Cartilage repair is a challenging research area because of the limited healing capacity of adult articular cartilage.1 Several methods to improve cartilage repair, including subchondral abrasion, pridie perforations, micro-fracture, transplantion of osteochondral plugs, and autologous chondrocyte implantation, have had limited success to date.2 Tissue engineering offers promising therapies for the treatment of osteoarticular disease, although some issues need to be addressed before they can be used in a clinical setting.3-6
For tissue-engineering strategies the scaffold is fundamentally important, as it provides structural support, mimicking the in vivo microenvironment and potentially influencing cellular differentiation.3 The unique combination, quantity, and organization of extracellular matrix (ECM) components give rise to tissue-specific structure and function. This is most obvious in cartilage tissue, whose ECM has properties ideally suited for the support and distribution of loads experienced by joints during daily activities. A broad range of scaffolds, both natural and synthetic, have been used for cartilage tissue engineering in vitro and in vivo.7-12 Recently, biological scaffolds derived from decellularized tissue and organs have been used for the repair or replacement of damaged tissue.13,14 Removal of the xenogeneic or allogenic cellular antigens and preservation of most of the structural and functional proteins results in acellular ECM scaffolds with relatively low immunogenicity, thereby providing an appropriate microenvironment for tissue engineering.15 However, the decellularization process remains technically difficult to accomplish because cartilage is very compact, thereby it does not allow for efficient penetration of reagents that can facilitate cell removal.16
The cell source is another important aspect of tissue engineering strategies.3,17 Bone marrow-derived mesenchymal stem cells (BMSCs) are considered promising candidates for tissue engineering because of their ability to rapidly proliferate, their differentiation potential, and the ease with which they can be harvested.6 BMSCs have been shown to have strong chondrogenic potential and have been widely used in combination with different scaffolds in vitro and in vivo.18,19
We have previously developed a natural, human cartilage ECM-derived 3D porous acellular scaffold for in vivo cartilage tissue engineering in nude mice.20 However, before these scaffolds can be used in clinical applications in vivo, the in vitro effects should be further explored. The aim of this study was to further characterize a novel decellularized cartilage ECM scaffold and to investigate the attachment, proliferation, and differentiation of PKH26-labeled BMSCs in long-term in vitro cultures. Morphology of the scaffold was determined by scanning electron microscopy (SEM), while biochemical components and biomechanical characteristics were detected by biochemical and biomechanical assays, respectively. The feasibility of using our scaffold for tissue engineering in vitro was also investigated.
Preparation of cartilage ECM-derived scaffold
Human cadaveric joints were collected following approval from the Chinese people's Liberation Army (PLA) General Hospital Ethics Committee, Beijing. Articular cartilage slices were cut from the joint within 6 hours post mortem and transported on ice in sterile phosphate-buffered saline (PBS; pH 7.6), then shattered and decellularized under aseptic conditions as described previously.20 All reagents were obtained from Sigma-Aldrich (Poole, UK), unless otherwise noted. Briefly, cartilage pieces were mixed in PBS containing 3.5% (w/v) phenylmethyl sulphonylfluoride (Merck, Darmstadt, Germany) and 0.1% (w/v) EDTA for 60 minutes. The resulting suspension was centrifuged at 500 × g, and the decellularized cartilage matrix microfilaments with diameters of approximately 500 nm to 5 mm were vigorously washed with sterile PBS and then made into a 3% (w/v) suspension in PBS. Cartilage ECM scaffolds were fabricated using a simple freeze-drying and cross-linking method. The suspension was frozen at -20°C for 1 hour, then at -80°C for 1 hour, then lyophilized for 48 hours in a freeze dryer (Boyikang, Beijing, China). All scaffolds were sterilized and cross-linked for 24 hours by dehydrothermal treatment. Further cross-linking was achieved by treating with a water-soluble carbodiimide. Scaffolds were immersed in a carbodiimide solution (14 mmol/L 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (EDAC), 5.5 mmol/L N-hydroxy sucinimide (NHS)) for 2 hours at 4°C. Excess EDAC was rinsed out using PBS. Scaffolds were then sterilized using γ-irradiation (60Co, 5 mrad).
