Preparation and characterization of (PCL-crosslinked-PEG)/ hydroxyapatite as bone tissue engineering scaffolds Narjes Koupaei,1 Akbar Karkhaneh,2* Morteza Daliri Joupari3 1
Department of Biomedical Engineering, Science and Research Branch, Islamic Azad University, Tehran, Iran Department of Biomedical Engineering, Amirkabir University of Technology (Tehran Polytechnic), Tehran, Iran 3 Department of Animal and Marine Biotechnology, National Institute of Genetic Engineering and Biotechnology, Tehran, Iran 2
Received 11 January 2015; revised 10 May 2015; accepted 19 May 2015 Published online 27 August 2015 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.a.35513 Abstract: In this study, interconnected porous bioactive scaffolds were synthesized for bone tissue engineering. At the first step, poly(E-caprolactone) (PCL) diols were diacrylated with acryloyl chloride. Then, the scaffolds were synthesized by radical crosslinking reaction of PCL and poly(ethyleneglycol) (PEG) diacrylates in the presence of hydroxyapatite (HA) particles. Morphological, swelling, thermal, and mechanical characteristics as well as degradability of the scaffolds were investigated. Results showed that increasing the ratio of PEG to PCL led to significant increase of swelling ratio and degradation rate, and decrease of crystallinity and compressive modulus of the networks, respectively. It was found that the incorporation of HA particles with the polymer matrices resulted in an augmented crystallinity, a decreased swelling ratio, and also a significantly increased compressive modulus
of the networks. Cytocompatability and osteoconductivity of the scaffolds were assessed by MTT and alkaline phosphatase (ALP) assays, respectively. The results confirmed the cytocompatible nature of PCL/PEG/HA scaffolds with no toxicity. MG-63 cells attached and spread on the pore walls offered by the scaffolds. PCL/PEG/HA scaffolds compared with PCL/PEG ones showed higher ALP activity. Thus, the results indicated that the PCL/PEG/HA scaffolds have the potential of being used as promising substrates in bone tisC 2015 Wiley Periodicals, Inc. J Biomed Mater Res sue engineering. V Part A: 103A: 3919–3926, 2015.
Key Words: polycaprolactone, polyethylene glycol, hydroxyapatite, biodegradable porous scaffolds, bone tissue engineering
How to cite this article: Koupaei N, Karkhaneh A, Daliri Joupari M. 2015. Preparation and characterization of (PCL-crosslinkedPEG)/hydroxyapatite as bone tissue engineering scaffolds. J Biomed Mater Res Part A 2015:103A:3919–3926.
INTRODUCTION
Bone tissue engineering is gaining popularity as alternative method for treatment of osseous defects.1 A number of biodegradable polymers have been explored for tissue engineering purposes.1,2 At present, poly(E-caprolactone) (PCL) has been considered for several biomedical applications such as scaffolds for supporting fibroblasts and osteoblasts growth.3 It is hydrophobic and insoluble in water but degradable through the hydrolytic attack of the ester bond.4 Because of its high crystallinity and hydrophilic/hydrophobic balance between ester and methylene groups, PCL has shown remarkably long in vivo degradation times.4,5 Moreover, the lack of bioactive functional groups and intrinsic hydrophobicity result in a poor cell adhesion, which is a critical issue for successful in vitro 3D cell culture and the subsequent tissue formation.6 Therefore, modification of PCL is of great necessity for a wider range of medical applications.3,5 Poly(ethyleneglycol) (PEG) has been an important type of hydrophilic polymers for biomedical applications because they have critical properties, such as good biocompatibility, hydrophilicity, lack of toxicity, and absence of antigenicity and immunogenicity.4,7 Hydroxyapatite (HA) is a
ceramic material with a composition and structure similar to natural bone mineral and has been demonstrated to have good biocompatibility, and even osteoconductivity and osteoinductivity.8 Some studies reported that inclusion of HA into PCL scaffold improved osteoconductivity and produced mineralized nodules.9 However, the hydrophobicity of PCL causes cell attached difficulties which may limit clinical applications.8 Therefore, PEG and HA are good candidates for modifying the property of PCL for use as bone tissue engineering scaffolds. Nowadays, although there are growing interests on the amphiphilic scaffolds based on PCL and a hydrophilic polymer such as PEG in tissue engineering, the most important disadvantage of this system is the release of short chain PEGs from the scaffold in absence of chemical crosslinking.10,11 On the other hand, the bioactivity of the scaffold in bone tissue engineering is a major challenge.12 The aim of this study was to introduce chemically crosslinked biodegradable scaffolds by crosslinking reaction of PCL and PEG diacrylates in the presence of HA. The scaffolds were characterized, MG63 cells were cultured onto the scaffolds, and the morphological and biological studies were carried out.
Correspondence to: A. Karkhaneh; e-mail: [email protected]
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TABLE I. Sample Code and Chemical Composition of PCLCrosslinked PEG Scaffolds
Sample Code
PCLDA/ PEGDA Weight Ratio
HA (wt/vol %)
Total Polymer Concentration (wt/vol %)
100PCLDA 100PCLDA/HA 75PCLDA 75PCLDA/HA 50PCLDA 50PCLDA/HA
100:0 100:0 75:25 75:25 50:50 50:50
0 5 0 5 0 5
30 30 30 30 30 30
(1H-NMR) spectra of PCL-DA were obtained on a Bruker Avance 400 (Bruker Biospin, Germany) spectrometer operating at 400 MHz. Also, attenuated total reflectance-Fourier transform-infrared (ATR-FT-IR, Jasco, Japan) spectra were recorded to analyze the chemical structure of the PCLcrosslinked PEG hydrogels. Differential scanning calorimetry (DSC). The DSC experiments were performed on a Perkin–Elmer Pyris DSC 6 under nitrogen atmosphere. All the samples were first heated from room temperature to 100 C and cooled to 0 C to erase the thermal history.15 Then a subsequent heating run was performed from 0 C to 100 C. The heating and cooling rate was of 10 C=min.
MATERIALS AND METHODS
Materials Poly(E-caprolactone) diol (PCL diol, Mn 5 2000), PEG diacrylate (PEGDA, Mn 5 700), and 2,2-azobisisobutyronitrile (AIBN) as a radical initiator were purchased from SigmaAldrich. Benzene, n-hexane, acryloyl chloride, triethylamine (TEA), and hydroxyapatite (HA) were purchased from Merck. 1,4-Dioxane was from AppliChem. Sodium chloride (NaCl) particulates with size in the range of 105 to 250 mm were used as the progen.
