Biphasic Mechanical Properties of in vivo Repaired Cartilage

Journal of Bionic Engineering 12 (2015) 473–482

Biphasic Mechanical Properties of in vivo Repaired Cartilage Qin Lian1, Cheng Chen1, Marie Chantal Uwayezu1, Weijie Zhang1, Weiguo Bian2, Junzhong Wang1, Zhongmin Jin1,3 1. State Key Laboratory for Manufacturing Systems Engineering, Xi’an Jiaotong University, Xi’an 710054, China 2. Orthopaedics Department, First Affiliated Hospital, Xi’an Jiaotong University, Xi’an 710061, China 3. Institute of Medical and Biological Engineering, School of Mechanical Engineering, University of Leeds, Leeds LS2 9JT, UK

Abstract In the fast growing field of scaffold-based tissue engineering, improvement on the mechanical properties of newly formed tissues, e.g. the repaired cartilage, has always been one of the core issues. Studies on the correlations among scaffold composition, in vivo morphological changes of the construct, and the finite deformation behaviors of new tissues (e.g. creep and stress-relaxation, and equilibrium response), have attracted increasing interests. In this paper, the correlations between the compressive biphasic mechanical properties (i.e., equilibrium elastic modulus E and permeability coefficient k) of 3D printing scaffold (consisting of collagen and β-tricalcium phosphate) and the proteoglycans (PGs) concentration of the repaired cartilages after 24 weeks, 36 weeks and 52 weeks of scaffold implantation were investigated. Results indicated that the repaired cartilage covered the entire cartilage surface of large cylindrical osteochondral defects (10 mm in diameter × 15 mm in depth) on the canine trochlea grooves after 24 weeks. The equilibrium elastic modulus of the repaired cartilage reached 22.4% at 24 weeks, 70.3% at 36 weeks, and 93.4% at 52 weeks of the native cartilage, respectively. Meanwhile, the permeability coefficient decreased with time and at 52 weeks was still inferior to that of the native cartilage in one order of magnitude. In addition, the amount of glycosaminoglycans (GAGs) of repaired cartilage increased constantly with time, which at 52 weeks approached to nearly 60% of that of native cartilage. 3D printed scaffolds have potential applications in repairing large-scale cartilage defects. Keywords: biphasic mechanical properties, PGs, repaired cartilage, osteochondral scaffolds, 3D printing Copyright © 2015, Jilin University. Published by Elsevier Limited and Science Press. All rights reserved. doi: 10.1016/S1672-6529(14)60138-4

1 Introduction As a load-bearing material, articular cartilage has unique composition and morphology as well as good biomechanical properties[1–3]. It is considered to be composed of an interstitial fluid phase (including 68% ~ 85% water and dissolved electrolytes) and a porous solid matrix phase (including 10% ~ 20% of collagen, 5% ~10% of PGs, chondrocytes and other proteins)[4–7]. It is known that the compressive stiffness of cartilage is positively related to the PGs concentration and negatively related to the water content, while the collagen is predominantly responsible for the tensile and shear properties[8]. The cartilage deformation under compression mainly depends on the flow of the interstitial fluid through the porous solid phase[4,7]. Under low-rate dynamic loading conditions, interstitial fluid carries the major responsibility on load bearing since it cannot be Corresponding author: Qin Lian E-mail: [email protected]

squeezed out immediately from the tissue due to the low permeability of cartilage, which depends on the PGs concentration[8]. As described by the biphasic theory[4,9–11], the fluid phase of cartilage is simplified as impressive and non-viscous while the solid phase is porous, permeable and incompressible, the biphasic mechanical properties (e.g. equilibrium elastic modulus E and hydraulic permeability coefficient k) can be calculated from the experimentally obtained compressive creep and stress-relaxation behaviors of cartilage via a nonlinear regression technique. A synovial joint model with biphasic cartilages can be used to address the role played by the interstitial fluid in the tribology and contact mechanics of the joint. Contact mechanics studies have further demonstrated that more than 80% of contact stresses are supported by the interstitial fluid pressurization under physiological loading conditions, which can protect the solid phase of cartilage from excessive