Characterization of decellularized cartilage matrix microfilaments
Decellularized cartilage matrix microfilaments were assessed by collagen II immunohistochemistry and by SEM (Hitachi S-520, Japan).
Characterization of the cartilage ECM-derived scaffold
Porosity of the scaffold was measured by the ethanol intrusion method. The scaffold was cut using a sharp blade and the samples were examined by SEM. To observe the inner microstructure of the scaffold, specimens were immersed in a 3% OsO4 solution to increase the X-ray attenuation and air dried before micro-CT (GE Medical Systems, London, Ontario, Canada).21
Histological and biochemical analysis of the scaffold
Scaffold specimens were fixed in 10% (v/v) neutral buffered formalin, dehydrated and embedded in paraffin. Specimen cross-sections of 10 mm thickness were stained with Toluidine blue and subjected to collagen II immunohistochemistry. DNA was stained using Hoechst H33258 dye (Molecular Probes, Eugene, OR, USA).
Freeze-dried samples were crushed and digested in 1 ml of papain buffer (1.25 μg/ml papain, 100 mmol/L PBS, 10 mmol/L EDTA, 10 mmol/L cysteine, pH 6.3) for 48 hours at 60°C and then stored at -20°C until analysis. Aliquots of the papain digest were assayed for glycosamoniglycans (GAGs) using a dimethylmethylene blue (DMMB) method that has been previously described.22 Briefly, a 100 μl aliquot of the papain digest was mixed with 2 ml of the DMMB dye (Sigma-Aldrich, St. Louis, MO, USA) and absorbance at 525 nm was measured. The amount of GAGs in the samples was extrapolated from a standard curve for shark chondroitin sulfate. Total collagen content was determined by measuring the hydroxyproline content of the scaffolds after acid hydrolysis and reaction with p-dimethylaminobenzaldehyde and chloramine-T, using 0.134 as the ratio of hydroxyproline to collagen.
Biomechanical analysis of the scaffold
The mechanical properties of the dry and hydrated scaffolds (n=6 for each) were investigated in unidirectional, unconfined compression tests.23 All scaffolds were machined to a height of 6 mm and diameter of 8 mm. Scaffolds were hydrated in PBS at 37°C for 12 hours, and were kept immersed in PBS during tests. All compression tests were performed perpendicular to the plane of the scaffold sheet. Mechanical tests on dry and hydrated scaffolds were performed with ElectroForce® 3100 Test Instruments (Bose Corporation, Eden Prairie, Minnesota, USA) at a constant speed of 1 mm/min using a 10 N load cell. Applied force and displacement were recorded continuously. The test was aborted when a maximum load of 10 N or 50% height compression of samples was reached. Force-displacement data was used to calculate the apparent compressive elastic modulus (E) of the scaffold. The elastic modulus was measured as the maximum slope using linear least squares regression to fit 20% segments of the stress-strain curve prior to the yield point. Data are expressed as means ± standard deviation. Differences between the dry and wet states were tested using Student's t-test (SPSS Inc., Chicago, IL, USA). A P-value less than 0.05 was considered statistically significant.