Measurement of gel fraction and swelling. To determine the gel fraction, three film-typed hydrogel samples for each solvent were immersed in excess CH2Cl2 and water for removing sol parts (soluble portion). After two days, undissolved part of samples were taken out and dried in a vacuum oven at 40 C to constant weight. The gel fraction (G) was calculated by the following Eq. (1): G ¼ ðWd =Wi Þ 3 100
Poly(E-caprolactone) diacrylate (PCLDA) synthesis PCL diol was end-capped with acrylate groups to form a polymerizable macromer.13 Briefly, 5 g (2.5 mmol) of PCL diol (Mn 5 2000) was dissolved in 40 mL of benzene and 0.866 mL (6.25 mmol) of triethylamine was added. Then, 0.505 mL (6.25 mmol) of acryloyl chloride, dissolved in a small amount of benzene, was added dropwise with constant magnetic stirring. The reaction mixture was stirred for 3 h at 80 C. Afterwards, it was filtered to remove triethylamine hydrochloride and then the filtrate was precipitated in an excessive amount of n-hexane. Finally, the precipitated product was collected and dried in a vacuum oven for 24 h. Preparation of PCL/PEG networks scaffold PCL/PEG networks were prepared by thermal polymerization method.14 PCL and PEG diacrylates with different weight ratios were dissolved in 1,4-dioxane to have a polymer concentration of 30% w/v. To this solution, the predetermined amounts of AIBN as initiator, HA, and NaCl particulates as porogen were added. Homogeneous PCLDA/ PEGDA/AIBN/HA/NaCl mixture was transferred into test tube and placed in an oven at 70 C for 12 h. The resulting gel was removed from the tube and cut into discs with thicknesses of 3 mm. Afterwards, the salt particulates were leached out by immersing the gel samples in distilled water for 4 days at room temperature, in order to completely remove the porogen. The sample codes and reactant compositions are shown in Table I. Characterization Analysis by FT-IR and 1H NMR. Fourier transform-infrared (FT-IR, Jasco, Japan) spectra were measured to confirm the formation of PCL-DA. 1H nuclear magnetic resonance
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(1)
where Wi is the initial weight of dried gel sample and Wd is the weight of dried gel sample after extraction in solvent. To determine the swelling ratio, the prepared samples were immersed in distilled water at 37 C. After reaching the equilibrium swelling state, the excessive water on the surface was removed by wiping superficially with a filter article. The swelling ratio (S) was calculated using the Eq. (2): S ¼ ½ðWs 2Wd Þ=Wd 3 100
(2)
where Ws and Wd are the weights of swollen and dried samples after extraction in solvent, respectively. The value is expressed as means 6 SD (n 5 3). Morphology observation by SEM. The cross-section morphology of the scaffolds was examined by scanning electron microscope (SEM, KYKY-EM-3200). The scaffolds were fractured in liquid nitrogen and coated with gold. Measurement of porosity. The porosity of the scaffold was determined using a density bottle based on Archimedes’ Principle according to a known technique.16–18 In brief, the porosity of scaffold was calculated by the following Eq. (3): Porosity ð%Þ ¼
ðW2 2W3 2WS Þ=qe 3100 ðW1 2W3 Þ=qe
(3)
where W1 is the weight of density bottle filled with ethanol, W2 the weight of density bottle including ethanol and scaffold, W3 the weight of density bottle taken out ethanolsaturated scaffold from W2, WS the weight of scaffold, qe the density of ethanol, ðW1 2W3 Þ=qe the total volume of the
PCL-CROSSLINKED-PEG HYDROXYAPATITE AS BONE TISSUE ENGINEERING SCAFFOLDS
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scaffold including pores, and ðW2 2W3 2WS Þ=qe volume in the scaffold.
the pore
Mechanical properties. The compression stress–strain curves of scaffolds were carried out by a universal testing machine (STM-20, Santam, Iran) at a cross-head speed of 1 mm min21. The specimens were stored in distilled water at 37 C for 48 h before test. Three specimens were tested for each sample, and the averages and standard deviations were reported. In vitro degradation. To assess the in vitro degradation behavior of the scaffolds, the samples were weighed (Wi), immersed in an excessive amount of distilled water, and incubated at 37 C with constant shaking for 6 months. Three samples of each group were tested. At preselected time point, the samples were removed from the water, freeze dried, and weighed (Wf). The percentage (%) degradation was calculated using the following formula: Degradation % ¼ ½ðWi 2Wf Þ=Wi 3100
(4)
Cell culture studies Osteosarcoma cell line MG63 (Pasteur Ins, NCBI, Iran) was chosen to evaluate the biological performance of the scaffolds synthesized in this study. Cell lines were cultured in Dulbecco’s Modified Eagle’s Medium (DMEM, GIBCO, Scotland)/supplemented with 10% (v/v) fetal bovine serum (FBS, Seromed, Germany), 100 U mL21 penicillin and 100 lg mL21 streptomycin. Cultured cells were detached by trypsinization, suspended in fresh culture medium, and used for investigating the cytocompatibility of the scaffolds. Before cell seeding, scaffolds were sterilized using UV treatment. The scaffolds were seeded at a density of 1 3 104 cells/scaffold and maintained in DMEM in an incubator at 378C in a humidified incubator with 5% CO2. MTT assay. The cytotoxicity of the scaffolds was determined using the colorimetric MTT assay. Extracts from the scaffolds were prepared by incubating the presterilized scaffolds incubated in culture medium for 14 days at 378C. At the end of this period, the scaffolds were removed, and the socalled extracts were obtained. For the MTT assay, cells were placed into a 96-well plate at a density of 1 3 104 cells/ well and were incubated under standard culturing conditions. After incubation for 24 h, the culture medium was removed and replaced by the extract and incubated for another 24 h, and then 100 mL of 0.5 mg mL21 MTT solution was added into each well. After that, the plate was incubated at 378C for 4 h. Then the medium was removed, 100 mL of isopropanol solution was added and further incubated for 15 min in order to enhance the dissolution of the formazan crystal. The optical density (OD) of formazan in the solution was detected by a multiwall microplate reader (STAT FAX 2100) at 545 nm. For the reference purpose, cells were seeded into a fresh culture medium (control)
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with same seeding condition. The percentage (%) viability was estimated as per the following formula: Viability % ¼ ðODs =ODc Þ3100
(5)
where ODs and ODc are the average optical densities of the sample and the control, respectively.19 Each assay was performed four times. Alkaline phosphatase activity. The osteocoductivity of scaffolds was evaluated by measuring the extraction of alkaline phosphatase (ALP) enzyme as an indicator of osteoblasts activity. The measurement of ALP activity was performed as follows. The cells were cultured on each scaffold in 24-well culture plates. After 7 and 14 days, the concentration of ALP released by cells into the medium in each well was measured via Autoanalyzer (Hitachi 717, Tokyo, Japan) at 405 nm. The culture medium with the same amount of cells was considered as control. Cell morphology on the scaffolds. Morphology and spreading pattern of cells on the scaffolds were evaluated 6 and 24 h after seeding using SEM. For SEM analysis, cell-seeded scaffolds were fixed with 4% glutaraldehyde, rinsed with PBS, dehydrated using graded series of ethanol (50–100%) and dried in vacuum overnight. The dry samples were coated with gold and observed by SEM. Statistical analysis Data were analyzed using one-way analysis of variance (ANOVA) with Tukey’s post hoc honestly significant difference (HSD) test to a confidence level of p < 0.05. RESULTS AND DISCUSSION
Characterization of PCL diacrylate The FT-IR spectra of PCL diol and PCL diacrylate were shown in Figure 1. PCL diacrylate shows absorption bands at 1635 and 813 cm21 assigned to the vinyl (H2C@CHA) groups due to acrylation of PCL diol.13,20 Those peaks were not observed in PCL diol itself [Fig. 1(a)]. The absorption bands at around 1725 and 2950 cm21, which were present in both PCL diol and PCL diacrylate, were attributed to ester carbonyl (AC@O) and methylene groups (ACH2A), respectively. The absorption peak around 3450 cm21 for hydroxyl (AOH) groups in PCL diol became weaker in PCL diacrylate as they were replaced by acrylate groups.13,20 The formation of PCL diacrylate was also confirmed through 1H-NMR spectrometer. As shown in 1H-NMR spectrum (Fig. 2), the vinyl groups of the PCL-DA appeared in the 5.7 to 6.5 ppm range in agreement with literature.14,20 From the above results, the terminal hydroxyl groups in the PCL diol were converted to acrylate groups by the reaction with acryloyl chloride. FT-IR studies of PCL/PEG hydrogel networks Biodegradable PCL/PEG hydrogel networks were synthesized by a radical crosslinking reaction of PEGDA and PCLDA in the presence of hydroxyapatite particulates. The
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FIGURE 3. ATR-FT-IR spectra of (a) 75PCLDA, (b) 75PCLDA/HA networks.
FIGURE 1. FT-IR spectra of (a) PCL diol, (b) PCLDA.
chemical structures of 75PCLDA and 75PCLDA/HA hydrogel networks were characterized using ATR-FT-IR spectra [Fig. 3(a,b)]. No evident peaks were found at 813 and 1635 cm21 for vinyl (H2C@CHA) groups in these networks, suggesting that the carbon-carbon double bonds in PCLDA were largely consumed in the crosslinking. The absorption peak for phosphate (PO4) group of HA at around 1029 cm21 (see Ref. 8) observed in the spectrum of 75PCLDA/HA hydrogel.
FIGURE 2. 1H NMR spectra of (a) PCL diol, (b) PCLDA.