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stresses and strains[9,12]. Meanwhile, the correlations between the cartilage biphasic mechanical properties and the biochemical and morphological characteristics have been attracting increasing interest. Using autologous implants for in vivo repairing of rabbits’ femoral condyle defects, Mow et al.[13] found that the repaired tissues after 8 weeks of implantation were similar to the normal articular cartilage in biphasic properties and composition although there was incomplete bonding between the repair tissue and the adjacent cartilage. The osteoarthritis studies have demonstrated that the biomechanical properties not only determine the functional behavior of articular cartilage, but also seem to be more sensitive to early degenerative variations than morphological properties, since biomechanics is determined by the composition and structural arrangement of the extracellular matrix[14]. Scaffold-based tissue engineering has made significant progress on the repair of cartilage and osteochondral defect[15–18]. In addition to the proofs of morphological and histological characteristics of the repaired tissue, biomechanical properties have also been investigated to demonstrate the restored functions. Using the biodegradable PLG (50:50 poly-DL-lactide-coglycolide) scaffolds coupled with different concentrations of TGFβ to repair the femoral osteochondral defects of adult goats (7 mm in diameter × 7 mm in depth), Athanasiou et al.[19] found that the biphasic creep biomechanical properties of neocartilages with high TGFβ resulted in better structural integrity, although incomplete grossly healing occurred at 16 weeks postoperatively. For the treatment of osteochondral defects (7 mm in diameter × 25 mm in depth) in the ovine knees using β-tricalcium phosphate (β-TCP) cylinders seeded with autologous chondrocytes, Mayr et al.[20] found that the mechanical properties of the TCP-chondrocytes constructs increased significantly with time after implantation and were similar to those of the native cartilage after 52 weeks, although the newly formed cartilage in the central transplanted area had a lower International Cartilage Repair Society Visual Assessment Scale Score (ICRS) than healthy cartilage. Hu et al.[21] used a self-assembling process to produce tissue engineered constructs over agarose and TCP in vitro and found that the self-assembled constructs reached over one-third of the stiffness of native cartilage, while the permeability values of constructs were close to those of the healthy

tissue. Moreover, Jansen et al.[22] put two kinds of porous poly(ethylene oxide terephthalate) and poly(butylene terephthalate) (PEOT/PBT) scaffolds with different biomechanical properties (equilibrium modulus, dynamic stiffness at 0.1 Hz) into the osteochondral defect sites of 6-month-old rabbit medial femoral condyles to compare their in vivo healing responses. They found that PEOT/PBT scaffolds with low mechanical properties were superior in 3-month of cartilage repair tissue formation in the defects. Therefore, the combination of morphological and biochemical assessment would help the surgeons and engineers to effectively determine the scaffold design and fabrication, specifically for cartilage and osteochondral repair. 3D printing has promising potential to prepare integrated multiple material scaffolds which would enhance the cartilage regeneration capability for large size cartilage and osteochondral defects[23,24]. In this study, we used the previously developed 3D printing scaffolds consisting of collagen and β-TCP (Col/β-TCP, with good bonding interfaces between materials)[25] to repair large osteochondral defects (10 mm in diameter × 15 mm in depth) in trochlea grooves of canine distal femur, which were followed up for 52 weeks postoperatively. The autogenous bone marrow Mesenchymal Stem Cells (MSCs) were infilled into the scaffolds during the implantation. This study aimed to quantitatively investigate the biphasic mechanical properties of the repaired cartilage via curve-fitting of the creep data obtained from unconfined compressive tests; to investigate whether there was a correlation among the biphasic mechanical properties and PGs concentration as well as the water content of the repaired cartilage.

2 Materials and methods 2.1 Material 2.1.1 Native cartilage sampling Distal femurs were obtained from 3 healthy adult canine joints which were freshly dissected, immediately wrapped by 0.15 M saline-wetted cloth and preserved at −20 ˚C. Nine cylindrical specimens (6 mm in diameter × 7 mm in depth) were excised at three positions of distal femurs using a manual trephine as shown in Fig. 1. All specimens were thawed in saline for approximately 2 hours prior to experimental testing.

