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Journal of Biomaterials Science, Polymer Edition Publication details, including instructions for authors and subscription information: http://www.tandfonline.com/loi/tbsp20
Solvent-free polymer/bioceramic scaffolds for bone tissue engineering: fabrication, analysis, and cell growth a
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Joshua Minton , Cara Janney , Rosa Akbarzadeh , Carlie Focke , c
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Aswati Subramanian , Tyler Smith , Joseph McKinney , Junyi Liu , b
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James Schmitz , Paul F. James & Azizeh-Mitra Yousefi
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Department of Chemical, Paper and Biomedical Engineering, Miami University, 650 E High Street, Oxford, OH 45056, USA b
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Center for Bone and Mineral Imaging, University of Texas Health Science Center at San Antonio, San Antonio, TX, USA c
Department of Biology, Miami University, Oxford, OH, USA Published online: 02 Sep 2014.
To cite this article: Joshua Minton, Cara Janney, Rosa Akbarzadeh, Carlie Focke, Aswati Subramanian, Tyler Smith, Joseph McKinney, Junyi Liu, James Schmitz, Paul F. James & Azizeh-Mitra Yousefi (2014) Solvent-free polymer/bioceramic scaffolds for bone tissue engineering: fabrication, analysis, and cell growth, Journal of Biomaterials Science, Polymer Edition, 25:16, 1856-1874, DOI: 10.1080/09205063.2014.953016 To link to this article: http://dx.doi.org/10.1080/09205063.2014.953016
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Journal of Biomaterials Science, Polymer Edition, 2014 Vol. 25, No. 16, 1856–1874, http://dx.doi.org/10.1080/09205063.2014.953016
Solvent-free polymer/bioceramic scaffolds for bone tissue engineering: fabrication, analysis, and cell growth Joshua Mintona,1, Cara Janneya, Rosa Akbarzadeha, Carlie Fockea, Aswati Subramanianc, Tyler Smithc, Joseph McKinneyc, Junyi Liua, James Schmitzb, Paul F. Jamesc and Azizeh-Mitra Yousefia*
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Department of Chemical, Paper and Biomedical Engineering, Miami University, 650 E High Street, Oxford, OH 45056, USA; bCenter for Bone and Mineral Imaging, University of Texas Health Science Center at San Antonio, San Antonio, TX, USA; cDepartment of Biology, Miami University, Oxford, OH, USA (Received 21 February 2014; accepted 6 August 2014) This study examines the potential use of porous polycaprolactone (PCL) and polycaprolocatone/hydroxyapatite (PCL/HA) scaffolds fabricated through melt molding and porogen leaching for bone tissue engineering. While eliminating organic solvents is desirable, the process steps proposed in this study for uniformly dispersing HA particles (~5 μm in size) within the scaffold can also contribute to homogeneous properties for these porous composites. Poly(ethylene oxide) (PEO) was chosen as a porogen due to its similar density and melting point as PCL. Pore size of the scaffold was controlled by limiting the size of PCL and PEO particles used in fabrication. The percent of HA in the fabricated scaffolds was quantified by thermogravimetric analysis (TGA). Mechanical testing was used to compare the modulus of the scaffolds to that of bone, and the pore size distribution was examined with microcomputed tomography (μCT). Scanning electron microscopy (SEM) was used to examine the effect on scaffold morphology caused by the addition of HA particles. Both μCT and SEM results showed that HA could be incorporated into PCL scaffolds without negatively affecting scaffold morphology or pore formation. Energy-dispersive X-ray spectroscopy (EDS) and elemental mapping demonstrated a uniform distribution of HA within PCL/HA scaffolds. Murine calvaria-derived MC3T3-E1 cells were used to determine whether cells could attach on scaffolds and grow for up to 21 days. SEM images revealed an increase in cell attachment with the incorporation of HA into the scaffolds. Similarly, DNA content analysis showed a higher cell adhesion to PCL/HA scaffolds. Keywords: bone; polycaprolactone; hydroxyapatite; scaffolds; tissue engineering; cell culture
1. Introduction Bone tissue engineering is a growing field that has the potential to impact the lives of millions of people by providing an alternative to the traditional methods used to treat bone defects. From 1998 to 2005, the number of joint arthroplasties and revision surgeries increased from 700,000 to over 1.1 million in the USA.[1] In 2003, it was estimated that over $20 billion of medical expenses were related to fracture, *Corresponding author. Email: yousefi[email protected] 1 Current affiliation: Associate Scientist, Clopay Plastic Products Co., Mason, OH, USA © 2014 Taylor & Francis
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reattachment, and replacement of hip and knee joints. This was predicted to increase to over $74 billion by 2015.[1,2] Bone is a specialized connective tissue that provides the body with both mechanical support and protection. It is able to undergo some level of regeneration depending on several factors including the size of the defect, nutrition, and vascularization.[3] Several treatment options are available when the body is not able to heal the damaged bone by itself. One of these options is autogenous bone grafting, which involves filling in the defect with bone tissue taken from another location in the body. Although this is the current gold standard, it comes with a risk of complications including donor site morbidity, pain, prolonged hospitalization and rehabilitation, increased risk of deep infection, hematoma, and inflammation.[4–10] An alternative option is allograft bone grafting, which uses bone tissue of other humans to fill in the defect. Possible complications of this procedure include donor-to-recipient infection, disease transmission, and host immune response.[11,12] Bone tissue engineering seeks to improve upon these limitations and involves transplanting a biofactor (such as cells) within a porous biodegradable three-dimensional (3D) scaffold. The scaffold should ideally match the physical properties and mechanical function of the tissue being replaced while having a porous structure that facilitates mass transport and aids in tissue regeneration.[13] The effectiveness of scaffolds for tissue engineering is determined by a number of factors including mechanical properties, pore size, pore interconnectivity, cell attachment, and growth. One of the major challenges in bone tissue engineering is the rapid restoration of tissue biomechanical function.[14] Therefore, mimicking the mechanical properties of native bone would allow the scaffolds to withstand the natural stresses put on them by the body at the implantation site. Porosity, pore size, and pore interconnectivity affect the function of the scaffold by facilitating cell loading and migration as well as the transport of oxygen, nutrients, and metabolic waste. A major obstacle in tissue engineering is that many conventional scaffold fabrication methods result in a random pore structure, inadequate pore size, and poor interconnectivity.[15] In general, a pore size greater than 100 μm has been recommended for bone regeneration (Table 1).[16–23] Scaffold fabrication technique has a critical effect on the microstructure and pore morphology of the final scaffold. Solvent casting/porogen leaching is one of the most widely used methods for scaffold fabrication, which involves casting a polymer solution containing a porogen in a mold, removal of the solvent, followed by porogen leaching.[24] Solvent toxicity is a major concern in this technique, so replacing solvent casting by melt-molding has been considered. Briefly, the melt-molding step consists of premixing polymer powder and porogen particles followed by compression molding at an elevated temperature. Then, the samples are subjected to the same porogen leaching
Table 1.
Optimum pore size range for bone regeneration (modified from Ref. [16]).
Optimum pore size (μm)
Refs.
