cytocompatibility, bone repair
Over the past few decades, metallic materials (eg, stainless steel, titanium, and cobalt–chromium, as well as their alloys) have been used extensively as implants for load-bearing bone repair, due to their high mechanical strength and biocompatibility.1 However, the elastic modulus of metallic materials is much higher than that of cortical bone in humans.1,2 After implantation in vivo, metallic implants with high elastic moduli produce a stress-shielding effect that causes bone resorption and bone atrophy, thereby leading to loosening/failure of the implants.3 In addition, corrosion of metals in a biological environment slowly releases corrosive products that might be toxic to the human body.4 In the past few decades, nondegradable polymers (NDPs), such as polyetheretherketone (PEEK), polyethylene (PE), polyamide (PA), polytetrafluoroethylene, polypropylene, and polyurethane, have been also widely applied for soft/hard-tissues repair, due to their atoxicity and biosafety in vivo.5–7 Moreover, NDPs have attracted much attention for load-bearing bone repair, because of superior biocompatibility and mechanical properties and easy processing, as well as corrosion resistance. Importantly, elastic moduli of NDPs are close to that of human cortical bone, which decreases the stress-shielding effect compared with metallic implants. Thereby, NDPs are considered preferred biomaterials to replace metallic materials for orthopedic applications.6,7
As an NDP, polyimide (PI) is a high-performance plastic with outstanding mechanical properties, chemical resistance, radiation resistance, and flexibility.7,8 PI is applied extensively in various fields, including aerospace, automobile, microelectronics, and sensors, as well as membranes.7,8 Nontoxic PI has also been proposed as a candidate for development of biomedical devices, such as nerve implants and energy-harvesting devices of microelectromechanical systems in vivo.9,10 PI exhibits good biocompatibility and little hemolysis in vitro and in vivo.11 Moreover, PI–carbon nanotube composites have improved performance over ultrahigh molecular–weight PE which might be a good candidate as cartilage materials to replace it.12 Therefore, PI might be a promising biomedical material for load-bearing bone repair in orthopedic applications, due to its outstanding biocompatibility, mechanical properties, corrosion resistance, and elastic moduli, which are close to human cortical bone.9,10 PI is a bioinert material that is indispensable in stimulating responses of osteoblasts and ultimately new bone (NB) regeneration.12 Therefore, enhancing bioactivity and accelerating NB formation and integration with host bone are the challenges for PI application in orthopedics.
Generally, integration with host bone tissue (osseointegration) for nondegradable biomaterials is considered a criterion for successful bone repair.13 After implantation in vivo, poor bioactivity of the biomaterial can cause inferior bone regeneration and osseointegration, which might reduce the initial fixation and long-term stability of the implant.14 Bioactive materials (eg, bioactive bioglass/ceramic) possess the ability to promote NB formation and osseointegration with host bone tissue.14,15 However, these inorganic bioactive materials exhibit low toughness because of their brittleness, making it difficult to repair load-bearing bone defects. Synthetic NDPs (eg, PA, PE, and PEEK) possess outstanding mechanical properties and machinability, as well as easy processing.5–7 Therefore, in the past few decades, inorganic bioactive materials–NDP composites (eg, hydroxyapatite–PA, calcium silicate–PEEK, and bioglass–PE) have been investigated for their great potential for load-bearing bone repair.16 Laponite (Lap), a synthetic layered silicate and bioactive material, has been developed as a new biomaterial for treatment of diseases and regenerative medicine, due to its biocompatibility.17,18 Lap not only stimulates the proliferation of human mesenchymal stem cells but also improves osteogenic differentiation, bone-related matrix–protein deposition, and matrix mineralization of MC3T3-E1 cells.19,20 Moreover, functionalized with lysine, arginine, and leucine, Lap improves the proliferation of human skin fibroblasts, indicating that it can be used as a wound dressing.21 Furthermore, Lap stimulates growth of mesenchymal stem cells in rats and treats bone defects in mice.22
In this study, Lap ceramic (LC) was prepared and LC–PI composites (LPCs) fabricated by melt blending as implantable biomaterials. The purpose of this study was to develop nondegradable PI-based composites with bioactivity for bone repair. We assumed that incorporation of LC with PI would create new composites with better biological performance than PI alone. To confirm this hypothesis, surface properties (roughness and hydrophilicity) and in vitro bioactivity of LPCs (apatite mineralization in simulated body fluid [SBF]) were evaluated. Rat bone-marrow stromal cell (BMSC) responses to LPCs were investigated in vitro. Osteogenesis and osseointegration of LPCs were evaluated in vivo by animal experiments.
