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Fiber diameter and seeding density influence chondrogenic differentiation of mesenchymal stem cells seeded on electrospun poly(-caprolactone) scaffolds
This content has been downloaded from IOPscience. Please scroll down to see the full text. 2015 Biomed. Mater. 10 015018 (http://iopscience.iop.org/1748-605X/10/1/015018) View the table of contents for this issue, or go to the journal homepage for more
Download details: IP Address: 129.2.19.100 This content was downloaded on 02/02/2015 at 15:01
Please note that terms and conditions apply.
Biomed. Mater. 10 (2015) 015018
doi:10.1088/1748-6041/10/1/015018
Paper
received
10 September 2014 re vised
14 December 2014
Fiber diameter and seeding density influence chondrogenic differentiation of mesenchymal stem cells seeded on electrospun poly(ε-caprolactone) scaffolds
accep ted for publication
5 January 2015 published
29 January 2015
Allison C Bean and Rocky S Tuan Center for Cellular and Molecular Engineering, Department of Orthopaedic Surgery, University of Pittsburgh School of Medicine, 450 Technology Drive, Room 221 Pittsburgh, PA 15219 USA E-mail: [email protected] Keywords: mesenchymal stem cells, chondrogenesis, electrospinning, fibrous scaffolds, biomaterials
Abstract Chondrogenic differentiation of mesenchymal stem cells is strongly influenced by the surrounding chemical and structural milieu. Since the majority of the native cartilage extracellular matrix is composed of nanofibrous collagen fibrils, much of recent cartilage tissue engineering research has focused on developing and utilizing scaffolds with similar nanoscale architecture. However, current literature lacks consensus regarding the ideal fiber diameter, with differences in culture conditions making it difficult to compare between studies. Here, we aimed to develop a more thorough understanding of how cell–cell and cell-biomaterial interactions drive in vitro chondrogenic differentiation of bone-marrow-derived mesenchymal stem cells (MSCs). Electrospun poly(εcaprolactone) microfibers (4.3 ± 0.8 µm diameter, 90 μm2 pore size) and nanofibers (440 ± 20 nm diameter, 1.2 μm2 pore size) were seeded with MSCs at initial densities ranging from 1 × 105 to 4 × 106 cells cm−3-scaffold and cultured under transforming growth factor-β (TGF-β) induced chondrogenic conditions for 3 or 6 weeks. Chondrogenic gene expression, cellular proliferation, as well as sulfated glycosaminoglycan and collagen production were enhanced on microfiber in comparison to nanofiber scaffolds, with high initial seeding densities being required for significant chondrogenic differentiation and extracellular matrix deposition. Both cell–cell and cell–material interactions appear to play important roles in chondrogenic differentiation of MSCs in vitro and consideration of several variables simultaneously is essential for understanding cell behavior in order to develop an optimal tissue engineering strategy.
1. Introduction Conservative estimates find that nearly 14% of the United States population over the age of 25 is affected by osteoarthritis and as the population continues to age, the prevalence is predicted to increase [1, 2]. The limited intrinsic healing capacity of articular cartilage and lack of successful reparative or regenerative techniques has left total joint arthroplasty as the mainstay of treatment for severe osteoarthritis over the past several decades. While relatively successful in reducing pain and improving overall quality of life [3–5], patients undergoing total joint procedures still suffer from decreased functional abilities in comparison to healthy individuals [3]. Furthermore, the failure rate of implants is 6% for 5 years and 12% for 10 years [6] and requires another surgery for revision. Due to © 2015 IOP Publishing Ltd
the large health and economic burden and limited efficacy of current treatment methods, development of a functional tissue engineered cartilage construct may lead to a significant improvement in outcome for patients with osteoarthritis. An often practiced paradigm for tissue engineering involves seeding cells onto biocompatible scaffolds and providing biochemical and/or biomechanical cues that drive cells toward a desired phenotype and stimulate them to secrete extracellular matrix (ECM) proteins to form a new and functional tissue replacement. In the native tissue environment, the complex and dynamic interactions that occur between cells and the ECM directly control tissue structure and composition. The ECM provides not only the tissue’s mechanical strength, but also essential topographical and biochemical cues that direct cell behaviors, including migration,
IOP Publishing
Biomed. Mater. 10 (2015) 015018
A C Bean and R S Tuan