Lian et al.: Biphasic Mechanical Properties of in vivo Repaired Cartilage

2.1.2 Repaired cartilage sampling Fifteen healthy mongrel dogs (2–3 years old skeletal maturity; 15 Kg – 30 Kg) used in this study were obtained from experimental animal center of Xi’an Jiaotong University. All animal experiments were approved by the Laboratory Animal Care Committee of Xi’an Jiaotong University. Under sterile conditions, thirty prepared 3D printing scaffolds[25] were implanted in the cylindrical osteochondral defects (10 mm in diameter × 15 mm in depth) in the bilateral femoral trochlea grooves. Animals were euthanized at 24 weeks, 36 weeks and 52 weeks postoperatively. The neotissues were sampled from canine distal femurs in the implantation areas for mechanical tests and biochemical as well as histology staining analyses (listed in Table 1). Moreover, the water content of repaired cartilage was tested using three samples at 24 weeks which is not listed in Table 1.

−20 ˚C cryopreservation. Before the mechanical experiments, the specimens were immersed in the PBS solution and thawed for 2 hours. A computer-controlled universal testing machine (CMT6503, SANS) was used to conduct the unconfined compression tests. The specimens and the indenter were entirely immersed in PBS solution as shown in Fig. 2. A 50 N pressure sensor was used with an impermeable stainless steel indenter. The compressive force (F) was set from 0.025 N to 5 N under 0.1 N·s−1 loading and then sustained for 1800 seconds to obtain the time-displacement data. In order to neglect the load increase during the ramp phase, the first experimental sampling was taken at 2 seconds before the load reached 5 N and then the measurements were continued until 1800 seconds under a constant creep load of 5 N.

2.2 Histological staining and microscopy Samples at 24 weeks, 36 weeks and 52 weeks were processed for histological evaluation. All samples were processed as previously described in Ref. [23]. In brief, all samples were fixed in 10% (v/v) buffered formalin for 48 hours, then decalcified in 10% EDTA, dehydrated, embedded in paraffin and sectioned at 4 μm, stained with safranin O/fast green staining protocol. 2.3 Mechanical experiments All cylindrical specimens for the unconfined compression tests were trimmed into a cylindrical shape with 6 mm in diameter and 7 mm in depth. The trimmed specimens had smooth and flat top surface parallel to the bottom. Three specimens were tested at 24 weeks, while six specimens were tested at 36 weeks and at 52 weeks, respectively. In addition, three specimens of the 3D printing scaffolds were trimmed for the unconfined compression test with the same procedure as the repaired tissue. Meanwhile, the shear experiment as described previously in Ref. [24] was conducted to investigate the bonding strength of the interface between the collagen and β-TCP of 3D printing scaffolds. Cartilage thickness of each specimen was measured by an optical image measuring instrument (VME040030020000, 3D FAMILY). All of the specimens were then immersed in Phosphate Buffer Saline (PBS) solution to prevent dehydration and preserved in

475

c

a

b

5 mm

Fig. 1 Sampling location of (a) medial condyle, (b) lateral condyle and (c) trochlea. Table 1 Detailed sampling protocol Implantation Time

Number of specimens for mechanical and biochemical tests

Number of specimens for histological staining

Defects number

Animal number

24 weeks

3

4

10

5

36 weeks

6

4

10

5

52 weeks

6

4

10

5

Fig. 2 Schematic diagram of unconfined compression test.