Preparation and Characterization of LC and LPCs
Nano-LC was prepared by sintering Lap (Altana, Germany) powders at 800°C for 6 hours in a roasting furnace (P300, Nabertherm, Germany). After that, LC was ground into powder using a planetary ball mill (QM-3SP2, Nanjing Nanda, China). LPCs — 20 w% LC (LPC20; weight/weight) and LPC40 — were prepared by melt processing. PI (Junhua, China) powders were added to absolute ethyl alcohol and stirred for 1 hour, and then LC powders were added into the PI dispersion and stirred continuously for 2 hours. The supernatant was decanted after the dispersion had stood for 4 hours and then dried in an oven (at 50°C) to obtain mixed powders. The mixtures were evenly mixed by the planetary ball mill to get the two kinds of powders (LPC20 and LPC40), which were compressed with a pressing machine (YP-40T, Jinfulun, China) into stainless-steel molds (Φ12×10 mm and Φ5×5 mm). The specimens obtained were sintered at 270°C for 4 hours to obtain LPC20 and LPC40. PI samples were also prepared by the same process as the controls.
Lap morphology was observed with transmission electron microscopy (JEM-2100F), and particle-size distribution of Lap and LC were measured by laser light–scattering spectrometry (CGS-5022F). LC morphology and composition were characterized using atomic force microscopy (DiMultiMode, Veeco, USA) and X-ray photoelectron spectroscopy (K-Alpha, Thermo Fisher Scientific, USA). Lap, LC, PI, LPC20, and LPC40 were characterized by X-ray diffraction (GeigerFlex, Rigaku, Japan) and Fourier-transform infrared spectroscopy (Magna-IR 550, Nicolet, USA), and LC, PI, LPC20, and LPC40 were analyzed by scanning electron microscopy (SEM; S-3400N, Hitachi, Japan) and energy-dispersive spectroscopy (EDS; S-4800, Hitachi).
Compressive Strength, Surface Roughness, and Hydrophilicity
Compressive strength and compressive moduli of (Φ12×10 mm) of PI, LPC20, and LPC40 were tested at 2 mm/min using a mechanical testing machine (CMT4000, MTS, China) with a 9 N load cell at room temperature, which was stopped when the compression ratio had reached 50% or chipping. Surface morphology and roughness of the samples (Φ12×2 mm) was determined using laser confocal microscopy 3-D microscopy (VK-X110, Keyence, Japan). Hydrophilicity of the specimens (Φ12×2 mm) were determined by water-contact angles (XG-CAMB3, Xuanyichuangxi Industrial Equipment, China).
Apatite Mineralization in SBF
Apatite mineralization of specimens (Φ12×2 mm) of PI, LPC20, and LPC40 was carried out in SBF (pH 7.4) at 37°C.15 Specimens were immersed in SBF (solid:liquid ratio 1 cm2:20 mL). After 7 days’ soaking, specimens were taken out, rinsed with deionized water, and dried for 24 hours at 60°C. Composition and surface morphology were determined by SEM and EDS. Concentrations of Li, Mg, Si, Ca, and P ions in SBF were detected with inductively coupled plasma atomic emission spectroscopy (IRIS 1000, Thermo Elemental, USA) after specimen soaking for different periods. pH values were measured after soaking.
Rat BMSCs were extracted from 4-week-old male Sprague Dawley rat limbs (Jiesijie, China). Isolation methods were based on previous research.23 Briefly, bilateral femora and tibiae of the rats were harvested and ravaged with αMEM supplemented with 10% FBS, HyClone, 100 U/mL penicillin, and 100 mg/L streptomycin (HyClone) under sterile conditions. After 3 days of culture, nonadherent cells were washed away with PBS, and the remaining adherent cells mostly comprised primary BMSCs (P0). The medium was refreshed every 3 days, and BMSCs were passaged when 80%–90% confluence had been reached. The third passages were used for the following cell experiments.