proliferation and differentiation [7]. Therefore, utilizing scaffolds that mimic the native ECM architecture may improve tissue formation in vitro. In articular cartilage, collagen makes up nearly 60% of the dry weight of the tissue and plays an important role in providing the mechanical properties needed in a load-bearing tissue [8]. The fibrils found in articular cartilage typically range from 40 to 640 nm in diameter [9] and therefore many recent investigations have focused on utilizing nanofibrous scaffolds derived either from synthetic or natural materials that have fiber diameters of less than 1 μm. However, it is currently unclear whether nanofibers provide optimal scaffold architecture or if other experimental factors are more important for successful cartilage tissue engineering. Therefore, examination of cell behavior on scaffolds with different fiber diameters in conjunction with other experimental variables may enhance our understanding of what conditions are ideal for generating a tissue engineered cartilage construct. Electrospinning is an inexpensive and versatile technique that has become popular for fabrication of nanofiber and microfiber scaffolds. Several studies have examined the effects of electrospun scaffold fiber diameter on different cell types and tissue applications including bone [10–13], nerve [14–17], ligament [18, 19], skin [20, 21], blood vessels [22] and cartilage [23– 26]. The results of these studies suggest that the optimal fiber diameter may be highly dependent on various experimental factors and it may be impossible to isolate a single ideal fiber diameter independent of other conditions. For example, one study concluded that differentiation of mesenchymal stem cells (MSCs) toward a cartilage phenotype was better on electrospun poly(εcaprolactone) (PCL) fibers that were 500 nm in diameter than on fibers 3 μm in diameter [24], while another group concluded that chondrogenesis was enhanced on poly(L-lactide) fibers with diameters of 5 or 9 μm in comparison to those between 300 nm and 1,400 nm in diameter [23]. While these studies appear to have conflicting results regarding the optimal fiber diameter for chondrogenesis, differences in other experimental parameters including differences in scaffold material, fiber alignment, medium composition and initial cell seeding density may have contributed to the differences in outcome. In vitro chondrogenesis is improved with 3D culture in comparison to 2D culture [27, 28], which is not surprising given that cell–cell interactions are essential to initiate the initial condensation phase during cartilage development in vivo [29]. In tissue engineering studies, wide ranges of initial seeding densities have been examined. Seeding densities as low as 1 × 104 cells cm−3scaffold up to 4 × 106 cells cm−3-scaffold have been used to examine MSC differentiation into a cartilage phenotype on fibrous scaffolds [23, 30]. Even higher seeding densities, up to 100 × 106 cells cm−3-scaffold, have been used when chondrocytes rather than stem cells were seeded, due to their reduced proliferation poten2
tial [31]. Higher seeding densities typically generate a greater amount of matrix sulfated glycosaminoglycan (GAG) and collagen produced per construct, but some studies have found that there is a saturation point where further increases in the number of cells seeded do not result in increased matrix deposition or mechanical properties [32, 33] and in some cases may even lead to a decrease in matrix production on a per cell basis [34, 35]. In this study, we examined the chondrogenic potential of human bone-marrow-derived MSCs seeded onto electrospun PCL nanofiber and microfiber scaffolds. In addition to comparing the effects of varying the scaffold diameter, the role of initial seeding density is also observed in order to determine if the density of cells on the scaffold, which impacts cell–cell interaction, influences the chondrogenic potential of the cells seeded onto scaffolds with fibers of different diameters.
2. Materials and methods 2.1. Scaffold fabrication Nanofiber and microfiber PCL scaffolds were created by electrospinning using a custom-made apparatus. PCL (80 kDa, Sigma-Aldrich, St. Louis, MO) was dissolved at 40 °C overnight in a 1:1 mixture of N,Ndimethylformamide and tetrahydrofuran (Fisher Scientific, Pittsburgh, PA) at a concentration of 11.5 and 22% w/v for nanofibers and microfibers, respectively. 0.06% w/v sodium chloride was added to the 11.5% solution to increase conductivity and improve homogeneity of fiber diameter. The polymer solution was drawn into a 20 ml syringe connected to a stainless steel needle. A syringe pump (Harvard Apparatus, Holliston, MA) was used to deliver the solution at a constant rate of 2.0 ml h−1. Microfibers were electrospun using a 12 inch, 18 G blunt-ended needle charged to 8 kV with a high voltage power supply (Gamma High Voltage Research Inc, Ormond, FL) at a distance of 23 cm from the collector. For nanofibrous scaffolds, a 4 inch, 22 G blunt-ended needle was charged to 13.5 kV at a distance of 15 cm from the collector. The collector was a custom-made grounded aluminum mandrel rotating at 0.75 m s−1. Additionally, aluminum shields were placed on either side of the mandrel and plate and charged to either 10 or 2 kV for nanofibers and microfibers, respectively, in order to guide the fibers onto the grounded surface. 2.2. Scaffold characterization Average scaffold thicknesses were measured using a digital micrometer in ten separate areas of the scaffolds. The structure of microfibers and nanofibers was examined using a scanning electron microscope (SEM) (JSM6335F; JEOL, Peabody, MA). Scaffolds were mounted onto aluminum studs, sputter-coated with 3.5 nm of palladium/gold alloy and imaged under an accelerating voltage of 5 kV and a working distance of 8 mm. Average fiber diameter, pore size and overall
IOP Publishing
Biomed. Mater. 10 (2015) 015018
A C Bean and R S Tuan
Table 1. Primers used for real-time RT-PCR.
