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Development of bioactive thermosensitive polymer–ceramic composite as bone substitute
Chemical Engineering Science 89 (2013) 133–141
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Development of bioactive thermosensitive polymer–ceramic composite as bone substitute Po-Liang Lai a,b,1, Carl Tsai-Yu Lin b,c,1, Ding-Wei Hong b, Shu-Rui Yang b, Yu-Han Chang a, Lih-Huei Chen a, Wen-Jer Chen a, I-Ming Chu b,n a
Department of Orthopedic Surgery, Chang Gung Memorial Hospital, College of Medicine Chang Gung University, Taoyuan, Taiwan Department of Chemical Engineering, National Tsing Hua University, Hsinchu, Taiwan c Biomedical Technology and Device Research Labs, Industrial Technology Research Institute, Hsinchu, Taiwan b
H I G H L I G H T S c c c c c
Thermosensitive composite gelcomposed of mPEG–PLGA and HAP/b-TCPwas made. The gel is capable of sol-to-gel transition between 4 1C to 70 1C and of in-situ gelling. The composite gels do not acidify the surrounding environment. Higher hydroxyapatite contents help raise bone union rate. Bone union is examined through radiographic method.
a r t i c l e i n f o
abstract
Article history: Received 15 July 2012 Received in revised form 13 September 2012 Accepted 16 November 2012 Available online 2 December 2012
The purpose of this study was to design a thermosensitive composite gel to be used as a bone graft substitute. This gel can provide a more suitable microenvironment by using the amphiphilic triblock copolymer (mPEG550PLGA1405) consisting of methoxy poly(ethylene glycol) (mPEG), poly(lactic-coglycolic acid) (PLGA). An aqueous dispersion of mPEG550PLGA1405 mixed with different ratios of HAP/bTCP (composite gel) underwent a sol–gel–sol transition as the temperature was increased from 4 to 70 1C. The particle size and critical micellization concentration (CMC) were increased by adding ceramics. During the in vitro degradation process, composite gels demonstrated a slight decrease in pH value, a slower degradation rate, less toxicity, and a higher cell survival rate. The biocompatibility of the composite gels was validated by hemolysis test. In vivo animal studies demonstrated both radiographic and gross bone union when the ratio of HAP/b-TCP was 7:3. Based on the results, we have developed novel thermosensitive composite gels as bone substitutes. & 2012 Elsevier Ltd. All rights reserved.
Keywords: Biomedical engineering Bone substitutes Composites Gels Polymers Thermosensitive
1. Introduction Hydrogels are now widely used as biomaterials. In recent years, in situ forming hydrogels have been used for various biomedical applications (Hatefi and Amsden, 2002). Thermosensitive hydrogel (Zhai et al., 2009), whose aqueous solution is a sol at room temperature or lower, and which forms a gel at body temperature (37 1C), can avoid the negative effects of high
n
Corresponding author. Tel.: þ886 3 5713 704; fax: þ886 3 5715 408. E-mail addresses: [email protected] (P.-L. Lai), [email protected] (C. Tsai-Yu Lin), [email protected] (D.-W. Hong), [email protected] (S.-R. Yang), [email protected] (Y.-H. Chang), [email protected] (L.-H. Chen), [email protected] (W.-J. Chen), [email protected] (I.-M. Chu). 1 contributed equally to this work. 0009-2509/$ - see front matter & 2012 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.ces.2012.11.019
temperature and organic solvent on drugs and cells (Hishikawa et al., 2004). Commercially available thermogelling polymers, poloxamers (Roques et al., 2007), poly N-isopropylacrylamide (PNIPAM) (Xu et al., 2008), and poly N-isopropylacrylamide– acrylic acid (PNIPAM-AAc) (Wang et al., 2011a) have been widely investigated for their potential biomedical applications. Such thermosensitive polymers have been studied extensively for separation, modulation of biofunction, catalysis, drug delivery, cellular transport (Perumal et al., 2008) and cell culture/tissue engineering applications (Wang et al., 1999). However, the use of these systems is limited because they are not biodegradable in vivo. A bioresorbable material demonstrates stability against biodegradation (Turner et al., 2005) and further resorption in vivo (Eglin et al., 2009). Some polyethylene glycol (PEG)-based compounds (Than et al., 2008) can form thermosensitive biodegradable
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hydrogels which have hydrophilic and hydrophobic monomer units. They are often used to build hydrophilic blocks because they are biocompatible, highly soluble, and hydrate in water (Zhang et al., 2006). The most prominent hydrophobic examples are poly(lactic acid) (Legrand et al., 2007) and poly(lacticco-glycolic acid) (PLGA) that are used as biocompatible nanocontainer materials in drug delivery devices (Muthu et al., 2009). The Food and Drug Administration (FDA) has approved several products comprised of PLA, PGA and their copolymer PLGA (Sander et al., 2004). However, their utility in tissue engineering applications is hampered due to the accumulation of acidic degradation products from bulk erosion of the polymers that can adversely affect biocompatibility and lead to the catastrophic loss in mechanical integrity (Taylor et al., 1994). Inorganic bioceramics such as hydroxyapatite (HAP) (Aboudzadeh et al., 2010), tricalcium phosphate (TCP), and biphasic calcium phosphates (BCP) are commonly used as fillers, which can confer osteoconductivity and strength to bone tissue engineering grafts (Moursi et al., 2002). Generally, TCP degrades too rapidly and HAP too slowly (Detsch et al., 2008). An additional problem encountered with HAP is that it contains 4–5 wt% of carbonate ions as well as magnesium, strontium and fluoride and, thus, is not an exact mimic of bone mineral (Boanini et al., 2010). This implies that HAP cannot be remodelled as efficiently as natural bone mineral (Zhang et al., 2009). Researchers investigating polymer–HAP mixtures have experienced difficulty with dispersion and re-agglomeration of HAP particles in the polymer matrix at high loads (Kothapalli et al., 2008). Other polymer/ceramic composites, such as peptide/ hydroxyapatite nanocomposites (Detsch et al., 2010), and PCL/ tricalcium phosphate composite (Arafat et al., 2011), showed more promise. Therefore, we mixed different weight ratios of HAP/b-TCP with mPEG–PLGA hydrogel, because HAP/b-TCP dissociation is known to release anions including PO34 and OH which can neutralize the acid produced by mPEG–PLGA. Our aim was to validate the application of thermogelling composite gels as bone substitutes. We examined sol–gel–sol transition, pH value change, in vitro toxicity, in vitro hemolysis test, and in vivo animal study.