BMSCs were cultured in αMEM supplemented with FBS (10%) containing 1% penicillin (100 U/mL) and streptomycin sulfate (100 μg/mL), and cell culture was carried out in an atmosphere of 5% CO2 at 37°C. The culture medium (Sigma-Aldrich) was exchanged every 3 days and normal passages performed, and third-generation rat BMSCs were used for cell experiments. Cells were detached to form a cell suspension using 0.25% trypsin and cell suspensions counted before being cocultured with samples, with cell density 2×104 cells per well.Specimens of PI, LPC20, and LPC40 (Φ12×2 mm) were sterilized at 50°C for 4 hours with ethylene oxide in a sterilizer (SQ-H120, Sanqiang Medical, China).
BMSCs were inoculated on the samples (PI, LPC20, and LPC40) in a 24-well plate and specimens taken out and treated with PBS to remove unattached cells at 1, 3 and 7 days after culturing. Cells adhering to sample surfaces were fixed using 2.5% glutaraldehyde for 40 minutes. Then, the glutaraldehyde solution was removed and specimens rinsed thrice with PBS, followed by sequential dehydration in graded ethanol (10%, 30%, 50%, 70%, 85%, 90%, and 100%) for 10 minutes. Specimens were placed in the air and dried for 24 hours, and BMSCs on the samples were observed with SEM. Samples were taken out and treated with PBS after culturing for 7 days. Cells adhering to sample surfaces were fixed in 2.5% glutaraldehyde for 40 minutes. Then, cell-seeded specimens were stained with 4ʹ,6-diamidino-2-phenylindole (Sigma-Aldrich) for 40 minutes. After being rinsed twice with PBS, cell-seeded specimens were stained with fluorescein isothiocyanate (Sigma-Aldrich) for 8 minutes. Cell morphology was observed utilizing confocal laser scanning microscopy (Hitachi).
For cell-proliferation assays, BMSC cultured on PI, LPC20, and LPC40 were tested with CCK8). At 1, 3, and 7 days after culturing, specimens were gently rinsed thrice with PBS to remove unattached cells. Then, cells on the specimens of each well were incubated with cell medium (400 μL) containing CCK8 (40 μL) for 6 hours. Optical density (OD) values of solution were tested at 405 nm utilizing a microplate reader (SpectraMax 384, Molecular Devices, USA)
Alkaline Phosphatase Activity
PI, LPC20, and LPC40 were placed in a 24-well cell-culture plate. BMSCs were seeded on sample surfaces. Cell-culture medium was replaced once a day. At 7, 10, and 14 days after culturing, the culture fluid was taken out, the specimens washed with PBS, and 500 μL of a 1% concentration of ethyl phenyl PE glycol solution (NP40) added. Then, 50 μL p-nitrophenyl phosphate solution (1 mg/mL) was added to each well for 16 minutes at 37°C. After that, 100 μL 0.1 M sodium hydroxide was added to terminate the color reaction. OD values were tested using the SpectraMax at 405 nm. BSA was applied as the standard protein. Total protein content in the lysate was determined using the BCA kit. ALP activity was expressed as OD per total protein content.
Osteogenesis and Osseointegration In Vivo
In vivo implantation was performed on eight male beagle dogs (weight 11–14 kg, age 18 months), which was approved by the institutional animal care and use committee of National Tissue Engineering Research Center (Shanghai, China) according to National Institutes of Health guidelines (publication 85–23 revised 1985) for the care and use of laboratory animals. The dogs were anesthetized by applying an intravenous injection of pentobarbital (1%, 80 mg/kg). Then midshafts of right femora were exposed by skin incisions (6 cm in length). Bone holes (Φ5×5 mm) spaced about 10 mm apart in each femur were made with a medical drill. Then, sterilized implants of PI, LPC20, and LPC40 (Φ5×5 mm) were implanted into the holes, and the overlying soft tissue and skin was closed. At 4 and 12 weeks after implantation, the dogs were euthanized and right femora obtained and fixed using glutaraldehyde solution (2.5%).