75–150 100–150 100–250 100–350 200–400 200–600
[17,18] [19] [20] [21] [22] [23]
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step as for solvent-casting.[25] However, the interconnectivity between the pores is directly affected by the spatial arrangement of the porogen particles. At low volume fractions of porogen, typically less than 65% for rigid spheres, the porogen particles that are not in direct contact with other particles may become entrapped in the polymer matrix.[25] To overcome the shortcomings associated with rigid porogens, some studies have examined melt blending of thermoplastic polymers followed by selective dissolution of the polymer that acts as a porogen.[16,26] Melt-blending is usually carried out in a twin-screw extruder to form a two-phase material with micron-sized domains, where a range of blend volume fractions can result in co-continuous networks of polymer and void spaces.[16,26] However, the pore size range obtained by this technique is usually below 100 μm. Incorporation of salt particles into the co-continuous network has been proposed so as to broaden the pore size range [25] although the salt particle broken down during extrusion could limit the control over pore size distribution. In this study a polymer/porogen/bioceramic system composed of two thermoplastic polymers (PCL and PEO) and hydroxyapatite (HA) particles has been examined. The blend was prepared using a vortex mixer and then heated above the melting point of both polymers, causing them to fuse together creating a continuous, interconnected structure. Leaching one of these polymers (PEO) from the structure using water created the interconnected pore network. Final pore size was controlled by limiting the size of the polymer and porogen particles used in this process. The scaffolds containing HA were subjected to additional process steps prior to blending with PEO, to enable a uniform dispersion of HA particles within the scaffolds. The produced scaffolds were seeded with MC3T3-E1 osteoblastic cells, and cell growth was evaluated in vitro for different time points up to 21 days using SEM imaging. In addition, DNA content analysis was performed on the scaffolds one day after cell seeding to quantify cell adhesion. Polycaprolactone (PCL) is a semicrystalline polymer that is biocompatible and biodegradable, and has been approved for some human clinical uses such as implantable devices.[26–28] The PCL homopolymer has a degradation time of 2 years, and PCL structures have been found to be morphologically stable for around 1 year.[28] Poly (ethylene oxide) (PEO) is a synthetic, nontoxic polymer that has been used as excipients and as carriers in pharmaceutical formulations, foods, and cosmetics by the US Food and Drug Administration (FDA). PEO is soluble in water, allowing it to be leached out from a scaffold without the use of toxic solvents, and has a similar density and melting temperature as PCL.[16] By itself, PCL is ductile and may not be ideal for matching the properties of bone tissue. Hence, a composite of PCL and HA has been considered in this study. Bioceramics can facilitate the differentiation of osteoprogenitors as well as the integration with host bone tissue. However, bioceramics are often too stiff and brittle, which limits their use in locations where they can be exposed to significant torsion, bending, or shear stress. Polymer/bioceramic composite scaffolds can combine the benefits of cell interaction and enhanced mineralization with the ease of processing and desirable mechanical strength of synthetic materials.[29–31] HA is a naturally occurring mineral form of calcium apatite, which is the primary mineral component of bone.[32] For bone tissue engineering, an HA concentration of ~20% w/w in the final scaffold appears to be efficient, since at higher concentrations HA might compromise the mechanical strength or the pore size range of the scaffold.[33–35]
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Current challenges in preparing homogeneous polymer/bioceramic systems have limited the commercial growth of these composites materials. Controlling the mixing between two dissimilar phases is a critical issue in the design of these inorganic– organic systems.[36] For example, well-dispersed bioceramic particles in polymer matrix have shown to enhance the compressive and tensile moduli, tensile strength at yield, and ultimate tensile strength more effectively than agglomerated bioceramics,[37] because the polymer/filler interfaces may affect the effectiveness of load transfer between the composite constituents. To improve the polymer/filler interfaces, the dispersion of bioceramic particles into polymer solution followed by subsequent consolidation (solvent casting) has been considered by a number of research groups.[38–41] However, the selection of a suitable solvent, polymer concentration, mixing mode, and stirring time are the key factors for even distribution of bioceramic particles in the final product.[36] In light of this, the present study can also contribute to improving the dispersion of bioceramic particles within a polymer matrix without the use of organic solvents. 2. Materials and methods 2.1. Materials PCL (Mw: 70,000–90,000) and PEO (Mw: 100,000) were purchased from SigmaAldrich. Some properties of these two polymers are listed in Table 2, including the glass transition temperature (Tg), melting point (Tm), and densities at two temperatures. HA with an approximate particle size of 5 μm was supplied by Plasma Biotal Ltd. MC3T3-E1 osteoblastic cells were purchased from Sigma-Aldrich. Table salt was used as a secondary porogen in some experiments. 2.2. Scaffold fabrication Only particles between 250 and 425 μm were used in scaffold fabrication to control the final pore size. The supplied PEO particles were less than 250 μm in size, therefore melt-molded PEO disks were fabricated in stainless steel molds with an internal cavity size of 20 mm (diameter) by 2 mm (depth). The molds were heated at 100 °C for 25 min inside an environmental chamber (CSZ ZPS-16, OH), removed, and cooled at room temperature. Then, the disks were ground in a SPEX SamplePrep 6770 Freezer/ Mill and sieved. PCL was supplied in pellet form, so the pellets were ground directly and sieved. Preliminary experiments were conducted to investigate the effect of scaffold composition on porosity and mechanical properties of the produced scaffolds (Table 3). The
Table 2.
Properties of PCL and PEO. Density (g/cm3)
Materials PCL PEO a
Tg (°C) −69 −67a
Provided by Sigma-Aldrich. From the literature [16].
b
b
Tm (°C) a
60 65a
25 °C
100 °C
a
1.036b 1.07b
1.145 1.13a
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Summary of the scaffold formulations.
Experiment
Scaffold
EXP-1
PCL40 PCL35 PCL30 PCL40S PCL40/HA10 PCL40/HA20 PCL40/HA30
EXP-2 EXP-3
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PEO (% w/w)
HA (% w/w)
NaCl (% w/w)
60 65 70 60 60 60 60
– – – – 10 20 30
– – – 40 – – –
scaffolds used for cell culture were selected based on these preliminary trials. In the first set of experiments (EXP-1 in Table 3), PCL scaffolds were produced with PCL/ PEO compositions of 40:60, 35:65, and 30:70% w/w (denoted as PCL40, PCL35, and PCL30 in Table 3). The rationale for the selected concentrations is that the continuity of the porogen phase can attain ~100% when the porogen content is kept between 60 and 70%.[16] The second set (EXP-2) involved sodium chloride (NaCl) salt particles so as to investigate the effect of additional porogen on scaffold architecture and mechanical properties. Table salt was sieved and particles ranging between 250 and 425 μm in size were incorporated into the PCL/PEO 40:60 blends prior to vortex mixing. The percent of NaCl/PCL was 40% w/w in these experiments, which is denoted as PCL40S in Table 3. In the third set of experiments (EXP-3), PCL/HA scaffolds were produced using the PCL/PEO 40:60 blends and three different HA concentrations: 10, 20, and 30% w/w (denoted as PCL40/HA10, PCL40/HA20, and PCL40/ HA30 in Table 3). Figure 1 shows the schematics of the scaffold fabrication process for the three experiments. In EXP-1, PCL scaffolds were produced by mixing PEO and PCL particles between 250 and 425 μm in size using a vortex mixer. Then ~0.5 g of the blend was placed in each stainless steel mold and heated for 1 h at 100 °C. The molds were then removed and cooled at room temperature. The resulting disks were placed in a bath of deionized water at 40 °C for 24 h to completely leach out PEO, and then airdried for 24 h at room temperature. PCL/PEO/NaCl blends were produced using the same procedure before leaching out PEO and NaCl in deionized water (EXP-2, Figure 1). The supplied HA had a particle size of 5 μm. Therefore, the first step consisted of preparing PCL/HA disks by mixing HA and PCL of particle sizes less than 250 μm using a vortex mixer (EXP-3, Figure 1). The PCL/HA blend was then heated at 100 °C for 1 h in stainless steel molds. The disks were cooled at room temperature, ground, and sieved (250–425 μm). After blending with PEO particles of the same size range, the scaffolds were fabricated using the same procedure as described above. The scaffolds were punched into smaller disks (5-mm in diameter) for scaffold characterization and cell culture trials. 2.3. Scaffold Characterization 2.3.1. Porosity measurement The porosity of the scaffolds was estimated using the following equation:
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Va Vt 100 Va
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where Va is the apparent volume of the scaffold estimated based on the geometry of each disk (thickness and diameter), and Vt is the true volume of each scaffold calculated based on the density of the matrix (qmatrix ) and scaffold mass (m) using Vt ¼ m=qmatrix . For scaffolds containing HA, the rule of mixtures was used to estimate qmatrix . The density of PCL used in the calculations was 1.145 g/cm3 (Table 2), whereas a density of 3.0 g/cm3 was used for HA. 2.3.2. Mechanical testing Scaffolds were tested under unconfined ramp compression using an Instron 3345 Material Testing System with a 1kN load cell. Compression tests to 60% strain were carried out using a displacement rate of 1 mm/min after a pre-load of 4.45 N. Three samples (n = 3) were tested for each scaffold formulation. The Young’s modulus and the modulus at 30% strain were calculated from the stress–strain curve. 2.3.3. Scanning electron microscopy (SEM) To visualize the scaffold morphology and pore structure, the specimens were sputtered with 20 nm gold (Denton Desk II Sputter Unit), and then the SEM images were acquired using a Zeiss Supra 35VP operating at 10 kV with a 7 mm working distance.