with a mechanical stirrer. An electric heater controlled the reactor temperature with the feedback sensor set to 160 1C. Stannous 2ethylhexanoate (34 mL) was added to the reactor to catalyze the polymerization process performed at 160 1C for 8 h. The resulting copolymer was dissolved in 60 mL dimethyl sulfoxide (DMSO). The solution was then purified by dialysis (MWCO¼1000) for 3 days at 4 1C and lyophilized at 20 1C for 3 days. 2.3. Characterization of mPEG–PLGA 1
H-nuclear magnetic resonance (NMR) spectroscopy, fourier transfer infrared spectra (FTIR) and gel permeation chromatography (GPC) were used to study the molecular structure and molecular weight of the mPEG–PLGA copolymer. NMR spectroscopy was performed on a 500-MHz NMR spectrometer (Varian Unityinova 500 NMR) at room temperature using CDCl3 as the solvent. 1H-NMR used a ratio of d ¼3.6 ppm, d ¼4.8 ppm, and d ¼5.2 ppm for the calculation of molecular weight and structure. FTIR measurements were obtained using the Perkin–Elmer system 2000 with KBr pellets and the transmittance data was collected. Molecular weight and molecular weight distributions were determined by GPC (RI-2031, PU-2080, JASCO), with tetrahydrofuran (THF) used as a solvent with a flow rate of 1 mL/min. 2.4. Fabrication of composite gel The mPEG–PLGA copolymer–ceramic composite gels were formed by adding 20 wt% mPEG–PLGA hydrogel with different weight ratios of HAP/b-TCP. The hydrogel (20 wt%) was obtained by mixing 200 mg of mPEG–PLGA hydrogel with 0.8 mL deionized H2O at 4 1C for 24 h. The mPEG–PLGA hydrogel (20 wt%) without ceramic was defined as GH0T0. Different types and amounts of ceramics were subsequently added to the hydrogel at 4 1C and stirred for 4 h to form composite gels. The weight ratio of polymer and ceramic was set at 2:1. The composite gels were named by the combination of HAP and b-TCP (i.e., HAP 70.00 mg and b-TCP 30.00 mg defined as CH7T3, HAP 50.00 mg and b-TCP 50.00 mg defined as CH5T5, and HAP 30.00 mg and b-TCP 70.00 mg defined as CH3T7). 2.5. Characterization of composite gel
2. Materials and methods 2.1. Materials Methoxy polyethylene glycol (mPEG) (Mn¼550 g/mole) was obtained from Aldrich Chem, Co. Inc. DL-lactide acid (LA) and glycolide acid were purchased from Purac. The b-TCP, HAP (nanopowder, o200 nm particle size) was from Sigma, the 1,6diphenyl-1,3,5-hexatriene (DPH) was purchased from Fluka, and stannous 2-ethylhexanoate (stannous octoate), which was used as a catalyst, was purchased from Aldrich Chem. 2.2. Synthesis of mPEG–PLGA Methoxy polyethylene glycol-co-poly(lactic-co-glycolic acid) (mPEG–PLGA) diblock copolymers were synthesized by ringopening polymerization of monomers and mPEG in the presence of stannous 2-ethylhexanoate. A typical synthetic procedure is shown in Fig. 1. According to our previous studies (Peng et al., 2010), mPEG–PLGA (molecular weight¼550–1405) copolymer is thermosensitive at physiological temperature and has the most suitable degradation time. To prepare the diblock copolymer, 24.02 g of mPEG was mixed with lactide acid (50.00 g) and glycolide acid (11.39 g) in a dry nitrogen atmosphere using a dry three-neck reactor equipped
The size of polymeric micelles was measured using dynamic light scattering (DLS). Measurements were carried out using a spectrophotometer (Nano Series Zeta Sizer; Malvern) equipped with a He–Ne laser at 633 nm, 25 1C, and a fixed scattering angle of 901. The nano-micelle solution, at a concentration of 0.1 wt%, was filtered through a 0.45-mm filter membrane before measurement. A 1 wt% composite gel was passed through a 0.45-mm filter and then serially diluted by a factor of 2 into 16 vials (1 wt%, 0.5 wt%, 0.025 wt%, y3.05 10 5 wt%). The fluorescent dye (DPH) was used at a concentration of 0.4 mM; 2 mL DPH solution was mixed with 100-mL copolymer solution at 4 1C overnight. The reaction proceeded in the dark. Using an enzyme-linked immunosorbent assay (ELISA) reader set at an excitation wavelength of 360 nm and emission wavelength of 480 nm, absorption of fluorescence intensity was measured. The ratio of fluorescence intensity was plotted against the logarithm of copolymer concentrations to determine the critical micellization concentration (CMC) (Hong et al., 2011). Different concentrations of mPEG–PLGA solution (i.e., 10%, 15%, 20%, 25%, 30%, and 35%) were prepared at 4 1C, and then, different weight ratios of HAP/b-TCP were sequentially added, followed by mixing at 4 1C for 12 h. 1-mL samples were prepared in Eppendorf tubes and incubated at 4 1C for 5 min until the temperature achieved equilibrium. The temperature was raised
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Fig. 1. Schematic diagram of the synthesis of the mPEG550–PLGA1405 diblock copolymer.
by intervals of 2 1C and maintained for 5 min before each sampling. The Eppendorf tubes were flipped upside down for 30 s to observe for any movement so as to determine the sol–gel status. The temperatures of sol-to-gel and then gel-to-sol transformation were recorded to produce a sol–gel–sol phase diagram using Gaussian regression. The mechanical properties of hydrogels with different weight ratios of HAP/b-TCP were measured using a rheometer (AR2000ex system; TA Instruments) with the temperature controller set from 5 1C to 45 1C and a heating rate of 2.2 1C/min. Polymeric solutions (0.65 mL, 20 wt%) were added to the instrument to analyze the rheological behavior of the sol–gel transition. Angular frequency (o) was set to 6.283 rad/s. G0 is defined as storage modulus, whereas G00 is the loss modulus. The temperature of initial G0 4G00 is defined as the phase transition temperature. The viscosity was measured along with the heating process. This method was used to determine the influence of HAP/b-TCP on the sol–gel–sol transition temperature and mechanical strength. Various composite copolymers (1 mL, 20 wt%) were prepared at 4 1C. Gel was formed in the release bottle of an incubator at 37 1C for 30 min until the gel stabilized; then, 9 mL phosphate buffered saline (PBS; pH ¼7.40) was added. The pH value of the solution containing degraded byproducts was measured by using a pH meter (Shindengen) over a 37-day period.
density measured at 570 nm. The results were analyzed using two-tailed Student’s t-test. Po0.05 was designated as statistically significant. Live and dead test was used to determine cell viability in 20 wt% composite gel. The cell culture medium was then withdrawn and mixed with 100 mL of fluorescent dye (2 mL Green (Calcein-AM), 1 mL Red (ethidium homodimer), and 1 mL of sterile PBS for 30 min in the dark. Inverted fluorescence microscopy (Axiovert 200, Zeiss) was used to determine cell viability, with green-stained cells indicating living cells, and red-stained cells indicating dead cells. The hemolytic test was also used to evaluate biocompatibility (Dinguizli et al., 2006). In brief, 10 wt% composite hydrogel solution was prepared by adding PBS. Then the diluted blood was added to ddH2O, composite hydrogel solution, and PBS (ddH2O and PBS were the positive and negative controls, respectively). After incubating the samples at 37 1C for 4 h, the supernatants of the solutions were collected by centrifugation. The greater the amount of precipitated hemoglobin in the solutions, the more severe the hemolysis, which has a negative impact on living organisms. In order to prevent the composite gels from forming into a gel phase in this experiment, 10 wt% composite gel was used in the hemolysis test. 2.8. Animal surgery
2.6. In vitro degradation of composite hydrogel The four copolymer mixture solutions (20 wt%, 0.2 mL) were injected into release bottles and incubated in a shaking bath at 37 1C. After 30 min, 1.8 mL of PBS solution (pH 7.4) was added to the formed gels, respectively. The bottles were then shaken at 100 rpm. At a predetermined time (0, 3, 6, 10, 15, 20, 25 and 30 day), three samples were taken out of the shaking bath for measuring weight loss. The remaining composite gels were freeze-dried until constant weight. The degradation rate was illustrated by weight loss. The residual weight percent was calculated via (Wd/Wo) 100%, where Wd was residual weight at the predetermined time, Wo was the original weight of the dried composite gel. 2.7. In vitro biocompatibility L929 fibroblast cells at a concentration of (5 103)/well in a 96-well plate were used to test the composite toxicity. The culture medium contained 10% fetal bovine serum mixed with 90% Dulbecco’s modified Eagle’s solution. After cell adhesion for 1 day, test samples were added to the wells. Because composite gels formed a gel phase at 37 1C, we could not test the biocompatibility of the gel directly. An experiment was conducted to test the medium indirectly. Briefly, 1 mL composite solution was formed into a gel at 37 1C in a release bottle. An aliquot of 9 mL of culture medium was added to the release bottle and incubated at 37 1C for 48 h. The medium was collected and sterilized with a 0.22 mm filter to isolate bacteria. The filtered culture medium was used to culture L929 cells for 48 h. MTT [3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide] solution at 2.5 mg/mL was used to test cytotoxicity. MTT is a yellow tetrazole, which is reduced by mitochondria in living cells to form formazan that has a purple color. The number of living cells can be inferred from the optical
In the preliminary study, GH0T0, CH7T3, CH5T5, and CH3T7 composite gels were injected subcutaneously into the NZW rabbits. There was serious foreign body reaction and fibrous tissue formation in GH0T0 group; the other three groups showed no obvious immunological reaction. Besides, the in vitro results demonstrated that GH0T0 was more acidic to environments, more toxic to cells and less biocompatible. Subsequent animal experiments enrolled CH7T3, CH5T5, and CH3T7, and excluded GH0T0. Eighteen 5-month-old New Zealand white rabbits (NZW, weighing 3.0 to 3.5 kg, Animal Health Research Institute, Miao-Li, Taiwan) were studied. The study was approved by the Animal Intuitional Review Board of Chung Gung Memorial Hospital and was conducted in compliance with the regulations for the care of laboratory animals. The NZWs were anesthetized by intramuscular injection of Zoletil 50 (25 mg/kg body weight) and 2% Rompon (0.15 mL/kg body weight). Under sterile conditions, the mid-shaft of the left femur was exposed through a lateral longitudinal skin incision and a 10-mm section of the diaphysis was osteotomized using an oscillating saw. The bone defect was filled with 0.8 mL composite gel (Lin et al., 2010). The 18 rabbits were divided into three groups (n¼6) according to the composite gels (CH7T3, CH5T5 and CH3T7, respectively). The dorsal side of the femur was stabilized with a stainless steel plate (DC-Plate, Synthes), which was fixed at the proximal and distal ends of the fractured femur with screws. After internal fixation, the deep muscle layer and skin were closed with interrupted sutures. The appetites and behaviors of the operative rabbits were normal. There was no mortality among the rabbits, postoperatively. 2.9. Radiologic assessment The rabbits underwent radiography at 4 and 8 weeks, postoperatively. Each radiograph was evaluated in a blinded fashion.