In vitrocell studies
Isolation, culture, and chondrogenic induction of BMSCs
The Institutional Review Board of the Chinese PLA General Hospital approved our study. Canine MSCs were obtained from aspirates (5-10 ml) of the iliac crest sampled from male dogs aged 2 years and isolated as described previously.26 Cells prior to passage three were used for our studies. Chondrogenesis was induced using defined chondrogenic medium consisting of Dulbecco's modified Eagle's medium (DMEM), 100 U/ml penicillin, 100 μg/ml streptomycin, 20% fetal bovine serum (FBS), 0.1 mmol/L ascorbic acid, 40 μg/ml L-hydroxyproline, 10-7 mol/L dexamethasone, 10 ng/ml transforming growth factor β1, 25 ng/ml basic fibroblast growth factor and ITS+1 (10 μg/ml insulin, 5.5 μg/ml transferrin, 5 ng/ml selenium, 0.5 mg/ml bovine serum albumin, 4.7 μg/ml linoleic acid). The medium was replenished twice a week and the cultures were harvested at Day 14. Before and after chondrogenic induction, BMSCs were observed using an inverted microscope. At Day 7, the induced BMSCs were characterized by hemotoxylin-eosin (HE) and toluidine blue staining and collagen II immunohistochemistry.
At the end of the expansion period (Day 14), chondrogenic BMSCs were labeled with the fluorescent dye PKH26 according to the manufacturer's protocol. Briefly, cultured cells were detached from the culture dish with use of trypsin, washed in serum-free medium and resuspended in 1 ml of dilution buffer from the manufacturer's labeling kit. The cell suspension was mixed with an equal volume of the labeling solution containing 4 10-6 mol/L PKH26 in the dilution buffer and incubated for 4 minutes at room temperature. The reaction was terminated by adding 2 ml of FBS.
Seeding and attachment assessment of BMSCS inside the scaffolds
Cartilage ECM-derived scaffolds (approximately 5 mm in diameter and 2 mm in thickness) were rinsed with medium for 20 minutes, then aliquots (50 ml) containing 1×106 cells were seeded onto each scaffold and allowed to attach for 2 hours at 37°C before being immersed in complete medium. Medium (5 ml) was added to the cell-scaffold construct every 30 minutes to avoid cell death. Cell attachment was confirmed by direct visualization of a construct after 48 hours using an inverted microscope and SEM. Samples were washed with PBS and fixed with 2.5% glutaraldehyde for 24 hours at 4°C, dehydrated in a graded ethanol series to 100% ethanol, treated with hexamethyl-disilazane, and sputter-coated with gold-palladium before electron microscopy. The viability of attached cells on the scaffold was confirmed using a LIVE/DEAD® Fixable Dead Cell Stain Kit (Molecular Probes; Eugene, OR).24 Live pieces of each cell-scaffold construct were prepared as previously described.24 Pieces were incubated for 30-45 minutes at room temperature in 2 mmol/L calcein AM (live cell stain) and 4 mmol/L EthD-1 (dead cell stain), then washed again with PBS. A laser-scanning confocal microscope (excitation/emission: 495/515 nm; Bio-Rad MRC-1024, USA) was used to view the samples.
Immunofluorescence examination of PKH26-labeled BMSC-scaffold constructs
After PKH26-labeled BMSC-scaffold constructs were cultured for 3 days, specimens were embedded in OCT compound (Tissue-Tek; Miles, USA). Cryosections (10-μm thickness) were cut and fixed in acetone for 30 minutes at 4°C, then washed with PBS. Sections were incubated twice in 50 mmol/L ammonium chloride made up in PBS, then blocked with blocking reagent, and incubated overnight at 4°C with a monoclonal antibody against collagen II. This was followed by extensive washes with PBS and further incubation for 1 hour at room temperature with a FITC-conjugated IgG antibody (Shiankexing, Beijing, China). After three washes in PBS, sections were mounted with Mowiol and observed using a fluorescence microscope.