Figure 1. Schematics of the scaffold fabrication process (EXP-1, EXP-2, and EXP-3, respectively).
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2.3.4. Energy-dispersive X-ray spectroscopy (EDS)
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Since HA (Ca10(PO4)6(OH)2) contains P and Ca, which are not present in PCL or PEO, the EDS analysis and elemental mapping were used to identify the presence of HA in PCL/HA scaffolds and to determine its dispersion in the final structure. 2.3.5. Microcomputed tomography (μCT) The morphology information for a selected group of scaffolds was gathered using μCT analysis, which was performed at the University of Texas Health Science Center at San Antonio. The scaffolds were scanned in air with the following settings: 40 kV, 250 μA beam intensity, 0.25 rotation step, 4 frame averaging, 800 ms exposure time, and 5 micron voxel size (Bruker SkyScan 1172, Kontich Belgium). A circular region of interest (ROI) of fixed diameter (4 mm) was centered within the scaffold. The pore size distributions were estimated for the ROI, as well as for the whole scaffold. 2.3.6. Thermogravimetric analysis (TGA) To quantitatively verify the final concentration of HA in PCL/HA scaffolds, a specimen was extracted from the scaffold (~15 mg in weight), heated in nitrogen atmosphere to 1000 ˚C at a rate of 2 ˚C/min (TGA Q500, TA Instruments), and the percent of weight loss over time was recorded. 2.4. Cell culture Murine calvaria-derived MC3T3-E1 cells were cultured in α-MEM medium supplemented with 10% fetal bovine serum (FBS, Gemini) and 1% penicillin/streptomycin (Pen/Strep, Gibco) at 37 °C in an incubator with 5% CO2. Media was replaced every 2–3 days and the subconfluent cells were detached using 0.25% trypsin EDTA (Gibco), centrifuged, re-suspended and counted. Scaffolds were sterilized prior to cell culture by soaking in 70% ethanol for 2 h and then washed and left overnight in phosphate-buffered saline (PBS) (HyClone®, Thermo Scientific, USA). The scaffolds were then washed with PBS two more times for 1.5 h each and then left overnight in 1 ml of culture media (MEM-α + 10% FBS + 1% Pen/Strep). The scaffolds were then transferred into 24-well culture plates. Cells were harvested and seeded on the scaffolds at a density of 7 × 105 cells/scaffold. Cells were allowed to adhere to the scaffolds for 24 h and then the scaffolds were transferred to new 24-well plates (denoted as day 0). The media was replaced every 2–3 days and the scaffolds were harvested on day 0, 7, 14, and 21. 2.5. Preparation for SEM imaging Cell-seeded scaffolds were removed from the media, washed twice in PBS, and fixed with 1% glutaraldehyde +2% paraformaldehyde in PBS for 1 h. The scaffolds were then put through serial dehydration in ethanol. In place of critical point drying (CPD), the scaffolds were prepared with hexamethyldisilazane (HMDS),[44] which was deemed more appropriate for PCL-based scaffolds following a series of experimental trials. After the final immersion in 100% ethanol, the scaffolds were immersed in a solution of 2:1 ethanol:HMDS for 30 min followed by a solution of 1:2 ethanol:HMDS
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for 30 min. Finally, the scaffolds were immersed in 100% HMDS three times for 30 min each. The scaffolds were air-dried for 24 h, silver-painted, and sputter-coated with gold. Images were taken using a Zeiss Supra 35 VP FEG scanning electron microscope.
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2.6. Proliferation assay DNA content after cell seeding (day 0) was measured using the CyQUANT cell proliferation assay (Life Technologies).[43] Cell-seeded scaffolds were washed twice with PBS and stored at −80 °C until the assay was performed. Scaffolds were thawed at room temperature and 250 μL of 1x CyQUANT cell-lysis buffer supplemented with 180 mM NaCl, 1 mM EDTA, and 0.75 Kunitz/ml RNAse was added to achieve cell lysis. Samples were sonicated followed by a 1 h incubation at room temperature. After incubation, samples were sonicated a second time and spun down. Then 100, 50, and 25 μL of cell lysate was brought to a total of 100 μL with cell-lysis buffer and mixed with 100 μL of 2x CyQUANT GR dye in lysis buffer in solid black, 96-well microplates. The fluorescence of the samples was measured with a NOVOstar cell-based fast kinetic microplate reader with a 482/50 excitation filter and a 528/20 emission filter. Standard curves were generated using known quantities of MC3T3-E1 cells. 2.7. Statistical analysis All quantitative data are expressed as mean ± standard deviation. The Student’s t-test was used to analyze the statistical significance of differences between the samples. Comparisons were made for two samples at a time, and the p values 10%). Given the lower porosity level for PCL40 (Figure 2(a)), these scaffolds show a considerably higher modulus (slope) than PCL30 and PCL35 in the entire strain range (Figure 2(b)). The trend for the scaffolds containing HA, shown by solid lines, suggests that the use of 30% HA could compromise the mechanical properties of the scaffolds. The stress–strain profile for PCL40 and PCL40/ HA20 is comparable, whereas PCL40/HA10 shows a higher modulus as the strain exceeds 30%. The unexpected results for PCL40S (produced by the PEO/NaCl dual porogen system) could be attributed to the traces of salt remaining in the structure.
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Figure 2. (a) Estimated porosity for the scaffolds (*p < 0.05); (b) comparison between the stress–strain curves obtained for the different scaffolds compositions. The scaffolds containing HA are presented in solid lines.
Similar observations for other scaffold compositions produced by salt-leaching support this hypothesis.[34] Figure 3(a) and (b) summarize the estimated moduli based on the stress–strain data in Figure 2(b). The scaffolds containing 10 and 30% HA appear to reduce the compressive Young’s modulus, estimated from the linear portion of the stress–strain curves (p < 0.05). Nevertheless, these results indicate a minor variation in the Young’s modulus as a function of scaffold composition (ranging between 1 and 2 MPa). Therefore, the modulus at a higher strain level (e.g. 30%) could serve as a more appropriate parameter for comparison purposes. Figure 3(b) shows that the modulus at 30% strain drops with reducing the PCL content (increase in porosity). Moreover, while 20% HA has no significant effect on the modulus at 30% strain, increasing the HA content to 30% drops the modulus significantly (p < 0.05), which is consistent with similar findings reported in the literature for HA contents greater than 20%.[33] This could be attributed to the poor polymer/filler interfaces at high HA contents. Given the osteogenic characteristics of HA, it is desirable to use a high concentration of HA in scaffold formulation, as long as it does not compromise the mechanical properties. Based on the preliminary results, the remaining characterizations were focused on the PCL40, PCL40S, and PCL40/HA20 scaffolds. Table 4 summarizes the results obtained from porosity calculations and mechanical testing for these scaffolds.
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Figure 3. Comparison between (a) the compressive Young’s modulus; and (b) the modulus at 30% strain for the different scaffold formulations (*p < 0.05).
SEM images showed that the use of salt particles increased the pore size of the scaffolds, as evidenced by larger pores observed in Figure 4(b) compared to those in Figure 4(a). In addition, it appeared that 20% HA could be incorporated into the scaffold without negatively affecting scaffold morphology or pore formation (Figure 4(a) and (c)). The EDS analysis confirmed the presence of P and Ca in the final PCL40/ HA20 scaffold (Figure 5(a)), while the elemental mapping identified the location of each element in the final scaffold (Figure 5(b)–(d)). By overlaying these images, it can be shown that HA is uniformly dispersed throughout the scaffold; not present in only isolated sections (Figure 5(e)). The μCT analysis determined the pore size distribution throughout the circular region of interest (ROI) of fixed diameter (4 mm), as shown in Figure 6(a) and (b). Similar to SEM images (Figure 4(a) and (b)), the results in Figure 6(a) suggest that salt Table 4.
Results of the porosity measurements and mechanical testing of the scaffolds.
Scaffold PCL40 PCL40S PCL40/HA20
Porosity (%)
Young’s Modulus (MPa)
Modulus at 30% strain (MPa)
58.4 ± 1.8 64.1 ± 0.5 58.8 ± 2.7
1.68 ± 0.37 1.26 ± 0.28 1.63 ± 0.55
3.66 ± 0.14 5.31 ± 0.30 3.62 ± 0.46
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SEM images of (a) PCL40; (b) PCL40S; and (c) PCL40/HA20 scaffolds.