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All films were evaluated by recording the radiographic characteristics of callus growth. A bone defect was considered healed if the radiograph demonstrated contiguous callus spanning the proximal and distal ends of the fracture site within the femur (Guo et al., 2011). The rabbits were sacrificed at 8 weeks after surgery and all femora were scanned with an animal computed tomography (Bioscan, Washington, DC). The slice image data was acquired using a high-resolution frame as setup in the system, with tube voltage of 55 keV, pitch of 1.0, and 180 projections. The axial scanning range was set as 3 cm with the bone defect at the center field of view. The three dimensional surface rendering images were generated by Image J (National Institutes of Health). 2.10. Gross bone union After harvesting the specimens and removing the implants, the femora were manually palpated to check for bone union. The femora were manipulated with forces small enough to avoid producing gross trauma but great enough to evaluate for gross osteotomy gap motion. There was no difficulty in differentiating solid union, since those with pseudoarthrosis had obvious motion across the osteotomy site. Only segments identified as no gross motion were considered to be gross union (Chen et al., 2002). 2.11. Histological analysis The femora were fixed in 10% neutral buffered formalin for 3 days and then decalcified with 10% EDTA for one week. A 3-cm segment in the mid-shaft of the diaphysis (encompassing the graft) was cut, embedded in paraffin and sectioned longitudinally (5 mm thick). The sections were stained with hematoxylin and eosin (H&E) and Masson’s trichrome for histologic analysis.
3. Results and discussion 3.1. Characterization of mPEG550–PLGA1405 polymer The molecular structures of copolymers were characterized by H-NMR (Fig. 2A). The methylene protons of the mPEG repeating segments were detected at 3.6 ppm, the methane protons of the PLA segments at 5.2 ppm, and the methylene protons of the PGA 1
segments at 4.8 ppm (Chen et al., 2010). In FTIR spectra (Fig. 2B), the major peaks assigned to the structure of mPEG–PLGA were 2900–3000 cm 1 (C–H stretching of CH3), 1760 cm 1 (ester C¼O stretching), and 1080 cm 1 (O–CH2 stretching) (Wang et al., 2011b). On the basis of this data, it was confirmed that the mPEG–PLGA diblock copolymers were successfully synthesized. The theoretical molecular weight ratio of mPEG to PLGA was 550:1405. The molecular weight averages (Mn and Mw) and molecular weight distribution (PDI; Mw/Mn) of mPEG–PLGA diblock copolymers were determined by GPC and 1H-NMR, as shown in Table 1. mPEG–PLGA copolymers had a narrow molecular weight distribution around 1.45 and the measured found mole ratio of mPEG–PLGA was close to the theoretical value. These results indicated a consistent fabrication procedure. 3.2. Characterization of composite gel Particle sizes measured at room temperature are shown in Table 2. The diameters of CH7T3, CH5T5, and CH3T7 were approximately 75 nm. In contrast, the diameter of GH0T0 was only 59 nm, which indicated that adding ceramics resulted in 27% increase in size. The ceramics slightly increased the micellization concentration of the self-assembling polymers. The polydispersity (PD) values of the 4 samples were less than 0.30, which meant that they all had a single distribution. Zeta potential values were all negative. These findings were consistent with those in the literature. HAP/b-TCP dissociation is known to release negatively charged ions, including PO34 and OH , thereby causing a negative zeta potential (Avgoustakis et al., 2003). The negative zeta potential increases the nanoparticles’ stability (Duan et al., 2006). A negatively-charged surface (negative zeta potential) is favorable for bone regeneration and osseo-integration (Smeets et al., 2009). CMC was obtained by plotting intensity ratio vs. logarithm of copolymer concentration. CMC values for the 4 materials were all under 6.00 10 2 mg/mL. Table 2 illustrates that CMC increased by adding HAP/b-TCP. The CMC value increased because HAP/bTCP nanoparticles interfered with the self-assembly of mPEG– PLGA. However, the value remained less than 400 ppm (Allen et al., 1999), and therefore, had no impact on biomedical applications. The standard free energy for micellization (DG) is calculated using DG ¼RT ln(XCMC) (Zhang et al., 2006), where R is the
Fig. 2. (A) 1H NMR spectra of mPEG550–PLGA1405 diblock copolymer. a–c represent the mPEG segment, d–e indicate the LA segment and the f peak belongs to GA. (B) FTIR of mPEG550–PLGA1405. Compared with mPEG, there are obvious peaks at 1080 cm 1, 1760 cm 1, and 2900–3000 cm 1, which indicate the functional groups of PLGA.
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Table 1 The molecular weight of mPEG–PLGA. Mna mPEG550PLGA1405 mPEG:PLGA
Mwa
2023 2941 Feed mole ratio 78:22
PDIb
Mnc
1.45 2275 Found mole ratiod 75:25
a
Determined by GPC measurement in THF. Polydispersion (Mw/Mn). c Calculated by 1H-NMR. d Calculated by 1H-NMR. b
Table 2 The properties of composite gels (n¼ 3). Sample Particle size Polydispersity (nm) (PD) (*10) GH0T0 CH7T3 CH5T5 CH3T7
59.49 7 0.87 71.66 7 1.86 77.62 7 2.33 78.95 7 3.66
2.94 7 0.39 2.71 7 0.47 2.76 7 0.98 2.57 7 0.26
Zeta potential (mV)
CMC (mg/ mL)
DG (KJ/ mol)
22.83 7 0.42 23.03 7 0.12 24.97 7 0.62 26.73 7 0.17
3.47 10 2 5.12 10 2 5.32 10 2 4.95 10 2
37.32 36.36 36.27 36.45
Fig. 3. Sol–gel–sol phase diagrams of mPEG550–PLGA1405 mixed with different weight ratios of HAP/b-TCP. When the concentration of the copolymers increases gradually from 10 wt% to 35 wt%, the sol–gel transition temperature decreases and the gel–sol transition temperature rises accordingly.
universal gas constant, T is the absolute temperature, and XCMC is the corresponding CMC, expressed as mole fraction. The DG values obtained in this study were all less than zero, which indicated that the 4 composite gels could all spontaneously form micelles at very low concentrations. Gelation temperatures and properties were determined through a sol–gel–sol phase transition experiment. Sol–gel transition caused an increase in viscosity because of the aggregation of polyester hydrophobic segments. Preliminary tests revealed that the 10 wt% gel was too fragile to store in the Eppendorf tube, regardless of gel composition. The sol–gel–sol transition curves of the four composite gels ran parallel to each other. Those results revealed that adding different weight ratios of HAP/b-TCP did not significantly change the diagram of the sol–gel–sol transition (Fig. 3). The variation of mechanical properties as affected by the temperature change was studied using a rheometer. The results showed that for all our samples tested, the range of temperature
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for G0 greater than G00 is between 25.2 1C and 26.6 1C. This means that the polymers were able to form gel within this temperature range. The range of gelation temperature (25.2–26.6 1C) could guarantee the composite materials were sol on shelf and gel in vivo. When the temperature elevated up to 37 1C in vivo, the gel became more stable. The storage modulus G0 and loss modulus G00 , as shown in Fig. 4B and C, were slightly increased by the adding of HAP/b-TCP. When the temperature approached 37 1C, the values of G0 and G00 were nearly identical for 20% mPEG–PLGA hydrogels with or without ceramics (GH0T0, CH3T7, CH5T7 and CH7T3). This means that there was no significant difference in mechanical properties for composite gels polymers at body temperature. The values of G0 and G00 of 10% mPEG–PLGA hydrogels were smaller than 20% mPEG–PLGA hydrogel. Besides, the viscosity of 10% and 15% mPEG–PLGA hydrogels were smaller than 20% mPEG–PLGA. This means that hydrogels at lower concentration ( o20%) were mechanically inferior and not suitable as scaffolds. 3.3. pH change and degradation rate of composite gel mPEG–PLGA degradation produces acidic substances, resulting in a significant decline in pH; this causes local inflammation, foreign body reaction, and low cell adhesion. HAP or b-TCP in aqueous solution can release alkaline ions and subsequently reduce the decline in pH, thereby improving the biomedical applicability of the material. The addition of HAP/b-TCP to the mPEG–PLGA copolymer significantly reduced the pH declination (Fig. 5A). The 37-day pH chart shows that pH declined quickly in the copolymer GH0T0 (without ceramic) with a pH value less than 3.0 within 7 days and maintained a pH of 2.0 after 15 days. However, the pH values of the other 3 composite gels (with ceramic) could be maintained at 5.0 or more within 15 days. The pH declinations of composite gels were similar to a zero-order decrease and the pH value was 3.3 until the 37th day. The degradation rate of polyester was significantly influenced by the pH value; at lower pH, PLGA had a faster degradation rate (Gopferich, 1996). However, adding ceramics can increase the pH value by a scale of two. Compared with the other composite gels, the GH0T0 gel degraded fastest within the first 15 days, while the other composite gels had slower degradation rates. This result showed that another benefit of adding HAP/b-TCP to mPEG–PLGA hydrogels was a reduction in the degradation rate and, thus, an increase in durability. The degradation profiles of the composite gels were determined in vitro. Hydrolytic degradation of composite gels were carried out in PBS at 37 1C. The weight losses of the four composite gels were measured for 30 days (Fig. 5B). The degradation of the mPEG–PLGA was mainly via hydrolysis of ester linkages. Up to 80% of the GH0T0 gel was degraded within the 30-day period. On the contrary, the composite gels (CH7T3, CH5T5, and CH3T7) showed slower degradation rate compared to GH0T0. There were more than 40% residual weights for the three composite gels. The trend of degradation was correlated with the pH change. The results confirmed that adding bioceramics into hydrogels could effectively slow down the degradation rate. 3.4. Cytotoxicity of composite gel Because the composite gels were sensitive to temperature, cell culture could not be performed by simply placing the gels in culture wells. The composite gels had to be placed in release bottles. The byproduct of gel degradation was then released into the medium, and this medium was used to culture the cells. The MTT assay for cell toxicity was performed by culture of L929 fibroblasts. Cell toxicity was determined on the basis of cell survival rate.
