Repair of articular cartilage defect in non-weight bearing areas using adipose derived stem cells loaded polyglycolic acid mesh

Biomaterials 30 (2009) 2683–2693

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Biomaterials journal homepage: www.elsevier.com/locate/biomaterials

Repair of articular cartilage defect in non-weight bearing areas using adipose derived stem cells loaded polyglycolic acid mesh Lei Cui a, b, c, *, Yaohao Wu a, Lian Cen b, Heng Zhou b, Shuo Yin c, Guangpeng Liu b, Wei Liu a, c, Yilin Cao a, b, c, * a b c

Department of Plastic and Reconstructive Surgery, Shanghai 9th People’s Hospital, Shanghai Jiao Tong University School of Medicine, 639 Zhi Zao Ju Road, Shanghai 200011, PR China National Tissue Engineering Center of China, 68 Jiangchuandong Road, Shanghai 200241, PR China Shanghai Tissue Engineering R&D Center, 100 Qin Zhou Road, Shanghai 200235, PR China

a r t i c l e i n f o

a b s t r a c t

Article history: Received 29 October 2008 Accepted 13 January 2009 Available online 12 February 2009

The current study was designed to observe chondrogenic differentiation of adipose derived stem cells (ASCs) on fibrous polyglycolic acid (PGA) scaffold stabilized with polylactic acid (PLA), and to further explore the feasibility of using the resulting cell/scaffold constructs to repair full thickness articular cartilage defects in non-weight bearing area in porcine model within a follow-up of 6 months. Autologous ASCs isolated from subcutaneous fat were expanded and seeded on the scaffold to fabricate ASCs/ PGA constructs. Chondrogenic differentiation of ASCs in the constructs under chondrogenic induction was monitored with time by measuring the expression of collagen type II (COL II) and glycosaminoglycan (GAG). The constructs after being in vitro induced for 2 weeks were implanted to repair full thickness articular cartilage defects (8 mm in diameter, deep to subchondral bone) in femur trochlea (the experimental group), while scaffold alone was implanted to serve as the control. Histologically, the generated neo-cartilage integrated well with its surrounding normal cartilage and subchondral bone in the defects of experimental group at 3 months post-implantation, whereas only fibrous tissue was filled in the defects of control group. Immunohistochemical and toluidine blue staining confirmed the similar distribution of COL II and GAG in the regenerated cartilage as the normal one. A vivid remolding process with post-operation time was also witnessed in the neo-cartilage as its compressive moduli increased significantly from 50.55% of the normal cartilage at 3 months to 88.05% at 6 months. The successful repair thus substantiates the potentiality of using chondrogenic induced ASCs and PGA/PLA scaffold for cartilage regeneration. Ó 2009 Elsevier Ltd. All rights reserved.

Keywords: Articular cartilage engineering Adipose derived stem cells Chondrogenic differentiation Polyglycolic acid mesh

1. Introduction Repair of full thickness articular cartilage defects remains a major challenge in the field of orthopedic surgery because of the unsatisfactory outcomes of current surgical strategies, mainly chondrectomy [1], subchondral drilling [2], periosteal [3] or perichonrial resurfacing [4] and transplantation of autochondrocytes [5]. Each of these treatments has its own limitations and in most cases the repaired tissues are fibrocartilage rather than hyaline cartilage, which usually results in degenerative changes with pain after a long-term follow-up. The poor outcomes in current clinical treatments of articular cartilage injuries are mainly due to the fact

* Corresponding authors. Department of Plastic and Reconstructive Surgery, Shanghai 9th People’s Hospital, Shanghai Jiao Tong University School of Medicine, 639 Zhi Zao Ju Road, Shanghai 200011, PR China. Tel.: þ86 21 63138341x5192; fax: þ86 21 53078128. E-mail addresses: [email protected] (L. Cui), [email protected] (Y. Cao). 0142-9612/$ – see front matter Ó 2009 Elsevier Ltd. All rights reserved. doi:10.1016/j.biomaterials.2009.01.045

that mature articular cartilage is avascular and the articular chondrocytes are highly differentiated cells that have seriously limited self-repair capacity. As an alternative to those current therapies, tissue engineering has been demonstrated to own promising therapeutic advantages in restoring both the structure and function of the damaged articular cartilage [6]. Bone marrow stromal cells (BMSCs) are the commonly used seed cells for cartilage engineering [7,8]. In our previous study, we have successfully repaired articular cartilage defects by transplanting engineered cartilage using chondrogenic induced BMSCs in a porcine model [9]. Compared with the engineered cartilage using mature chondrocytes as seed cells, the one with BMSCs not only led to the regeneration of cartilage layer but also the subchondral bone, indicating that the implanted stem cells underwent both chondrogenic and osteogenic differentiation in vivo. However, due to the limitation of the obtainable amount of autologous bone marrow, extensive in vitro culture is routinely required to achieve a therapeutic cell dose, whereas long-term in vitro expansion is not