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Figure 4.
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Figure 5. (a) EDS spectra for a PCL40/HA20 scaffold; (b–d) individual elemental mapping images of O, P, and Ca, respectively; (e) corresponding elemental mapping image of O, P, and Ca overlaid onto a single image.
particles broadened the pore size distribution by increasing the percent of larger pores (>428 μm). Salt particles in direct contact with each other or with PEO particles may have generated these larger pores. As for the effect of HA, Figure 6(b) suggests that
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Figure 6. Comparison between the pore size distribution of (a) PCL40 and PCL40S, and (b) PCL40 and PCL40/HA20 scaffolds. The distribution represents the percent of pores with the indicated pore size range for the region of interest (ROI, 4 mm in diameter).
PCL40/HA20 contained a slightly higher percent of smaller pores (247 μm). Nevertheless, it appears that PCL40/HA20 had approximately 80% of pores greater than 106 μm in size (~92% for PCL40 scaffolds). A summary of these results is provided in Table 5. The corresponding Table 5.
Pore size distribution of PCL40, PCL40S, and PCL40/HA20 scaffolds. ROI (4 mm)
Scaffold PCL40 PCL40S PCL40/HA20
Whole scaffold (5 mm)
>106 μm (%)
247–428 μm (%)
>106 μm (%)
247–428 μm (%)
91.8 90.9 79.5
53.5 44.9 30.2
96.3 95.1 91.8
33.6 40.9 37.3
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results for the whole scaffolds (5 mm in diameter) are also listed in Table 5, which indicate that all the three scaffold formulations had over 90% of pores greater than 106 μm in size. Finally, the TGA analysis of PCL40/HA20 determined a final concentration of HA to be ~17% (Figure 7). The difference between the initial concentration of 20% and the final measured concentration of 17% could be the result of HA being removed during porogen leaching or of human error during the fabrication process. Based on these findings, the PCL40 and PCL40/HA20 scaffolds were chosen for the cell culture trials. 3.2. Cell-seeded scaffolds SEM images of the cell-seeded PCL40 and PCL40/HA20 scaffolds showed evidence of continued cell growth until day 14 and then a decrease was seen at day 21 (Figure 8(a)–(h)). Cells appeared to grow by forming flat, multi-layered structures on
Figure 7.
A TGA thermogram showing the weight loss for a PCL40/HA20 scaffold.
Figure 8. SEM images of (a–d) PCL40 and (e–h) PCL40/HA20 scaffolds at various time points after cell seeding.
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Figure 9. The results of DNA content analysis for PCL40 and PCL40/HA20 scaffolds demonstrating improved cell adhesion for the composite scaffolds (*p < 0.05).
the scaffold. SEM images showed a higher number of cells attaching to the PCL40/ HA20 scaffolds compared to PCL40 scaffolds. Cell bridging of scaffold pores can be observed at day 14 in PCL40/HA20 scaffolds (Figure 8(g)) suggesting that smaller pores could be covered by cells as proliferation continues over time. Therefore, these results suggest that HA enhances the cell attachment and growth on the scaffolds. The results of the DNA content performed on the seeded PCL40 and PCL40/HA20 scaffolds are presented in Figure 9. These results support the SEM results and indicate that approximately 36% of the seeded cells adhered to the PCL40/HA20 scaffolds, whereas the corresponding number for PCL40 scaffolds was significantly lower (~18%). 4. Discussion Effectiveness of the scaffolds used in bone tissue engineering is influenced by their porosity, interconnectivity of the pores, and mechanical properties. Sufficient porosity and pore interconnectivity throughout the scaffold is required to promote cell loading and migration, tissue and vasculature growth, and adequate mass transport. In general, a major drawback with conventional scaffold fabrication methods is that they result in a random pore structure, leading to poor permeability and interconnectivity.[46] Moreover, the majority of scaffold fabrication techniques involve the application of an organic solvent. Although solvent toxicity may not be considered critical for in vitro trials, such approaches will be subject to scrutiny once the technology is advanced to clinical trials.[47] Scaffold fabrication by melt processing may involve the use of two immiscible polymers mixed together, which fuse to form an interconnected structure when exposed to temperatures above the melting point of the two polymers.[34,43] Generally, this cocontinuous structure occurs when there is a 40:60–60:40 composition of the mixture.[45] In this state, both polymers are continuous in the structure. This allows for the formation of an interconnected pore network if one of the polymers can be leached out from the scaffold using a solvent. However, the channels generated by the interpenetrated polymer network are usually too small (100 μm). Pore size was controlled by limiting the size of the PCL and PEO particles used in the fabrication process. Based on SEM imaging and μCT analysis, HA was successfully incorporated into the scaffolds without negatively affecting scaffold morphology or pore formation. SEM images also showed that the incorporation of HA into the scaffolds increased the cell attachment seen over time on scaffolds seeded with MC3T3-E1 cells. These results suggest that the melt molded scaffolds were capable of maintaining cell growth over time. The results of DNA content analysis showed a significant enhancement in cell adhesion for PCL40/HA20 scaffolds. Carefully designed composites of bioceramics and biodegradable polymers may offer adequate mechanical property, desired biocompatibility, and resorption rate, as well as improved tissue interaction and osteoconductivity. Acknowledgments This work was partially supported by the Ohio Board of Regents and the Ohio Third Frontier Program grant entitled: ‘Ohio Research Scholars in Layered Sensing’. The authors wish to thank Dr. Gilbert Pacey for providing support with the TGA measurements and Dr. Roberto Fajardo for the μCT analysis. The authors also acknowledge the technical assistance of Dr. Byran Smucker, Dr. Richard Edelmann, Matt Duley, Doug Hart, Bill Lack, Barry Landrum, Lynn Johnson, Jayson Alexander, Kevin Harris, Ryan Walczak, and the administrative assistance of Laurie Guest.
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[30] Gleeson JP, Brien FJO. Composite scaffolds for orthopaedic regenerative medicine. Adv. Comp. Mater. Med. Nanotechnol. 2011;1–27. doi:10.5772/15033. Attaf B, editor; ISBN 978-953-307-235-7. [31] Shikinami Y, Okuno M. Bioresorbable devices made of forged composites of hydroxyapatite (HA) particles and poly-L-lactide (PLLA): Part I. Basic Characteristics. Biomaterials. 1999;20:859–877. [32] Eshraghi S, Das S. Micromechanical finite-element modeling and experimental characterization of the compressive mechanical properties of polycaprolactone-hydroxyapatite composite scaffolds prepared by selective laser sintering for bone tissue engineering. Acta Biomater. 2012;8:3138–3143. [33] Wagoner Johnson AJ, Herschler BA. A review of the mechanical behavior of cap and cap/ polymer composites for applications in bone replacement and repair. Acta biomaterialia. 2001;7:16–30. [34] Minton J, Janney C, Focke C, Yousefi AM. Design and fabrication of polymer/ceramic scaffolds for bone tissue engineering. In: Proceedings of the SPE-ANTEC technical conference, Cincinnati, OH, April 22–24, 2013; 6 p. [35] Akbarzadeh R, Yousefi AM. Effects of processing parameters in thermally induced phase separation technique on porous architecture of scaffolds for bone tissue engineering. J. Biomed. Mater. Res: Part B-Appl. Biomater. 2014;102:1304–1315. [36] Supová M. Problem of hydroxyapatite dispersion in polymer matrices: a review. J. Mater. Sci. Mater. Med. 2009;20:1201–1213. [37] Liu H, Webster TJ. Mechanical properties of dispersed ceramic nanoparticles in polymer composites for orthopedic applications. Int. J. Nanomedicine. 2010;5:299–313. [38] Wei G, Ma PX. Structure and properties of nano-hydroxyapatite/polymer composite scaffolds for bone tissue engineering. Biomaterials. 2004;25:4749–4757. [39] Mathieu LM, Bourban PE, Månson J-AE. Processing of homogeneous ceramic/polymer blends for bioresorbable composites. Compos. Sci. Technol. 2006;66:1606–1614. [40] Wang M, Chen LJ, Ni J, Weng J, Yue CY. Manufacture and evaluation of bioactive and biodegradable materials and scaffolds for tissue engineering. J. Mater. Sci. Mater. Med. 2001;12:855–860. [41] Zhang R, Ma PX. Poly(α-hydroxyl acids)/hydroxyapatite porous composites for bone-tissue engineering. I. Preparation and morphology. J. Biomed. Mater. Res. 1999;44A:446–455. [42] Yousefi AM, Gauvin C, Sun L, DiRaddo RW, Fernandes J. Design and fabrication of 3Dplotted polymeric scaffolds in functional tissue engineering. Polym. Eng. Sci. 2007;47:608– 618. [43] St-Pierre J-P, Gauthier M, Lefebvre L-P, Tabrizian M. Three-dimensional growth of differentiating MC3T3-E1 pre-osteoblasts on porous titanium scaffolds. Biomaterials. 2005;26:7319–7328. [44] Shively S, Miller WR. The use of HMDS (hexamethyldisilazane) to replace Critical Point Drying (CPD) in the preparation of tardigrades for SEM (Scanning Electron Microscope) imaging. Trans. Kansas Acad. Sci. 2009;112:198–200. [45] Guarino V, Guaccio A, Ambrosio L. Manipulating co-continuous polymer blends to create PCL scaffolds with fully interconnected and anisotropic pore architecture. J. Appl. Biomater. Biomech. 2011;9:34–39. [46] Dhandayuthapani B, Yoshida Y, Maekawa T, Kumar DS. Polymeric scaffolds in tissue engineering application: a review. Int. J. Polym. Sci. 2011;2011:1–19. [47] Dalton PD, Klinkhammer K, Salber J, Klee D, Möller M. Direct in vitro electrospinning with polymer melts. Biomacromolecules. 2006;7:686–690. [48] Meyers MA, Chen P-Y, Lin AY-M, Seki Y. Biological materials: structure and mechanical properties. Prog. Mater. Sci. 2008;53:1–206. [49] Goulet RW, Goldstein SA, Ciarelli MJ, Kuhn JL, Brown MB, Feldkamp LA. The relationship between the structural and orthogonal compressive properties of trabecular bone. J. Biomech. 1994;27:375–389. [50] Athanasiou KA, Zhu C, Lanctot DR, Agrawal CM. Wang X Fundamentals of biomechanics in tissue engineering of bone. Tissue Eng. 2000;6:361–381. [51] Yang S, Leong K, Du Z, Chua C. The design of scaffolds for use in tissue engineering. part i. traditional factors. Tissue Eng. 2001;7:679–689.