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Fig. 4. The mechanical properties of hydrogels (A) viscosity (B) storage modulus G0 , (C) loss modulus G00 . The values of viscosity of 10% and 15% mPEG–PLGA are smaller than 20% mPEG–PLGA. The values of G0 and G00 of 10% mPEG–PLGA are smaller than the others.
Fig. 5. (A) The pH profiles of composite gels (in PBS) during a 37-day period. Initially, the pH value of GH0T0 decreases significantly, but the other composite gels show only a slight decline in pH value. (B) In vitro degradation of the composite gels. The weight losses of composite gels with different weight ratios of HAP/b-TCP were measured in a 30-day period. GH0T0 had the fastest degradation rate. Adding bioceramic can effectively slow down the degradation rate.
Fig. 6. (A) Cytotoxicity test (MTT assay) of the four composite gels at various concentrations. Regardless of the concentration, the cell survival rate of GH0T0 is significantly lower than the other composite gels. (B) The results of live and dead tests show the amount of cell survival (green fluorescence) of GH0T0 is significantly less than the other composite gels. The scale bar indicates 100 mm. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
The survival rates of GH0T0 at a concentration of 10%, 15%, 20%, 25% and 30% were 66.0%, 51.0%, 45.7%, 28.0% and 17.5%, respectively (Fig. 6A). With increasing concentration, the cell survival rate of GH0T0 decreased significantly. The survival rates of GH0T0, CH3T7, CH5T5 and CH7T3 at a concentration of 20%
were 45.7%, 72.3%, 79.7% and 81.0%, respectively. At any concentration, hydrogels with ceramics had better cell survival rates than hydrogels without ceramics (p o0.05). Therefore, adding HAP/b-TCP to mPEG–PLGA can neutralize the formation of acidic substances and enhance cell survival.
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Live and dead tests simulated the toxic effect of composite gels on neighboring cells in vivo. Fig. 6B shows that the amount of green fluorescence was significantly higher in 20 wt% CH7T3 composite gel, when each well had the same initial cell number. This suggests that the addition of different weight ratios of HAP/ b-TCP to mPEG–PLGA increases cell survival, thereby facilitating the biomedical applications of composite gels. In the hemolysis test, the supernatant was the aggregation of hemoglobin released from ruptured red blood cells. After centrifugation, a larger amount of supernatant represents a greater degree of hemolysis. The polymer with the greater degree of hemolysis had poorer blood oxygen transport capacity and more adverse biological effects. PBS has fixed osmotic pressure and good biocompatibility; therefore its hemolytic index is zero. Composite gels also had similar hemolytic indices, as shown in Table 3. A material having a hemolytic index less than 2% is considered non-hemolytic (Cencetti et al., 2011). The hydrogel GH0T0 with hemolytic index 3.1% was graded as hemolytic; while composite gels (C3H3T7, CH5T5 and CH7T3) with hemolytic index less than 2% were graded as non-hemolytic. The results indicated that the composites gels did not cause hemolysis. Considering the above in vitro results, 20 wt% composite gels were chosen for in vivo animal experiments, as the mechanical strength was too fragile when the concentration was less than 20 wt%. In addition, the hydrogel produced too much acid (affecting cell survival) when the concentration was higher than 20 wt%. 3.5. Radiographic and gross union Six rabbits underwent radiographic studies at 4 and 8 weeks after surgery. At 4 weeks, none of the osteotomy sites of the three groups showed radiographic union, but there were vague signs of
callus formation. The composite gels, CH7T3 and CH5T5, appeared to promote more callus formation than CH3T7. At 8 weeks, radiographic union was shown in 5 out of 6 osteotomy sites in CH7T3, 2 out of 6 in CH5T5 and 2 out of 6 in CH3T7 (Fig. 7). The radiographic union rate of CH7T3 was significantly higher compared with CH5T5 and CH3T7 (Table 4). After sacrifice at 8 weeks, the appearance of CH7T3 and CH5T5 demonstrated more callus formation, while CH3T7 showed more fibrous tissues in the gap. The stability of the osteotomy sites was evaluated by manual palpation. Gross union was noted in 4 out of 6 osteotomy sites in CH7T3, 2 out of 6 in CH5T5 and 1 out of 6 in CH3T7. Micro CT showed CH5T5 and CH3T7 with less new bone formation. CH7T3 triggered apparently more new bone formation, which bridged the bone defects. Using the composite gels as bone grafts, there were obvious new bone formations in the animal experiments. The imaging findings correlated with the gross appearance. 3.6. Histological analysis The implantation of composite gels resulted in new bone formation within the osteotomy sites at eight weeks after surgery. Sections through the osteotomy sites stained with H&E are shown in Fig. 8. CH7T3 showed obvious bone bridging across the fractured ends. CH5T5 exhibited pseudoarthrosis but there were signs of bone matrix deposition and mineralization. CH3T7 contained fibrous tissues with less bone matrix deposition and mineralization than CH5T5. Using the combination of radiologic assessment, gross bone union evaluation and histologic analysis, CH7T3 had better new Table 4 Result of radiographic union and gross union (n¼6).
Radiographic union
Table 3 Hemolysis test. ddH2O
PBS
mPEG550PLGA1405 solution (10 wt%) GH0T0
CH7T3
CH5T5
CH3T7
3.1% Hemolytic
2.2% 2.2% Non-hemolytic
1.1%
Gross union Hemolytic index Hemolytic grade
105.9%
0%
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Sample
Time (weeks)
Number of bone union
CH7T3 CH5T5 CH3T7 CH7T3 CH5T5 CH3T7 CH7T3 CH5T5 CH3T7
4 4 4 8 8 8 8 8 8
0 0 0 5 2 2 4 2 1
Fig. 7. Representative radiographs and gross appearance of the NZW rabbit femora at 8 weeks, postoperatively. The X-ray images were taken with the animals in supine position and with the right limb abducted and externally rotated. ((A), (D) and (G)) CH7T3 shows solid union and bone bridging, ((B), (E) and (H)) CH5T5 shows partial new bone formation and ((C), (F) and (I)) CH3T7 shows a fibrous gap.
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Fig. 8. Histological analysis (H&E staining, 40) of the composites gels. (A1) CH7T3 shows new bone formation bridging the osteotomy site at 8 weeks, while (B1) CH5T5 shows partial new bone formation, and (C1) CH3T7 shows fibrous tissue at the gap. Masson’s trichrome stains illustrate decalcified mineralized bone (blue) formation at the osteotomy sites (A2, B2 and C2). The scale bar indicates 1000 mm. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
bone formation compared to CH5T5 and CH7T7. Since b-TCP degrades faster than HAP, Blokhuis et al. (2000) the composite gels with higher ratios of b-TCP dissolved theoretically faster than those with higher ratios of HAP. The rapid dissolution of composite gels could then potentially lead to cyst formation within the osteotomy sites. The subsequent fibrous tissue incorporated into the cyst would then block further bone bridging. As shown in Fig. 8, the fibrous scar tissue blocked the osteogenic process in CH3T7 group. More osseous tissue formation was noted at the osteotomy sites of CH7T3 group than of CH5T3 and CH3T7 groups, as shown in Masson’s trichrome stain.
4. Conclusions We successfully prepared thermosensitive aqueous solutions of mPEG–PLGA diblock copolymers mixed with different weight ratios of HAP/b-TCP with suitable sol–gel transition temperatures. Amphiphilic mPEG–PLGA diblock gels were formed via aggregation of copolymer micelles when the temperature exceeded the gelling temperature. Adding ceramics did not affect micellization in the aqueous solution. The composite gels increased pH value, decreased hydrogel degradation rate and also enhanced in vivo cell viability. According to the results of our in vivo animal study, all the polymer–ceramic composite gels showed callus formation. With higher ratios of hydroxyapatite, both radiographic and gross union rates increased. The results suggested that the novel composite gels are suitable as bone substitutes for osteogenesis.
Acknowledgements This work was supported by grant CMRPG3B0541 from Chang Gung Memorial Hospital. We would like to thank the Laboratory Animal Center, Molecular Imaging Center, and Expensive Advanced
Instrument Core Laboratory, Chang Gung Memorial Hospital, Linkou for their support in this project.