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Calculation of cartilage biphasic mechanical parameters The finite element analysis was used to predict the deformation response of the osteochondral specimens under the unconfined compression test. The predicted deformations (i.e. the creep displacement) for the unconfined compression tests using the commercial package ABAQUS (Version 6.10-1, Dassault Systems, France) were based on the biphasic theory[4–6,11,26,27]. An axisymmetric osteochondral cylinder model with the biphasic cartilage layer against a rigid impermeable flat-ended indenter was implemented in the ABAQUS. The flat-ended indenter was set as a rigid body with a Reference Point (RP) at its center. The biphasic cartilage layer was modeled as poroelastic with a void ratio of 4 (that is, the ratio of the 80 % interstitial fluid phase of cartilage to 20 % solid matrix phase). Cartilage was set with the initial elastic modulus E0 = 1 MPa, Poisson's ratio of 0 and the initial permeability coefficient k0 = 1.5 × 10−15 m4·(N·s)−1. The underlying bone firmly fixed to the cartilage was assumed to have a Young’s modulus of 2 GPa and a Poisson's ratio of 0.2[26]. Cartilage was meshed as CAX4P (Four-node bilinear displacement and pore pressure element). In the vertical direction (i.e., in the Y or axial direction), uniform elements were set with an element size of 0.1 mm. In the horizontal direction (i.e., in the X direction), non-uniform elements were set. The seeds were set by number and single bias, the horizontal number of elements was 5 times of the cartilage diameter and the bias ratio of 4. Bone was meshed as CAX4 (Four-node bilinear axisymmetric quadrilateral element). The seeds were set by number with the vertical number of elements of 5 and the horizontal number was set in the same way as the cartilage. The bottom nodes of the lower bone were constrained in both the horizontal and vertical directions. The nodes on the axis of the osteochondral samples were constrained in the horizontal direction. The pore pressure on the nodes of cartilage outer edge was set as zero so that the interstitial fluid flowed freely. The RP of the indenter was imposed with all constraints in the horizontal direction and a constraint in the axial rotation. The indenter was moved vertically downwards to impose a compressive displacement in the cartilage over a ramp time of 2 seconds and then held with the constant load of 5 N for further 1800 seconds.

The cartilage biphasic parameters were determined by matching the finite element prediction with the experimental displacement[26,28]. Briefly, by comparing and minimizing squared errors between the finite element prediction and the experimental data, the equilibrium elastic modulus E and permeability coefficient k were determined by curve-fitting the final 30% of experimental deformation data representing at the equilibrium state according to the previous studies[10,29–31]. Nonlinear least-squares were solved by Matlab (Version R2009a) Lsqnonlin function (Eq. (1)) and the degree of fitting r2 function (Eq. (2)).

min f (x) 2 = min ( f12 ( x ) + f 22 ( x ) + ... + f n2 ( x ) ) , (1) 2

x

x

⎡E j ⎤ where, f i = yi − si , x = ⎢ ⎥ ; n is the number of dis⎣kj ⎦

placements acquired from the experiment for a creep test; yi (mm) is the experimental deformation value (the creep displacement in this study); si is the finite element predicted value (mm); Ej (MPa) and kj (m4·(N·s)−1) are the equilibrium elastic modulus and permeability coefficient inputs respectively in the jth iteration computing. The degree of fitting r2 function is given as n

r2 = 1−

∑( y i =1 n

i

− si )

2

− ⎛ ⎞ ⎜ yi − y ⎟ ∑ ⎝ ⎠ i =1

2

.

(2)

ABAQUS command was iteratively called by Matlab function. That is, the minimized function was terminated when the variance between the experimental data and the FE prediction was minimal or the number of iteration reached the setting value. Then the optimal values of equilibrium elastic modulus Ej, permeability coefficient kj and the degree of fitting r2 were determined. 2.5 Biochemical and water content assessments After mechanical tests, the cartilage peeled from the samples was further processed for biochemical tests as previously describled in Ref. [23]. Briefly, the cartilage was digested and tested with BlyscanTM Glycosaminoglycan Assay Kit (Biocolor) and SircolTM assay Kit (Biocolor) to investigate the contents of GAGs and collagen. The content of GAGs was determined as PGs are macromolecules that consist of a core protein car-

Lian et al.: Biphasic Mechanical Properties of in vivo Repaired Cartilage

rying a number of covalently linked GAG chains[6]. Each cartilage block with more than 100 mg of weight was taken from the specimen after the unconfined compression test in order to investigate its water content. Briefly, each block was soaked in 0.9% saline for 48 hours. The superficial water was then removed using an absorbent paper, prior to weighing the block on an electronic balance (ES225SM-DR, Precisa). The block was then dried for 48 hours in a vacuum drying oven at 60 ˚C following by a further weighing. The water content was calculated by the wet weight minus dry weight and divided by the wet weight.

477

Fig. 3 Repaired femurs at different time of (a) 24 weeks, (b) 36 weeks and (c) 52 weeks.

(a)

(b)

(c)

(d)

3 Results

3.2 Biochemical and water content assessment Fig. 5 demonstrates that the GAGs amount of repaired cartilage increased constantly with time, and approached to nearly 60% of that of native cartilage at 52 weeks. Meanwhile, the collagen content of the repaired cartilage decreased as time went on during the initial 36 weeks, and was approximately close to the native level at 36 weeks and 52 weeks. For the water content, Table 2 indicates that both the native and the repaired cartilage were in a range from 74% to 82%. No significant difference was found between the two groups.