Histological and immunohistochemical analysis of cell-scaffold constructs
Cell-scaffold constructs were cultured in vitro for 1 and 3 weeks and then assessed by histological and immunohistochemical analysis to observe the growth, proliferation, and differentiation of seeded BMSCs. Samples were fixed with 4% paraformaldehyde, dehydrated through a graded series of ethanol and embedded in paraffin. Serial sections (5-mm thickness) were cut, deparaffinized in xylene, rehydrated through a graded ethanol series, and stained with HE, Safranin-O, and toluidine blue. Immunohistochemistry was conducted to screen for the expression of collagen II. Sections were sequentially washed in 70% ethanol in PBS, and treated with 3% H2O2 in PBS containing 0.15% (v/v) Triton X-100. Following blocking with a 1% (w/v) bovine serum albumin solution made up in PBS, sections were incubated with a mouse monoclonal anti-collagen II antibody (1:200 dilution; Beijing Zhongshan Biotechnology Corporation) at 4°C overnight, followed by the addition of a biotinylated secondary antibody. All immunohistochemistry samples were counter-stained with hematoxylin for visualization of cell nuclei before observation with a microscope.
Characterization of decellularized cartilage matrix microfilaments
Normal cartilage slices were white with a glass-like appearance (Figure 1A). Decellularized cartilage matrix microfilaments exhibited plume-like structures (Figure 1B) with collagen II being present (Figure 1C). From the high magnification SEM images, decellularized cartilage matrix microfilaments were found to have a random alignment on the nanometer scale (Figure 1D).
Gross morphology, microstructure, and histology of the scaffold
The cartilage ECM-derived scaffold was white, sponge-like, porous and interconnected (Figure 2A), with a pore diameter of (231.6±57.2) μm. In agreement with the SEM data, the micro-CT revealed the same porous and 3-D interconnected microstructure with the same pore diameter (Figure 2B, C and D). Toluidine blue staining was positive, with GAGs remaining after decellularization (Figure 3A). Immunohistochemistry showed that collagen II components were present in the scaffold (Figure 3B), and Hoechst 33258 stained the nuclei of cells in normal human cartilage (Figure 3C). In contrast, after decellularization, cells, cell fragments, and nuclei were absent from the scaffold (Figure 3D).
Biochemical and biomechanical analyses
For the acellular scaffold, the dry weight of total GAG and collagen was (254.7±25.9) μg/mg and (708.2±44.7) μg/mg, respectively. The lengthways modulus of elasticity for the dehydrated scaffold was (1.226±0.288) MPa, which was reduced to (0.052±0.007) MPa after hydration. There was a statistically significant difference for the longitudinal elastic modulus between the dry and wet scaffolds (P <0.05).
Chondrogenic induction of BMSCs
BMSCs had a long fusiform shape, however after chondrogenic induction, cells had a polygonal or elliptic morphology (Figure 4A and B). After 7 days of culture, cells were chondrocyte-like as indicated by safranin O, tuoluidine blue and collagen II staining (Figure 4C, D, and E).
Cell attachment and viability of cell-scaffold constructs
The BMSCs adhered to the cartilage ECM scaffold after 48 hours in culture (Figure 5A). SEM results showed that cells induced towards a chondrogenic fate attached to the scaffold and covered the pore space (Figure 5B and C). Majority of cells had an elliptic or round morphology and were surrounded by a matrix. Viability assessment of BMSC-scaffold constructs showed that nearly all live cells adhered to the rim of pores and were uniformly distributed inside scaffolds (Figure 5D). The viability assessments we conducted were in agreement with the microscopy results we obtained.
Immunofluorescence examination of cell-scaffold constructs
Fluorescing cells labeled with PKH26 adhered to the rims of pores in cartilage ECM scaffolds (Figure 6A). Within the same field of view, collagen II was observed around the inner cells and the cartilage ECM scaffold (Figure 6B), with certain cells positive for both PKH26 and collagen II (Figure 6C).