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Contents lists available at SciVerse ScienceDirect
Chemical Engineering Science journal homepage: www.elsevier.com/locate/ces
Development of bioactive thermosensitive polymer–ceramic composite as bone substitute Po-Liang Lai a,b,1, Carl Tsai-Yu Lin b,c,1, Ding-Wei Hong b, Shu-Rui Yang b, Yu-Han Chang a, Lih-Huei Chen a, Wen-Jer Chen a, I-Ming Chu b,n a
Department of Orthopedic Surgery, Chang Gung Memorial Hospital, College of Medicine Chang Gung University, Taoyuan, Taiwan Department of Chemical Engineering, National Tsing Hua University, Hsinchu, Taiwan c Biomedical Technology and Device Research Labs, Industrial Technology Research Institute, Hsinchu, Taiwan b
H I G H L I G H T S c c c c c
Thermosensitive composite gelcomposed of mPEG–PLGA and HAP/b-TCPwas made. The gel is capable of sol-to-gel transition between 4 1C to 70 1C and of in-situ gelling. The composite gels do not acidify the surrounding environment. Higher hydroxyapatite contents help raise bone union rate. Bone union is examined through radiographic method.
a r t i c l e i n f o
abstract
Article history: Received 15 July 2012 Received in revised form 13 September 2012 Accepted 16 November 2012 Available online 2 December 2012
The purpose of this study was to design a thermosensitive composite gel to be used as a bone graft substitute. This gel can provide a more suitable microenvironment by using the amphiphilic triblock copolymer (mPEG550PLGA1405) consisting of methoxy poly(ethylene glycol) (mPEG), poly(lactic-coglycolic acid) (PLGA). An aqueous dispersion of mPEG550PLGA1405 mixed with different ratios of HAP/bTCP (composite gel) underwent a sol–gel–sol transition as the temperature was increased from 4 to 70 1C. The particle size and critical micellization concentration (CMC) were increased by adding ceramics. During the in vitro degradation process, composite gels demonstrated a slight decrease in pH value, a slower degradation rate, less toxicity, and a higher cell survival rate. The biocompatibility of the composite gels was validated by hemolysis test. In vivo animal studies demonstrated both radiographic and gross bone union when the ratio of HAP/b-TCP was 7:3. Based on the results, we have developed novel thermosensitive composite gels as bone substitutes. & 2012 Elsevier Ltd. All rights reserved.
Keywords: Biomedical engineering Bone substitutes Composites Gels Polymers Thermosensitive
1. Introduction Hydrogels are now widely used as biomaterials. In recent years, in situ forming hydrogels have been used for various biomedical applications (Hatefi and Amsden, 2002). Thermosensitive hydrogel (Zhai et al., 2009), whose aqueous solution is a sol at room temperature or lower, and which forms a gel at body temperature (37 1C), can avoid the negative effects of high
n
Corresponding author. Tel.: þ886 3 5713 704; fax: þ886 3 5715 408. E-mail addresses: [email protected] (P.-L. Lai), [email protected] (C. Tsai-Yu Lin), [email protected] (D.-W. Hong), [email protected] (S.-R. Yang), [email protected] (Y.-H. Chang), [email protected] (L.-H. Chen), [email protected] (W.-J. Chen), [email protected] (I.-M. Chu). 1 contributed equally to this work. 0009-2509/$ - see front matter & 2012 Elsevier Ltd. All rights reserved. http://dx.doi.org/10.1016/j.ces.2012.11.019
temperature and organic solvent on drugs and cells (Hishikawa et al., 2004). Commercially available thermogelling polymers, poloxamers (Roques et al., 2007), poly N-isopropylacrylamide (PNIPAM) (Xu et al., 2008), and poly N-isopropylacrylamide– acrylic acid (PNIPAM-AAc) (Wang et al., 2011a) have been widely investigated for their potential biomedical applications. Such thermosensitive polymers have been studied extensively for separation, modulation of biofunction, catalysis, drug delivery, cellular transport (Perumal et al., 2008) and cell culture/tissue engineering applications (Wang et al., 1999). However, the use of these systems is limited because they are not biodegradable in vivo. A bioresorbable material demonstrates stability against biodegradation (Turner et al., 2005) and further resorption in vivo (Eglin et al., 2009). Some polyethylene glycol (PEG)-based compounds (Than et al., 2008) can form thermosensitive biodegradable
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hydrogels which have hydrophilic and hydrophobic monomer units. They are often used to build hydrophilic blocks because they are biocompatible, highly soluble, and hydrate in water (Zhang et al., 2006). The most prominent hydrophobic examples are poly(lactic acid) (Legrand et al., 2007) and poly(lacticco-glycolic acid) (PLGA) that are used as biocompatible nanocontainer materials in drug delivery devices (Muthu et al., 2009). The Food and Drug Administration (FDA) has approved several products comprised of PLA, PGA and their copolymer PLGA (Sander et al., 2004). However, their utility in tissue engineering applications is hampered due to the accumulation of acidic degradation products from bulk erosion of the polymers that can adversely affect biocompatibility and lead to the catastrophic loss in mechanical integrity (Taylor et al., 1994). Inorganic bioceramics such as hydroxyapatite (HAP) (Aboudzadeh et al., 2010), tricalcium phosphate (TCP), and biphasic calcium phosphates (BCP) are commonly used as fillers, which can confer osteoconductivity and strength to bone tissue engineering grafts (Moursi et al., 2002). Generally, TCP degrades too rapidly and HAP too slowly (Detsch et al., 2008). An additional problem encountered with HAP is that it contains 4–5 wt% of carbonate ions as well as magnesium, strontium and fluoride and, thus, is not an exact mimic of bone mineral (Boanini et al., 2010). This implies that HAP cannot be remodelled as efficiently as natural bone mineral (Zhang et al., 2009). Researchers investigating polymer–HAP mixtures have experienced difficulty with dispersion and re-agglomeration of HAP particles in the polymer matrix at high loads (Kothapalli et al., 2008). Other polymer/ceramic composites, such as peptide/ hydroxyapatite nanocomposites (Detsch et al., 2010), and PCL/ tricalcium phosphate composite (Arafat et al., 2011), showed more promise. Therefore, we mixed different weight ratios of HAP/b-TCP with mPEG–PLGA hydrogel, because HAP/b-TCP dissociation is known to release anions including PO34 and OH which can neutralize the acid produced by mPEG–PLGA. Our aim was to validate the application of thermogelling composite gels as bone substitutes. We examined sol–gel–sol transition, pH value change, in vitro toxicity, in vitro hemolysis test, and in vivo animal study.