porosity were quantified from images using ImageJ analysis software (National Institutes of Health, Bethesda, MD). 2.3. Cell culture and seeding MSCs were isolated from bone marrow obtained from the femoral heads of patients undergoing total hip arthroplasty with approval from the Institutional Review Board (University of Washington, Seattle, WA). Trabecular bone was isolated and the marrow was harvested using a bone curet and washed with DMEM containing 10% fetal bovine serum (FBS) (Gibco, Carlsbad, CA). The marrow solution was then passed through a 40 μm filter into a 50 mL conical tube and centrifuged for 5 min at 1,100 RPM. The pellet that formed was washed with phosphate buffered saline (PBS) and cells were plated onto tissue culture-treated flasks. Cells were maintained in growth medium consisting of DMEM supplemented with 10% FBS and 1% antibiotic–antimycotic (anti–anti) (Gibco) at 37 °C and 5% CO2. Cells were passaged when they reached approximately 80% confluency using 0.5% trypsin/ EDTA (Gibco) and all experiments were performed using passage 5 MSCs. Scaffolds were cut using a 1 cm diameter punch (McMaster–Carr, Elmhurst, IL) and then hydrated in 70% ethanol for 2 h. Scaffolds were then rinsed with PBS twice for 30 min each and left overnight in growth medium. To seed MSCs onto scaffolds, cells were first trypsinized, resuspended and 50 μl of the solution was pipetted onto scaffolds to give initial cell seeding densities of 100 000 (100 k), 500 000 (500 k), 2 000 000 (2 000 k) or 4 000 000 (4 000 k) cells cm −3scaffold. Cells were allowed to attach for 2 h prior to adding 1 ml of growth medium to each well of a 24well plate coated with Sigmacote (Sigma). After 24 h of culture, growth medium was replaced with 2 ml of chondrogenic medium containing serum-free DMEM supplemented with 1% anti–anti, 50 μg ml−1 ascorbic acid (Sigma), 50 μg ml−1 L-proline (Sigma), 0.1 μM dexamethasone (Sigma), 1% insulin–transferrin– selenium (Gibco) and 10 ng ml−1 recombinant human transforming growth factor-β3 (TGF-β3; Peprotech, Rocky Hill, NJ). Medium was changed every 3 d for 3 to 6 weeks prior to harvesting. 2.4. RNA isolation and quantitative real-time polymerase chain reaction (qPCR) Constructs from each group were rinsed with PBS, minced and placed into tubes. Six scaffolds were pooled into each tube in order to obtain sufficient RNA from low seeding density scaffolds. 700 μl of TRIZOL (Invitrogen, Carlsbad, CA) was then added to each tube and the scaffolds were manually homogenized before storing at −80 °C until the next step was performed. After adding 140 μl of chloroform (Fisher Scientific), each solution was incubated at room temperature for 5 min prior to centrifuging at 12 000 rpm and 4 °C for 15 min. One volume of 70% ethanol was added to each sample followed by completing the RNA extraction 3
Gene
Primer Sequences
Annealing Temp (°C)
ACANa
Forward: 5′-AGG GGC GAG TGG AAT GAT GTT-3′
56
Reverse: 5′-GGT GGC TGT GCC CTT TTT-3′ COL1b
Forward: 5′-CGA AGA CAT CCC ACC AAT CAC-3′
60
Reverse: 5′-CAT CGC ACA ACA CCT TGC C-3’ COL2c
Forward: 5′-GGC AAT AGC AGG TTC ACG TAC A-3′
52
Reverse: 5′-CGA TAA CAG TCT TGC CCC ACT T-3′ a
Aggrecan Collagen, type I, α2 c Collagen, type II, α 1 b
using the RNeasy Micro Kit (Qiagen, Valencia, CA) following the manufacturer’s protocol. The concentration and purity of the RNA was determined using a Nanodrop 2000 c spectrophotometer (Thermo Scientific, Wilmington, DE). Each RNA sample (200 ng) was reverse transcribed into cDNA using the SuperScript III First-Strand Synthesis Super Mix (Invitrogen) following the manufacturer’s protocol using Oligo(dT)20 primers. Gene-specific primers (table 1) were designed using Primer–Blast software (NIH, Bethesda, MD) [36] and purchased from Integrated DNA Technologies (IDT, Coralville, IA). The primer concentrations and annealing temperatures were optimized for efficiency (95– 105% efficiency) and specificity (melting curve analysis and product size by electrophoresis). cDNA (10 ng) was added to 10 μl SYBR® Green PCR Master Mix (Applied Biosystems, Carlsbad, CA) and 150 nM primers to a final volume of 20 μl and qPCR was performed using the StepOnePlus Real-Time PCR system (Applied Biosystems). After initial denaturation at 95 °C for 10 min, 40 cycles of 15 sec at 95 °C and 1 min at the primer-specific annealing temperature were completed to amplify the DNA. Quantification of the target gene expression, relative to the 100 k nanofiber sample at each time point, was performed using the ΔΔCT method with cyclophilin A used as an endogenous control. 2.5. Biochemical assays After washing with PBS, constructs were lyophilized and digested for 18 h in 200 μl of a papain solution (125 μg ml−1 papain, 50 mM sodium phosphate buffer, 2 mM N-acetyl cysteine (Sigma), pH 6.5). Following digestion, samples were stored at −20 °C. The DNA content in the scaffolds was quantified using the QuantiT PicoGreen dsDNA Assay Kit (Invitrogen), following the manufacturer’s protocol. The fluorescence of the PicoGreen-DNA solution was measured using