Micro-CT scanning (SkyScan 1076) was conducted to determine NB tissue surrounding the implants (40 kV voltage, 250 μA current, and 35 mm resolution). The scanning area (Φ5×5 mm) containing the implants was selected as the region of interest for micro-CT evaluation. NRecon and CTVol (Skyscan Company, Belgium) were utilized to construct 3-D images of the samples, and the threshold was optimized to isolate bone tissue and implants. Bone-mineral density, bone volume/total volume, trabecular number, and trabecular thickness were determined using CTAn (Skyscan).
Evaluation of Histology and Pushout Loads
Samples containing bone tissue and implants were dehydrated using ethanol (from 75% to 100%) and fixed using PMMA. Thy were embedded and cut into sections (150 μm) using a Leica SP1600 saw microtome and ground to thinner sections (50 μm), which were then stained by Van Gieson’s picrofuchsin. Inverted microscopy was used to observe these stained sections, and ImagePro Plus 6.0 (Media Cybernetics, Silver Spring, USA) was utilized to determine the percentage of bone-implant contact (BIC). A universal material-testing system (Instron) from High Wycombe (UK) was used to test the pushout loads of implants with bone tissue at a load rate of 5 mm/min. A load–displacement curve was recorded during pushing, and peak of the curve defined as failure load.
Quantitative data obtained from the experiments performed in triplicate are expressed as means ± SD. Data were analyzed with Origin 9.0. ANOVA was used to detect statistical differences, and p<0.05 was regarded as significant.
Characterization of Lap and LC
Transmission electron microscopy of Lap is shown in Figure 1A. Lap particles sized around 30 nm exhibited spherical morphology. SEM of LC is shown in Figure 1B. The LC particles sized around 300 nm revealed irregular morphology. Particle-size distribution of Lap and LC is shown in Figure 1C and D, respectively. Particle-size distribution of Lap was20–40 nm, while LC was 200–500 nm. Atomic force microscopy of LC is shown in Figure 1E. LC-particle size was 500–1,000 nm. X-ray photoelectron spectra of LC are shown in Figure 1F. LC contained Li, Si, and Mg elements.
Figure 1 TEM (A) of Lap, SEM (B) of LC, particle-size distribution of Lap (C) and LC (D), and 3-D AFM (E) and XPS spectra (F) of LC.
Characterization of LPC
X-ray diffraction patterns of Lap, LC, PI, LPC20, and LPC40 are shown in Figure 2A. Lap peaks appeared at 2θ= 20.0°, 28.0°, 35.0°, and 60.6° while LC peaks were found at 2θ=20.0°, 22.0°, 28.0°, 35.0°, 47.0°, and 62.0°.20 The standard powder-diffractioncard number of Lap was 25–1385. A broad peak at around 21.3° was attributed to PI, exhibiting the amorphous phase. Xharacteristic peaks of LC were found in both LPC20 and LPC40. Fourier-transform infrared spectra of Lap, LC, PI, LPC20, and LPC40 are shown in Figure 2B. For Lap, the peak at 3,415 cm−1 represented hydroxyl (O–H) stretching vibration, and those at 1,629 cm−1 and 1,010 cm−1 represented silicon–oxygen bonds (Si–O–Si) and asymmetric stretching of the silicate layer. Compared to Lap, hydroxyl peaks in LC decreased after sintering at high temperature. For PI, the peak at 1,775 cm−1 was an asymmetrical vibration of C=O, 1,720 cm−1 was ascribed to C=O symmetrical stretching vibration, and 1,373 cm−1 was assigned to C–N–C stretching vibration in the imide bond. In addition, peaks at 1,610 cm−1 and 1,508 cm−1 represented absorption peaks of a benzene ring. Characteristic peaks of both LC and PI were detected in both LPC20 and LPC40, while no new peak was observed. Stress–strain curves of PI, LPC20, and LPC40 are shown in Figure 2C. Stress of the samples increased with increased strain for PI, LPC20, and LPC40.Compressive strengths of PI, LPC20, and LPC40 are shown in Figure 2D, and were 80.28±2.1 MPa, 98.48±3.4 MPa, and 136.76±4.6 MPa, respectively. Compressive moduli of PI, LPC20 and LPC40 were 2.9±0.2 GPa, 3.7±0.1 GPa, and 5.6±0.3 GPa, respectively.