with a mechanical stirrer. An electric heater controlled the reactor temperature with the feedback sensor set to 160 1C. Stannous 2ethylhexanoate (34 mL) was added to the reactor to catalyze the polymerization process performed at 160 1C for 8 h. The resulting copolymer was dissolved in 60 mL dimethyl sulfoxide (DMSO). The solution was then purified by dialysis (MWCO¼1000) for 3 days at 4 1C and lyophilized at 20 1C for 3 days. 2.3. Characterization of mPEG–PLGA 1
H-nuclear magnetic resonance (NMR) spectroscopy, fourier transfer infrared spectra (FTIR) and gel permeation chromatography (GPC) were used to study the molecular structure and molecular weight of the mPEG–PLGA copolymer. NMR spectroscopy was performed on a 500-MHz NMR spectrometer (Varian Unityinova 500 NMR) at room temperature using CDCl3 as the solvent. 1H-NMR used a ratio of d ¼3.6 ppm, d ¼4.8 ppm, and d ¼5.2 ppm for the calculation of molecular weight and structure. FTIR measurements were obtained using the Perkin–Elmer system 2000 with KBr pellets and the transmittance data was collected. Molecular weight and molecular weight distributions were determined by GPC (RI-2031, PU-2080, JASCO), with tetrahydrofuran (THF) used as a solvent with a flow rate of 1 mL/min. 2.4. Fabrication of composite gel The mPEG–PLGA copolymer–ceramic composite gels were formed by adding 20 wt% mPEG–PLGA hydrogel with different weight ratios of HAP/b-TCP. The hydrogel (20 wt%) was obtained by mixing 200 mg of mPEG–PLGA hydrogel with 0.8 mL deionized H2O at 4 1C for 24 h. The mPEG–PLGA hydrogel (20 wt%) without ceramic was defined as GH0T0. Different types and amounts of ceramics were subsequently added to the hydrogel at 4 1C and stirred for 4 h to form composite gels. The weight ratio of polymer and ceramic was set at 2:1. The composite gels were named by the combination of HAP and b-TCP (i.e., HAP 70.00 mg and b-TCP 30.00 mg defined as CH7T3, HAP 50.00 mg and b-TCP 50.00 mg defined as CH5T5, and HAP 30.00 mg and b-TCP 70.00 mg defined as CH3T7). 2.5. Characterization of composite gel
2. Materials and methods 2.1. Materials Methoxy polyethylene glycol (mPEG) (Mn¼550 g/mole) was obtained from Aldrich Chem, Co. Inc. DL-lactide acid (LA) and glycolide acid were purchased from Purac. The b-TCP, HAP (nanopowder, o200 nm particle size) was from Sigma, the 1,6diphenyl-1,3,5-hexatriene (DPH) was purchased from Fluka, and stannous 2-ethylhexanoate (stannous octoate), which was used as a catalyst, was purchased from Aldrich Chem. 2.2. Synthesis of mPEG–PLGA Methoxy polyethylene glycol-co-poly(lactic-co-glycolic acid) (mPEG–PLGA) diblock copolymers were synthesized by ringopening polymerization of monomers and mPEG in the presence of stannous 2-ethylhexanoate. A typical synthetic procedure is shown in Fig. 1. According to our previous studies (Peng et al., 2010), mPEG–PLGA (molecular weight¼550–1405) copolymer is thermosensitive at physiological temperature and has the most suitable degradation time. To prepare the diblock copolymer, 24.02 g of mPEG was mixed with lactide acid (50.00 g) and glycolide acid (11.39 g) in a dry nitrogen atmosphere using a dry three-neck reactor equipped
The size of polymeric micelles was measured using dynamic light scattering (DLS). Measurements were carried out using a spectrophotometer (Nano Series Zeta Sizer; Malvern) equipped with a He–Ne laser at 633 nm, 25 1C, and a fixed scattering angle of 901. The nano-micelle solution, at a concentration of 0.1 wt%, was filtered through a 0.45-mm filter membrane before measurement. A 1 wt% composite gel was passed through a 0.45-mm filter and then serially diluted by a factor of 2 into 16 vials (1 wt%, 0.5 wt%, 0.025 wt%, y3.05 10 5 wt%). The fluorescent dye (DPH) was used at a concentration of 0.4 mM; 2 mL DPH solution was mixed with 100-mL copolymer solution at 4 1C overnight. The reaction proceeded in the dark. Using an enzyme-linked immunosorbent assay (ELISA) reader set at an excitation wavelength of 360 nm and emission wavelength of 480 nm, absorption of fluorescence intensity was measured. The ratio of fluorescence intensity was plotted against the logarithm of copolymer concentrations to determine the critical micellization concentration (CMC) (Hong et al., 2011). Different concentrations of mPEG–PLGA solution (i.e., 10%, 15%, 20%, 25%, 30%, and 35%) were prepared at 4 1C, and then, different weight ratios of HAP/b-TCP were sequentially added, followed by mixing at 4 1C for 12 h. 1-mL samples were prepared in Eppendorf tubes and incubated at 4 1C for 5 min until the temperature achieved equilibrium. The temperature was raised
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Fig. 1. Schematic diagram of the synthesis of the mPEG550–PLGA1405 diblock copolymer.
by intervals of 2 1C and maintained for 5 min before each sampling. The Eppendorf tubes were flipped upside down for 30 s to observe for any movement so as to determine the sol–gel status. The temperatures of sol-to-gel and then gel-to-sol transformation were recorded to produce a sol–gel–sol phase diagram using Gaussian regression. The mechanical properties of hydrogels with different weight ratios of HAP/b-TCP were measured using a rheometer (AR2000ex system; TA Instruments) with the temperature controller set from 5 1C to 45 1C and a heating rate of 2.2 1C/min. Polymeric solutions (0.65 mL, 20 wt%) were added to the instrument to analyze the rheological behavior of the sol–gel transition. Angular frequency (o) was set to 6.283 rad/s. G0 is defined as storage modulus, whereas G00 is the loss modulus. The temperature of initial G0 4G00 is defined as the phase transition temperature. The viscosity was measured along with the heating process. This method was used to determine the influence of HAP/b-TCP on the sol–gel–sol transition temperature and mechanical strength. Various composite copolymers (1 mL, 20 wt%) were prepared at 4 1C. Gel was formed in the release bottle of an incubator at 37 1C for 30 min until the gel stabilized; then, 9 mL phosphate buffered saline (PBS; pH ¼7.40) was added. The pH value of the solution containing degraded byproducts was measured by using a pH meter (Shindengen) over a 37-day period.
density measured at 570 nm. The results were analyzed using two-tailed Student’s t-test. Po0.05 was designated as statistically significant. Live and dead test was used to determine cell viability in 20 wt% composite gel. The cell culture medium was then withdrawn and mixed with 100 mL of fluorescent dye (2 mL Green (Calcein-AM), 1 mL Red (ethidium homodimer), and 1 mL of sterile PBS for 30 min in the dark. Inverted fluorescence microscopy (Axiovert 200, Zeiss) was used to determine cell viability, with green-stained cells indicating living cells, and red-stained cells indicating dead cells. The hemolytic test was also used to evaluate biocompatibility (Dinguizli et al., 2006). In brief, 10 wt% composite hydrogel solution was prepared by adding PBS. Then the diluted blood was added to ddH2O, composite hydrogel solution, and PBS (ddH2O and PBS were the positive and negative controls, respectively). After incubating the samples at 37 1C for 4 h, the supernatants of the solutions were collected by centrifugation. The greater the amount of precipitated hemoglobin in the solutions, the more severe the hemolysis, which has a negative impact on living organisms. In order to prevent the composite gels from forming into a gel phase in this experiment, 10 wt% composite gel was used in the hemolysis test. 2.8. Animal surgery
2.6. In vitro degradation of composite hydrogel The four copolymer mixture solutions (20 wt%, 0.2 mL) were injected into release bottles and incubated in a shaking bath at 37 1C. After 30 min, 1.8 mL of PBS solution (pH 7.4) was added to the formed gels, respectively. The bottles were then shaken at 100 rpm. At a predetermined time (0, 3, 6, 10, 15, 20, 25 and 30 day), three samples were taken out of the shaking bath for measuring weight loss. The remaining composite gels were freeze-dried until constant weight. The degradation rate was illustrated by weight loss. The residual weight percent was calculated via (Wd/Wo) 100%, where Wd was residual weight at the predetermined time, Wo was the original weight of the dried composite gel. 2.7. In vitro biocompatibility L929 fibroblast cells at a concentration of (5 103)/well in a 96-well plate were used to test the composite toxicity. The culture medium contained 10% fetal bovine serum mixed with 90% Dulbecco’s modified Eagle’s solution. After cell adhesion for 1 day, test samples were added to the wells. Because composite gels formed a gel phase at 37 1C, we could not test the biocompatibility of the gel directly. An experiment was conducted to test the medium indirectly. Briefly, 1 mL composite solution was formed into a gel at 37 1C in a release bottle. An aliquot of 9 mL of culture medium was added to the release bottle and incubated at 37 1C for 48 h. The medium was collected and sterilized with a 0.22 mm filter to isolate bacteria. The filtered culture medium was used to culture L929 cells for 48 h. MTT [3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide] solution at 2.5 mg/mL was used to test cytotoxicity. MTT is a yellow tetrazole, which is reduced by mitochondria in living cells to form formazan that has a purple color. The number of living cells can be inferred from the optical
In the preliminary study, GH0T0, CH7T3, CH5T5, and CH3T7 composite gels were injected subcutaneously into the NZW rabbits. There was serious foreign body reaction and fibrous tissue formation in GH0T0 group; the other three groups showed no obvious immunological reaction. Besides, the in vitro results demonstrated that GH0T0 was more acidic to environments, more toxic to cells and less biocompatible. Subsequent animal experiments enrolled CH7T3, CH5T5, and CH3T7, and excluded GH0T0. Eighteen 5-month-old New Zealand white rabbits (NZW, weighing 3.0 to 3.5 kg, Animal Health Research Institute, Miao-Li, Taiwan) were studied. The study was approved by the Animal Intuitional Review Board of Chung Gung Memorial Hospital and was conducted in compliance with the regulations for the care of laboratory animals. The NZWs were anesthetized by intramuscular injection of Zoletil 50 (25 mg/kg body weight) and 2% Rompon (0.15 mL/kg body weight). Under sterile conditions, the mid-shaft of the left femur was exposed through a lateral longitudinal skin incision and a 10-mm section of the diaphysis was osteotomized using an oscillating saw. The bone defect was filled with 0.8 mL composite gel (Lin et al., 2010). The 18 rabbits were divided into three groups (n¼6) according to the composite gels (CH7T3, CH5T5 and CH3T7, respectively). The dorsal side of the femur was stabilized with a stainless steel plate (DC-Plate, Synthes), which was fixed at the proximal and distal ends of the fractured femur with screws. After internal fixation, the deep muscle layer and skin were closed with interrupted sutures. The appetites and behaviors of the operative rabbits were normal. There was no mortality among the rabbits, postoperatively. 2.9. Radiologic assessment The rabbits underwent radiography at 4 and 8 weeks, postoperatively. Each radiograph was evaluated in a blinded fashion.