IOP Publishing
Biomed. Mater. 10 (2015) 015018
A C Bean and R S Tuan
Figure 1. SEM image of nanofiber and microfiber scaffolds. (a) Nanofiber scaffolds with average diameter of 440 ± 20 nm, average pore size of 1.2 ± 0.2 μm2 and overall porosity of 88 ± 3%. (b) Microfiber scaffolds with average diameter of 4.3 ± 0.8 μm, average pore size of 90 ± 10 μm2 and overall porosity of 90 ± 2%. Scale bar = 10 μm.
a Synergy HT plate reader (BioTek, Highland Park, VT) at an excitation/emission of 485/528 nm and DNA content in the samples was estimated using the DNA standard provided in the kit. Quantification of sulfated GAGs in the scaffolds was performed using a commercial Blyscan Glycosaminoglycan Assay kit (Accurate Chemical and Scientific, Westbury, NY). The absorbance at 656 nm was measured using a plate reader and a chondroitin-4-sulfate solution provided with the kit was used as a standard. The lower detection threshold used was 0.25 μg GAG/100 μl sample, as defined by the manufacturer. The collagen content of the scaffolds was determined using a modified hydroxyproline assay [37] using bovine collagen type I as a standard. Briefly, 125 μl of each sample and standard were hydrolyzed with 125 μl of 4 N sodium hydroxide (Fisher) at 121 °C for 20 min. 125 μl of 4 N HCl (Fisher) was added and the solution was titrated to a neutral pH. 187.5 μl of chloramine-T solution (Sigma) (14.1 g L−1 chloramine-T, 50 g L−1 citric acid, 120 g L−1 sodium acetate trihydrate, 34 g NaOH, 0.21 M acetic acid) was incubated with the sample at room temperature for 25 min. Then, 187.5 μl of 15 g L−1 p-dimethylaminobenzaldehyde in 2:1 isopropanol:perchloric acid was added and the solution was placed in a 60 °C water bath for 20 min. Finally, 200 μl of each sample in triplicate was added into a 96 well plate and absorbance was read at 550 nm. The lower threshold of detection was defined as 7.5 μg collagen/125 μl of sample, the first non-zero standard that was utilized to form the linear standard curve. 2.6. Immunofluorescence Whole-mount constructs were utilized for detailed examination of matrix components. Briefly, constructs were fixed in 4% paraformaldehyde for 20 min, washed and stored in PBS at 4 °C until staining was performed. Prior to antibody application, enzymatic antigen retrieval was performed, with constructs incubated in 0.1% pepsin in 0.01 HCl, pH 3 for 15 min at 37 °C to unmask antigens. Following retrieval, non-specific 4
antibody binding was blocked with 5% normal goat serum (Vector) for 1 h. Following blocking, samples were incubated overnight at 4 °C with primary antibodies against collagen type I (Col I; 1:200; ab292, Abcam, Cambridge, MA) and collagen type II (Col II; 1:100; MAB8887; EMD Millipore, Billerica, MA). Samples were then washed and incubated in Cy3conjugated goat anti-rabbit secondary to detect Col I and AlexaFluor 488-conjugated goat anti-mouse secondary antibodies to detect Col II for 1 h followed by nuclear staining with 40,6-diamidino-2-phenylindole (DAPI; 1:1000; Invitrogen) for 2 min. Negative controls included omission of the primary antibody as well as species-specific isotype controls. Constructs were imaged using laser scanning confocal microscopy (Fluoview 1000, Olympus, Center Valley, PA) and processed using NIS Elements software (Nikon Instruments, Inc, Melville, NY). 2.7. Statistical analysis All data were expressed as mean ± standard deviation and statistical analysis was performed using two-way analysis of variance (ANOVA) followed by Tukey’s test for post hoc comparisons. A threshold of p
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My IOPscience
Fiber diameter and seeding density influence chondrogenic differentiation of mesenchymal stem cells seeded on electrospun poly(-caprolactone) scaffolds
This content has been downloaded from IOPscience. Please scroll down to see the full text. 2015 Biomed. Mater. 10 015018 (http://iopscience.iop.org/1748-605X/10/1/015018) View the table of contents for this issue, or go to the journal homepage for more
Download details: IP Address: 129.2.19.100 This content was downloaded on 02/02/2015 at 15:01
Please note that terms and conditions apply.