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favorable for cells to maintain their genomic stability [10]. Therefore, it is of great interest to find an alternative cell source which is abundant, easy to access, easy to expand in culture and possesses chondrogenic differentiation potential. Recently, a type of multipotent MSCs were identified by several studies to exist in fat tissue, and designated as adipose derived stem cell (ASC) [11]. Compared with BMSCs, ASCs obviously exhibit advantages including ease of isolation, relative abundance, rapidity of expansion, and multipotency that is independent upon serum source and quality [11–13]. Chondrogenic differentiation potential of ASCs has been extensively documented in vivo as well as in vitro. According to Zheng et al. [14], ASCs possess the capacity of multiple differentiation but primarily osteogenic and chondrogenic differentiation in vivo. Moreover, Dragoo et al. [15] have shown that hyaline-like cartilage formed in an animal model with subcutaneous implantation of a chondrogenic induced ASCs nodule. More recently, it was documented that chondrogenic potential of ASCs could be achieved to a comparable level as that of BMSCs using a combination of TGF-b1 and IGF-1 for chondrogenic induction in vitro [16]. Thus, it is possible to use chondrogenic induced ASCs as seed cells to engineer cartilage tissue. A variety of biomaterials, either naturally derived or artificial, have been investigated as scaffolds for cartilage tissue engineering, such as compositions of proteins (collagen, fibrin), polysaccharides (agarose, alginate, hyaluronic acid, chitosan) and synthetic polymers (polyethylene glycol, polylactic acid) [17]. Among these, scaffolds fabricated from woven or non-woven polyglycolic acid (PGA) have been widely used for anchoring chondrocytes or stem cells to facilitate the generation of cartilage tissues in vitro or to restore damaged articular cartilage in vivo [18,19]. Suitable biocompatibility of PGA scaffold to those seed cells has been demonstrated. Moreover, in order to prepare scaffolds with specific anatomical shape and to improve the mechanical stabilities, certain amount of polylactic acid (PLA) solution was usually added during the processing of PGA scaffold. Such kind of PGA/PLA composite has been successfully used in healing cartilage defects by tissue engineering approach in our previous studies [20]. However, to our knowledge, no report has been available to date regarding to the feasibility of in vitro chondrogenic differentiation of ASCs in a 3 dimensional (3-D) microenvironment of PGA scaffold and further repairing of articular cartilage defects using ASCs combined PGA/PLA composites. Thus, to investigate whether ASCs could repair full thickness articular cartilage defects in immuno-competent large animals with the follow-up of long-term outcomes, autologous ASCs were first expanded in vitro, seeded onto PGA/PLA scaffold and then chondrogenically induced. The cell–scaffold complexes after being induced in vitro for 2 weeks were implanted back to treat the critical-sized articular cartilage defects. Successful repair was achieved at 6 months after transplantation, indicating that autologous ASCs, along with the PGA/PLA scaffold, could be potentially used for cartilage regeneration. 2. Materials and methods 2.1. Isolation and culture of ASCs The experimental protocol was approved by the Animal Care and Experiment Committee of Shanghai Jiao Tong University School of Medicine. Autologous subcutaneous adipose tissue in the nape was harvested from 8-week old hybrid pigs, either male (10 animals) or female (4 animals) weighing 10–15 kg, after anesthetization through intramuscular injection of ketamine (10 mg/kg). ASCs were isolated from the adipose tissue as described by Valina et al. [21]. Briefly, adipose tissue was washed three times with 0.1 M phosphate-buffered saline (PBS, pH 7.4) and treated with 0.075% type I collagenase (Washington Biochemical Corp., USA) at 37  C for 30 min. Enzymatic activity was neutralized with low glucose Dulbecco’s Modified Eagle’s Medium (LG-DMEM, Gibco, USA), containing 10% fetal bovine serum (FBS, HyClone, USA), and the digested solution was then centrifuged at 1200 g for 10 min. The yielding cells were resuspended in LG-DMEM culture

medium containing 10% FBS, 100 mg/mL streptomycin and 100 U/mL penicillin (defined as the growth medium), and plated at 4  104cells/cm2 in F 100 mm culture dishes (Falcon, USA) with the medium changed twice a week. When reached 70–80% confluence, cells were passaged and ASCs prior to passage 3 were used in the following study.

2.2. Preparation of PGA/PLA scaffold and cell seeding As previously reported [9], 30 mg of unwoven PGA fibers (Shanghai Ju Rui Biomaterials Company Inc., China) with diameters of 20–30 mm were pressed into a silicone rubber mold which was 8 mm in diameter and 6 mm in thickness. The shape of PGA scaffold in mold was stabilized by adding 0.3 mL polylactic acid (PLA, Sigma–Aldrich, USA) solution (1.5% in dichloromethane) via instillation. The scaffold within the mold was then kept in air for 10 min to facilitate the evaporation of the solvent. The PGA/PLA composites were then removed from the mold and sterilized by soaking in 75% alcohol for 1 h and washed 3 times with PBS followed by washing with the growth medium. The excessive medium left within scaffold was removed by extensive suction, and the scaffolds were then air-dried for 30 min under ultraviolet light. The microstructure of the scaffold was characterized by scanning electron microscope (SEM, Philips XL-30, Netherlands). ASCs were harvested and resuspended in the culture medium at a density of 5.0  107cells/mL. Aliquots of 0.3 mL cell suspensions were then evenly seeded by instillation into PGA/PLA scaffolds to form cell–scaffold constructs. After being incubated for 3 h to allow cell attachment, the constructs placed in culture dishes were added either with 6 mL of the growth medium or chondrogenic inducing medium (the growth medium further supplemented with 10 ng/mL transforming growth factor-b1 (TGF-b1), 100 ng/ml insulin-like growth factor-1 (IGF-1), 40 ng/mL dexamethasone and 6.25 mg/ml transferrin (all from Sigma–Aldrich)). The cell–PGA constructs in both media were subsequently cultured in vitro for 5 weeks, respectively, with the media changed twice a week, while the respective growth and chondrogenic differentiation of ASCs were monitored with time. On the other hand, cell–PGA constructs after being chondrogenic induced in vitro for 2 weeks were implanted back to repair articular cartilage defects as will be described later.