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All films were evaluated by recording the radiographic characteristics of callus growth. A bone defect was considered healed if the radiograph demonstrated contiguous callus spanning the proximal and distal ends of the fracture site within the femur (Guo et al., 2011). The rabbits were sacrificed at 8 weeks after surgery and all femora were scanned with an animal computed tomography (Bioscan, Washington, DC). The slice image data was acquired using a high-resolution frame as setup in the system, with tube voltage of 55 keV, pitch of 1.0, and 180 projections. The axial scanning range was set as 3 cm with the bone defect at the center field of view. The three dimensional surface rendering images were generated by Image J (National Institutes of Health). 2.10. Gross bone union After harvesting the specimens and removing the implants, the femora were manually palpated to check for bone union. The femora were manipulated with forces small enough to avoid producing gross trauma but great enough to evaluate for gross osteotomy gap motion. There was no difficulty in differentiating solid union, since those with pseudoarthrosis had obvious motion across the osteotomy site. Only segments identified as no gross motion were considered to be gross union (Chen et al., 2002). 2.11. Histological analysis The femora were fixed in 10% neutral buffered formalin for 3 days and then decalcified with 10% EDTA for one week. A 3-cm segment in the mid-shaft of the diaphysis (encompassing the graft) was cut, embedded in paraffin and sectioned longitudinally (5 mm thick). The sections were stained with hematoxylin and eosin (H&E) and Masson’s trichrome for histologic analysis.
3. Results and discussion 3.1. Characterization of mPEG550–PLGA1405 polymer The molecular structures of copolymers were characterized by H-NMR (Fig. 2A). The methylene protons of the mPEG repeating segments were detected at 3.6 ppm, the methane protons of the PLA segments at 5.2 ppm, and the methylene protons of the PGA 1
segments at 4.8 ppm (Chen et al., 2010). In FTIR spectra (Fig. 2B), the major peaks assigned to the structure of mPEG–PLGA were 2900–3000 cm 1 (C–H stretching of CH3), 1760 cm 1 (ester C¼O stretching), and 1080 cm 1 (O–CH2 stretching) (Wang et al., 2011b). On the basis of this data, it was confirmed that the mPEG–PLGA diblock copolymers were successfully synthesized. The theoretical molecular weight ratio of mPEG to PLGA was 550:1405. The molecular weight averages (Mn and Mw) and molecular weight distribution (PDI; Mw/Mn) of mPEG–PLGA diblock copolymers were determined by GPC and 1H-NMR, as shown in Table 1. mPEG–PLGA copolymers had a narrow molecular weight distribution around 1.45 and the measured found mole ratio of mPEG–PLGA was close to the theoretical value. These results indicated a consistent fabrication procedure. 3.2. Characterization of composite gel Particle sizes measured at room temperature are shown in Table 2. The diameters of CH7T3, CH5T5, and CH3T7 were approximately 75 nm. In contrast, the diameter of GH0T0 was only 59 nm, which indicated that adding ceramics resulted in 27% increase in size. The ceramics slightly increased the micellization concentration of the self-assembling polymers. The polydispersity (PD) values of the 4 samples were less than 0.30, which meant that they all had a single distribution. Zeta potential values were all negative. These findings were consistent with those in the literature. HAP/b-TCP dissociation is known to release negatively charged ions, including PO34 and OH , thereby causing a negative zeta potential (Avgoustakis et al., 2003). The negative zeta potential increases the nanoparticles’ stability (Duan et al., 2006). A negatively-charged surface (negative zeta potential) is favorable for bone regeneration and osseo-integration (Smeets et al., 2009). CMC was obtained by plotting intensity ratio vs. logarithm of copolymer concentration. CMC values for the 4 materials were all under 6.00 10 2 mg/mL. Table 2 illustrates that CMC increased by adding HAP/b-TCP. The CMC value increased because HAP/bTCP nanoparticles interfered with the self-assembly of mPEG– PLGA. However, the value remained less than 400 ppm (Allen et al., 1999), and therefore, had no impact on biomedical applications. The standard free energy for micellization (DG) is calculated using DG ¼RT ln(XCMC) (Zhang et al., 2006), where R is the
Fig. 2. (A) 1H NMR spectra of mPEG550–PLGA1405 diblock copolymer. a–c represent the mPEG segment, d–e indicate the LA segment and the f peak belongs to GA. (B) FTIR of mPEG550–PLGA1405. Compared with mPEG, there are obvious peaks at 1080 cm 1, 1760 cm 1, and 2900–3000 cm 1, which indicate the functional groups of PLGA.
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Table 1 The molecular weight of mPEG–PLGA. Mna mPEG550PLGA1405 mPEG:PLGA
Mwa
2023 2941 Feed mole ratio 78:22
PDIb
Mnc
1.45 2275 Found mole ratiod 75:25
a
Determined by GPC measurement in THF. Polydispersion (Mw/Mn). c Calculated by 1H-NMR. d Calculated by 1H-NMR. b
Table 2 The properties of composite gels (n¼ 3). Sample Particle size Polydispersity (nm) (PD) (*10) GH0T0 CH7T3 CH5T5 CH3T7
59.49 7 0.87 71.66 7 1.86 77.62 7 2.33 78.95 7 3.66
2.94 7 0.39 2.71 7 0.47 2.76 7 0.98 2.57 7 0.26
Zeta potential (mV)
CMC (mg/ mL)
DG (KJ/ mol)
22.83 7 0.42 23.03 7 0.12 24.97 7 0.62 26.73 7 0.17
3.47 10 2 5.12 10 2 5.32 10 2 4.95 10 2
37.32 36.36 36.27 36.45
Fig. 3. Sol–gel–sol phase diagrams of mPEG550–PLGA1405 mixed with different weight ratios of HAP/b-TCP. When the concentration of the copolymers increases gradually from 10 wt% to 35 wt%, the sol–gel transition temperature decreases and the gel–sol transition temperature rises accordingly.
universal gas constant, T is the absolute temperature, and XCMC is the corresponding CMC, expressed as mole fraction. The DG values obtained in this study were all less than zero, which indicated that the 4 composite gels could all spontaneously form micelles at very low concentrations. Gelation temperatures and properties were determined through a sol–gel–sol phase transition experiment. Sol–gel transition caused an increase in viscosity because of the aggregation of polyester hydrophobic segments. Preliminary tests revealed that the 10 wt% gel was too fragile to store in the Eppendorf tube, regardless of gel composition. The sol–gel–sol transition curves of the four composite gels ran parallel to each other. Those results revealed that adding different weight ratios of HAP/b-TCP did not significantly change the diagram of the sol–gel–sol transition (Fig. 3). The variation of mechanical properties as affected by the temperature change was studied using a rheometer. The results showed that for all our samples tested, the range of temperature
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for G0 greater than G00 is between 25.2 1C and 26.6 1C. This means that the polymers were able to form gel within this temperature range. The range of gelation temperature (25.2–26.6 1C) could guarantee the composite materials were sol on shelf and gel in vivo. When the temperature elevated up to 37 1C in vivo, the gel became more stable. The storage modulus G0 and loss modulus G00 , as shown in Fig. 4B and C, were slightly increased by the adding of HAP/b-TCP. When the temperature approached 37 1C, the values of G0 and G00 were nearly identical for 20% mPEG–PLGA hydrogels with or without ceramics (GH0T0, CH3T7, CH5T7 and CH7T3). This means that there was no significant difference in mechanical properties for composite gels polymers at body temperature. The values of G0 and G00 of 10% mPEG–PLGA hydrogels were smaller than 20% mPEG–PLGA hydrogel. Besides, the viscosity of 10% and 15% mPEG–PLGA hydrogels were smaller than 20% mPEG–PLGA. This means that hydrogels at lower concentration ( o20%) were mechanically inferior and not suitable as scaffolds. 3.3. pH change and degradation rate of composite gel mPEG–PLGA degradation produces acidic substances, resulting in a significant decline in pH; this causes local inflammation, foreign body reaction, and low cell adhesion. HAP or b-TCP in aqueous solution can release alkaline ions and subsequently reduce the decline in pH, thereby improving the biomedical applicability of the material. The addition of HAP/b-TCP to the mPEG–PLGA copolymer significantly reduced the pH declination (Fig. 5A). The 37-day pH chart shows that pH declined quickly in the copolymer GH0T0 (without ceramic) with a pH value less than 3.0 within 7 days and maintained a pH of 2.0 after 15 days. However, the pH values of the other 3 composite gels (with ceramic) could be maintained at 5.0 or more within 15 days. The pH declinations of composite gels were similar to a zero-order decrease and the pH value was 3.3 until the 37th day. The degradation rate of polyester was significantly influenced by the pH value; at lower pH, PLGA had a faster degradation rate (Gopferich, 1996). However, adding ceramics can increase the pH value by a scale of two. Compared with the other composite gels, the GH0T0 gel degraded fastest within the first 15 days, while the other composite gels had slower degradation rates. This result showed that another benefit of adding HAP/b-TCP to mPEG–PLGA hydrogels was a reduction in the degradation rate and, thus, an increase in durability. The degradation profiles of the composite gels were determined in vitro. Hydrolytic degradation of composite gels were carried out in PBS at 37 1C. The weight losses of the four composite gels were measured for 30 days (Fig. 5B). The degradation of the mPEG–PLGA was mainly via hydrolysis of ester linkages. Up to 80% of the GH0T0 gel was degraded within the 30-day period. On the contrary, the composite gels (CH7T3, CH5T5, and CH3T7) showed slower degradation rate compared to GH0T0. There were more than 40% residual weights for the three composite gels. The trend of degradation was correlated with the pH change. The results confirmed that adding bioceramics into hydrogels could effectively slow down the degradation rate. 3.4. Cytotoxicity of composite gel Because the composite gels were sensitive to temperature, cell culture could not be performed by simply placing the gels in culture wells. The composite gels had to be placed in release bottles. The byproduct of gel degradation was then released into the medium, and this medium was used to culture the cells. The MTT assay for cell toxicity was performed by culture of L929 fibroblasts. Cell toxicity was determined on the basis of cell survival rate.