Surface Morphology, Composition, and Hydrophilicity
Photos and SEM images of surface morphology of PI, LPC20, and LPC40 are shown in Figure 3. Compared to PI (Figure 3A and D), with its smooth surface, both LPC20 and LPC40 exhibited rough surfaces (Figure 3B and C). At higher magnification, (Figure 3E and F), many LC particles were found on the surfaces of LPC20 and LPC40. EDS spectra of PI, LPC20, and LPC40 are shown in Figure 4A–C. LPCs showed peaks for Na, Mg, and Si, which increased with increased LC content. EDS mapping is shown in Figure 4D–I. C, Mg, and Si were detected in both LPC20 (Figure 4D–F) and LPC40 (Figure 3G–I), confirming that LC particles were dispersed in the PI matrix. Laser confocal microscopy 3-D images of PI, LPC20, and LPC40 are shown in Figure 5. PI (Figure 5A) showed a smooth surface, while both LPC20 (Figure 5B) and LPC40 (Figure 5C) displayed rough surfaces. Surface roughness (Figure 5D) of PI was 1.05±0.03μm, while LPC20 and LPC40 were 1.21±0.12μm and 1.55±0.27μm, respectively. Hydrophilicity of PI, LPC20, and LPC40 was determined by testing water-contact angles shown in Figure 5E, which were were 78.0°, 65.5°, and 50.5°, respectively.
Figure 3 Photos (top) and SEM of PI (A and D), LPC20 (B and E), and LPC40 (C and F).
Figure 4 EDS spectra of PI (A), LPC20 (B), and LPC40 (C), EDS mapping of LPC20 (D–, E Mg, F Si) and LPC40 (C, G, H Mg, I Si).
Figure 5 Laser confocal microscopy 3-D images of PI (A), LPC20 (B), and LPC40 (C) and surface roughness (D) and water-contact angles (E) of samples. *p<0.05 for LPC40 and LPC20 vs PI.
Apatite Mineralization in SBF
Figure 6A–C shows SEM of surface morphology of PI, LPC20, and LPC40 after immersion in SBF for 7 days. PI exhibited a smooth surface without precipitates, while a large number of microsphere precipitates sized about 2 μm were found on both LPC20 and LPC40 surfaces. Microsphere precipitates on LPC40 surface were greater in number than LPC20. Figure 6D–F shows the EDS of PI, LPC20, and LPC40 after soaking in SBF for 7 days. Peaks of Ca and P were detected. Ca:P ratios for LPC20 and LPC40 were 1.55 and 1.60, respectively (approach to Ca:P ratio of hydroxyapatite of 1.67), revealing the formation of apatite on both LPC20 and LPC40.
Figure 6G and H shows changes in Ca, Mg, P, Li and Si ion concentrations with time after LPC20 and LPC40 soaking in SBF for different periods. Concentrations of Mg, Si, and Li ions kept increasing, while Ca and P ions decreased during the mineralization process. Figure 6I shows changes in pH values with time after PI, LPC20, and LPC40 soaking for different periods. During immersion, pH values for PI showed no change, while those for LPC20 (7.79–7.89) and LPC40 (7.88–8.02) increased gradually with time (1–7 days).
Figure 7 shows SEM of morphology of BMSCs cultured on PI, LPC20, and LPC40 for different periods. Cells on LPC20 and LPC40 surfaces had adhered better than PI at 1, 3, and 7 days. More cells were found to adhere and spread better on LPC40 than LPC20 and LPC20 than PI at 3 and 7 days. Confocal laser–scanning microscopy images of cytoskeleton (BMSCs) on samples stained with FITC and DAPI are shown in Figure 8. Cytoplasm of the cells was stained green by FITC, while cell nucleiwere labeled blue by DAPI. Adhesion and spread of the cells on LPC40 were more than LPC20, and LPC20 more than PI.