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Journal of Biomaterials Science, Polymer Edition Publication details, including instructions for authors and subscription information: http://www.tandfonline.com/loi/tbsp20
Solvent-free polymer/bioceramic scaffolds for bone tissue engineering: fabrication, analysis, and cell growth a
a
a
a
Joshua Minton , Cara Janney , Rosa Akbarzadeh , Carlie Focke , c
c
c
a
Aswati Subramanian , Tyler Smith , Joseph McKinney , Junyi Liu , b
c
James Schmitz , Paul F. James & Azizeh-Mitra Yousefi
a
a
Department of Chemical, Paper and Biomedical Engineering, Miami University, 650 E High Street, Oxford, OH 45056, USA b
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Center for Bone and Mineral Imaging, University of Texas Health Science Center at San Antonio, San Antonio, TX, USA c
Department of Biology, Miami University, Oxford, OH, USA Published online: 02 Sep 2014.
To cite this article: Joshua Minton, Cara Janney, Rosa Akbarzadeh, Carlie Focke, Aswati Subramanian, Tyler Smith, Joseph McKinney, Junyi Liu, James Schmitz, Paul F. James & Azizeh-Mitra Yousefi (2014) Solvent-free polymer/bioceramic scaffolds for bone tissue engineering: fabrication, analysis, and cell growth, Journal of Biomaterials Science, Polymer Edition, 25:16, 1856-1874, DOI: 10.1080/09205063.2014.953016 To link to this article: http://dx.doi.org/10.1080/09205063.2014.953016
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Journal of Biomaterials Science, Polymer Edition, 2014 Vol. 25, No. 16, 1856–1874, http://dx.doi.org/10.1080/09205063.2014.953016
Solvent-free polymer/bioceramic scaffolds for bone tissue engineering: fabrication, analysis, and cell growth Joshua Mintona,1, Cara Janneya, Rosa Akbarzadeha, Carlie Fockea, Aswati Subramanianc, Tyler Smithc, Joseph McKinneyc, Junyi Liua, James Schmitzb, Paul F. Jamesc and Azizeh-Mitra Yousefia*
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a
Department of Chemical, Paper and Biomedical Engineering, Miami University, 650 E High Street, Oxford, OH 45056, USA; bCenter for Bone and Mineral Imaging, University of Texas Health Science Center at San Antonio, San Antonio, TX, USA; cDepartment of Biology, Miami University, Oxford, OH, USA (Received 21 February 2014; accepted 6 August 2014) This study examines the potential use of porous polycaprolactone (PCL) and polycaprolocatone/hydroxyapatite (PCL/HA) scaffolds fabricated through melt molding and porogen leaching for bone tissue engineering. While eliminating organic solvents is desirable, the process steps proposed in this study for uniformly dispersing HA particles (~5 μm in size) within the scaffold can also contribute to homogeneous properties for these porous composites. Poly(ethylene oxide) (PEO) was chosen as a porogen due to its similar density and melting point as PCL. Pore size of the scaffold was controlled by limiting the size of PCL and PEO particles used in fabrication. The percent of HA in the fabricated scaffolds was quantified by thermogravimetric analysis (TGA). Mechanical testing was used to compare the modulus of the scaffolds to that of bone, and the pore size distribution was examined with microcomputed tomography (μCT). Scanning electron microscopy (SEM) was used to examine the effect on scaffold morphology caused by the addition of HA particles. Both μCT and SEM results showed that HA could be incorporated into PCL scaffolds without negatively affecting scaffold morphology or pore formation. Energy-dispersive X-ray spectroscopy (EDS) and elemental mapping demonstrated a uniform distribution of HA within PCL/HA scaffolds. Murine calvaria-derived MC3T3-E1 cells were used to determine whether cells could attach on scaffolds and grow for up to 21 days. SEM images revealed an increase in cell attachment with the incorporation of HA into the scaffolds. Similarly, DNA content analysis showed a higher cell adhesion to PCL/HA scaffolds. Keywords: bone; polycaprolactone; hydroxyapatite; scaffolds; tissue engineering; cell culture
1. Introduction Bone tissue engineering is a growing field that has the potential to impact the lives of millions of people by providing an alternative to the traditional methods used to treat bone defects. From 1998 to 2005, the number of joint arthroplasties and revision surgeries increased from 700,000 to over 1.1 million in the USA.[1] In 2003, it was estimated that over $20 billion of medical expenses were related to fracture, *Corresponding author. Email: yousefi[email protected] 1 Current affiliation: Associate Scientist, Clopay Plastic Products Co., Mason, OH, USA © 2014 Taylor & Francis
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reattachment, and replacement of hip and knee joints. This was predicted to increase to over $74 billion by 2015.[1,2] Bone is a specialized connective tissue that provides the body with both mechanical support and protection. It is able to undergo some level of regeneration depending on several factors including the size of the defect, nutrition, and vascularization.[3] Several treatment options are available when the body is not able to heal the damaged bone by itself. One of these options is autogenous bone grafting, which involves filling in the defect with bone tissue taken from another location in the body. Although this is the current gold standard, it comes with a risk of complications including donor site morbidity, pain, prolonged hospitalization and rehabilitation, increased risk of deep infection, hematoma, and inflammation.[4–10] An alternative option is allograft bone grafting, which uses bone tissue of other humans to fill in the defect. Possible complications of this procedure include donor-to-recipient infection, disease transmission, and host immune response.[11,12] Bone tissue engineering seeks to improve upon these limitations and involves transplanting a biofactor (such as cells) within a porous biodegradable three-dimensional (3D) scaffold. The scaffold should ideally match the physical properties and mechanical function of the tissue being replaced while having a porous structure that facilitates mass transport and aids in tissue regeneration.[13] The effectiveness of scaffolds for tissue engineering is determined by a number of factors including mechanical properties, pore size, pore interconnectivity, cell attachment, and growth. One of the major challenges in bone tissue engineering is the rapid restoration of tissue biomechanical function.[14] Therefore, mimicking the mechanical properties of native bone would allow the scaffolds to withstand the natural stresses put on them by the body at the implantation site. Porosity, pore size, and pore interconnectivity affect the function of the scaffold by facilitating cell loading and migration as well as the transport of oxygen, nutrients, and metabolic waste. A major obstacle in tissue engineering is that many conventional scaffold fabrication methods result in a random pore structure, inadequate pore size, and poor interconnectivity.[15] In general, a pore size greater than 100 μm has been recommended for bone regeneration (Table 1).[16–23] Scaffold fabrication technique has a critical effect on the microstructure and pore morphology of the final scaffold. Solvent casting/porogen leaching is one of the most widely used methods for scaffold fabrication, which involves casting a polymer solution containing a porogen in a mold, removal of the solvent, followed by porogen leaching.[24] Solvent toxicity is a major concern in this technique, so replacing solvent casting by melt-molding has been considered. Briefly, the melt-molding step consists of premixing polymer powder and porogen particles followed by compression molding at an elevated temperature. Then, the samples are subjected to the same porogen leaching
Table 1.