Gross morphology, immunohistochemistry, and histology
At Day 7 the cell-scaffold constructs had become opaque with a smooth, pink surface (Figure 7A). The size of the constructs remained similar for over 3 weeks of culture. By Day 21, the cell-scaffold construct was pink, smooth, and translucent, with tenacity and flexibility similar to that of autologous cartilage (Figure 7B). From Days 7 to 21, histological staining demonstrated that the number of seeded cells increased significantly inside and on the cartilage ECM scaffold (Figure 8). After 7 days of culture, the distribution of seeded cells was predominant at the periphery of the scaffold pores, with 2-3 layers of cells residing on the surface of scaffold pores (Figure 8 A, C, E, and G). After 21 days of culture, more cells were present inside the pores of the cartilage ECM scaffold and on the surface, with the formation of cartilaginous tissue. All specimens were intensely stained with safranin O and tuoluidine blue (Figure 8 C, D, E, and F). Collagen II immunohistochemical staining was positive for both seeded cells and ECM (Figure 8 G and H). During 3 weeks of in vitro culture, no shrinkage or degradation of the cartilage ECM scaffold was observed.
Although various materials, including synthetic polymers and naturally derived materials, have been used in cartilage tissue engineering, no ideal scaffold has been identified. The ideal cartilage scaffold should mimic ECM composition and the morphological structure of the cartilage. Recently, decellularized chondrocyte-derived ECM scaffolds for cartilage tissue engineering have been researched. It has been found that these scaffolds strongly support developing cartilage tissue.25-27 Unfortunately, these scaffolds have limited resources and lower mechanical properties. We previously developed a novel method, using a combined physical pulverization and modified decellularization protocol, to obtain cartilage ECM scaffolds, which may be strong candidates for cartilage tissue engineering.20 In this study, we further characterized a cartilage ECM-derived scaffold and explored the feasibility of applying this type of scaffold seeded with BMSCs for in vitro cartilage tissue engineering.
We chose cartilage ECM as a scaffold material, as it contains native articular cartilage ECM components, which can provide signals to drive undifferentiated cells toward chondrogenesis.15 Scaffolds developed for cartilage tissue engineering to date have predominantly focused on replicating the physicochemical properties of the natural cartilage via the incorporation of xenogenically purified ECM components. These components include collagen type II, chondroitin-6-sulfate, hyaluronic acid, and/or aggrecan. The chondrogenesis of MSCs has also been enhanced by culturing in type II collagen-based hydrogels,28 further suggesting that native ECM components, at least in part, guide the differentiation process. However, potential shortcomings of these scaffolds include their inability to retain sulfated GAGs that are produced by seeded cells, the inability to incorporate whole aggrecan molecules, and a lack of minor collagen components typically found in cartilage. Furthermore, natural cartilage is a dense and compact tissue, consisting of a highly organized network of collagen and large aggregating proteoglycans. It is characterized by an effective pore size of only a few nanometers, preventing efficient penetration of decellularization reagents. Kelley et al29 attempted to subject rabbit nasal septa cartilage to several conventional physicochemical processing methods, but cell fragments remained in the treated cartilage, with the possibility that they could induce immune reactions. Schwarz et al30 attempted to decellularize whole cartilage to develop a scaffold. They were able to successfully remove cells and cell membranes with collagen fibrils remaining intact, however this scaffold had lower porosity (44.14±6.84%) and a pore diameter of 10-40 μm, therefore cells struggled to migrate deeper into the scaffold.
In our study, a simple method combining physical pulverization and decellularization was developed to fabricate cartilage ECM scaffold. During the process, decellularized cartilage matrix microfilaments had a plume-like microstructure, were positive for collagen II, and randomly aligned at the nanofibre scale. These results demonstrate that cartilage matrix microfilaments retain their microstructure and collagen II components during processing. Our cartilage ECM-derived scaffold had greater porosity and larger pore diameter than those described by Kelley et al and Schwarz et al. Interconnected pore structure facilitates cell migration into internal pores, favors flow transport of nutrients and metabolic waste, and enhances the communication of cells in different pores during cell culture. In our study, the seeded cells had higher viability and penetration, uniform cell attachment, and distribution.