Biomed. Mater. 10 (2015) 015018
doi:10.1088/1748-6041/10/1/015018
Paper
received
10 September 2014 re vised
14 December 2014
Fiber diameter and seeding density influence chondrogenic differentiation of mesenchymal stem cells seeded on electrospun poly(ε-caprolactone) scaffolds
accep ted for publication
5 January 2015 published
29 January 2015
Allison C Bean and Rocky S Tuan Center for Cellular and Molecular Engineering, Department of Orthopaedic Surgery, University of Pittsburgh School of Medicine, 450 Technology Drive, Room 221 Pittsburgh, PA 15219 USA E-mail: [email protected] Keywords: mesenchymal stem cells, chondrogenesis, electrospinning, fibrous scaffolds, biomaterials
Abstract Chondrogenic differentiation of mesenchymal stem cells is strongly influenced by the surrounding chemical and structural milieu. Since the majority of the native cartilage extracellular matrix is composed of nanofibrous collagen fibrils, much of recent cartilage tissue engineering research has focused on developing and utilizing scaffolds with similar nanoscale architecture. However, current literature lacks consensus regarding the ideal fiber diameter, with differences in culture conditions making it difficult to compare between studies. Here, we aimed to develop a more thorough understanding of how cell–cell and cell-biomaterial interactions drive in vitro chondrogenic differentiation of bone-marrow-derived mesenchymal stem cells (MSCs). Electrospun poly(εcaprolactone) microfibers (4.3 ± 0.8 µm diameter, 90 μm2 pore size) and nanofibers (440 ± 20 nm diameter, 1.2 μm2 pore size) were seeded with MSCs at initial densities ranging from 1 × 105 to 4 × 106 cells cm−3-scaffold and cultured under transforming growth factor-β (TGF-β) induced chondrogenic conditions for 3 or 6 weeks. Chondrogenic gene expression, cellular proliferation, as well as sulfated glycosaminoglycan and collagen production were enhanced on microfiber in comparison to nanofiber scaffolds, with high initial seeding densities being required for significant chondrogenic differentiation and extracellular matrix deposition. Both cell–cell and cell–material interactions appear to play important roles in chondrogenic differentiation of MSCs in vitro and consideration of several variables simultaneously is essential for understanding cell behavior in order to develop an optimal tissue engineering strategy.
1. Introduction Conservative estimates find that nearly 14% of the United States population over the age of 25 is affected by osteoarthritis and as the population continues to age, the prevalence is predicted to increase [1, 2]. The limited intrinsic healing capacity of articular cartilage and lack of successful reparative or regenerative techniques has left total joint arthroplasty as the mainstay of treatment for severe osteoarthritis over the past several decades. While relatively successful in reducing pain and improving overall quality of life [3–5], patients undergoing total joint procedures still suffer from decreased functional abilities in comparison to healthy individuals [3]. Furthermore, the failure rate of implants is 6% for 5 years and 12% for 10 years [6] and requires another surgery for revision. Due to © 2015 IOP Publishing Ltd
the large health and economic burden and limited efficacy of current treatment methods, development of a functional tissue engineered cartilage construct may lead to a significant improvement in outcome for patients with osteoarthritis. An often practiced paradigm for tissue engineering involves seeding cells onto biocompatible scaffolds and providing biochemical and/or biomechanical cues that drive cells toward a desired phenotype and stimulate them to secrete extracellular matrix (ECM) proteins to form a new and functional tissue replacement. In the native tissue environment, the complex and dynamic interactions that occur between cells and the ECM directly control tissue structure and composition. The ECM provides not only the tissue’s mechanical strength, but also essential topographical and biochemical cues that direct cell behaviors, including migration,
IOP Publishing
Biomed. Mater. 10 (2015) 015018
A C Bean and R S Tuan
proliferation and differentiation [7]. Therefore, utilizing scaffolds that mimic the native ECM architecture may improve tissue formation in vitro. In articular cartilage, collagen makes up nearly 60% of the dry weight of the tissue and plays an important role in providing the mechanical properties needed in a load-bearing tissue [8]. The fibrils found in articular cartilage typically range from 40 to 640 nm in diameter [9] and therefore many recent investigations have focused on utilizing nanofibrous scaffolds derived either from synthetic or natural materials that have fiber diameters of less than 1 μm. However, it is currently unclear whether nanofibers provide optimal scaffold architecture or if other experimental factors are more important for successful cartilage tissue engineering. Therefore, examination of cell behavior on scaffolds with different fiber diameters in conjunction with other experimental