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DSC studies Figure 4 shows DSC thermograms of PCL diol, PCLDA, and PCL/PEG networks (with and without HA) from the heating run. The melting temperature (Tm), heat of fusion (DHm), and crystallinity (Xc) are shown in Table II. Tm for PCL diol and PCLDA were determined using the highest peak temperature among multiple exothermal peaks, which corresponded to the different arm lengths in them.20 The Xc of the PCL was calculated from DH assuming proportionality to the experimental enthalpy. The reported enthalpy of fusion (135.44 J/g) was used as a reference of 100% Xc of PCL.21 The Xc for PCL moiety in the PCL/PEG networks was calculated using the Eq. (5): c Xc % ¼ DHm = wPCL DHm 3 100 (6)
FIGURE 4. DSC curves of PCL diol, PCLDA, and PCL/PEG networks (with or without HA).
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TABLE II. Melting Point, Enthalpy of Fusion, and Crystallinity of PCL Diol, PCLDA, and PCL/PEG Networks (with and Without HA)
PCL diol PCLDA 100PCLDA 100PCLDA/HA 75PCLDA 75PCLDA/HA 50PCLDA 50PCLDA/HA
Tm (8C)
DHm ðJ=gÞ
Xc(%)
39.5 37 33.5 34.3 28.8 29.8 21.5 26.7
74.5 68.2 44.9 40.5 31.3 28.4 19.3 18.7
55 50.3 33.2 34.8 31.8 32.8 27.04 31.4
TABLE III. Gel Fraction and Swelling Ratio of PCL/PEG Networks (with and Without HA)
100PCLDA 100PCLDA/HA 75PCLDA 75PCLDA/HA 50PCLDA 50PCLDA/HA
Gel Fraction (%) (n = 3)
Swelling Ratio (%) (n = 3)
99.69 6 0.18 99.75 6 0.01 99.53 6 0.19 99.59 6 0.36 99 6 0.41 99.43 6 0.31
7.64 6 0.93 6.92 6 2.36 20.38 6 0.47 15.63 6 1.68 45.82 6 2.42 36.03 6 3.41
hydrogel that is not able to swell as much as a network without HA.24 where, DHm is the melting enthalpy associated to the PCL c peak in the DSC thermograms, DHm represents the melting enthalpy of the completely crystalline PCL and wPCL is the weight fraction of PCL in PCL/PEG networks. As shown in Table II, both Tm (378C) and Xc (50.3%) of PCLDA showed lower value than those of PCL diol itself (39.58C and 55%, respectively) due to incorporation of the acrylate group into PCL diol.21 The crosslinked PCL networks showed the lowest Tm (33.58C) and Xc (33.2%), when compared with that of PCL diol and PCLDA because the crystallization of PCL segments was strongly restricted by the polymer network.22 The Xc increased when semicrystalline PCLDA were crosslinked in the presence of HA. This behavior is in agreement with the results obtained by Cai and Wang,20 that could be attributed to HA particles that provided nucleation sites for PCL segments. The Xc of PCL networks decreased after incorporation of PEG chains. Because of this impairment to the crystalline domains, Tm also decreased for PCL/PEG networks.
Gel fraction and swelling studies The gel fraction and swelling ratio of the PCL/PEG networks (with and without HA) were measured and the results are illustrated in Table III. As can be seen, all networks had a gel fraction 99%. A significant difference was not observed in the gel fraction of these networks. The gel fractions of all networks increased when HA was incorporated into the structure, indicating that crosslinking occurred efficiently in the presence of HA and the crosslinked networks had good integrity with HA particles embedded inside.22 The ability of a scaffold to preserve water is an important aspect to be investigated for tissue engineering.23 The swelling ratios of the PCL/PEG networks were observed to be 7.64, 20.38, and 45.82%, respectively, in the order of increasing PEG content. It is reasonable that the network with higher PEG content would demonstrate significantly higher hydrophilicity and thus, a higher swelling ratio value. This result is in agreement with what was reported by Im et al.14 The swelling ratios of all networks decreased after incorporating HA with the polymer networks. This effect may be attributed to the HA that was physically filled in the polymeric network, which in return leads to a compact
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SEM studies Scaffolds for bone tissue engineering must have a highly porous and interconnected pore structure.25 The porous structure of our scaffolds was achieved through a process of salt leaching. The cross-sections of NaCl-leached 75PCLDA (a and b) and 75PCLDA/HA (c and d) scaffolds are shown in the SEM micrographs presented in Figure 5. The highly interconnected porous morphologies with macropores larger than 100 mm and irregular shapes are observed in the 75PCLDA scaffold as well as the 75PCLDA/HA scaffold. Figure 5(b,d) show the porous structures at higher magnification. The porosities of the scaffolds ranged from 73% to 76% (Table IV). All the scaffolds had rather high porosities (over 73%), which was considered to be beneficial to cell growth and survival. Mechanical properties Figure 6 shows the compressive stress–strain curves of PCL/PEG and PCL/PEG/HA networks scaffold in the hydrated state. All curves exhibit the typical characteristics of porous polymeric foam, in which a linear elastic region appears at small strain followed by a plateau region at larger strain and finally a densification region where the stress increased sharply at very large strain. The calculated compressive modulus from the stress–strain data are listed in Table IV. The results show that the compressive modulus decreased significantly after incorporating PEGDA short chains with the PCL networks. Moreover, the compressive modulus decreased significantly with the increasing content of PEGDA. This is most likely due to their lower crystallinity and higher swelling ratio. As shown in Table IV, inclusion of HA increased significantly the compressive modulus from 4.15 to 10.58 MPa for 100PCLDA porous scaffolds and from 1.88 to 4.09 MPa for 75PCLDA porous scaffolds. Inclusion of HA significantly improved the compressive modulus of the PCL and PCL/PEG scaffolds. Because crystallites serving as physical fillers and forming a physical network, 100PCLDA/ HA scaffold with the highest crystallinity had the highest compressive modulus among these scaffolds.20 In vitro degradation In vitro degradation study was carried out to investigate the effect of PEG content and incorporation of HA on the
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FIGURE 5. SEM microphotographs of PCL/PEG network scaffolds. (a, b) 75PCLDA and (c, d) 75PCLDA/HA after leaching out of NaCl particulates. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
degradability of the PCL scaffold. The weight loss of different scaffolds after incubation in distilled water for six months is listed in Table IV. A slight weight loss was observed for the 100PCLDA scaffold, which lost about 4.9% of its initial weight after immersion in distilled water for six months. However, the PCL/PEG networks scaffold presented significantly faster degradation during the 6-month immersion in distilled water. Furthermore, the weight loss of the PCL/PEG scaffolds increased significantly with the increasing amount of PEG in the network. After immersion in distilled water for 6 months, the weight loss for the 75PCLDA and 50PCLDA scaffolds were 15.1% and 18%, respectively. It is known that the degradation of PCL comes from the hydrolysis of ester bonds. The higher hydrophilicity of
50PCLDA scaffolds due to hydrophilic PEG may facilitate water to infiltrate into the scaffolds, resulting in a faster degradation of PCL in the 50PCLDA scaffolds. Moreover, the results show that the weight loss of the PCL/PEG scaffolds decreased by incorporating HA particles in networks, although not statistically significant. This might be attributed to a lower hydrophilicity and a higher degree of crystallinity caused by the presence of HA. The hydrolytic rate of aliphatic polyesters depends primarily on the kinetics of the cleavage of the ester bonds and the degradation is much faster in amorphous domains than in crystalline domains, as
TABLE IV. Compressive Modulus, Weight Loss, and Percent Porosity of PCL/PEG Scaffolds (with and Without HA)
100PCLDA 100PCLDA/HA 75PCLDA 75PCLDA/HA 50PCLDA 50PCLDA/HA
Compressive Modulus (MPa) (n = 3)
Weight Loss After 6 Months (%) (n = 3)
Porosity (%) (n = 3)
4.15 6 1.44 10.58 6 2 1.88 6 0.27 4.09 6 0.77 0.33 6 0.1 0.58 6 0.06
4.9 6 1.7 3.7 6 0.7 15.1 6 3.4 11.7 6 1.1 18 6 2.9 13.6 6 1.5
73.15 6 1.90 74.11 6 1 75.06 6 1.82 75.77 6 1.09 73.8 6 2.25 76.04 6 1.49 FIGURE 6. Compressive stress–strain curves of the different scaffolds.
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FIGURE 7. MG63 cell viability results for various types of scaffolds. Data were used as mean 6 SD for n 5 4; *p < 0.05 compared with control.