Larson et al31 demonstrated that culturing chondrocytes with a native pericellular matrix greatly enhanced the production of ECM macromolecules compared with chondrocytes lacking a pericellular matrix. One important characteristic of the cartilage ECM scaffolds used in our study is that their native biochemical makeup remained unchanged after decellularization. This can be seen from our histological and biochemical findings. We also speculate that most of the minor collagen components, such as collagen types IX and XI, remain in the cartilage ECM scaffold. Also, no cell fragments or DNA remain, as shown by Hoechst staining, indicating that the decellularization process was successful.
One concern regarding the use of natural tissue-derived scaffold in tissue-engineering application is their poor mechanical properties. In our study, after physical and chemical cross-linking, the lengthways modulus of elasticity for the dehydrated cartilage ECM scaffold was (1.226±0.288) MPa, similar to that of normal human cartilage. The mechanical characteristics of cartilage are closely related to the composition of the ECM. We believe that our cartilage ECM scaffold possesses appropriate biomechanical properties as it retains most of the normal cartilage ECM components. Although the mechanical characteristics in the wet state were reduced, there is a greater possibility of practical use as it secretes the required ECM molecules to a greater degree than in the dry state, thereby enhancing its mechanical properties.
Several methods, such as autoclaving, dry heat, exposure to UV light or gamma radiation, and immersion of chemical reagents are commonly used in the sterilization of natural tissues during the production of bioimplants. The current gold standard for sterilization is autoclaving; however it is not suitable for many tissues, especially cartilage, as it leads to denaturation of the tissue and adversely affects tissue biomechanics. Chemical sterilization can be highly toxic because of the reagents used, and was also deemed inappropriate for our cartilage ECM scaffold; therefore we used γ-irradiation for sterilization in our study.
PKH26 is a fluorescent dye that labels membranes, and has been widely used for cell tracking in vitro and in vivo.32,33 Its fluorescence intensity can be used as an indirect measure of proliferation.33,34 Because the cells in samples displayed red fluorescence, a feature of the PKH26 dye, we inferred that the cells in neocartilage were derived from the implanted and induced BMSCs. Green fluorescence was seen in the same areas that we observed red fluorescence, indicative of the presence of collagen II in cells and the ECM. Collagen II-expressing cells also displayed red fluorescence. These results indicated that chondrocyte-like cells in constructs were derived from seeded BMSCs and excreted ECM at high levels.
Our findings show that our novel cartilage ECM-derived porous scaffold combined with BMSCs had the ability to form cartilage-like tissues after long-term culture in vitro. For in vitro production of functional neocartilage tissue, substantial neosynthesis of GAGs and collagen is required. In this study, 3D in vitro cultures showed that BMSCs adhered to the cartilage ECM scaffold surface and were able to migrate into the matrix. The capacity of BMSCs to synthesize new cartilage-specific ECM after long-term 3D culture was demonstrated by positive safranin O, toluidine blue, and collagen II staining. It is also worth noting that the number of seeded cells and the amount of secreted cartilage ECM increased significantly after 3 weeks of in vitro culture. Our histology results showed that a large number of seeded cells clustered on the surface of the cartilage ECM scaffold. Because providing an adequate nutrient supply may be necessary for enhanced cell proliferation and ECM production, the use of bioreactors to improve nutrient transport and scale-up constructs might be beneficial for further development of tissue engineered cartilage.35
We thank HUANG Jing-xiang, TIAN Yue, and SUI Xiang for kind assistance in cell culture and histology.
1. Hunziker EB. Articular cartilage repair: basic science and clinical progress. A review of the current status and prospects. Osteoarthritis Cartilage 2002; 10: 432-463.
2. Simon TM, Jackson DW. Articular cartilage: injury pathways and treatment options. Sports Med Arthrosc 2006; 14: 146-154.
3. Danisovic L, Varga I, Zamborsky R, Bohmer D. The tissue engineering of articular cartilage: cells, scaffolds and stimulating factors. Exp Biol Med (Maywood) 2012; 237: 10-17.
4. Kock L, van Donkelaar CC, Ito K. Tissue engineering of functional articular cartilage: the current status. Cell Tissue Res 2012; 347: 613-627.