Fig. 4 Repaired cartilage and native knee cartilage in safranin-O staining (× 200): (a) Repaired cartilage at 24 weeks; (b) repaired cartilage at 36 weeks; (c) repaired cartilage at 52 weeks; (d) native cartilage.

GAGs amount ( g·mg−1)

0.5 0.4 0.3 0.2 0.1 0.0

Collagen amount ( g·mg−1)

3.1 Gross morphological and histological observation No obvious immunologic or infectious complications were observed during the whole experiment. As shown in Fig. 3, the defects at 24 weeks were entirely covered with a white tissue which filled 50% defect volume, although the incomplete bonding interface between the repaired tissue and the adjacent native cartilage was largely visible. At 36 weeks, nearly 70% of the defect was filled with the white opaque tissue but with rough surface. The color of the repaired cartilage at 52 weeks was close to that of normal one, but the filling volume was little changed compared with that of 36 weeks. No obvious cartilage delamination of samples was observed during the sampling process. PGs were stained as red by Safranin-O staining. As shown in Fig. 4, the PGs concentration (red area) was significantly increased with time from 24 weeks to 52 weeks in vivo. The intensity of PGs staining of the repaired cartilage at 52 weeks was approaching to that of native cartilage.

24 weeks

36 weeks 52 weeks Groups (a)

Native

60

40

20

0

24 weeks

36 weeks 52 weeks Groups (b)

Native

Fig. 5 (a) GAGs and (b) collagen contents of repaired cartilage (mean ± rms).

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3.3 Biomechanical properties 3.3.1 Bonding strength of osteochondral scaffolds The bonding strength of osteochondral composite scaffolds was estimated through maximum shear force and interfacial integration strength[24,32] which are listed in Table 3. The interfacial integration strength of our Col/β-TCP osteochondral scaffolds under the wet condition was approximately half of that under the dry condition, but nearly 60 folds higher than that of the previously osteochondral composite scaffolds integrated by fibrin[32].

3.3.2 Cartilage biphasic properties Biphasic parameters of native articular cartilage are listed in Table 4. Experimental curves and the fitting curves of the cartilage specimens were in good agreement as shown in Fig. 6. The degrees of fitting r2 (the minimal value of all of specimen in one group is set as the degree of fitting of the group) in media condyle, lateral condyle and trochlea were greater than 0.99, 0.96 and 0.95, respectively. The biphasic mechanics characteristics of the repaired cartilage are shown in Table 5. The experimental curves and fitting curves are in good match. The degrees of fitting at 24 week, 36 week, 52 week of repaired cartilage are greater than 0.95, 0.88 and 0.92, respectively.

4 Discussion The mechanical properties of articular cartilage determine its ability to perform important mechanical functions in vivo[5]. According to the biphasic theory, the frictional resistance generated by the interstitial fluid flow through the porous matrix solid ensures the compressive load bearing ability of articular cartilage[4–7]. Previous studies have pointed out the difference of cartilage biphasic parameters values cannot be attributed to the size of specimens but can come from various compressive tests (i.e., indentation, unconfined and confined compression experiments)[33]. Using unconfined compression experiment, the equilibrium elastic modulus and permeability coefficient of native canine cartilage in this study were E = 2.47 ± 1.02 MPa, k=(2.89 ± 1.31) ×10−15 m4·(N·s)−1 ×10−15 m4·(N·s)−1 for lateral condyle; E =0.79 ± 0.28 MPa, k = (2.47±1.02) ×10−15 m4·(N·s)−1 for medial condyle; E=1.17 ± 0.21 MPa, k=(2.14 ± 0.31) ×10−15 m4·(N·s)−1 for trochlea, respectively. Compared with the previous study by indentation[34] (i.e., E=0.81±0.32 MPa,

Table 2 Cartilage water contents (mean ± rms) Samples

Water content

Significance level

Trochlea

76% ± 1%

P > 0.05

Medial femoral condyle

77% ± 2%

P > 0.05

Lateral femoral condyle

76% ± 2%

P > 0.05

Repaired cartilage at 24 Weeks

78% ± 4%

P > 0.05

Table 3 Bonding strength of osteochondral scaffolds (mean ± rms)