2.3. Proliferation and chondrogenic differentiation of ASCs on scaffold Cell numbers on the scaffold at 1, 2, 3, 4 and 5 weeks post-seeding were quantified by DNA assay using Hoechst 33258 dye (Sigma–Aldrich) as previously described [22], while extracellular matrix (ECM) deposition by ASCs on PGA/PLA scaffold was observed by SEM examination at 2, 4 weeks after seeding. Furthermore, to visualize the growth and spatial distribution of ASCs on the PGA/PLA scaffold, ASCs were pre-labeled before seeding with fluorescent 3,30 -dioc tadecyloxacarbocyanine perchlorate (DiO) dye (Molecular Probes, USA) at 37  C for 20 min following the manufacturer’s protocol. The labeled cells were then seeded on PGA/PLA scaffold as described above. The cells grown on the scaffold were then observed by a fluorescence microscope (Nikon Y-FL, Japan) at weeks 1 and 4 after seeding. Expression of collagen type II (COL II) by ASCs on the scaffold and glycosaminoglycan (GAG) content of the cell–PGA construct was also quantified weekly. Firstly, the cell–scaffold complex was rinsed with doubly-distilled H2O, lyophilized for 12 h followed by adding 1 mL cold H2O, and incubated at 4  C in a microcentrifuge tube overnight. After lysis with repeated freeze thawing and sonication cycles, samples were centrifuged at 10,000 rpm for 3 min. The supernatant was collected for GAG determination, while the precipitate was subjected to COL II assay using Native Type II Collagen Detection Kit (Chondrex, USA) according to the manufacturer’s instructions. For the COL II assay, the precipitate was resuspended in 0.8 mL acetic acid of 0.05 M containing 0.5 M NaCl, and digested with pepsin (1 mg/ mL) at 4  C for 48 h. After adjusting the pH to 8.0 with 1 N NaOH, the above solution was added with 0.1 mL pancreatic elastase (1 mg/mL dissolved in 1  TBS (0.1 M Tris–0.2 M NaCl–5 mM CaCl2), pH 7.8–8.0) and then mixed thoroughly at 4  C overnight. After that, a quantitative sandwich enzyme immunoassay technique was applied by adding the digested mixture to the microplate pre-coated with monoclonal antibody specific for COL II. After the color was developed, the absorption at 490 nm was measured spectrophotometrically using a Varioskan multi-mode detection reader (Thermo Electron Corporation, USA). A standard curve of known concentrations of COL II was generated concurrently and used to determine unknown sample concentrations. For the GAG determination, aliquots of 25 mL supernatant were mixed with 5 ml NaCl of 2.3 M and 200 mL 9-dimethylmethylene blue chloride solution (DMMB, Sigma–Aldrich). The absorption at 520 nm was then measured spectrophotometrically. A standard curve of known concentrations of chondroitin sulfate B was run concurrently and used to determine unknown sample concentrations. ASCs from three individual pigs were subjected to the above in vitro tests. For each animal, three pieces of adipose tissue were obtained to yield a mean number at each time point. An average value of three individual animals was further obtained from the above mean number. Both the non-induced ASCs seeded on the PGA/PLA scaffold and the PGA/PLA scaffold without any cells but cultured under the same condition were also measured.

L. Cui et al. / Biomaterials 30 (2009) 2683–2693 2.4. Surgical procedures

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Table 2 Histological grading scale for cartilage regeneration.a

As described previously [9], an articular cartilage defect deep to subchondral bone of 8 mm in diameter and 6 mm in depth was created by drilling using a trephine at non-weight bearing area of femur trochlea in one randomly chosen knee of each animal. The same procedure was performed at the contralateral knee joint and thus a total of 2 defects were created in one pig. After removal of blood clots, one defect in each animal was repaired with the autologous chondrogenically induced cell–PGA/ PLA construct as the experimental group, while the other was repaired by PGA/PLA scaffold alone as the control group. To ensure that a good integration of the implant with its surrounding native tissue was achieved, the implant was stabilized by a crossing trans-osseous fixation using biodegradable sutures [23]. Post-operatively, all animals were allowed to move freely. All of them regained their normal gaits and locomotion within 1 week after the operation. Animals were euthanized at 3 and 6 months post-surgery for sample harvest, respectively.

2.5. Macrographic examination and grading The harvested samples which include the cartilage defect and the underlying cancellous bone were first examined grossly and then sawed through the middle line of the defects to see the longitudinal cross-section as well as the interface between the repaired and its adjacent normal osteochondral tissue. The repair was graded according to the criteria reported previously as presented in Table 1 [24]. After the gross view examination, some samples were used for histological examination and immunohistochemical staining, while others were for biomechanical analysis followed by the GAG content determination.

2.6. Histological examination and grading The harvested samples were fixed, decalcified, embedded in paraffin, and sectioned in 5 mm slices. The slices were stained with hematoxylin and eosin (H&E) and toluidine blue, respectively. According to the scoring system described previously by Wakitani et al. [25], the histological grading was performed in accordance with the histological structure, matrix staining, surface regularity, thickness of the repaired cartilage, and integration of the repaired tissue with its surrounding normal one as presented in Table 2. The grading was performed as a blind test by three independent individuals.

2.7. Immunohistochemical staining Expression of COL II in the engineered cartilage was examined by immunohistochemical staining at 3 and 6 months post-implantation, respectively. Briefly, paraffin sections were deparaffinized followed by hydration in ethanol solutions of decreasing concentrations (100% to 70%). Endogenous peroxidase activity was stopped by incubation with 3% hydrogen peroxide for 15 min. Nonspecific binding was blocked by incubation in PBS with 10% horse serum. The section was then incubated at 37  C for 1 h with mouse anti-collagen-II monoclonal antibody (Santa Cruz, CA) diluted in PBS (1:200), followed by incubation with 1:100 diluted horseradish peroxidase (HRP)-conjugated anti-mouse antibody (DAKO, Carpinteria, CA) for 30 min. Color was developed with diaminobenzidine tetrahydrochloride (DAB). Table 1 Grading scale for gross appearance.a Description

Points

Intraarticular adhesions None Minimal/fine loose fibrous tissue Major/dense fibrous tissue

2 1 0

Description

Points

Cell morphology Hyaline cartilage Mostly hyaline cartilage Mostly fibrocartilage Mostly non-cartilage Non-cartilage only

0 1 2 3 4

Matrix staining (metachromasia) Normal (compared with host adjacent cartilage) Slightly reduced Markedly reduced No metachromatic stain

0 1 2 3

Surface regularity Smooth (>3/4) Moderate (>1/2–3/4) Irregular (1/4–1/2) Severely irregular (<1/4)

0 1 2 3

Thickness of cartilage >2/3 1/3–2/3 <1/3

0 1 2

Integration of donor with host adjacent cartilage Both edges integrated One edge integrated Neither edge integrated Maximum total

0 1 2 14

a This table was cited from the histological grading scale of Wakitani et al. [25] without modification.