Figure 7 SEM of BMSC morphology with PI (A–C), LPC20 (D–F), and LPC40 (G–I) at 1 (A, D, G), 3 (B, E, H), and 7 days (C, F, I) after culturing (yellow arrows represent the cells).
Figure 8 CLSM of BMSC morphology on PI (A), LPC20 (B), and LPC40 (C) at 7 days after culturing.
Cell Proliferation and ALP Activity
OD values (revealing cell proliferation) of BMSCs on PI, LPC20, and LPC40 at 1, 3, and 7 days after culturing are shown in Figure 9A. Values of cells on samples increased with time. In addition, OD values of cells on LPC40 were the highest while values for PI lowest. ALP activity (indicating cell osteogenic differentiation) of cells on PI, LPC20, and LPC40 at 7, 10, and 14 days after culturing is shown in Figure 9B. ALP activity of cells on samples increased with time. ALP activity of cells on LPC40was highest and PI lowest.
Figure 9 OD values (A) and ALP activity (B) of BMSCs in samples after culturing for different periods . *p<0.05 for LPC40 and LPC20 vs PI).
Osteogenesis and Osseointegration in vivo
Figure 10A shows reconstructed 3-D images of NB tissue around the samples (PI, LPC20, and LPC40) at different time points after implantation. At 4 and 12 weeks, NB for LPC40 was highest and NB for PI . In addition, NB around LPC40 and LPC20 increased with implanted time, while no significant change was detected for PI during the entire implantation period.
Figure 10BE shows quantitative analysis of bone-mineral density (BMD), bone volume/total volume (BV/TV), trabecular number (Tb.N), and trabecular thickness (Tb.Th) of the implanted samples. BMD, BV/TV, Tb.N and Tb.Th for LPC40 were highest, while PI exhibited the lowest at 4 and 12 weeks. BMD, BV/TV, Tb.N and Tb.Th for LPC40 and LPC20 increased with time, while there no significant change for PI during the entire implantation period.
Histological and Pushout-Load Evaluation
Figure 11A shows images of histological sections. At interfaces between NB and implants, NB on the surface of both LPC40 and LPC20 were more than PI at 4 and 12 weeks. Compared with LPC20 and PI, LPC40 obtained the best contact with NB. Almost no gaps between NB and implants were found for LPC40, while obvious gaps had appeared for PI at 12 weeks. Figure 11B shows the percentage of BIC. BIC for LPC40 and LPC20 increased with time, with no obvious change for PI during the entire implantation period. At 12 weeks, BIC for LPC40 (88.6%) was highest compared with LPC20 (53.2%) and PI (12.7%). Figure 11C shows the pushout loads of the implants. Loads for LPC40 (4W 286.3 N and 12W 427.6 N) were obviously higher than LPC20 (4W 193.4 N and 12W 301.5 N) and PI (4W 69.2 N and 12W 87.1 N) at 4 and 12 weeks.
Generally, implantable materials with biocompatibility and osteogenic activity can promote bone regeneration and be integrated with host bone (osseointegration) that obtains initial fixation and long-term stability, confirming successful bone repair.24 In this study, nano–LC was prepared and incorporated into PI to obtain bioactive LPCs. The results showed that LC particles were distributed in the PI matrix for both LPC20 and LPC40, which exhibited rough surfaces compared to PI, which was smooth. Surface roughness of the biomaterial not only influences cell behavior and function (eg, attachment, proliferation, and osteogenic differentiation) but also increases bone-tissue anchorage and growth, and thus enhances osseointegration of the implant.25 In this study, the surface roughness of LPC40 (1.55 μm) was higher than LPC20 (1.21 μm) and PI (1.05 μm). Therefore, improvements of surface roughness of the composites were attributed to the increase in LC content.