variables may enhance our understanding of what conditions are ideal for generating a tissue engineered cartilage construct. Electrospinning is an inexpensive and versatile technique that has become popular for fabrication of nanofiber and microfiber scaffolds. Several studies have examined the effects of electrospun scaffold fiber diameter on different cell types and tissue applications including bone [10–13], nerve [14–17], ligament [18, 19], skin [20, 21], blood vessels [22] and cartilage [23– 26]. The results of these studies suggest that the optimal fiber diameter may be highly dependent on various experimental factors and it may be impossible to isolate a single ideal fiber diameter independent of other conditions. For example, one study concluded that differentiation of mesenchymal stem cells (MSCs) toward a cartilage phenotype was better on electrospun poly(εcaprolactone) (PCL) fibers that were 500 nm in diameter than on fibers 3 μm in diameter [24], while another group concluded that chondrogenesis was enhanced on poly(L-lactide) fibers with diameters of 5 or 9 μm in comparison to those between 300 nm and 1,400 nm in diameter [23]. While these studies appear to have conflicting results regarding the optimal fiber diameter for chondrogenesis, differences in other experimental parameters including differences in scaffold material, fiber alignment, medium composition and initial cell seeding density may have contributed to the differences in outcome. In vitro chondrogenesis is improved with 3D culture in comparison to 2D culture [27, 28], which is not surprising given that cell–cell interactions are essential to initiate the initial condensation phase during cartilage development in vivo [29]. In tissue engineering studies, wide ranges of initial seeding densities have been examined. Seeding densities as low as 1 × 104 cells cm−3scaffold up to 4 × 106 cells cm−3-scaffold have been used to examine MSC differentiation into a cartilage phenotype on fibrous scaffolds [23, 30]. Even higher seeding densities, up to 100 × 106 cells cm−3-scaffold, have been used when chondrocytes rather than stem cells were seeded, due to their reduced proliferation poten2
tial [31]. Higher seeding densities typically generate a greater amount of matrix sulfated glycosaminoglycan (GAG) and collagen produced per construct, but some studies have found that there is a saturation point where further increases in the number of cells seeded do not result in increased matrix deposition or mechanical properties [32, 33] and in some cases may even lead to a decrease in matrix production on a per cell basis [34, 35]. In this study, we examined the chondrogenic potential of human bone-marrow-derived MSCs seeded onto electrospun PCL nanofiber and microfiber scaffolds. In addition to comparing the effects of varying the scaffold diameter, the role of initial seeding density is also observed in order to determine if the density of cells on the scaffold, which impacts cell–cell interaction, influences the chondrogenic potential of the cells seeded onto scaffolds with fibers of different diameters.
2. Materials and methods 2.1. Scaffold fabrication Nanofiber and microfiber PCL scaffolds were created by electrospinning using a custom-made apparatus. PCL (80 kDa, Sigma-Aldrich, St. Louis, MO) was dissolved at 40 °C overnight in a 1:1 mixture of N,Ndimethylformamide and tetrahydrofuran (Fisher Scientific, Pittsburgh, PA) at a concentration of 11.5 and 22% w/v for nanofibers and microfibers, respectively. 0.06% w/v sodium chloride was added to the 11.5% solution to increase conductivity and improve homogeneity of fiber diameter. The polymer solution was drawn into a 20 ml syringe connected to a stainless steel needle. A syringe pump (Harvard Apparatus, Holliston, MA) was used to deliver the solution at a constant rate of 2.0 ml h−1. Microfibers were electrospun using a 12 inch, 18 G blunt-ended needle charged to 8 kV with a high voltage power supply (Gamma High Voltage Research Inc, Ormond, FL) at a distance of 23 cm from the collector. For nanofibrous scaffolds, a 4 inch, 22 G blunt-ended needle was charged to 13.5 kV at a distance of 15 cm from the collector. The collector was a custom-made grounded aluminum mandrel rotating at 0.75 m s−1. Additionally, aluminum shields were placed on either side of the mandrel and plate and charged to either 10 or 2 kV for nanofibers and microfibers, respectively, in order to guide the fibers onto the grounded surface. 2.2. Scaffold characterization Average scaffold thicknesses were measured using a digital micrometer in ten separate areas of the scaffolds. The structure of microfibers and nanofibers was examined using a scanning electron microscope (SEM) (JSM6335F; JEOL, Peabody, MA). Scaffolds were mounted onto aluminum studs, sputter-coated with 3.5 nm of palladium/gold alloy and imaged under an accelerating voltage of 5 kV and a working distance of 8 mm. Average fiber diameter, pore size and overall
IOP Publishing
Biomed. Mater. 10 (2015) 015018
A C Bean and R S Tuan
Table 1. Primers used for real-time RT-PCR.