2.8. Biomechanical and biochemical analysis of the engineered cartilage The biomechanical property of the engineered cartilage was tested by measuring the compressive modulus as described previously [9]. Briefly, the harvested samples were trimmed to fit in a test chamber (5 mm in diameter) of the biomechanical analyzer (Instron5542, USA). The force-displacement curve was obtained by applying a constant compressive strain at a rate of 1 mm/min until reaching the maximal force at 450 N. The compressive modulus of the tested tissue was automatically calculated by the auxiliary software in the machine and further verified by manual calculation with the formula: DP/A  L/DL (DP: the compressive force margin of the two points on linear segment of the curve before the first break point, DL: the displacement margin of the corresponding two points, A: the area of tested tissue, L: the thickness of tested tissue). As the defects in the control group at 3 months were only filled with small volume of fibrous tissue which further collapsed at 6 months, biomechanical tests on the regenerated tissue can not be performed properly. After the biomechanical analysis, the samples were collected and their GAG contents were analyzed spectrophotometrically in a similar manner as described in the previous section. Since no neo-cartilage was generated in the control group which was treated with scaffold alone, only samples of the experimental group were randomly chosen from three animals for this analysis at both 3 and 6 months postimplantation. Normal articular cartilage in the same non-weight bearing area from animals at ages equal to those of 3 and 6 months post-surgery (three individuals at each time point), respectively, was served as a normal control in the biomechanical and biochemical assays. 2.9. Statistical analyses

Restoration of articular surface Complete restoration Partial restoration No Restoration

2 1 0

Erosion of cartilage None Defect site/site border Defect site and adjacent normal cartilage

2 1 0

Appearance of cartilage Translucent Opaque Discolored or irregular Maximum total

2 1 0 8

a This table was cited from the grading scale of Cook et al. [24] without modification.

The grading of the macroscopic and histological examination was analyzed using the Mann–Whitney U signed-rank test for Two-Independent-Samples. Compressive modulus and GAG content among different groups were analyzed with ANOVA. A p-value of less than 0.05 was considered statistically significant.

3. Results 3.1. Growth and chondrogenic differentiation of ASCs on PGA/PLA scaffold Grossly, the PGA/PLA scaffold appeared as a soft and porous spongy-like disk, and the shape was stable upon evaporation of the

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solvent added with PLA (Fig. 1A). The porosity of this scaffold was around 93%. By SEM examination, it was shown that the distribution of PLA in scaffold was either in a fashion of meshy membrane between PGA fibers or in a way of wrapping PGA fibers together (Fig. 1B and C). To evaluate the proliferation of ASCs on the PGA/PLA scaffold, quantitative Hoechst 33258 assay was performed. As shown in Fig. 1D, it was found that cell numbers of the cell–PGA constructs both in the growth and chondrogenic media began to increase dramatically with time and reached a peak value both at the measuring point of 3 weeks after cell seeding. Moreover, a similar proliferation trend was observed, although at 2, 3, 4 weeks the DNA amount of the non-induced group was significantly higher than the respective one of the induced group (p < 0.05, at all of the three time points). To further visualize the growth and distribution of cells on the scaffolds, cells were pre-labeled with fluorescent DiO dye and seeded on the PGA/PLA scaffold. As shown in Fig. 1E and F, labeled cells distributed in a dispersed way along PGA fibers at 1 week and became a dense assembly over the fibers at 4 weeks postseeding, which confirmed the above dramatic increase in cell numbers by quantitative assay. By SEM observation, abundant ECM deposition was found evidently surrounding PGA fibers at 2 and 4 weeks after seeding (Fig. 2A and B). To further evaluate chondrogenic differentiation of ASCs on the scaffold, content of cartilage-specific ECM was measured at 1, 2, 3, 4 and 5 weeks post-seeding. As shown in Fig. 2C and D, both COL II and GAG contents per sample in the induced group underwent a robust increase during the first week and maintained at a high level within the following test duration, whereas those in the non-induced group still kept at a similar low level as those measured at 1 week post-seeding. 3.2. Macrographic examination and grading At 3 months post-implantation, the defects in the experimental group were mostly repaired with engineered cartilage tissue which

DNA content (ug/sample)

A

50

D

Induced Non-induced Blank

40

exhibited a whitish appearance (Fig. 3A). After 6 months, the engineered cartilage became more mature as a smooth joint surface appeared, and it bore great similarity in color and texture to the surrounding native cartilage (Fig. 3B). However, a distinguishable border between the reparative cartilage and surrounding normal tissue still existed even at 6 months post-implantation (Fig. 3B). According to the cross-sectional morphology (Fig. 3A1 and B1), a nice interface with excellent healing was achieved between the engineered cartilage with both its adjacent native cartilage and subchondral bone at either 3 or 6 months post-implantation. In contrast to reparative conditions of the experimental group, all defects in the control group were occupied mainly by fibrous tissue with little cartilage-like tissue regeneration at 3 months postimplantation (Fig. 3C and C1). Moreover, the defect became deeper with the collapse of its adjacent subchondral bone when the time reached 6 months (Fig. 3D and D1). The corresponding grading of all cases in the experimental and control groups was shown in Table 3. The global macroscopic scores obtained from the experimental group were statistically better than those in the control group (Table 3).

3.3. Histological examination and grading The histological examination further confirmed the above observation. At 3 months post-surgery, the defects of the experimental group were repaired by newly generated hyaline cartilage featured by the formation of lacunas and cell clusters (Fig. 4A). The engineered cartilage integrated well with its adjacent cartilage (Fig. 4A and A1) and its underlying subchondral bone (Fig. 4A2 and A3), although the interface between engineered and normal cartilage could still be identified. However, the cell density remained quite high and cellular arrangement was disordered in the neocartilage as shown in Fig. 4A1. Moreover, the surface and thickness of engineered cartilage was uneven with penetration of some cartilage-like tissue into the subchondral bone (Fig. 4A).