Wet Col/β-TCP

4.08 ± 1.26

Interfacial integration strength (MPa) 0.052 ± 0.016

Dry Col/β-TCP

10.76 ± 1.96

0.137 ± 0.025

10

Fibrin integration[32]

0.08 ± 0.01

0.00084 ± 0.00011

11

Interface integration

Maximum shear force (N)

Diameter (mm) 10

Table 4 Biphasic characteristics and thickness of native cartilages (mean ± rms) Specimen location (Specimen number, n) Medial condyle (n = 3) Lateral Condyle (n = 3) Trochlea (n = 3)

Equilibrium elastic modulus E (MPa)

Permeability coefficient k (×10−15m4·(N·s)−1)

Thickness of cartilage h (mm)

0.79 ± 0.28

2.47 ± 1.02

0.66 ± 0.22

0.76 ± 0.18

2.89 ± 1.31

0.64 ± 0.28

1.17 ± 0.21

2.14 ± 0.31

0.75 ± 0.23

r2 > 0.95

0.07 0.06 Displacement (mm)

478

0.05 0.04 0.03

Experiment curve FE fitting curve

0.02 0.01 0.00

0

200

400 600 800 1000 1200 1400 1600 1800 2000 Time (s)

Fig. 6 A typical curve-fitting result of the canine trochlea cartilage.

k = (0.77 ± 0.56 ) ×10−15 m4·(N·s)−1 for lateral condyle; E = 1.59 ± 0.39 MPa, k = (0.80 ± 0.78) × 10-15 m4(N·s)-1 for medial condyle; E = 0.57 ± 0.15 MPa, k = (0.93 ± 0.84) × 10-15 m4(N·s)-1 for trochlea), the permeability coefficient k in our study is slightly higher than that by indentation from different locations but both are of the same order of magnitude, while there is slight difference between elastic modulus values. Except from the variety of animal species, these differences may occur due to the various compression tests. So we evaluated the repaired cartilage compared to the native one under the same unconfined experiment.

Lian et al.: Biphasic Mechanical Properties of in vivo Repaired Cartilage

In this study, we used the previously developed 3D printing scaffolds consisting of collagen and β-TCP (Col/β-TCP, with good bonding interface between the two materials)[25] to repair the large osteochondral defect (10 mm in diameter × 15 mm in depth) in the trochlea groove of canine distal femur for 52 weeks postoperatively. In order to further examine the trend of the biphasic properties of the repaired cartilage, the experimental data from Table 4 and Table 5 were plotted together as a function of implantation time as shown in Fig. 7. The equilibrium elastic modulus of repaired cartilage reached 22.4% at 24 weeks, 70.3% at 36 weeks, 93.4% at 52 weeks of the native cartilage, respectively. Although the equilibrium elastic modulus of repaired cartilage at 24 weeks was lower than native cartilage, the equilibrium elastic modulus gradually increased with time and repaired cartilage at 52 weeks had reached to the level of native cartilage. It is clear that the repaired cartilage biphasic properties are significantly improved with time, i.e., the elastic modulus is increased while the permeability is decreased. Interestingly, the elastic modulus, ~ 0.2 MPa of the repaired cartilage at 24 weeks is similar to that (~ 0.3 MPa) of the collagen-gel top of scaffolds, but the repaired cartilage at 52 weeks with an elastic modulus of ~0.9 MPa is close to the native tissue (~1.2 MPa). Moreover, permeability coefficient of the repaired cartilage at 52 weeks was inferior to that of the native cartilage by one order of magnitude, but was significantly superior to that of the collagen-gel top of scaffolds. This indicates that reasonable elastic modulus and high permeability of collagen-scaffold can be beneficial for the cartilage repairing, though the current Col/β-TCP construct (the Col/β-TCP scaffolds infilled with bone marrow MSCs) yet lacks the ability to fully constrain the fluid inside the solid matrix under a load bearing condition. The water content of cartilage determines the ratio of solid phase volume to fluid phase volume and plays an essential role in maintaining the compressive stiffness of cartilage[8]. No water content data of repaired cartilage at 36 weeks and 52 weeks were available in this study due to the lack of samples. Nevertheless, as demonstrated in Table 2, the water content of the native and repaired cartilage at 24 weeks postoperatively fell into a similar range (74% to 82%) with no statistic difference. Based on the abovementioned facts, the inter-