FIGURE 8. Alkaline phosphatase activity of MG63 cells seeded onto the different scaffolds; *p < 0.05 compared with the 100PCLDA scaffold; **p < 0.05 compared with the scaffolds without HA.
with control. The ALP activity of cultured MG63 cells on various scaffolds at days 7 and 14 of cell culture is demonstrated in Figure 8. Apparently, for all of the scaffolds investigated, the ALP activity at day 14 was much greater than that at day 7. A higher level of ALP activity was recorded for 75PCLDA/HA scaffolds. The data reveal that addition of 5 wt % HA particles and 7.5 wt % PEG polymers to the PCL scaffolds seem to exert a significant impact on osteoblastic phenotype and enhance cellular responses. The favorite of HA to the stimulation of bone cell response and bone formation has been reported.26 In our case, the possible mechanism for HA and PEG to promote the differentiation of MG63 cells might be ascribed to the increased affinity of MG63 cells to the scaffolds and the improved osteoconductive property of scaffolds because of the incorporation of HA and PEG components. After 14 days of incubation, a significant difference was observed in every case when comparing the expression of alkaline phosphatase of the corresponding scaffolds without HA to the scaffolds with HA. Figure 9 presents SEM images of the morphology features of MG63 cells cultured on the 75PCLDA/HA scaffolds after 6 and 24 hours of culture. It is observed that many nearly round cells are attached to the pore walls and over the scaffolds after 6 h of culture, which is the result of the favorable physicochemical properties and cytocompatibility of scaffolds. After 24 h, the MG63 cells coalesce to form a large and flat layer of cells and cover on the scaffolds. These results suggest that cells could attach and spread on the 75PCLDA/HA scaffolds.
water penetration is easier within a disordered network of polymer chains.18 CONCLUSION
Cell studies An ideal scaffold should not release toxic products or produce adverse reactions, which could be evaluated through in vitro cytotoxic tests. Figure 7 shows the percentage cell viability of MG63 cells after 24 h of incubation with the 14 days extracted medium from various types of scaffolds. The results from the MTT assay of all scaffold extracts showed >95% cell viability after 24 h of incubation. Statistical results showed that there was no significant decrease in the viability values in cells treated with 14 days extracted from 75PCLDA/HA and 50 PCLDA/HA scaffolds in comparison
Biodegradable PEG/PCL networks were synthesized by a radical crosslinking reaction of PEGDA and PCLDA in the presence of sodium chloride salt and hydroxyapatite particulates. From the following salt leaching process, a highly porous structure could be generated in the network. The incorporation of PEG polymer led to an enhanced hydrophilicity of the PCL scaffolds. The results of the in vitro degradation study indicate that the degradation rate of the PCL scaffolds could be significantly increased by increasing the amount of PEG polymer introduced into the PCL matrix. In addition, inclusion of HA significantly improved the
FIGURE 9. SEM photographs of the MG63 cells cultured on the top surface of the 75PCLDA/HA scaffolds. (a) Cells attached to the pore walls and over the scaffolds at 6 h, (b) Higher magnification image showing initial signs of cell spreading, and (c) cell spreading after 24 h. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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compressive modulus of the PCL/PEG scaffolds. 75PCLDA/ HA scaffolds were found to be cytocompatibility in MTT assay. A higher level of ALP activity was also recorded for these scaffolds. Cells seeded on the 75PCLDA/HA scaffolds attached to the pore walls and over the scaffolds and spread overtime. These findings suggest that the 75PCLDA/HA scaffolds may have potential for applications in bone tissue engineering. REFERENCES 1. Peter M, Binulal NS, Soumya S, Nair SV, Furuike T, Tamura H, Jayakumar R. Nanocomposite scaffolds of bioactive glass ceramic nanoparticles disseminated chitosan matrix for tissue engineering applications. J Carbohydr Polym 2010; 79:284–289. 2. Karkhaneh A, Naghizadeh Z, Shokrgozar MA, Bonakdar S, Solouk A, Haghighipour N. Effects of hydrostatic pressure on biosynthetic activity during chondrogenic differentiation of MSCs in hybrid scaffolds. Int J Artif Organs 2014; 36:142–148. 3. Ayala GG, Pace ED, Laurienzo P, Pantalena D, Sommab E, Nobile MR. Poly(E-caprolactone) modified by functional groups: preparation and chemical–physical investigation. Eur Polym J 2009; 45: 3217–3229. 4. Guo Q, Slavov S, Halley PJ. Phase behavior, crystallization, and morphology in thermosetting blends of a biodegradable poly(ethylene glycol)-type epoxy resin and poly(E-caprolactone). J Polym Sci Part B: Polym Phys 2004; 42:2833–2843. 5. Yan J, Zhang Y, Xiao Y, Zhang Y, Lang M. Novel poly(E-caprolactone)s bearing amino groups: Synthesis, characterization and biotinylation. J React Funct Polym 2010; 70:400–407. 6. Chen M, Le DQS, Baatrup A, Nygaard JV, Hein S, Bjerre L, Kassem M, Zou X, B€ unger C. Self-assembled composite matrix in a hierarchical 3-D scaffold for bone tissue engineering. J Acta Biomater 2011; 7:2244–2255. 7. Zhu J. Bioactive modification of poly(ethylene glycol) hydrogels for tissue engineering. J Biomater 2010; 31:4639–4656. 8. Jian CP, Chen YY, Hsieh MF. Biofabrication and in vitro study of hydroxyapatite/mPEG–PCL–mPEG scaffolds for bone tissue engineering using air pressure-aided deposition technology. J Mater Sci Eng C 2013; 33:680–690. 9. Chuenjitkuntaworn B, Inrung W, Damrongsri D, Mekaapiruk K, Supaphol P, Pavasant P. Polycaprolactone/hydroxyapatite composite scaffolds: Preparation, characterization, and in vitro and in vivo biological responses of human primary bone cells. J Biomed Mater Res Part A 2010; 94:241–251. 10. Jiang Y, Mao K, Cai X, Lai S, Chen X. Poly(ethyl glycol) assisting water sorption enhancement of poly(E-caprolactone) blend for drug delivery. J Appl Polym Sci 2011; 122:2309–2316. 11. Cho CS, Han SY, Ha JH, Kim SH, Lim DY. Clonazepam release from bioerodible hydrogels based on semi-interpenetrating polymer networks composed of poly(E-caprolactone) and poly(ethylene glycol) macromer. Int J Pharm 1999; 181:235–242.
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12. Amini AR, Laurencin CT, Nukavarapu1 SP. Bone tissue engineering: Recent advances and challenges. J Crit Rev Biomed Eng 2012; 40:363–408. 13. Kweona HY, Yoob MK, Park IK, Kimb TH, Lee HC, Lee HS, Oh JS, Akaiked T, Cho CS. A novel degradable polycaprolactone networks for tissue engineering. J Biomater 2003; 24:801–808. 14. Im SJ, Choi YM, Subramanyam E, Huh KM. Synthesis and characterization of biodegradable elastic hydrogels based on poly(ethylene glycol) and poly(E-caprolactone) blocks. J Macromol Res 2007; 15:363–369. 15. Neves SC, Teixeira LSM, Moroni L, Reis RL, Blitterswijk CAV, Alves NM, Karperien M, Mano JF. Chitosan/poly(E-caprolactone) blend scaffolds for cartilage repair. J Biomater 2011; 32:1068– 1079. 16. Yang J, Shi G, Bei J, Wang S, Cao Y, Shang Q, Yang G, Wang W. Fabrication and surface modification of macroporous poly(L-lactic acid) and poly(L-lactic-co-glycolic acid) (70/30) cell scaffolds for human skin fibroblast cell culture. J Biomed Mater Res 2002; 62: 438–446. 17. Wu H, Wan Y, Dalai S, Zhang R. Response of rat osteoblasts to polycaprolactone/chitosan blend porous scaffolds. J Biomed Mater Res Part A 2010; 92:238–245. 18. Wan Y, Wu H, Cao X, Dalai S. Compressive mechanical properties and biodegradability of porous poly(caprolactone)/chitosan scaffolds. J Polym Degrad Stabil 2008; 93:1736–1741. 19. Karkhaneh A, Naghizadeh Z, Shokrgozar MA, Bonakdar S. Evaluation of the chondrogenic differentiation of mesenchymal stem cells on hybrid biomimetic scaffolds. J Appl Polym Sci 2014; 131: 40635, doi: 10.1002/app.40635. 20. Cai L, Wang S. Poly(E-caprolactone) acrylates synthesized using a facile method for fabricating networks to achieve controllable physicochemical properties and tunable cell responses. J Polym 2010; 51:164–177. 21. Lee WF, Cheng TS. Studies on preparation and properties of porous biodegradable poly(NIPAAm) hydrogels. J Appl Polym Sci 2008; 109:1982–1992. 22. Cai L, Guinn AS, Wang S. Exposed hydroxyapatite particles on the surface of photo-crosslinked nanocomposites for promoting mc3t3 cell proliferation and differentiation. J Acta Biomater 2011; 7:2185–2199. 23. Prabaharan M, Rodriguez-Perez MA, Saja JA, Mano JF. Preparation and characterization of poly(L-lactic acid)-chitosan hybrid scaffolds with drug release capability. J Biomed Mater Res Part B 2007; 81:427–434. 24. Gaharwar AK, Dammu SA, Canter JM, Wu CJ, Schmidt G. Highly extensible, tough, and elastomeric nanocomposite hydrogels from poly(ethylene glycol) and hydroxyapatite nanoparticles. J Biomacromol 2011; 12:1641–1650. 25. Li Z, Ramay HR, Hauch KD, Xiao D, Zhang M. Chitosan–alginate hybrid scaffolds for bone tissue engineering. J Biomater 2005; 26: 3919–3928. 26. Prabhakaran MP, Venugopal J, Ramakrishna S. Electrospun nanostructured scaffolds for bone tissue engineering. J Acta Biomater 2009; 5:2884–2893.