B

C

E

F

30 20 10 0

0w

1w

2w 3w Time

4w

5w

Fig. 1. Characterization of PGA/PLA scaffold and proliferation of ASCs on it. (A) Gross view of PGA/PLA scaffold in a size of around 8 mm in diameter and 6 mm in thickness. SEM observation showed that PLA bonded PGA fibers together in a fashion of meshy membrane between PGA fibers (B) or in a way of wrapping PGA fibers together (C). (D) Proliferation of induced and non-induced ASCs on PGA/PLA scaffold determined by DNA assay using Hoechst 33258 dye at 1,2,3,4 and 5 weeks post-seeding, respectively. Scaffolds alone cultured in the chondrogenic inducing medium were served as blank values at each time point. Representative confocal images of DiO labeled induced ASCs at 1 (E) and 4 (F) weeks postseeding, respectively (Bar scales: 100 mm for B; 40 mm for C; 250 mm for E and F.).

L. Cui et al. / Biomaterials 30 (2009) 2683–2693

A

B

C

SEI 20.0kV

×500

10µm WD10mm

0.15

Induced Non-induced

4 3 2 1 0

1w

SICCA

GAG (ug / ug DNA)

COL II content (ng / ug DNA)

SICCA

5

2687

2w

3w Time

4w

5w

D

×500

SEI 20.0kV

10µm WD11mm

Induced Non-induced

0.10

0.05

0.00

1w

2w

3w Time

4w

5w

Fig. 2. ECM deposition of chondrogenic differentiated ASCs on PGA/PLA scaffold. SEM evaluation of matrix deposition by chondrogenic induced ASCs on PGA/PLA scaffold at 2 (A) and 4 weeks (B) post-seeding, respectively. Chondrogenic differentiation of induced and non-induced ASCs on PGA/PLA scaffold was quantified by amounts of the deposited collagen type II (C) and GAG (D) at 1, 2, 3, 4 and 5 weeks post-seeding, respectively. PGA/PLA scaffolds alone cultured in chondrogenic induced medium were also assayed as blank values which were then subtracted from the corresponding samples at each time point (Bar scales: 10 mm for A and B.).

When it came to 6 months post-surgery, the engineered cartilage became more mature as the neo-tissue was characterized with a similar cell density and cartilage thickness to those of the surrounding native cartilage tissue (Fig. 4B and B1). Cells in

the superficial area exhibited an exclusively single cell unit and a flat profile paralleling to the articular cartilage surface (Fig. 4B2). Cells in the deeper zone were well organized in either columns or clusters aligned perpendicularly to the articular

Fig. 3. Gross and cross-sectional appearance of the repaired cartilage at 3 and 6 months post-implantation. From the gross (A) and cross-sectional (A1) appearance, most of the defects were repaired with neo-generated cartilage in the experimental group at 3 months post-surgery. At 6 months after implantation, the engineered cartilage integrated well with its surrounding normal cartilage and subchondral bone in the experimental group as observed by the gross (B) and cross-sectional (B1) view. However, defects in the control group (PGA/PLA scaffold alone) were only filled with fibrous tissue at 3 months after implantation (C and C1). Collapse of the surrounding normal cartilage and subchondral bone was observed in the control group with the further increase in time upon 6 months (D and D1).

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Table 3 Results of grading scale for gross appearance. Grading

Intraarticular adhesions Restoration of articular surface Erosion of cartilage Appearance of cartilage Total

3 months (n ¼ 7)

6 months (n ¼ 7)

Exp

Control

Exp

Control

1.14(1–2) 1.28(1–2) 1.14(1–2) 0.71(0–2) 4.28(3–6)#

0.43(0–1) 0.57(0–1) 0.57(0–1) 0.28(0–1) 1.86(1–3)

1.57(1–2) 1.57(1–2) 1.71(1–2) 1.42(1–2) 6.28(5–8)#,*

0.14(0–1) 0.42(0–1) 0.42(0–1) 0.14(0–1) 1.14(1–2)

Each data represents mean and range (parentheses) of the scales. p < 0.01, versus the control group; *p < 0.05, versus the experimental group of 3 M.

#

surface (Fig. 4B2). The integration of engineered cartilage with its subchondrol bone became more harmonious at this time point compared to that of the engineered tissue at 3 months (Fig. 4B3 versus Fig. 4A3). On the contrary, the defect treated with PGA/PLA scaffold alone (the control group) was only filled with fibrous tissue at 3 months post-surgery (Fig. 4C). Normal cartilage also collapsed surrounding

the central region of the generated fibrous tissue (Fig. 4C1). Furthermore, at 6 months, in addition to the fibrous tissue at the defect area, some underneath cancellous bone also collapsed (Fig. 4D). On the other hand, the polymeric scaffolding materials were almost completely resorbed in all implanted constructs of the two groups as they were not distinguishable in the defect regions as early as 3 months.

Fig. 4. Histology of the repaired tissue at 3 and 6 months post-implantation. (A) Regeneration of neo-cartilage with a layer of fibrous tissue on the surface was observed in the defect of experimental group at 3 months post-surgery. (B) At 6 months after repairing, defect in the experimental group was repaired with hyaline cartilage which was completely integrated into surrounding normal cartilage and subchondral bone. (C) Defect in the control group (PGA/PLA scaffold alone) was filled with fibrous tissue at 3 months postimplantation. (D) Collapse of subchondral bone occurred in the control group after 6 months of repairing. The reparative area was indicated by arrowheads. A1, B1, C1 and D1 are high-magnification images of the interface between the repaired tissue and adjacent normal cartilage. A2, B2, C2 and D2 are central area of reparative tissue in full thickness from superficial cartilage to subchondral bone. A3, B3, C3 and D3 are interface between reparative tissue and subchondral bone (Bar scales: 750 mm for A and B, 1000 mm for C and D; 200 mm for others.).