Decreased water-contact angles of the biomaterial surface are an indicator ofimprovement in hydrophilicity.26 In this study, water-contact angles of the composites gradually decreased with increased LC content, indicating improvement in hydrophilicity. The hydrophilicity of LPC40 was highest, while PI was lowest. Clearly, the increased hydrophilicity of LPC40 was due to increased content of hydrophilic LC. In the design of bone implants for maxillofacial and cranial reconstruction, as well as in interbody cages for spinal fusion, the structure needs to sustain interlumbar pressure.27 Therefore, implantable materials must have enough mechanical strength for implantation without plastic fragmentation or deformation in vivo. In the present study, the compressive strength of LPCs gradually increased with increased LC content, indicating that the LC played a key role in improvement in mechanical strength of the composites. In addition, the compressive strength of LPC40 was 137 MPa, which was close to cortical bone of the human body (55–190 MPa).28 Furthermore, the compressive modulus of LPC40 was 5.6 GPa, close to that of human bone (3–20 GPa).29
Bioactive materials, such as bioglasses/ceramics, have the ability to form mineral apatite layers on their surfaces both in vitro and in vivo, which can be integrated with bone tissue by the apatite layer.29,30 In this study, apatite formed on both LPC20 and LPC40 after immersion in SBFm while PI did not. Moreover, the formation of apatite on LPC40 was more than that for LPC20. Clearly, the improvement in apatite formation on LPCs was due to the increase in LC content. After LPC immersion into SBF, Mg, Si, and Li ion concentrations gradually increased with time, indicating that these ions continuously released from LPCs, due to the slight dissolution of LC in the composites. However, Ca and P ion concentrations obviously declined because ofapatite formation, which consumed the Ca and P ions in SBF. The apatite-formation mechanism on the composites can be concluded as follows. Due to the presence of saturated Ca ions in SBF and rich Si–OH (caused by LC) on the composite surface, Si–OH groups were negatively charged, which attracted Ca ions through electrostatic interactions, and then P ions were attracted by Ca ions, which caused the formation of apatite.31 In addition, after the samples had soaked in SBF, pH values in the solution for PI showed no change. However, pH values in solutions for both LPC20 and LPC40 gradually increased, which led to a weak alkaline microenvironment due to the presence of LC that released Li, Mg, and Si ions during the soaking period. Previous studies have shown that a weak alkaline environment and bioactive ions (eg, Li, Mg, and Si) are very useful for cell proliferation and differentiation.32
Cellular responses (including adhesion, proliferation, and osteogenic differentiation) to biomaterials are strongly influenced by not only surface physical properties (eg, composition, morphology, roughness, and hydrophilicity) but also chemical properties (eg, ion release and acid–base microenvironment), which are also closely related to cytocompatibility.33 In the present study, more cells adhered and spread better on the surfaces of LPC40 and LPC20 than PI, demonstrating that the composites promoted cell adhesion and spreading. Compared with LPC20 and PI, the obvious increase in cell adhesion and spreading for LPC40 was ascribed to the presence of higher LC content. Bioactive implantable materials can promote not only cell adhesion but also cell proliferation and osteogenic differentiation, which could be closely correlated with NB regeneration.27 In this study, LPC40 with high LC content significantly promoted the proliferation of BMSCs compared with LPC20 and PI. Therefore, the remarkable improvement in cell proliferation was ascribed to the increase of LC content in LPCs. In addition, ALP activity is usually regarded as an early marker for osteogenic differentiation, and the increase in ALP activity of cells on biomaterials can generally be considered an indicator to promote osteogenic differentiation.34 In the present study, LPC40 remarkably improved ALP activity in cells compared with LPC20 and PI. Therefore, the significant promotion of osteogenic differentiation of BMSCs was due to high LC content, which improved the surface properties of LPC40.
The in vivo results from micro-CT images and quantitative analysis showed that the amount of NB around LPC40 was higher than LPC20 and LPC20 was higher than PI. Therefore, compared with PI and LPC20, LPC40 with high LC content obviously promoted osteogenesis in vivo. Osseointegration is considered the bonding of bone tissue to implant that can achieve early loading and long-term stability of the implant.35 A significant increase in pullout loads in implant detachment from bone can be used to indicate improvement in osseointegration.36 In the present study, the pushout loads of both LPC40 and LPC20 significantly increased from 4 to 12 weeks, while no significant change for PI was found. In addition, LPC40 possessed the highest pushout loads (indicating superior osseointegration), while PI exhibited the lowest (indicating poor osseointegration). Therefore, compared with PI and LPC20, LPC40 with high LC content obviously improved NB formation and osseointegration.