5. Gobbi A, Nunag P, Malinowski K. Treatment of full thickness chondral lesions of the knee with microfracture in a group of athletes. Knee Surg Sports Traumatol Arthrosc 2005; 13: 213-221.
6. Raghunath J, Salacinski HJ, Sales KM, Butler PE, Seifalian AM. Advancing cartilage tissue engineering
: the application of stem cell technology. Curr Opin Biotechnol 2005; 16: 503-509.
7. Jia S, Liu L, Pan W, Meng G, Duan C, Zhang L, et al. Oriented cartilage extracellular matrix
for cartilage tissue engineering
. J Biosci Bioeng 2012; 113: 647-653.
8. Dai XS, Cai YZ. Matrix-induced autologous chondrocyte implantation addressing focal chondral defect in adolescent knee. Chin Med J 2012; 125: 4130-4133.
9. Yang Q, Peng J, Lu SB, Guo QY, Zhao B, Zhang L, et al. Evaluation of an extracellular matrix
-derived acellular biphasic scaffold
/cell construct in the repair of a large articular high-load-bearing osteochondral defect in a canine model. Chin Med J 2011; 124: 3930-3938.
10. Li JW, Guo XL, He CL, Tuo YH, Wang Z, Wen J, et al. In vitro
chondrogenesis of the goat bone marrow mesenchymal stem cells directed by chondrocytes in monolayer and 3-dimetional indirect co-culture system. Chin Med J 2011; 124: 3080-3086.
11. Sun S, Ren Q, Wang D, Zhang L, Wu S, Sun XT. Repairing cartilage defects using chondrocyte and osteoblast composites developed using a bioreactor. Chin Med J 2011; 124: 758-763.
12. Moutos FT, Freed LE, Guilak F. A biomimetic three-dimensional woven composite scaffold
for functional tissue engineering of cartilage. Nat Mater 2007; 6: 162-167.
13. Turner NJ, Yates AJ Jr., Weber DJ, Qureshi IR, Stolz DB, Gilbert TW, et al. Xenogeneic extracellular matrix
as an inductive scaffold
for regeneration of a functioning musculotendinous junction. Tissue Eng Part A 2010; 16: 3309-3317.
14. Gilbert TW, Sellaro TL, Badylak SF. Decellularization of tissues and organs. Biomaterials 2006; 27: 3675-3683.
15. Badylak SF. The extracellular matrix
as a biologic scaffold
material. Biomaterials 2007; 28: 3587-3593.
16. Gibson JS, Milner PI, White R, Fairfax TP, Wilkins RJ. Oxygen and reactive oxygen species in articular cartilage: modulators of ionic homeostasis. Pflugers Arch 2008; 455: 563-573.
17. Wang L, Tran I, Seshareddy K, Weiss ML, Detamore MS. A comparison of human bone marrow-derived mesenchymal stem cells
and human umbilical cord-derived mesenchymal stromal cells for cartilage tissue engineering
. Tissue Eng Part A 2009; 15: 2259-2266.
18. Oliveira JT, Correlo VM, Sol PC, Costa-Pinto AR, Malafaya PB, Salgado AJ, et al. Assessment of the suitability of chitosan/polybutylene succinate scaffolds seeded with mouse mesenchymal progenitor cells for a cartilage tissue engineering
approach. Tissue Eng Part A 2008; 14: 1651-1661.
19. Salgado AJ, Oliveira JT, Pedro AJ, Reis RL. Adult stem cells in bone and cartilage tissue engineering
. Curr Stem Cell Res Ther 2006; 1: 345-364.
20. Yang Q, Peng J, Guo Q, Huang J, Zhang L, Yao J, et al. A cartilage ECM-derived 3-D porous acellular matrix scaffold
for in vivo cartilage tissue engineering
with PKH26-labeled chondrogenic bone marrow-derived mesenchymal stem cells
. Biomaterials 2008; 29: 2378-2387.