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According to the histological grading, scores of the experimental group were all significantly better than the respective ones of the control group at either 3 or 6 months (Table 4). Furthermore, scores of the engineered cartilage at 6 months post-implantation are significantly better over those corresponding ones at 3 months as summarized in Table 4, indicating that the engineered cartilage underwent a vivid remolding process with post-operation time to gradually enhance its reparative effect. At the end of 6 months, 4 of 7 samples in the experimental group showed complete regeneration as exhibited in images chosen for figures. In the left three samples, more than 80% of the area was covered with neo-cartilage. However, only 2 of 7 samples in control group were observed to exhibit sparse distribution of fibrocartilage-like tiny island in the area adjacent exactly to the surrounding normal cartilage.

also compared with that of the normal one at the similar age. At 3 months post-operation, the compressive modulus of engineered cartilage was 14.01  2.46 Mpa, reaching 50.55% of that of the normal cartilage (p < 0.05). At 6 months post-operatively, the compressive modulus of engineered cartilage increased to 38.17  2.77 Mpa, reaching 88.05% of the corresponding value of the normal one, but no significant difference was observed between the values of the engineered and normal cartilage (p > 0.05). On the other hand, the compressive modulus of the experimental group at 6 months exhibited significant improvement over that at 3 months (p < 0.05). The enhancement in the compressive modulus over time suggested that remodeling in both histology and biochemistry occurred constantly with the increase in repairing time, which is usually required for the recovery of articular cartilage not only structurally but also functionally.

3.4. Expression of cartilage ECM in engineered cartilage

4. Discussion

At 3 months post-implantation, expression of COL II in the engineered cartilage was revealed by immunohistochemical staining (Fig. 5A and A1). However, a thin layer of fibrous tissue which was negative for COL II staining was also observed in the superficial surface of regenerated cartilage. With the maturation of engineered cartilage, the expression and distribution of COL II was found to be in a similar fashion as that of the neighboring native cartilage at 6 months after surgery (Fig. 5B and B1). The superficial fibrous-like tissue could no longer be observed. However, no expression of COL II was observed in all specimens of the control group in which the defects were occupied by fibrous tissue at either 3 or 6 months (Fig. 5C and D). Deposition of GAG in the engineered cartilage was confirmed by Toluidine blue staining at 3 (Fig. 6A) and 6 months (Fig. 6B) postsurgery, respectively. The deposition profile of GAG with time is in well consistence with that of COL II. Similarly, little deposition of GAG was detected in the control group after 3 or 6 months (Fig. 6C and D) since implantation.

Compared with differentiated chondrocytes isolated from cartilage, adult stem cells derived from mesenchymal tissue represent a more promising source for cartilage engineering as they are relatively abundant and easily accessible with minimal donor site morbidity. At the same time, the identification of multipotent stem cells, ASCs, within subcutaneous adipose tissue further promoted the development of tissue engineering as they have higher cell yield and more rapid proliferation ability during in vitro expansion when compared with BMSCs. It was reported that one gram of human adipose tissue could yield over 70,000 ASCs within 24 h of in vitro culture [26]. In the current study, we also found that 5  105 cells could be easily obtained within 5 days of primary culture from 1 g of pig adipose tissue These favorable features of ASCs make them more practicable for clinical treatments. Furthermore, it was shown that ASCs are capable of differentiating into chondrocyte-like cells in vitro under specific culture conditions and further maintaining the chondrogenic phenotype in vivo [14]. A recent report documented an interesting finding that chondrogenic induced ASCs in combination with fibrin glue could be used to repair articular cartilage defect in the weight bearing area in a rabbit model [27]. The combination of TGF-b1 and IGF-1 was adopted in the current work to induce the chondrogenic differentiation of ASCs. It has been reported by Longobardi et al. [28] that IGF-1 induced chondrogenic potential of bone marrow mesenchymal stem cells by stimulating proliferation, regulating apoptosis and inducing expression of chondrogenic markers in an independent way from that of TGF-b1. When combined together, the two growth factors have synergistic effects in stimulating chondrogenic differentiation of MSCs [28]. Although it remains controversial whether the chondrogenic ability of ASCs is comparable to that of BMSCs, it has been documented that a high-dose combination of the two growth factors can stimulate chondrogenic differentiation of ASCs to a similar extent as that of BMSCs [16].

3.5. Biochemical and biomechanical evaluation The GAG deposition of engineered cartilage was further quantified by biochemical analysis and also compared with that of the normal one. As shown in Fig. 7A, it was found that the amount of GAG in the engineered cartilage at 3 months reached w70% of that in the corresponding normal cartilage (p < 0.05). However, with the maturation of engineered cartilage till 6 months, GAG content in the regenerated cartilage was increased reaching w90% of that in corresponding native cartilage (p < 0.05). Significant difference was also observed between the GAG amount of the engineered cartilage at 3 months and 6 months (p < 0.05). The biomechanical property of the engineered cartilage was estimated by measuring its compressive modulus. The value was

Table 4 Results of histological grading for cartilage regeneration. Grading

Cell morphology Matrix staining Surface regularity Thickness of cartilage Integration with host cartilage Total

3 months (n ¼ 7) Control

Exp

Control

2.00(1–3) 0.86(0–2) 1.57(0–2) 0.57(0–1) 0.86(0–1) 5.86(4–8)#

2.29(2–3) 1.43(1–2) 2.43(2–3) 1.43(1–2) 1.14(0–2) 8.71(7–11)

1.14(0–2) 1.14(0–2) 0.43(0–1) 0.43(0–1) 0.28(0–1) 3.43(2–5)#,*

3.43(2–4) 2.57(2–3) 2.57(2–3) 1.71(1–2) 1.71(1–2) 12.00(8–14)

Each data represents mean and range (parentheses) of the scales. p < 0.01, versus the control group; *p < 0.05, versus the experimental group of 3 M.