Optimum pore size range for bone regeneration (modified from Ref. [16]).
Optimum pore size (μm)
Refs.
75–150 100–150 100–250 100–350 200–400 200–600
[17,18] [19] [20] [21] [22] [23]
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step as for solvent-casting.[25] However, the interconnectivity between the pores is directly affected by the spatial arrangement of the porogen particles. At low volume fractions of porogen, typically less than 65% for rigid spheres, the porogen particles that are not in direct contact with other particles may become entrapped in the polymer matrix.[25] To overcome the shortcomings associated with rigid porogens, some studies have examined melt blending of thermoplastic polymers followed by selective dissolution of the polymer that acts as a porogen.[16,26] Melt-blending is usually carried out in a twin-screw extruder to form a two-phase material with micron-sized domains, where a range of blend volume fractions can result in co-continuous networks of polymer and void spaces.[16,26] However, the pore size range obtained by this technique is usually below 100 μm. Incorporation of salt particles into the co-continuous network has been proposed so as to broaden the pore size range [25] although the salt particle broken down during extrusion could limit the control over pore size distribution. In this study a polymer/porogen/bioceramic system composed of two thermoplastic polymers (PCL and PEO) and hydroxyapatite (HA) particles has been examined. The blend was prepared using a vortex mixer and then heated above the melting point of both polymers, causing them to fuse together creating a continuous, interconnected structure. Leaching one of these polymers (PEO) from the structure using water created the interconnected pore network. Final pore size was controlled by limiting the size of the polymer and porogen particles used in this process. The scaffolds containing HA were subjected to additional process steps prior to blending with PEO, to enable a uniform dispersion of HA particles within the scaffolds. The produced scaffolds were seeded with MC3T3-E1 osteoblastic cells, and cell growth was evaluated in vitro for different time points up to 21 days using SEM imaging. In addition, DNA content analysis was performed on the scaffolds one day after cell seeding to quantify cell adhesion. Polycaprolactone (PCL) is a semicrystalline polymer that is biocompatible and biodegradable, and has been approved for some human clinical uses such as implantable devices.[26–28] The PCL homopolymer has a degradation time of 2 years, and PCL structures have been found to be morphologically stable for around 1 year.[28] Poly (ethylene oxide) (PEO) is a synthetic, nontoxic polymer that has been used as excipients and as carriers in pharmaceutical formulations, foods, and cosmetics by the US Food and Drug Administration (FDA). PEO is soluble in water, allowing it to be leached out from a scaffold without the use of toxic solvents, and has a similar density and melting temperature as PCL.[16] By itself, PCL is ductile and may not be ideal for matching the properties of bone tissue. Hence, a composite of PCL and HA has been considered in this study. Bioceramics can facilitate the differentiation of osteoprogenitors as well as the integration with host bone tissue. However, bioceramics are often too stiff and brittle, which limits their use in locations where they can be exposed to significant torsion, bending, or shear stress. Polymer/bioceramic composite scaffolds can combine the benefits of cell interaction and enhanced mineralization with the ease of processing and desirable mechanical strength of synthetic materials.[29–31] HA is a naturally occurring mineral form of calcium apatite, which is the primary mineral component of bone.[32] For bone tissue engineering, an HA concentration of ~20% w/w in the final scaffold appears to be efficient, since at higher concentrations HA might compromise the mechanical strength or the pore size range of the scaffold.[33–35]
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Current challenges in preparing homogeneous polymer/bioceramic systems have limited the commercial growth of these composites materials. Controlling the mixing between two dissimilar phases is a critical issue in the design of these inorganic– organic systems.[36] For example, well-dispersed bioceramic particles in polymer matrix have shown to enhance the compressive and tensile moduli, tensile strength at yield, and ultimate tensile strength more effectively than agglomerated bioceramics,[37] because the polymer/filler interfaces may affect the effectiveness of load transfer between the composite constituents. To improve the polymer/filler interfaces, the dispersion of bioceramic particles into polymer solution followed by subsequent consolidation (solvent casting) has been considered by a number of research groups.[38–41] However, the selection of a suitable solvent, polymer concentration, mixing mode, and stirring time are the key factors for even distribution of bioceramic particles in the final product.[36] In light of this, the present study can also contribute to improving the dispersion of bioceramic particles within a polymer matrix without the use of organic solvents. 2. Materials and methods 2.1. Materials PCL (Mw: 70,000–90,000) and PEO (Mw: 100,000) were purchased from SigmaAldrich. Some properties of these two polymers are listed in Table 2, including the glass transition temperature (Tg), melting point (Tm), and densities at two temperatures. HA with an approximate particle size of 5 μm was supplied by Plasma Biotal Ltd. MC3T3-E1 osteoblastic cells were purchased from Sigma-Aldrich. Table salt was used as a secondary porogen in some experiments. 2.2. Scaffold fabrication Only particles between 250 and 425 μm were used in scaffold fabrication to control the final pore size. The supplied PEO particles were less than 250 μm in size, therefore melt-molded PEO disks were fabricated in stainless steel molds with an internal cavity size of 20 mm (diameter) by 2 mm (depth). The molds were heated at 100 °C for 25 min inside an environmental chamber (CSZ ZPS-16, OH), removed, and cooled at room temperature. Then, the disks were ground in a SPEX SamplePrep 6770 Freezer/ Mill and sieved. PCL was supplied in pellet form, so the pellets were ground directly and sieved. Preliminary experiments were conducted to investigate the effect of scaffold composition on porosity and mechanical properties of the produced scaffolds (Table 3). The
Table 2.
Properties of PCL and PEO. Density (g/cm3)
Materials PCL PEO a
Tg (°C) −69 −67a
Provided by Sigma-Aldrich. From the literature [16].
b
b
Tm (°C) a
60 65a
25 °C
100 °C
a
1.036b 1.07b
1.145 1.13a
1860 Table 3.
Summary of the scaffold formulations.