21. van Lenthe GH, Hagenmuller H, Bohner M, Hollister SJ, Meinel L, Muller R. Nondestructive micro-computed tomography for biological imaging and quantification of scaffold
-bone interaction in vivo. Biomaterials 2007; 28: 2479-2490.
22. Hoemann CD, Sun J, Chrzanowski V, Buschmann MD. A multivalent assay to detect glycosaminoglycan, protein, collagen, RNA, and DNA content in milligram samples of cartilage or hydrogel-based repair cartilage. Anal Biochem 2002; 300: 1-10.
23. Converse GL, Conrad TL, Roeder RK. Mechanical properties of hydroxyapatite whisker reinforced polyetherketoneketone composite scaffolds. J Mech Behav Biomed Mater 2009; 2: 627-635.
24. Li X, Jin L, Balian G, Laurencin CT, Greg Anderson D. Demineralized bone matrix gelatin as scaffold
for osteochondral tissue engineering. Biomaterials 2006; 27: 2426-2433.
25. Choi KH, Choi BH, Park SR, Kim BJ, Min BH. The chondrogenic differentiation of mesenchymal stem cells on an extracellular matrix scaffold
derived from porcine chondrocytes. Biomaterials 2010; 31: 5355-5365.
26. Jin CZ, Choi BH, Park SR, Min BH. Cartilage engineering using cell-derived extracellular matrix scaffold in vitro
. J Biomed Mater Res A 2010; 92: 1567-1577.
27. Jin CZ, Park SR, Choi BH, Park K, Min BH. In vivo cartilage tissue engineering
using a cell-derived extracellular matrix scaffold
. Artif Organs 2007; 31: 183-192.
28. Zheng L, Fan HS, Sun J, Chen XN, Wang G, Zhang L, et al. Chondrogenic differentiation of mesenchymal stem cells induced by collagen-based hydrogel: an in vivo
study. J Biomed Mater Res A 2010; 93: 783-792.
29. Kelley TF, Sutton FM, Wallace VP, Wong BJ. Chondrocyte repopulation of allograft cartilage: a preliminary investigation and strategy for developing cartilage matrices for reconstruction. Otolaryngol Head Neck Surg 2002; 127: 265-270.
30. Schwarz S, Koerber L, Elsaesser AF, Goldberg-Bockhorn E, Seitz AM, Dürselen L, et al. Dooellularized cartilage matrix as a novel biomatrix for cartilage tissue-engineering applications. Tissue Eng Part A 2012; 18(21-22): 2195-2209.
31. Larson CM, Kelley SS, Blackwood AD, Banes AJ, Lee GM. Retention of the native chondrocyte pericellular matrix results in significantly improved matrix production. Matrix Biol 2002; 21: 349-359.
32. Kamelger FS, Marksteiner R, Margreiter E, Klima G, Wechselberger G, Hering S, et al. A comparative study of three different biomaterials in the engineering of skeletal muscle using a rat animal model. Biomaterials 2004; 25: 1649-1655.
33. Kang EJ, Byun JH, Choi YJ, Maeng GH, Lee SL, Kang DH, et al. In vitro
and in vivo
osteogenesis of porcine skin-derived mesenchymal stem cell-like cells with a demineralized bone and fibrin glue scaffold
. Tissue Eng Part A 2010; 16: 815-827.
34. Chen J, Wang C, Lu S, Wu J, Guo X, Duan C, et al. In vivo
chondrogenesis of adult bone-marrow-derived autologous mesenchymal stem cells. Cell Tissue Res 2005; 319: 429-438.
35. Mabvuure N, Hindocha S, Khan WS. The role of bioreactors in cartilage tissue engineering
. Curr Stem Cell Res Ther 2012; 7: 287-292.
Keywords:© 2013 Chinese Medical Association
bone marrow-derived mesenchymal stem cells; cartilage tissue engineering; extracellular matrix; in vitro; scaffold