PCL-CROSSLINKED-PEG HYDROXYAPATITE AS BONE TISSUE ENGINEERING SCAFFOLDS
Department of Biomedical Engineering, Science and Research Branch, Islamic Azad University, Tehran, Iran Department of Biomedical Engineering, Amirkabir University of Technology (Tehran Polytechnic), Tehran, Iran 3 Department of Animal and Marine Biotechnology, National Institute of Genetic Engineering and Biotechnology, Tehran, Iran 2
Received 11 January 2015; revised 10 May 2015; accepted 19 May 2015 Published online 27 August 2015 in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.a.35513 Abstract: In this study, interconnected porous bioactive scaffolds were synthesized for bone tissue engineering. At the first step, poly(E-caprolactone) (PCL) diols were diacrylated with acryloyl chloride. Then, the scaffolds were synthesized by radical crosslinking reaction of PCL and poly(ethyleneglycol) (PEG) diacrylates in the presence of hydroxyapatite (HA) particles. Morphological, swelling, thermal, and mechanical characteristics as well as degradability of the scaffolds were investigated. Results showed that increasing the ratio of PEG to PCL led to significant increase of swelling ratio and degradation rate, and decrease of crystallinity and compressive modulus of the networks, respectively. It was found that the incorporation of HA particles with the polymer matrices resulted in an augmented crystallinity, a decreased swelling ratio, and also a significantly increased compressive modulus
of the networks. Cytocompatability and osteoconductivity of the scaffolds were assessed by MTT and alkaline phosphatase (ALP) assays, respectively. The results confirmed the cytocompatible nature of PCL/PEG/HA scaffolds with no toxicity. MG-63 cells attached and spread on the pore walls offered by the scaffolds. PCL/PEG/HA scaffolds compared with PCL/PEG ones showed higher ALP activity. Thus, the results indicated that the PCL/PEG/HA scaffolds have the potential of being used as promising substrates in bone tisC 2015 Wiley Periodicals, Inc. J Biomed Mater Res sue engineering. V Part A: 103A: 3919–3926, 2015.
Key Words: polycaprolactone, polyethylene glycol, hydroxyapatite, biodegradable porous scaffolds, bone tissue engineering
How to cite this article: Koupaei N, Karkhaneh A, Daliri Joupari M. 2015. Preparation and characterization of (PCL-crosslinkedPEG)/hydroxyapatite as bone tissue engineering scaffolds. J Biomed Mater Res Part A 2015:103A:3919–3926.
INTRODUCTION
Bone tissue engineering is gaining popularity as alternative method for treatment of osseous defects.1 A number of biodegradable polymers have been explored for tissue engineering purposes.1,2 At present, poly(E-caprolactone) (PCL) has been considered for several biomedical applications such as scaffolds for supporting fibroblasts and osteoblasts growth.3 It is hydrophobic and insoluble in water but degradable through the hydrolytic attack of the ester bond.4 Because of its high crystallinity and hydrophilic/hydrophobic balance between ester and methylene groups, PCL has shown remarkably long in vivo degradation times.4,5 Moreover, the lack of bioactive functional groups and intrinsic hydrophobicity result in a poor cell adhesion, which is a critical issue for successful in vitro 3D cell culture and the subsequent tissue formation.6 Therefore, modification of PCL is of great necessity for a wider range of medical applications.3,5 Poly(ethyleneglycol) (PEG) has been an important type of hydrophilic polymers for biomedical applications because they have critical properties, such as good biocompatibility, hydrophilicity, lack of toxicity, and absence of antigenicity and immunogenicity.4,7 Hydroxyapatite (HA) is a
ceramic material with a composition and structure similar to natural bone mineral and has been demonstrated to have good biocompatibility, and even osteoconductivity and osteoinductivity.8 Some studies reported that inclusion of HA into PCL scaffold improved osteoconductivity and produced mineralized nodules.9 However, the hydrophobicity of PCL causes cell attached difficulties which may limit clinical applications.8 Therefore, PEG and HA are good candidates for modifying the property of PCL for use as bone tissue engineering scaffolds. Nowadays, although there are growing interests on the amphiphilic scaffolds based on PCL and a hydrophilic polymer such as PEG in tissue engineering, the most important disadvantage of this system is the release of short chain PEGs from the scaffold in absence of chemical crosslinking.10,11 On the other hand, the bioactivity of the scaffold in bone tissue engineering is a major challenge.12 The aim of this study was to introduce chemically crosslinked biodegradable scaffolds by crosslinking reaction of PCL and PEG diacrylates in the presence of HA. The scaffolds were characterized, MG63 cells were cultured onto the scaffolds, and the morphological and biological studies were carried out.
Correspondence to: A. Karkhaneh; e-mail: [email protected]
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TABLE I. Sample Code and Chemical Composition of PCLCrosslinked PEG Scaffolds
Sample Code
PCLDA/ PEGDA Weight Ratio
HA (wt/vol %)
Total Polymer Concentration (wt/vol %)
100PCLDA 100PCLDA/HA 75PCLDA 75PCLDA/HA 50PCLDA 50PCLDA/HA
100:0 100:0 75:25 75:25 50:50 50:50
0 5 0 5 0 5
30 30 30 30 30 30
(1H-NMR) spectra of PCL-DA were obtained on a Bruker Avance 400 (Bruker Biospin, Germany) spectrometer operating at 400 MHz. Also, attenuated total reflectance-Fourier transform-infrared (ATR-FT-IR, Jasco, Japan) spectra were recorded to analyze the chemical structure of the PCLcrosslinked PEG hydrogels. Differential scanning calorimetry (DSC). The DSC experiments were performed on a Perkin–Elmer Pyris DSC 6 under nitrogen atmosphere. All the samples were first heated from room temperature to 100 C and cooled to 0 C to erase the thermal history.15 Then a subsequent heating run was performed from 0 C to 100 C. The heating and cooling rate was of 10 C=min.
MATERIALS AND METHODS
Materials Poly(E-caprolactone) diol (PCL diol, Mn 5 2000), PEG diacrylate (PEGDA, Mn 5 700), and 2,2-azobisisobutyronitrile (AIBN) as a radical initiator were purchased from SigmaAldrich. Benzene, n-hexane, acryloyl chloride, triethylamine (TEA), and hydroxyapatite (HA) were purchased from Merck. 1,4-Dioxane was from AppliChem. Sodium chloride (NaCl) particulates with size in the range of 105 to 250 mm were used as the progen.