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Fig. 4. The mechanical properties of hydrogels (A) viscosity (B) storage modulus G0 , (C) loss modulus G00 . The values of viscosity of 10% and 15% mPEG–PLGA are smaller than 20% mPEG–PLGA. The values of G0 and G00 of 10% mPEG–PLGA are smaller than the others.
Fig. 5. (A) The pH profiles of composite gels (in PBS) during a 37-day period. Initially, the pH value of GH0T0 decreases significantly, but the other composite gels show only a slight decline in pH value. (B) In vitro degradation of the composite gels. The weight losses of composite gels with different weight ratios of HAP/b-TCP were measured in a 30-day period. GH0T0 had the fastest degradation rate. Adding bioceramic can effectively slow down the degradation rate.
Fig. 6. (A) Cytotoxicity test (MTT assay) of the four composite gels at various concentrations. Regardless of the concentration, the cell survival rate of GH0T0 is significantly lower than the other composite gels. (B) The results of live and dead tests show the amount of cell survival (green fluorescence) of GH0T0 is significantly less than the other composite gels. The scale bar indicates 100 mm. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
The survival rates of GH0T0 at a concentration of 10%, 15%, 20%, 25% and 30% were 66.0%, 51.0%, 45.7%, 28.0% and 17.5%, respectively (Fig. 6A). With increasing concentration, the cell survival rate of GH0T0 decreased significantly. The survival rates of GH0T0, CH3T7, CH5T5 and CH7T3 at a concentration of 20%
were 45.7%, 72.3%, 79.7% and 81.0%, respectively. At any concentration, hydrogels with ceramics had better cell survival rates than hydrogels without ceramics (p o0.05). Therefore, adding HAP/b-TCP to mPEG–PLGA can neutralize the formation of acidic substances and enhance cell survival.
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Live and dead tests simulated the toxic effect of composite gels on neighboring cells in vivo. Fig. 6B shows that the amount of green fluorescence was significantly higher in 20 wt% CH7T3 composite gel, when each well had the same initial cell number. This suggests that the addition of different weight ratios of HAP/ b-TCP to mPEG–PLGA increases cell survival, thereby facilitating the biomedical applications of composite gels. In the hemolysis test, the supernatant was the aggregation of hemoglobin released from ruptured red blood cells. After centrifugation, a larger amount of supernatant represents a greater degree of hemolysis. The polymer with the greater degree of hemolysis had poorer blood oxygen transport capacity and more adverse biological effects. PBS has fixed osmotic pressure and good biocompatibility; therefore its hemolytic index is zero. Composite gels also had similar hemolytic indices, as shown in Table 3. A material having a hemolytic index less than 2% is considered non-hemolytic (Cencetti et al., 2011). The hydrogel GH0T0 with hemolytic index 3.1% was graded as hemolytic; while composite gels (C3H3T7, CH5T5 and CH7T3) with hemolytic index less than 2% were graded as non-hemolytic. The results indicated that the composites gels did not cause hemolysis. Considering the above in vitro results, 20 wt% composite gels were chosen for in vivo animal experiments, as the mechanical strength was too fragile when the concentration was less than 20 wt%. In addition, the hydrogel produced too much acid (affecting cell survival) when the concentration was higher than 20 wt%. 3.5. Radiographic and gross union Six rabbits underwent radiographic studies at 4 and 8 weeks after surgery. At 4 weeks, none of the osteotomy sites of the three groups showed radiographic union, but there were vague signs of
callus formation. The composite gels, CH7T3 and CH5T5, appeared to promote more callus formation than CH3T7. At 8 weeks, radiographic union was shown in 5 out of 6 osteotomy sites in CH7T3, 2 out of 6 in CH5T5 and 2 out of 6 in CH3T7 (Fig. 7). The radiographic union rate of CH7T3 was significantly higher compared with CH5T5 and CH3T7 (Table 4). After sacrifice at 8 weeks, the appearance of CH7T3 and CH5T5 demonstrated more callus formation, while CH3T7 showed more fibrous tissues in the gap. The stability of the osteotomy sites was evaluated by manual palpation. Gross union was noted in 4 out of 6 osteotomy sites in CH7T3, 2 out of 6 in CH5T5 and 1 out of 6 in CH3T7. Micro CT showed CH5T5 and CH3T7 with less new bone formation. CH7T3 triggered apparently more new bone formation, which bridged the bone defects. Using the composite gels as bone grafts, there were obvious new bone formations in the animal experiments. The imaging findings correlated with the gross appearance. 3.6. Histological analysis The implantation of composite gels resulted in new bone formation within the osteotomy sites at eight weeks after surgery. Sections through the osteotomy sites stained with H&E are shown in Fig. 8. CH7T3 showed obvious bone bridging across the fractured ends. CH5T5 exhibited pseudoarthrosis but there were signs of bone matrix deposition and mineralization. CH3T7 contained fibrous tissues with less bone matrix deposition and mineralization than CH5T5. Using the combination of radiologic assessment, gross bone union evaluation and histologic analysis, CH7T3 had better new Table 4 Result of radiographic union and gross union (n¼6).
Radiographic union
Table 3 Hemolysis test. ddH2O
PBS
mPEG550PLGA1405 solution (10 wt%) GH0T0
CH7T3
CH5T5
CH3T7
3.1% Hemolytic
2.2% 2.2% Non-hemolytic
1.1%
Gross union Hemolytic index Hemolytic grade
105.9%
0%
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Sample
Time (weeks)
Number of bone union
CH7T3 CH5T5 CH3T7 CH7T3 CH5T5 CH3T7 CH7T3 CH5T5 CH3T7
4 4 4 8 8 8 8 8 8
0 0 0 5 2 2 4 2 1
Fig. 7. Representative radiographs and gross appearance of the NZW rabbit femora at 8 weeks, postoperatively. The X-ray images were taken with the animals in supine position and with the right limb abducted and externally rotated. ((A), (D) and (G)) CH7T3 shows solid union and bone bridging, ((B), (E) and (H)) CH5T5 shows partial new bone formation and ((C), (F) and (I)) CH3T7 shows a fibrous gap.
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Fig. 8. Histological analysis (H&E staining, 40) of the composites gels. (A1) CH7T3 shows new bone formation bridging the osteotomy site at 8 weeks, while (B1) CH5T5 shows partial new bone formation, and (C1) CH3T7 shows fibrous tissue at the gap. Masson’s trichrome stains illustrate decalcified mineralized bone (blue) formation at the osteotomy sites (A2, B2 and C2). The scale bar indicates 1000 mm. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
bone formation compared to CH5T5 and CH7T7. Since b-TCP degrades faster than HAP, Blokhuis et al. (2000) the composite gels with higher ratios of b-TCP dissolved theoretically faster than those with higher ratios of HAP. The rapid dissolution of composite gels could then potentially lead to cyst formation within the osteotomy sites. The subsequent fibrous tissue incorporated into the cyst would then block further bone bridging. As shown in Fig. 8, the fibrous scar tissue blocked the osteogenic process in CH3T7 group. More osseous tissue formation was noted at the osteotomy sites of CH7T3 group than of CH5T3 and CH3T7 groups, as shown in Masson’s trichrome stain.
4. Conclusions We successfully prepared thermosensitive aqueous solutions of mPEG–PLGA diblock copolymers mixed with different weight ratios of HAP/b-TCP with suitable sol–gel transition temperatures. Amphiphilic mPEG–PLGA diblock gels were formed via aggregation of copolymer micelles when the temperature exceeded the gelling temperature. Adding ceramics did not affect micellization in the aqueous solution. The composite gels increased pH value, decreased hydrogel degradation rate and also enhanced in vivo cell viability. According to the results of our in vivo animal study, all the polymer–ceramic composite gels showed callus formation. With higher ratios of hydroxyapatite, both radiographic and gross union rates increased. The results suggested that the novel composite gels are suitable as bone substitutes for osteogenesis.
Acknowledgements This work was supported by grant CMRPG3B0541 from Chang Gung Memorial Hospital. We would like to thank the Laboratory Animal Center, Molecular Imaging Center, and Expensive Advanced
Instrument Core Laboratory, Chang Gung Memorial Hospital, Linkou for their support in this project.
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