porosity were quantified from images using ImageJ analysis software (National Institutes of Health, Bethesda, MD). 2.3. Cell culture and seeding MSCs were isolated from bone marrow obtained from the femoral heads of patients undergoing total hip arthroplasty with approval from the Institutional Review Board (University of Washington, Seattle, WA). Trabecular bone was isolated and the marrow was harvested using a bone curet and washed with DMEM containing 10% fetal bovine serum (FBS) (Gibco, Carlsbad, CA). The marrow solution was then passed through a 40 μm filter into a 50 mL conical tube and centrifuged for 5 min at 1,100 RPM. The pellet that formed was washed with phosphate buffered saline (PBS) and cells were plated onto tissue culture-treated flasks. Cells were maintained in growth medium consisting of DMEM supplemented with 10% FBS and 1% antibiotic–antimycotic (anti–anti) (Gibco) at 37 °C and 5% CO2. Cells were passaged when they reached approximately 80% confluency using 0.5% trypsin/ EDTA (Gibco) and all experiments were performed using passage 5 MSCs. Scaffolds were cut using a 1 cm diameter punch (McMaster–Carr, Elmhurst, IL) and then hydrated in 70% ethanol for 2 h. Scaffolds were then rinsed with PBS twice for 30 min each and left overnight in growth medium. To seed MSCs onto scaffolds, cells were first trypsinized, resuspended and 50 μl of the solution was pipetted onto scaffolds to give initial cell seeding densities of 100 000 (100 k), 500 000 (500 k), 2 000 000 (2 000 k) or 4 000 000 (4 000 k) cells cm −3scaffold. Cells were allowed to attach for 2 h prior to adding 1 ml of growth medium to each well of a 24well plate coated with Sigmacote (Sigma). After 24 h of culture, growth medium was replaced with 2 ml of chondrogenic medium containing serum-free DMEM supplemented with 1% anti–anti, 50 μg ml−1 ascorbic acid (Sigma), 50 μg ml−1 L-proline (Sigma), 0.1 μM dexamethasone (Sigma), 1% insulin–transferrin– selenium (Gibco) and 10 ng ml−1 recombinant human transforming growth factor-β3 (TGF-β3; Peprotech, Rocky Hill, NJ). Medium was changed every 3 d for 3 to 6 weeks prior to harvesting. 2.4. RNA isolation and quantitative real-time polymerase chain reaction (qPCR) Constructs from each group were rinsed with PBS, minced and placed into tubes. Six scaffolds were pooled into each tube in order to obtain sufficient RNA from low seeding density scaffolds. 700 μl of TRIZOL (Invitrogen, Carlsbad, CA) was then added to each tube and the scaffolds were manually homogenized before storing at −80 °C until the next step was performed. After adding 140 μl of chloroform (Fisher Scientific), each solution was incubated at room temperature for 5 min prior to centrifuging at 12 000 rpm and 4 °C for 15 min. One volume of 70% ethanol was added to each sample followed by completing the RNA extraction 3
Gene
Primer Sequences
Annealing Temp (°C)
ACANa
Forward: 5′-AGG GGC GAG TGG AAT GAT GTT-3′
56
Reverse: 5′-GGT GGC TGT GCC CTT TTT-3′ COL1b
Forward: 5′-CGA AGA CAT CCC ACC AAT CAC-3′
60
Reverse: 5′-CAT CGC ACA ACA CCT TGC C-3’ COL2c
Forward: 5′-GGC AAT AGC AGG TTC ACG TAC A-3′
52
Reverse: 5′-CGA TAA CAG TCT TGC CCC ACT T-3′ a
Aggrecan Collagen, type I, α2 c Collagen, type II, α 1 b
using the RNeasy Micro Kit (Qiagen, Valencia, CA) following the manufacturer’s protocol. The concentration and purity of the RNA was determined using a Nanodrop 2000 c spectrophotometer (Thermo Scientific, Wilmington, DE). Each RNA sample (200 ng) was reverse transcribed into cDNA using the SuperScript III First-Strand Synthesis Super Mix (Invitrogen) following the manufacturer’s protocol using Oligo(dT)20 primers. Gene-specific primers (table 1) were designed using Primer–Blast software (NIH, Bethesda, MD) [36] and purchased from Integrated DNA Technologies (IDT, Coralville, IA). The primer concentrations and annealing temperatures were optimized for efficiency (95– 105% efficiency) and specificity (melting curve analysis and product size by electrophoresis). cDNA (10 ng) was added to 10 μl SYBR® Green PCR Master Mix (Applied Biosystems, Carlsbad, CA) and 150 nM primers to a final volume of 20 μl and qPCR was performed using the StepOnePlus Real-Time PCR system (Applied Biosystems). After initial denaturation at 95 °C for 10 min, 40 cycles of 15 sec at 95 °C and 1 min at the primer-specific annealing temperature were completed to amplify the DNA. Quantification of the target gene expression, relative to the 100 k nanofiber sample at each time point, was performed using the ΔΔCT method with cyclophilin A used as an endogenous control. 