Measurement of gel fraction and swelling. To determine the gel fraction, three film-typed hydrogel samples for each solvent were immersed in excess CH2Cl2 and water for removing sol parts (soluble portion). After two days, undissolved part of samples were taken out and dried in a vacuum oven at 40 C to constant weight. The gel fraction (G) was calculated by the following Eq. (1): G ¼ ðWd =Wi Þ 3 100
Poly(E-caprolactone) diacrylate (PCLDA) synthesis PCL diol was end-capped with acrylate groups to form a polymerizable macromer.13 Briefly, 5 g (2.5 mmol) of PCL diol (Mn 5 2000) was dissolved in 40 mL of benzene and 0.866 mL (6.25 mmol) of triethylamine was added. Then, 0.505 mL (6.25 mmol) of acryloyl chloride, dissolved in a small amount of benzene, was added dropwise with constant magnetic stirring. The reaction mixture was stirred for 3 h at 80 C. Afterwards, it was filtered to remove triethylamine hydrochloride and then the filtrate was precipitated in an excessive amount of n-hexane. Finally, the precipitated product was collected and dried in a vacuum oven for 24 h. Preparation of PCL/PEG networks scaffold PCL/PEG networks were prepared by thermal polymerization method.14 PCL and PEG diacrylates with different weight ratios were dissolved in 1,4-dioxane to have a polymer concentration of 30% w/v. To this solution, the predetermined amounts of AIBN as initiator, HA, and NaCl particulates as porogen were added. Homogeneous PCLDA/ PEGDA/AIBN/HA/NaCl mixture was transferred into test tube and placed in an oven at 70 C for 12 h. The resulting gel was removed from the tube and cut into discs with thicknesses of 3 mm. Afterwards, the salt particulates were leached out by immersing the gel samples in distilled water for 4 days at room temperature, in order to completely remove the porogen. The sample codes and reactant compositions are shown in Table I. Characterization Analysis by FT-IR and 1H NMR. Fourier transform-infrared (FT-IR, Jasco, Japan) spectra were measured to confirm the formation of PCL-DA. 1H nuclear magnetic resonance
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(1)
where Wi is the initial weight of dried gel sample and Wd is the weight of dried gel sample after extraction in solvent. To determine the swelling ratio, the prepared samples were immersed in distilled water at 37 C. After reaching the equilibrium swelling state, the excessive water on the surface was removed by wiping superficially with a filter article. The swelling ratio (S) was calculated using the Eq. (2): S ¼ ½ðWs 2Wd Þ=Wd 3 100
(2)
where Ws and Wd are the weights of swollen and dried samples after extraction in solvent, respectively. The value is expressed as means 6 SD (n 5 3). Morphology observation by SEM. The cross-section morphology of the scaffolds was examined by scanning electron microscope (SEM, KYKY-EM-3200). The scaffolds were fractured in liquid nitrogen and coated with gold. Measurement of porosity. The porosity of the scaffold was determined using a density bottle based on Archimedes’ Principle according to a known technique.16–18 In brief, the porosity of scaffold was calculated by the following Eq. (3): Porosity ð%Þ ¼
ðW2 2W3 2WS Þ=qe 3100 ðW1 2W3 Þ=qe
(3)
where W1 is the weight of density bottle filled with ethanol, W2 the weight of density bottle including ethanol and scaffold, W3 the weight of density bottle taken out ethanolsaturated scaffold from W2, WS the weight of scaffold, qe the density of ethanol, ðW1 2W3 Þ=qe the total volume of the
PCL-CROSSLINKED-PEG HYDROXYAPATITE AS BONE TISSUE ENGINEERING SCAFFOLDS
ORIGINAL ARTICLE
scaffold including pores, and ðW2 2W3 2WS Þ=qe volume in the scaffold.
the pore
Mechanical properties. The compression stress–strain curves of scaffolds were carried out by a universal testing machine (STM-20, Santam, Iran) at a cross-head speed of 1 mm min21. The specimens were stored in distilled water at 37 C for 48 h before test. Three specimens were tested for each sample, and the averages and standard deviations were reported. In vitro degradation. To assess the in vitro degradation behavior of the scaffolds, the samples were weighed (Wi), immersed in an excessive amount of distilled water, and incubated at 37 C with constant shaking for 6 months. Three samples of each group were tested. At preselected time point, the samples were removed from the water, freeze dried, and weighed (Wf). The percentage (%) degradation was calculated using the following formula: Degradation % ¼ ½ðWi 2Wf Þ=Wi 3100
(4)
Cell culture studies Osteosarcoma cell line MG63 (Pasteur Ins, NCBI, Iran) was chosen to evaluate the biological performance of the scaffolds synthesized in this study. Cell lines were cultured in Dulbecco’s Modified Eagle’s Medium (DMEM, GIBCO, Scotland)/supplemented with 10% (v/v) fetal bovine serum (FBS, Seromed, Germany), 100 U mL21 penicillin and 100 lg mL21 streptomycin. Cultured cells were detached by trypsinization, suspended in fresh culture medium, and used for investigating the cytocompatibility of the scaffolds. Before cell seeding, scaffolds were sterilized using UV treatment. The scaffolds were seeded at a density of 1 3 104 cells/scaffold and maintained in DMEM in an incubator at 378C in a humidified incubator with 5% CO2. MTT assay. The cytotoxicity of the scaffolds was determined using the colorimetric MTT assay. Extracts from the scaffolds were prepared by incubating the presterilized scaffolds incubated in culture medium for 14 days at 378C. At the end of this period, the scaffolds were removed, and the socalled extracts were obtained. For the MTT assay, cells were placed into a 96-well plate at a density of 1 3 104 cells/ well and were incubated under standard culturing conditions. After incubation for 24 h, the culture medium was removed and replaced by the extract and incubated for another 24 h, and then 100 mL of 0.5 mg mL21 MTT solution was added into each well. After that, the plate was incubated at 378C for 4 h. Then the medium was removed, 100 mL of isopropanol solution was added and further incubated for 15 min in order to enhance the dissolution of the formazan crystal. The optical density (OD) of formazan in the solution was detected by a multiwall microplate reader (STAT FAX 2100) at 545 nm. For the reference purpose, cells were seeded into a fresh culture medium (control)
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with same seeding condition. The percentage (%) viability was estimated as per the following formula: Viability % ¼ ðODs =ODc Þ3100
(5)
where ODs and ODc are the average optical densities of the sample and the control, respectively.19 Each assay was performed four times. Alkaline phosphatase activity. The osteocoductivity of scaffolds was evaluated by measuring the extraction of alkaline phosphatase (ALP) enzyme as an indicator of osteoblasts activity. The measurement of ALP activity was performed as follows. The cells were cultured on each scaffold in 24-well culture plates. After 7 and 14 days, the concentration of ALP released by cells into the medium in each well was measured via Autoanalyzer (Hitachi 717, Tokyo, Japan) at 405 nm. The culture medium with the same amount of cells was considered as control. Cell morphology on the scaffolds. Morphology and spreading pattern of cells on the scaffolds were evaluated 6 and 24 h after seeding using SEM. For SEM analysis, cell-seeded scaffolds were fixed with 4% glutaraldehyde, rinsed with PBS, dehydrated using graded series of ethanol (50–100%) and dried in vacuum overnight. The dry samples were coated with gold and observed by SEM. Statistical analysis Data were analyzed using one-way analysis of variance (ANOVA) with Tukey’s post hoc honestly significant difference (HSD) test to a confidence level of p < 0.05. RESULTS AND DISCUSSION
Characterization of PCL diacrylate The FT-IR spectra of PCL diol and PCL diacrylate were shown in Figure 1. PCL diacrylate shows absorption bands at 1635 and 813 cm21 assigned to the vinyl (H2C@CHA) groups due to acrylation of PCL diol.13,20 Those peaks were not observed in PCL diol itself [Fig. 1(a)]. The absorption bands at around 1725 and 2950 cm21, which were present in both PCL diol and PCL diacrylate, were attributed to ester carbonyl (AC@O) and methylene groups (ACH2A), respectively. The absorption peak around 3450 cm21 for hydroxyl (AOH) groups in PCL diol became weaker in PCL diacrylate as they were replaced by acrylate groups.13,20 The formation of PCL diacrylate was also confirmed through 1H-NMR spectrometer. As shown in 1H-NMR spectrum (Fig. 2), the vinyl groups of the PCL-DA appeared in the 5.7 to 6.5 ppm range in agreement with literature.14,20 From the above results, the terminal hydroxyl groups in the PCL diol were converted to acrylate groups by the reaction with acryloyl chloride. FT-IR studies of PCL/PEG hydrogel networks Biodegradable PCL/PEG hydrogel networks were synthesized by a radical crosslinking reaction of PEGDA and PCLDA in the presence of hydroxyapatite particulates. The
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FIGURE 3. ATR-FT-IR spectra of (a) 75PCLDA, (b) 75PCLDA/HA networks.