#

6 months (n ¼ 7)

Exp

2690

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Fig. 5. Immunohistochemical staining for collagen type II in reparative tissue at 3 and 6 months post-implantation, respectively. The arrowheads indicate the interface between the reparative tissue and adjacent host cartilage. (A) Expression of collagen type II was found in part of neo-cartilage at 3 months after surgery. (B) Expression of collagen type II was revealed through the regenerated cartilage layer at 6 months after surgery. (A1) and (B1) are high-magnification views of central area of reparative tissue in full thickness from superficial cartilage to subchondral bone. Little expression of collagen type II was detected in the control group at 3 (C) and 6 months (D) post-surgery (Bar scales: 500 mm for A, B, 1000 mm for C and D; 200 mm for A1 and B1.).

Therefore, we applied the combination of TGF-b1 (10 ng/mL) and IGF-1 (100 ng/mL) in our study to induce chondrogenic differentiation of ASCs with the achievement of a good chondrogenesis of ASCs on PGA/PLA scaffold. Of note, it was also report by Hennig et al. [29] that a combination of TGF-b3 and BMP-6 significantly increased the reduced chondrogenic potential of ASCs treated with TGF-b3 alone. However, it was also found that chondrogenesis of ASCs induced with combination of TGF-b3 and BMP-6 was associated with hypertrophy in vitro and calcification in vivo [29]. Anyway, further work need to be done to explore the underlying mechanism related to chondrogenesis of ASCs in order to find a more efficient way. PGA is one of the most commonly used synthetic polymers in cartilage tissue engineering. To maintain its dimensional stability and enhance its mechanical properties, fibrous PGA meshes are always coated with solutions of PLA. Evaporation of the solvent for PLA would thus result in PGA/PLA composites with specific shapes, such as auricular [18], nasal [30] and trachea [31]. The feasibility of using PGA/PLA composite as scaffold to engineer cartilage tissue has been well documented in a variety of studies. It was also shown that adhesion and proliferation of chondrocytes on PGA fibers was significantly suppressed with the increase in the amount of PLA

added [32]. Therefore, the concentration of PLA solution to be added has to be lowered but sufficient to function as glue to maintain the structural stability of the PGA 3-D scaffold. In our study, 1.5% PLA in dichloromethane was used as the further lowering would result in an unstable configuration of the resultant scaffold (results not shown). It was observed by SEM that PLA of this concentration can wrap PGA fibers together and the shape of scaffold could be maintained well when they were kept in culture medium for as long as 5 weeks in this study. Zuk et al. [12] reported that chondrogenic differentiation of ASCs could only be induced in micromass culture but not in monolayer cultures. Thus, to get chondrogenic differentiated ASCs as seed cells in cartilage engineering, a procedure of micromass culture which was followed by enzymatic digestion for breaking apart aggregated cells has to be taken. This made the utilization of ASCs in cartilage tissue engineering more complicated and time consuming. On the other hand, it has been speculated that 3-D microenvironment offered by PGA scaffold rather than 2D monolayer culture conditions allowed more cell–matrix and cell–cell contact that is required for the enhanced chondrogenic differentiation of cells seeded on them [33]. Therefore, we seeded expanded ASCs directly in the PGA/PLA scaffold without growing the cells in

Fig. 6. Toluidine blue staining of repaired tissue at 3 and 6 months after implantation, respectively. The arrowheads indicate the interface between the reparative tissue and adjacent host cartilage. (A) At 3 months post-surgery, repaired tissue exhibited a metachromatic matrix deposition in the experimental group, while no obvious tidemark was observed. (B) At 6 months after repairing, distribution of GAG in the engineered cartilage appeared in a similar way as that in adjacent normal cartilage. A1 and B1 are highmagnification views of central area of reparative tissue in full thickness from superficial cartilage to subchondral bone in (A) and (B), while partial appearance of tide mark in B1 was observed. Little deposition of GAG could be detected in the defect area of the control group at 3 (C) and 6 months (D) post-surgery (Bar scales: 500 mm for A, B, 1000 mm for C and D; 200 mm for A1 and B1.).

L. Cui et al. / Biomaterials 30 (2009) 2683–2693

15

B

Experimental Normal

10

Compressive Moduli (Mpa)

GAG content (mg/g)

A

#* #

5

0

3M

6M Time

50

2691

Experimental Normal

*

40 30 20

#

10 0

3M

6M Time

Fig. 7. GAG content (wet weight) and biomechanical test of the engineered cartilage (Experimental) at 3 and 6 months, respectively. (A) GAG content of the engineered cartilage at 3 months reached about 70% (p < 0.05) of that of the corresponding normal ones, and further increased to nearly 90% (p < 0.05). (B) The compressive moduli of engineered cartilage at 3 months post-repair reached about 50% of that of the corresponding native one (p < 0.05), and further increased to nearly 90% (p < 0.05). A significant increase in both GAG content and compressive moduli of engineered cartilage at 6 months post-repair was observed compared with those of 3 months (p < 0.05) (*means significantly different compared with engineered cartilage at 3 months; #means significantly different compared with the corresponding normal cartilage).

micromass culture before seeding in order to observe whether ASCs could undergo chondrogenic differentiation in the 3-D microenvironment offered by this scaffold. By fluorescent DiO labeling and Hoechst 33258 quantitative determination, it was shown that the amount of ASCs on PGA/PLA scaffold kept on increasing up to 3 weeks. Deposition of abundant ECM was also detected by SEM observation. In addition, amounts of COL II and GAG, the two critical cartilage-specific ECM, were found to be significantly increased at 2 weeks post-seeding and maintained at a stable high level in the following 3 weeks. In a previous study, we have demonstrated that chondrogenic induced BMSCs proliferated well on PGA/PLA composite scaffold which was fabricated in the same way as the current one [9]. Taken together, these results suggested that PGA/ PLA composite is well suited for initial adhesion and subsequent proliferation of chondrogenic induced adult MSCs as well as their cartilage-specific ECM deposition. One major advantage of using PGA/PLA scaffold for cartilage tissue engineering is its suitable degradation rate which matches the kinetics of new cartilage formation in vivo [30,32,33]. The complete degradation of non-woven PGA scaffold is reported to be finished over a period of 2 months in vivo [30,33]. In the present study, no remnant of un-degraded PGA fibers could be observed histologically either in the experimental or control groups at 3 month post-implantation. Due to its fast degradation, PGA scaffold was found to be able to accelerate chondrogenesis of constructs prepared from dedifferentiated chondrocytes and PGA, as the accumulation of the deposited cartilage-specific ECM and expression of marker genes both in vitro and in vivo were significantly enhanced compared with those of constructs prepared from PLGA (copolymer of L-lactide and glycolide, 90/10) and same cells [33]. It was also proposed that early degradation of PGA fibers might render a positive effect on chondrogenesis by leaving new spaces for cells to further fill in and produce new intercellular matrix which in turn may facilitate the formation of more cell–matrix and cell–cell contact. The secure fixation of cell–scaffold constructs in defects by transosseous biodegradable sutures is another critical issue for success in cartilage regeneration in vivo as it ensured good integration of engineered cartilage with its surrounding native cartilage and subchondral bone. According to Knecht et al. [23], a trans-osseous