2.5. Biochemical assays After washing with PBS, constructs were lyophilized and digested for 18 h in 200 μl of a papain solution (125 μg ml−1 papain, 50 mM sodium phosphate buffer, 2 mM N-acetyl cysteine (Sigma), pH 6.5). Following digestion, samples were stored at −20 °C. The DNA content in the scaffolds was quantified using the QuantiT PicoGreen dsDNA Assay Kit (Invitrogen), following the manufacturer’s protocol. The fluorescence of the PicoGreen-DNA solution was measured using
IOP Publishing
Biomed. Mater. 10 (2015) 015018
A C Bean and R S Tuan
Figure 1. SEM image of nanofiber and microfiber scaffolds. (a) Nanofiber scaffolds with average diameter of 440 ± 20 nm, average pore size of 1.2 ± 0.2 μm2 and overall porosity of 88 ± 3%. (b) Microfiber scaffolds with average diameter of 4.3 ± 0.8 μm, average pore size of 90 ± 10 μm2 and overall porosity of 90 ± 2%. Scale bar = 10 μm.
a Synergy HT plate reader (BioTek, Highland Park, VT) at an excitation/emission of 485/528 nm and DNA content in the samples was estimated using the DNA standard provided in the kit. Quantification of sulfated GAGs in the scaffolds was performed using a commercial Blyscan Glycosaminoglycan Assay kit (Accurate Chemical and Scientific, Westbury, NY). The absorbance at 656 nm was measured using a plate reader and a chondroitin-4-sulfate solution provided with the kit was used as a standard. The lower detection threshold used was 0.25 μg GAG/100 μl sample, as defined by the manufacturer. The collagen content of the scaffolds was determined using a modified hydroxyproline assay [37] using bovine collagen type I as a standard. Briefly, 125 μl of each sample and standard were hydrolyzed with 125 μl of 4 N sodium hydroxide (Fisher) at 121 °C for 20 min. 125 μl of 4 N HCl (Fisher) was added and the solution was titrated to a neutral pH. 187.5 μl of chloramine-T solution (Sigma) (14.1 g L−1 chloramine-T, 50 g L−1 citric acid, 120 g L−1 sodium acetate trihydrate, 34 g NaOH, 0.21 M acetic acid) was incubated with the sample at room temperature for 25 min. Then, 187.5 μl of 15 g L−1 p-dimethylaminobenzaldehyde in 2:1 isopropanol:perchloric acid was added and the solution was placed in a 60 °C water bath for 20 min. Finally, 200 μl of each sample in triplicate was added into a 96 well plate and absorbance was read at 550 nm. The lower threshold of detection was defined as 7.5 μg collagen/125 μl of sample, the first non-zero standard that was utilized to form the linear standard curve. 2.6. Immunofluorescence Whole-mount constructs were utilized for detailed examination of matrix components. Briefly, constructs were fixed in 4% paraformaldehyde for 20 min, washed and stored in PBS at 4 °C until staining was performed. Prior to antibody application, enzymatic antigen retrieval was performed, with constructs incubated in 0.1% pepsin in 0.01 HCl, pH 3 for 15 min at 37 °C to unmask antigens. Following retrieval, non-specific 4
antibody binding was blocked with 5% normal goat serum (Vector) for 1 h. Following blocking, samples were incubated overnight at 4 °C with primary antibodies against collagen type I (Col I; 1:200; ab292, Abcam, Cambridge, MA) and collagen type II (Col II; 1:100; MAB8887; EMD Millipore, Billerica, MA). Samples were then washed and incubated in Cy3conjugated goat anti-rabbit secondary to detect Col I and AlexaFluor 488-conjugated goat anti-mouse secondary antibodies to detect Col II for 1 h followed by nuclear staining with 40,6-diamidino-2-phenylindole (DAPI; 1:1000; Invitrogen) for 2 min. Negative controls included omission of the primary antibody as well as species-specific isotype controls. Constructs were imaged using laser scanning confocal microscopy (Fluoview 1000, Olympus, Center Valley, PA) and processed using NIS Elements software (Nikon Instruments, Inc, Melville, NY). 2.7. Statistical analysis All data were expressed as mean ± standard deviation and statistical analysis was performed using two-way analysis of variance (ANOVA) followed by Tukey’s test for post hoc comparisons. A threshold of p