FIGURE 1. FT-IR spectra of (a) PCL diol, (b) PCLDA.
chemical structures of 75PCLDA and 75PCLDA/HA hydrogel networks were characterized using ATR-FT-IR spectra [Fig. 3(a,b)]. No evident peaks were found at 813 and 1635 cm21 for vinyl (H2C@CHA) groups in these networks, suggesting that the carbon-carbon double bonds in PCLDA were largely consumed in the crosslinking. The absorption peak for phosphate (PO4) group of HA at around 1029 cm21 (see Ref. 8) observed in the spectrum of 75PCLDA/HA hydrogel.
FIGURE 2. 1H NMR spectra of (a) PCL diol, (b) PCLDA.
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DSC studies Figure 4 shows DSC thermograms of PCL diol, PCLDA, and PCL/PEG networks (with and without HA) from the heating run. The melting temperature (Tm), heat of fusion (DHm), and crystallinity (Xc) are shown in Table II. Tm for PCL diol and PCLDA were determined using the highest peak temperature among multiple exothermal peaks, which corresponded to the different arm lengths in them.20 The Xc of the PCL was calculated from DH assuming proportionality to the experimental enthalpy. The reported enthalpy of fusion (135.44 J/g) was used as a reference of 100% Xc of PCL.21 The Xc for PCL moiety in the PCL/PEG networks was calculated using the Eq. (5): c Xc % ¼ DHm = wPCL DHm 3 100 (6)
FIGURE 4. DSC curves of PCL diol, PCLDA, and PCL/PEG networks (with or without HA).
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TABLE II. Melting Point, Enthalpy of Fusion, and Crystallinity of PCL Diol, PCLDA, and PCL/PEG Networks (with and Without HA)
PCL diol PCLDA 100PCLDA 100PCLDA/HA 75PCLDA 75PCLDA/HA 50PCLDA 50PCLDA/HA
Tm (8C)
DHm ðJ=gÞ
Xc(%)
39.5 37 33.5 34.3 28.8 29.8 21.5 26.7
74.5 68.2 44.9 40.5 31.3 28.4 19.3 18.7
55 50.3 33.2 34.8 31.8 32.8 27.04 31.4
TABLE III. Gel Fraction and Swelling Ratio of PCL/PEG Networks (with and Without HA)
100PCLDA 100PCLDA/HA 75PCLDA 75PCLDA/HA 50PCLDA 50PCLDA/HA
Gel Fraction (%) (n = 3)
Swelling Ratio (%) (n = 3)
99.69 6 0.18 99.75 6 0.01 99.53 6 0.19 99.59 6 0.36 99 6 0.41 99.43 6 0.31
7.64 6 0.93 6.92 6 2.36 20.38 6 0.47 15.63 6 1.68 45.82 6 2.42 36.03 6 3.41
hydrogel that is not able to swell as much as a network without HA.24 where, DHm is the melting enthalpy associated to the PCL c peak in the DSC thermograms, DHm represents the melting enthalpy of the completely crystalline PCL and wPCL is the weight fraction of PCL in PCL/PEG networks. As shown in Table II, both Tm (378C) and Xc (50.3%) of PCLDA showed lower value than those of PCL diol itself (39.58C and 55%, respectively) due to incorporation of the acrylate group into PCL diol.21 The crosslinked PCL networks showed the lowest Tm (33.58C) and Xc (33.2%), when compared with that of PCL diol and PCLDA because the crystallization of PCL segments was strongly restricted by the polymer network.22 The Xc increased when semicrystalline PCLDA were crosslinked in the presence of HA. This behavior is in agreement with the results obtained by Cai and Wang,20 that could be attributed to HA particles that provided nucleation sites for PCL segments. The Xc of PCL networks decreased after incorporation of PEG chains. Because of this impairment to the crystalline domains, Tm also decreased for PCL/PEG networks.
Gel fraction and swelling studies The gel fraction and swelling ratio of the PCL/PEG networks (with and without HA) were measured and the results are illustrated in Table III. As can be seen, all networks had a gel fraction 99%. A significant difference was not observed in the gel fraction of these networks. The gel fractions of all networks increased when HA was incorporated into the structure, indicating that crosslinking occurred efficiently in the presence of HA and the crosslinked networks had good integrity with HA particles embedded inside.22 The ability of a scaffold to preserve water is an important aspect to be investigated for tissue engineering.23 The swelling ratios of the PCL/PEG networks were observed to be 7.64, 20.38, and 45.82%, respectively, in the order of increasing PEG content. It is reasonable that the network with higher PEG content would demonstrate significantly higher hydrophilicity and thus, a higher swelling ratio value. This result is in agreement with what was reported by Im et al.14 The swelling ratios of all networks decreased after incorporating HA with the polymer networks. This effect may be attributed to the HA that was physically filled in the polymeric network, which in return leads to a compact
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SEM studies Scaffolds for bone tissue engineering must have a highly porous and interconnected pore structure.25 The porous structure of our scaffolds was achieved through a process of salt leaching. The cross-sections of NaCl-leached 75PCLDA (a and b) and 75PCLDA/HA (c and d) scaffolds are shown in the SEM micrographs presented in Figure 5. The highly interconnected porous morphologies with macropores larger than 100 mm and irregular shapes are observed in the 75PCLDA scaffold as well as the 75PCLDA/HA scaffold. Figure 5(b,d) show the porous structures at higher magnification. The porosities of the scaffolds ranged from 73% to 76% (Table IV). All the scaffolds had rather high porosities (over 73%), which was considered to be beneficial to cell growth and survival. Mechanical properties Figure 6 shows the compressive stress–strain curves of PCL/PEG and PCL/PEG/HA networks scaffold in the hydrated state. All curves exhibit the typical characteristics of porous polymeric foam, in which a linear elastic region appears at small strain followed by a plateau region at larger strain and finally a densification region where the stress increased sharply at very large strain. The calculated compressive modulus from the stress–strain data are listed in Table IV. The results show that the compressive modulus decreased significantly after incorporating PEGDA short chains with the PCL networks. Moreover, the compressive modulus decreased significantly with the increasing content of PEGDA. This is most likely due to their lower crystallinity and higher swelling ratio. As shown in Table IV, inclusion of HA increased significantly the compressive modulus from 4.15 to 10.58 MPa for 100PCLDA porous scaffolds and from 1.88 to 4.09 MPa for 75PCLDA porous scaffolds. Inclusion of HA significantly improved the compressive modulus of the PCL and PCL/PEG scaffolds. Because crystallites serving as physical fillers and forming a physical network, 100PCLDA/ HA scaffold with the highest crystallinity had the highest compressive modulus among these scaffolds.20 In vitro degradation In vitro degradation study was carried out to investigate the effect of PEG content and incorporation of HA on the
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FIGURE 5. SEM microphotographs of PCL/PEG network scaffolds. (a, b) 75PCLDA and (c, d) 75PCLDA/HA after leaching out of NaCl particulates. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
degradability of the PCL scaffold. The weight loss of different scaffolds after incubation in distilled water for six months is listed in Table IV. A slight weight loss was observed for the 100PCLDA scaffold, which lost about 4.9% of its initial weight after immersion in distilled water for six months. However, the PCL/PEG networks scaffold presented significantly faster degradation during the 6-month immersion in distilled water. Furthermore, the weight loss of the PCL/PEG scaffolds increased significantly with the increasing amount of PEG in the network. After immersion in distilled water for 6 months, the weight loss for the 75PCLDA and 50PCLDA scaffolds were 15.1% and 18%, respectively. It is known that the degradation of PCL comes from the hydrolysis of ester bonds. The higher hydrophilicity of
50PCLDA scaffolds due to hydrophilic PEG may facilitate water to infiltrate into the scaffolds, resulting in a faster degradation of PCL in the 50PCLDA scaffolds. Moreover, the results show that the weight loss of the PCL/PEG scaffolds decreased by incorporating HA particles in networks, although not statistically significant. This might be attributed to a lower hydrophilicity and a higher degree of crystallinity caused by the presence of HA. The hydrolytic rate of aliphatic polyesters depends primarily on the kinetics of the cleavage of the ester bonds and the degradation is much faster in amorphous domains than in crystalline domains, as
TABLE IV. Compressive Modulus, Weight Loss, and Percent Porosity of PCL/PEG Scaffolds (with and Without HA)
100PCLDA 100PCLDA/HA 75PCLDA 75PCLDA/HA 50PCLDA 50PCLDA/HA
Compressive Modulus (MPa) (n = 3)
Weight Loss After 6 Months (%) (n = 3)
Porosity (%) (n = 3)
4.15 6 1.44 10.58 6 2 1.88 6 0.27 4.09 6 0.77 0.33 6 0.1 0.58 6 0.06
4.9 6 1.7 3.7 6 0.7 15.1 6 3.4 11.7 6 1.1 18 6 2.9 13.6 6 1.5
73.15 6 1.90 74.11 6 1 75.06 6 1.82 75.77 6 1.09 73.8 6 2.25 76.04 6 1.49 FIGURE 6. Compressive stress–strain curves of the different scaffolds.