Histological evaluation demonstrated that at the interface between NB and implants, more NB tissue was in contact with the surfaces of both LPC40 and LPC20 than PI at both 4 and 12 weeks. No obvious gap at the interface between NB and implants was found for LPC40, while obvious gaps appeared for PI at 12 weeks. Furthermore, at 12 weeks, compared with LPC20 and PI, LPC40 achieved the best contact with NB, displaying excellent osseointegration. BIC, defined as the percentage of NB tissue covering the implant surface, is one of the critical measurements applied to quantify the degree of osseointegration.37 In the present study, the BIC of both LPC40 and LPC20 increased with time, while no obvious change for PI during entire implantation period. At 12 weeks, the BIC of PI was lowest and LPC40 highest, indicating the best osseointegration. It has been reported that bioactive materials not only stimulate NB formation but also improve osseointegration.38
Surface characteristics (eg, chemical composition, roughness, hydrophilicity, and ion release) of biomaterials greatly influence osteoblast responses (adhesion, proliferation, and osteogenic differentiation) and NB formation around the implants, both in vitro and in vivo.39,40 In the present study, compared with PI and LPC20, LPC40 containing higher LC content with improved surface properties (eg, roughness and hydrophilicity) not only stimulated osteoblast responses in vitro but also promoted NB formation in vivo. Previous studies have revealed that dissolution products containing Si, Mg, and Li ions from bioactive ceramics/glasses stimulated osteoblast proliferation and osteogenic differentiation.41,42 In the present study, compared with PI and LPC20, LPC40 containing higher LC content released more Li, Mg, and Si ions, which was critical to stimulating cellular responses and promoting NB regeneration in vivo. Moreover, formation of apatite exists between the bioactive implant surface and NB in vivo, which closely resembles bone mineral in composition, as well as a structure that serves as a foundation for NB formation and osseointegration.29,30 In this study, compared with PI and LPC20, LPC40 containing higher LC content induced more mineralized apatite formation, which stimulated osteoblast responses and promoted osteogenesis and osseointegration. Therefore, compared with PI and LPC20, improvements osteoblast responses in vitro and osteogenesis and osseointegration in vivo for LPC40 contributed to the presence of higher LC content. In summary, the positive responses of BMSC/bone tissue to LPC40 were ascribed to the presence of higher LC content, which caused improvements in surface properties (roughness and hydrophilicity), ion (Li, Mg and Si) release, and apatite formation, as well as the formation of weak alkaline microenvironment, which might have been synergistic effects on final results. As a result, incorporation of LC into PI significantly enhanced surface physicochemical properties, thereby improving the biological performance of LPCs. LPC40 would be an alternative candidate for bone repair.
Bioactive composites containing LC and PI (LPC) were prepared by melt processing for potential orthopedic applications. Compared with PI, the addition of LC into PI obviously improved compressive strength, surface hydrophilicity, and roughness of LPCs, which increased with LC content increasing. Moreover, in comparison with LPC20 and PI, LPC40 with high LC content significantly improved apatite mineralization, displaying good bioactivity. In addition, compared with LPC20 and PI, LPC40 with high LC content significantly stimulated responses (attachment, proliferation, and ALP activity) of BMSCs in vitro. Furthermore, in comparison with LPC20 and PI, LPC40 with high LC content obviously promoted osteogenesis and osseointegration in vivo. In short, incorporation of LC into PI significantly enhanced surface physicochemical properties, thereby improving the biological performance of LPCs. Consequently, LPC40 with better biocompatibility and bioactivity might have great potential as bone implants for bone repair.
This work was supported by Jilin Province Science and Technology Development Project (20190201244JC).
The authors report no conflicts of interest in this work.
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