Experiment
Scaffold
EXP-1
PCL40 PCL35 PCL30 PCL40S PCL40/HA10 PCL40/HA20 PCL40/HA30
EXP-2 EXP-3
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PEO (% w/w)
HA (% w/w)
NaCl (% w/w)
60 65 70 60 60 60 60
– – – – 10 20 30
– – – 40 – – –
scaffolds used for cell culture were selected based on these preliminary trials. In the first set of experiments (EXP-1 in Table 3), PCL scaffolds were produced with PCL/ PEO compositions of 40:60, 35:65, and 30:70% w/w (denoted as PCL40, PCL35, and PCL30 in Table 3). The rationale for the selected concentrations is that the continuity of the porogen phase can attain ~100% when the porogen content is kept between 60 and 70%.[16] The second set (EXP-2) involved sodium chloride (NaCl) salt particles so as to investigate the effect of additional porogen on scaffold architecture and mechanical properties. Table salt was sieved and particles ranging between 250 and 425 μm in size were incorporated into the PCL/PEO 40:60 blends prior to vortex mixing. The percent of NaCl/PCL was 40% w/w in these experiments, which is denoted as PCL40S in Table 3. In the third set of experiments (EXP-3), PCL/HA scaffolds were produced using the PCL/PEO 40:60 blends and three different HA concentrations: 10, 20, and 30% w/w (denoted as PCL40/HA10, PCL40/HA20, and PCL40/ HA30 in Table 3). Figure 1 shows the schematics of the scaffold fabrication process for the three experiments. In EXP-1, PCL scaffolds were produced by mixing PEO and PCL particles between 250 and 425 μm in size using a vortex mixer. Then ~0.5 g of the blend was placed in each stainless steel mold and heated for 1 h at 100 °C. The molds were then removed and cooled at room temperature. The resulting disks were placed in a bath of deionized water at 40 °C for 24 h to completely leach out PEO, and then airdried for 24 h at room temperature. PCL/PEO/NaCl blends were produced using the same procedure before leaching out PEO and NaCl in deionized water (EXP-2, Figure 1). The supplied HA had a particle size of 5 μm. Therefore, the first step consisted of preparing PCL/HA disks by mixing HA and PCL of particle sizes less than 250 μm using a vortex mixer (EXP-3, Figure 1). The PCL/HA blend was then heated at 100 °C for 1 h in stainless steel molds. The disks were cooled at room temperature, ground, and sieved (250–425 μm). After blending with PEO particles of the same size range, the scaffolds were fabricated using the same procedure as described above. The scaffolds were punched into smaller disks (5-mm in diameter) for scaffold characterization and cell culture trials. 2.3. Scaffold Characterization 2.3.1. Porosity measurement The porosity of the scaffolds was estimated using the following equation:
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Va Vt 100 Va
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where Va is the apparent volume of the scaffold estimated based on the geometry of each disk (thickness and diameter), and Vt is the true volume of each scaffold calculated based on the density of the matrix (qmatrix ) and scaffold mass (m) using Vt ¼ m=qmatrix . For scaffolds containing HA, the rule of mixtures was used to estimate qmatrix . The density of PCL used in the calculations was 1.145 g/cm3 (Table 2), whereas a density of 3.0 g/cm3 was used for HA. 2.3.2. Mechanical testing Scaffolds were tested under unconfined ramp compression using an Instron 3345 Material Testing System with a 1kN load cell. Compression tests to 60% strain were carried out using a displacement rate of 1 mm/min after a pre-load of 4.45 N. Three samples (n = 3) were tested for each scaffold formulation. The Young’s modulus and the modulus at 30% strain were calculated from the stress–strain curve. 2.3.3. Scanning electron microscopy (SEM) To visualize the scaffold morphology and pore structure, the specimens were sputtered with 20 nm gold (Denton Desk II Sputter Unit), and then the SEM images were acquired using a Zeiss Supra 35VP operating at 10 kV with a 7 mm working distance.
Figure 1. Schematics of the scaffold fabrication process (EXP-1, EXP-2, and EXP-3, respectively).
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2.3.4. Energy-dispersive X-ray spectroscopy (EDS)
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Since HA (Ca10(PO4)6(OH)2) contains P and Ca, which are not present in PCL or PEO, the EDS analysis and elemental mapping were used to identify the presence of HA in PCL/HA scaffolds and to determine its dispersion in the final structure. 2.3.5. Microcomputed tomography (μCT) The morphology information for a selected group of scaffolds was gathered using μCT analysis, which was performed at the University of Texas Health Science Center at San Antonio. The scaffolds were scanned in air with the following settings: 40 kV, 250 μA beam intensity, 0.25 rotation step, 4 frame averaging, 800 ms exposure time, and 5 micron voxel size (Bruker SkyScan 1172, Kontich Belgium). A circular region of interest (ROI) of fixed diameter (4 mm) was centered within the scaffold. The pore size distributions were estimated for the ROI, as well as for the whole scaffold. 2.3.6. Thermogravimetric analysis (TGA) To quantitatively verify the final concentration of HA in PCL/HA scaffolds, a specimen was extracted from the scaffold (~15 mg in weight), heated in nitrogen atmosphere to 1000 ˚C at a rate of 2 ˚C/min (TGA Q500, TA Instruments), and the percent of weight loss over time was recorded. 2.4. Cell culture Murine calvaria-derived MC3T3-E1 cells were cultured in α-MEM medium supplemented with 10% fetal bovine serum (FBS, Gemini) and 1% penicillin/streptomycin (Pen/Strep, Gibco) at 37 °C in an incubator with 5% CO2. Media was replaced every 2–3 days and the subconfluent cells were detached using 0.25% trypsin EDTA (Gibco), centrifuged, re-suspended and counted. Scaffolds were sterilized prior to cell culture by soaking in 70% ethanol for 2 h and then washed and left overnight in phosphate-buffered saline (PBS) (HyClone®, Thermo Scientific, USA). The scaffolds were then washed with PBS two more times for 1.5 h each and then left overnight in 1 ml of culture media (MEM-α + 10% FBS + 1% Pen/Strep). The scaffolds were then transferred into 24-well culture plates. Cells were harvested and seeded on the scaffolds at a density of 7 × 105 cells/scaffold. Cells were allowed to adhere to the scaffolds for 24 h and then the scaffolds were transferred to new 24-well plates (denoted as day 0). The media was replaced every 2–3 days and the scaffolds were harvested on day 0, 7, 14, and 21. 2.5. Preparation for SEM imaging Cell-seeded scaffolds were removed from the media, washed twice in PBS, and fixed with 1% glutaraldehyde +2% paraformaldehyde in PBS for 1 h. The scaffolds were then put through serial dehydration in ethanol. In place of critical point drying (CPD), the scaffolds were prepared with hexamethyldisilazane (HMDS),[44] which was deemed more appropriate for PCL-based scaffolds following a series of experimental trials. After the final immersion in 100% ethanol, the scaffolds were immersed in a solution of 2:1 ethanol:HMDS for 30 min followed by a solution of 1:2 ethanol:HMDS
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for 30 min. Finally, the scaffolds were immersed in 100% HMDS three times for 30 min each. The scaffolds were air-dried for 24 h, silver-painted, and sputter-coated with gold. Images were taken using a Zeiss Supra 35 VP FEG scanning electron microscope.
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2.6. Proliferation assay DNA content after cell seeding (day 0) was measured using the CyQUANT cell proliferation assay (Life Technologies).[43] Cell-seeded scaffolds were washed twice with PBS and stored at −80 °C until the assay was performed. Scaffolds were thawed at room temperature and 250 μL of 1x CyQUANT cell-lysis buffer supplemented with 180 mM NaCl, 1 mM EDTA, and 0.75 Kunitz/ml RNAse was added to achieve cell lysis. Samples were sonicated followed by a 1 h incubation at room temperature. After incubation, samples were sonicated a second time and spun down. Then 100, 50, and 25 μL of cell lysate was brought to a total of 100 μL with cell-lysis buffer and mixed with 100 μL of 2x CyQUANT GR dye in lysis buffer in solid black, 96-well microplates. The fluorescence of the samples was measured with a NOVOstar cell-based fast kinetic microplate reader with a 482/50 excitation filter and a 528/20 emission filter. Standard curves were generated using known quantities of MC3T3-E1 cells. 2.7. Statistical analysis All quantitative data are expressed as mean ± standard deviation. The Student’s t-test was used to analyze the statistical significance of differences between the samples. Comparisons were made for two samples at a time, and the p values 10%). Given the lower porosity level for PCL40 (Figure 2(a)), these scaffolds show a considerably higher modulus (slope) than PCL30 and PCL35 in the entire strain range (Figure 2(b)). The trend for the scaffolds containing HA, shown by solid lines, suggests that the use of 30% HA could compromise the mechanical properties of the scaffolds. The stress–strain profile for PCL40 and PCL40/ HA20 is comparable, whereas PCL40/HA10 shows a higher modulus as the strain exceeds 30%. The unexpected results for PCL40S (produced by the PEO/NaCl dual porogen system) could be attributed to the traces of salt remaining in the structure.
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Figure 2. (a) Estimated porosity for the scaffolds (*p < 0.05); (b) comparison between the stress–strain curves obtained for the different scaffolds compositions. The scaffolds containing HA are presented in solid lines.