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stitial fluid phase of cartilage was set as 80% and the void rate was set to be 4 in this study. Moreover, since the ratio of solid phase volume to fluid phase volume affects the analytic value of E, k under creep compressive test[4,6,11], the measuring technique for water content of cartilage will be developed in the future work by the improvement of measuring precision and reduction of sample dosage. Previous studies have pointed out that collagen is predominantly responsible for the tensile and shear properties, while the PGs concentration is positively related to the compressive stiffness of cartilage[6,8]. Biochemical results (Fig. 5) demonstrates that the collagen content of repaired cartilage decreased but the GAGs amount of repaired cartilage increased constantly with time, and at 52 weeks approached to nearly 60% of that of native cartilage. Collagen contents of the repaired cartilage at 36 weeks and 52 weeks were close to those of the native tissue. However, the tensile and shear properties of repaired cartilage were not included in this study because of the lack of samples. Consistent with GAGs results, the staining intensity of PGs in red area increased significantly with time as shown in the Table 5 Biphasic parameters and thickness of repaired cartilages (mean ± rms) Implantation of time (Specimen number, n) 0 week (n = 3)

Equilibrium elastic modulus E (MPa)

Permeability coefficient k (×10−15m4·(N·s)−1)

Thickness of cartilage h (mm)

0.32 ± 0.02

1171.10 ± 182.32

2.7 ± 0.06

24 weeks (n = 3)

0.20 ± 0.04

133.93 ± 58.41

1.59 ± 0.36

36 weeks (n = 6)

0.64 ± 0.40

29.05 ± 19.96

1.13 ± 0.54

52weeks (n = 6)

0.85 ± 0.43

23.46 ± 12.30

1.52 ± 0.40

Fig. 7 A typical curve-fitting result of the canine trochlea cartilage.

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histological results (Fig. 4), approaching to that of native tissue at 52 weeks postoperatively. Combined with the biphasic mechanical and biochemical results of the repaired cartilage, i.e. the elastic modulus increased while permeability coefficient decreased at equilibrium with time and the mechanical results at 52 weeks postoperatively were close to those of native cartilage, the PGs concentration was positively related to the equilibrium elastic modulus of repaired cartilage. The repaired cartilage has progressively restored the ability to bear the compressive load by improving both the composition and the biphasic properties with time. The combination of morphological and mechanical assessment would help the surgeons and engineers to effectively select proper scaffolds for the cartilage and osteochondral repair. In this study, the gross appearance and histological results were found not to be fully restored as shown in Fig. 3 and Fig. 4. However, the surface color of the repaired cartilage at 52 weeks was close to that of the native one, and no cartilage delamination of samples was observed. In addition, the biomechanical properties of repaired cartilage gradually tended to those of the native cartilage with implantation time, the compressive load bearing capability of repaired cartilage was preliminarily restored. It implies the 3D printing osteochondral scaffolds can enhance the functional cartilage restoration. It has been widely accepted that the stable interface between the cartilage part and the bone part is critical to the integrative cartilage repair using osteochondral scaffolds[5,24,27,32,35,36]. The interfacial integration strength of our 3D printed scaffold was nearly 62 folds higher than that of fibrin integrated ones[32] under wet condition, and 163 folds higher under dry condition. Hence, our 3D printing osteochondral scaffolds with the enhanced bonding interface might be beneficial to the repair process of cartilage.

paired cartilage increases with time while the permeability decreases; consistent with the equilibrium elastic modulus, PGs content of the repaired cartilage also increases with time. Therefore, 3D printed scaffolds have potential applications in repairing large-scale cartilage defects.

Acknowledgment This work was supported by grants from the Native Science Foundation of China (Nos. 51323007, 51375371 and 51075320), the National High Technology Research and Development Program of China (No. 2015AA020303) and the Fundamental Research Funds for the Central Universities. The authors would like to acknowledge the contributions of Dichen Li, Manyi Wang, Yongmei Chen and Yusheng Qiu of Xi’an Jiaotong University.

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