suture fixation of chondrocyte–PGA complex in situ could dramatically improve its ability to withstand high mechanical loading. Thus, we also stabilized the implant by two-crossed trans-osseous fixation using biodegradable sutures and no dislocation of the implant occurred in the following study. However, the preparation of non-woven PGA scaffold in this study was hand-made without standardization, which resulted in unorganized distribution of individual fibers. In turn, initial distribution of cells in the construct would also be random, which might contribute to the formation of engineered cartilage with uneven thickness. Further development on the processing techniques is required to improve the uniformity of scaffolds made from PGA unwoven fibers. Although most of clinical injuries occurred in the weight bearing area of articular cartilage, it was also indicated that damage to cartilage even in low weight bearing area can result in osteoarthritis [34]. Due to the loss of cartilage layer, its articulating, lubricating and load distributing effects also lost. Thus, as evidenced from the results in the control group in this study that the subchondral bone as well as its surrounding normal cartilage collapsed due to the fact that the subchondral bone would then directly be subjected to load. On the other hand, it was reported that morphology, proliferation and differentiation of ASCs could be essentially regulated by biomechanical loading [35]. Thus, to elucidate whether ASCs could be used to regenerate cartilage in a microenvironment with low biomechanical stimulation, we transplanted constructs of chondrogenic induced ASCs and PGA/ PLA scaffolds to repair articular cartilage defects in the non-weight bearing area. At 6 months post-implantation, the defects had been well repaired by engineered cartilage with a similar appearance to the surrounding native tissue in color and smoothness. The gross cross-sectional morphology also demonstrated an ideal interface between the engineered cartilage and its adjacent native cartilage/ cancellous bone tissue. In contrast, defects repaired with PGA/PLA alone were filled with fibrous tissue at 3 months and collapse of the surrounding native cartilage together with subchondral bone was observed at 6 months post-implantation. Moreover, it was also demonstrated in our previous work that no obvious repair or only fibrotic tissue was observed in similar defects without any treatments within 6 months [9]. Hence, such an additional control group was not included in this work.

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As shown by histological morphology, the defect treated with ASCs based engineered cartilage at 6 months post-implantation has an architecture that is characteristic of mature hyaline cartilage. Chondrocytes within the superficial zone are exclusively single cell units and have a flat profile orientated parallel to the articular cartilage surface, whereas those in the deeper zone arranged themselves in distinct clusters or cell columns aligned vertically to the cartilage surface. As one of the most important components of extracellular matrix in hyaline cartilage, COL II was found to be expressed in the engineered tissue at 3 months after surgery and its expression was enhanced when observed at 6 months. The result of toluidine blue staining also documented the deposition of GAG in the engineered cartilage, which is another major component responsible for maintaining the water content and biomechanical property of hyaline cartilage. Of note, the engineered cartilage showed good integration with its subchondral bone, which is not only essential for the fixation of implant but also strengthens its ability to resist shear force resulted from the joint movement in vivo. Since the biochemical composition (GAG quantification) and biomechanical property (compressive modulus) of engineered cartilage at 6 months post-implantation were similar to those of a corresponding normal one, it indicated that the defects were repaired both structurally and functionally. Taken together, these results suggest that regeneration of a specific hyaline cartilage to repair a respective defect could be achieved using chondrogenic induced ASCs with a biodegradable scaffold via tissue engineering techniques. Finally, our results revealed that the repair process took a relatively long time to undergo tissue remolding in cellular and compositional re-organization for achieving satisfactory repair. After 3 months of repair, cells in the repaired area distributed singly with isotropic organization, while at 6 months cellular clusters and columns with highly anisotropic arrangement were observed. Such structural development is accompanied with a decrease in the numerical density of cells and in the thickness of the engineered articular cartilage layer. At the same time, an increase in GAG content in the engineered cartilage should be mentioned from 3 to 6 months after implantation. Accordingly, a better biomechanical property (compress modulus) of the repaired cartilage was achieved at 6 months post-surgery. 5. Conclusion The current study thus demonstrated that an engineered cartilage composed of autologous chondrogenic differentiated ASCs on fibrous PGA/PLA scaffold could be successfully obtained and further applied to repair an articular cartilage defect of 8 mm in diameter and 6 mm in depth at non-weight bearing area in a porcine model. The chondrogenic differentiation of ASCs on the scaffold was ascertained by the enhanced synthesis of two critical cartilagespecific ECM, COL II and GAG. At 3 months post-implantation, the generated neo-cartilage in the defect repaired with the autologous chondrogenically induced ASCs/scaffold construct integrated well with its surrounding normal cartilage and subchondral bone. At 6 months post-surgery, the engineered cartilage became more mature with a similar appearance to the surrounding native tissue in color, smoothness and histological microstructure. The distribution of COL II and GAG in the regenerated cartilage, and the compressive moduli of the neo-tissue all underwent a vivid remolding process with in vivo repairing time. In contrast, defects repaired with scaffold alone were filled with fibrous tissue at 3 months and collapse of the surrounding tissue was observed at 6 months post-implantation. Hence, the successful repair substantiates the potential of using chondrogenic induced ASCs and PGA/PLA scaffold for cartilage regeneration.

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