Similar observations for other scaffold compositions produced by salt-leaching support this hypothesis.[34] Figure 3(a) and (b) summarize the estimated moduli based on the stress–strain data in Figure 2(b). The scaffolds containing 10 and 30% HA appear to reduce the compressive Young’s modulus, estimated from the linear portion of the stress–strain curves (p < 0.05). Nevertheless, these results indicate a minor variation in the Young’s modulus as a function of scaffold composition (ranging between 1 and 2 MPa). Therefore, the modulus at a higher strain level (e.g. 30%) could serve as a more appropriate parameter for comparison purposes. Figure 3(b) shows that the modulus at 30% strain drops with reducing the PCL content (increase in porosity). Moreover, while 20% HA has no significant effect on the modulus at 30% strain, increasing the HA content to 30% drops the modulus significantly (p < 0.05), which is consistent with similar findings reported in the literature for HA contents greater than 20%.[33] This could be attributed to the poor polymer/filler interfaces at high HA contents. Given the osteogenic characteristics of HA, it is desirable to use a high concentration of HA in scaffold formulation, as long as it does not compromise the mechanical properties. Based on the preliminary results, the remaining characterizations were focused on the PCL40, PCL40S, and PCL40/HA20 scaffolds. Table 4 summarizes the results obtained from porosity calculations and mechanical testing for these scaffolds.
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Figure 3. Comparison between (a) the compressive Young’s modulus; and (b) the modulus at 30% strain for the different scaffold formulations (*p < 0.05).
SEM images showed that the use of salt particles increased the pore size of the scaffolds, as evidenced by larger pores observed in Figure 4(b) compared to those in Figure 4(a). In addition, it appeared that 20% HA could be incorporated into the scaffold without negatively affecting scaffold morphology or pore formation (Figure 4(a) and (c)). The EDS analysis confirmed the presence of P and Ca in the final PCL40/ HA20 scaffold (Figure 5(a)), while the elemental mapping identified the location of each element in the final scaffold (Figure 5(b)–(d)). By overlaying these images, it can be shown that HA is uniformly dispersed throughout the scaffold; not present in only isolated sections (Figure 5(e)). The μCT analysis determined the pore size distribution throughout the circular region of interest (ROI) of fixed diameter (4 mm), as shown in Figure 6(a) and (b). Similar to SEM images (Figure 4(a) and (b)), the results in Figure 6(a) suggest that salt Table 4.
Results of the porosity measurements and mechanical testing of the scaffolds.
Scaffold PCL40 PCL40S PCL40/HA20
Porosity (%)
Young’s Modulus (MPa)
Modulus at 30% strain (MPa)
58.4 ± 1.8 64.1 ± 0.5 58.8 ± 2.7
1.68 ± 0.37 1.26 ± 0.28 1.63 ± 0.55
3.66 ± 0.14 5.31 ± 0.30 3.62 ± 0.46
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SEM images of (a) PCL40; (b) PCL40S; and (c) PCL40/HA20 scaffolds.
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Figure 4.
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Figure 5. (a) EDS spectra for a PCL40/HA20 scaffold; (b–d) individual elemental mapping images of O, P, and Ca, respectively; (e) corresponding elemental mapping image of O, P, and Ca overlaid onto a single image.
particles broadened the pore size distribution by increasing the percent of larger pores (>428 μm). Salt particles in direct contact with each other or with PEO particles may have generated these larger pores. As for the effect of HA, Figure 6(b) suggests that
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Figure 6. Comparison between the pore size distribution of (a) PCL40 and PCL40S, and (b) PCL40 and PCL40/HA20 scaffolds. The distribution represents the percent of pores with the indicated pore size range for the region of interest (ROI, 4 mm in diameter).
PCL40/HA20 contained a slightly higher percent of smaller pores (247 μm). Nevertheless, it appears that PCL40/HA20 had approximately 80% of pores greater than 106 μm in size (~92% for PCL40 scaffolds). A summary of these results is provided in Table 5. The corresponding Table 5.
Pore size distribution of PCL40, PCL40S, and PCL40/HA20 scaffolds. ROI (4 mm)
Scaffold PCL40 PCL40S PCL40/HA20
Whole scaffold (5 mm)
>106 μm (%)
247–428 μm (%)
>106 μm (%)
247–428 μm (%)
91.8 90.9 79.5
53.5 44.9 30.2
96.3 95.1 91.8
33.6 40.9 37.3
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results for the whole scaffolds (5 mm in diameter) are also listed in Table 5, which indicate that all the three scaffold formulations had over 90% of pores greater than 106 μm in size. Finally, the TGA analysis of PCL40/HA20 determined a final concentration of HA to be ~17% (Figure 7). The difference between the initial concentration of 20% and the final measured concentration of 17% could be the result of HA being removed during porogen leaching or of human error during the fabrication process. Based on these findings, the PCL40 and PCL40/HA20 scaffolds were chosen for the cell culture trials. 3.2. Cell-seeded scaffolds SEM images of the cell-seeded PCL40 and PCL40/HA20 scaffolds showed evidence of continued cell growth until day 14 and then a decrease was seen at day 21 (Figure 8(a)–(h)). Cells appeared to grow by forming flat, multi-layered structures on
Figure 7.
A TGA thermogram showing the weight loss for a PCL40/HA20 scaffold.
Figure 8. SEM images of (a–d) PCL40 and (e–h) PCL40/HA20 scaffolds at various time points after cell seeding.
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Figure 9. The results of DNA content analysis for PCL40 and PCL40/HA20 scaffolds demonstrating improved cell adhesion for the composite scaffolds (*p < 0.05).
the scaffold. SEM images showed a higher number of cells attaching to the PCL40/ HA20 scaffolds compared to PCL40 scaffolds. Cell bridging of scaffold pores can be observed at day 14 in PCL40/HA20 scaffolds (Figure 8(g)) suggesting that smaller pores could be covered by cells as proliferation continues over time. Therefore, these results suggest that HA enhances the cell attachment and growth on the scaffolds. The results of the DNA content performed on the seeded PCL40 and PCL40/HA20 scaffolds are presented in Figure 9. These results support the SEM results and indicate that approximately 36% of the seeded cells adhered to the PCL40/HA20 scaffolds, whereas the corresponding number for PCL40 scaffolds was significantly lower (~18%). 4. Discussion Effectiveness of the scaffolds used in bone tissue engineering is influenced by their porosity, interconnectivity of the pores, and mechanical properties. Sufficient porosity and pore interconnectivity throughout the scaffold is required to promote cell loading and migration, tissue and vasculature growth, and adequate mass transport. In general, a major drawback with conventional scaffold fabrication methods is that they result in a random pore structure, leading to poor permeability and interconnectivity.[46] Moreover, the majority of scaffold fabrication techniques involve the application of an organic solvent. Although solvent toxicity may not be considered critical for in vitro trials, such approaches will be subject to scrutiny once the technology is advanced to clinical trials.[47] Scaffold fabrication by melt processing may involve the use of two immiscible polymers mixed together, which fuse to form an interconnected structure when exposed to temperatures above the melting point of the two polymers.[34,43] Generally, this cocontinuous structure occurs when there is a 40:60–60:40 composition of the mixture.[45] In this state, both polymers are continuous in the structure. This allows for the formation of an interconnected pore network if one of the polymers can be leached out from the scaffold using a solvent. However, the channels generated by the interpenetrated polymer network are usually too small (100 μm). Pore size was controlled by limiting the size of the PCL and PEO particles used in the fabrication process. Based on SEM imaging and μCT analysis, HA was successfully incorporated into the scaffolds without negatively affecting scaffold morphology or pore formation. SEM images also showed that the incorporation of HA into the scaffolds increased the cell attachment seen over time on scaffolds seeded with MC3T3-E1 cells. These results suggest that the melt molded scaffolds were capable of maintaining cell growth over time. The results of DNA content analysis showed a significant enhancement in cell adhesion for PCL40/HA20 scaffolds. Carefully designed composites of bioceramics and biodegradable polymers may offer adequate mechanical property, desired biocompatibility, and resorption rate, as well as improved tissue interaction and osteoconductivity. Acknowledgments This work was partially supported by the Ohio Board of Regents and the Ohio Third Frontier Program grant entitled: ‘Ohio Research Scholars in Layered Sensing’. The authors wish to thank Dr. Gilbert Pacey for providing support with the TGA measurements and Dr. Roberto Fajardo for the μCT analysis. The authors also acknowledge the technical assistance of Dr. Byran Smucker, Dr. Richard Edelmann, Matt Duley, Doug Hart, Bill Lack, Barry Landrum, Lynn Johnson, Jayson Alexander, Kevin Harris, Ryan Walczak, and the administrative assistance of Laurie Guest.
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