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FIGURE 7. MG63 cell viability results for various types of scaffolds. Data were used as mean 6 SD for n 5 4; *p < 0.05 compared with control.
FIGURE 8. Alkaline phosphatase activity of MG63 cells seeded onto the different scaffolds; *p < 0.05 compared with the 100PCLDA scaffold; **p < 0.05 compared with the scaffolds without HA.
with control. The ALP activity of cultured MG63 cells on various scaffolds at days 7 and 14 of cell culture is demonstrated in Figure 8. Apparently, for all of the scaffolds investigated, the ALP activity at day 14 was much greater than that at day 7. A higher level of ALP activity was recorded for 75PCLDA/HA scaffolds. The data reveal that addition of 5 wt % HA particles and 7.5 wt % PEG polymers to the PCL scaffolds seem to exert a significant impact on osteoblastic phenotype and enhance cellular responses. The favorite of HA to the stimulation of bone cell response and bone formation has been reported.26 In our case, the possible mechanism for HA and PEG to promote the differentiation of MG63 cells might be ascribed to the increased affinity of MG63 cells to the scaffolds and the improved osteoconductive property of scaffolds because of the incorporation of HA and PEG components. After 14 days of incubation, a significant difference was observed in every case when comparing the expression of alkaline phosphatase of the corresponding scaffolds without HA to the scaffolds with HA. Figure 9 presents SEM images of the morphology features of MG63 cells cultured on the 75PCLDA/HA scaffolds after 6 and 24 hours of culture. It is observed that many nearly round cells are attached to the pore walls and over the scaffolds after 6 h of culture, which is the result of the favorable physicochemical properties and cytocompatibility of scaffolds. After 24 h, the MG63 cells coalesce to form a large and flat layer of cells and cover on the scaffolds. These results suggest that cells could attach and spread on the 75PCLDA/HA scaffolds.
water penetration is easier within a disordered network of polymer chains.18 CONCLUSION
Cell studies An ideal scaffold should not release toxic products or produce adverse reactions, which could be evaluated through in vitro cytotoxic tests. Figure 7 shows the percentage cell viability of MG63 cells after 24 h of incubation with the 14 days extracted medium from various types of scaffolds. The results from the MTT assay of all scaffold extracts showed >95% cell viability after 24 h of incubation. Statistical results showed that there was no significant decrease in the viability values in cells treated with 14 days extracted from 75PCLDA/HA and 50 PCLDA/HA scaffolds in comparison
Biodegradable PEG/PCL networks were synthesized by a radical crosslinking reaction of PEGDA and PCLDA in the presence of sodium chloride salt and hydroxyapatite particulates. From the following salt leaching process, a highly porous structure could be generated in the network. The incorporation of PEG polymer led to an enhanced hydrophilicity of the PCL scaffolds. The results of the in vitro degradation study indicate that the degradation rate of the PCL scaffolds could be significantly increased by increasing the amount of PEG polymer introduced into the PCL matrix. In addition, inclusion of HA significantly improved the
FIGURE 9. SEM photographs of the MG63 cells cultured on the top surface of the 75PCLDA/HA scaffolds. (a) Cells attached to the pore walls and over the scaffolds at 6 h, (b) Higher magnification image showing initial signs of cell spreading, and (c) cell spreading after 24 h. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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compressive modulus of the PCL/PEG scaffolds. 75PCLDA/ HA scaffolds were found to be cytocompatibility in MTT assay. A higher level of ALP activity was also recorded for these scaffolds. Cells seeded on the 75PCLDA/HA scaffolds attached to the pore walls and over the scaffolds and spread overtime. These findings suggest that the 75PCLDA/HA scaffolds may have potential for applications in bone tissue engineering. REFERENCES 1. Peter M, Binulal NS, Soumya S, Nair SV, Furuike T, Tamura H, Jayakumar R. Nanocomposite scaffolds of bioactive glass ceramic nanoparticles disseminated chitosan matrix for tissue engineering applications. J Carbohydr Polym 2010; 79:284–289. 2. Karkhaneh A, Naghizadeh Z, Shokrgozar MA, Bonakdar S, Solouk A, Haghighipour N. Effects of hydrostatic pressure on biosynthetic activity during chondrogenic differentiation of MSCs in hybrid scaffolds. Int J Artif Organs 2014; 36:142–148. 3. Ayala GG, Pace ED, Laurienzo P, Pantalena D, Sommab E, Nobile MR. Poly(E-caprolactone) modified by functional groups: preparation and chemical–physical investigation. Eur Polym J 2009; 45: 3217–3229. 4. Guo Q, Slavov S, Halley PJ. Phase behavior, crystallization, and morphology in thermosetting blends of a biodegradable poly(ethylene glycol)-type epoxy resin and poly(E-caprolactone). J Polym Sci Part B: Polym Phys 2004; 42:2833–2843. 5. Yan J, Zhang Y, Xiao Y, Zhang Y, Lang M. Novel poly(E-caprolactone)s bearing amino groups: Synthesis, characterization and biotinylation. J React Funct Polym 2010; 70:400–407. 6. Chen M, Le DQS, Baatrup A, Nygaard JV, Hein S, Bjerre L, Kassem M, Zou X, B€ unger C. Self-assembled composite matrix in a hierarchical 3-D scaffold for bone tissue engineering. J Acta Biomater 2011; 7:2244–2255. 7. Zhu J. Bioactive modification of poly(ethylene glycol) hydrogels for tissue engineering. J Biomater 2010; 31:4639–4656. 8. Jian CP, Chen YY, Hsieh MF. Biofabrication and in vitro study of hydroxyapatite/mPEG–PCL–mPEG scaffolds for bone tissue engineering using air pressure-aided deposition technology. J Mater Sci Eng C 2013; 33:680–690. 9. Chuenjitkuntaworn B, Inrung W, Damrongsri D, Mekaapiruk K, Supaphol P, Pavasant P. Polycaprolactone/hydroxyapatite composite scaffolds: Preparation, characterization, and in vitro and in vivo biological responses of human primary bone cells. J Biomed Mater Res Part A 2010; 94:241–251. 10. Jiang Y, Mao K, Cai X, Lai S, Chen X. Poly(ethyl glycol) assisting water sorption enhancement of poly(E-caprolactone) blend for drug delivery. J Appl Polym Sci 2011; 122:2309–2316. 11. Cho CS, Han SY, Ha JH, Kim SH, Lim DY. Clonazepam release from bioerodible hydrogels based on semi-interpenetrating polymer networks composed of poly(E-caprolactone) and poly(ethylene glycol) macromer. Int J Pharm 1999; 181:235–242.
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PCL-CROSSLINKED-PEG HYDROXYAPATITE AS BONE TISSUE ENGINEERING SCAFFOLDS
