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Tissue Engineering and Regenerative Medicine

A biomimetic bacterial cellulose-enhanced double-network hydrogel with excellent mechanical properties applied for the osteochondral defect repair. Xiangbo Zhu, Taijun Chen, Bo Feng, Jie Weng, Ke Duan, Jianxin Wang, and Xiaobo Lu ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.8b00682 • Publication Date (Web): 09 Sep 2018 Downloaded from http://pubs.acs.org on September 9, 2018

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ACS Biomaterials Science & Engineering

A biomimetic bacterial cellulose-enhanced doublenetwork hydrogel with excellent mechanical properties applied for the osteochondral defect repair. †

†

†

†

†

Xiangbo Zhu , Taijun Chen , Bo Feng , Jie Weng ,Ke Duan , Jianxin Wang* †

，†

and Xiaobo Lu*

，‡

Key Laboratory of Advanced Technologies of Materials, Ministry of Education, School of

Materials Science and Engineering, Southwest Jiaotong University, Chengdu 610031, P.R. China ‡

Department of Bone and Joint Surgery, Affiliated Hospital of Southwest Medical University,

Luzhou, Sichuan, 646000

ABSTRACT

Although hydrogels based on biopolymers show many advantages, the low mechanical properties limit their applications in osteochondral tissue engineering. In this study, one part of our work aimed at preparing a high strength bio-hydrogel by using a double-network (DN) hydrogel system, which consisted of two interpenetrating polymer networks composed of γglutamic acid, lysine and alginate, and meanwhile by incorporating bacterial cellulose into the

*

Correspondence to: [email protected].

J.

Wang;

E-mail:

[email protected]

or

X.

Lu;

E-mail:

ACS Paragon Plus Environment

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DN structures. The results showed that compression modulus of the resultant hydrogel (0.322MPa) was comparable with that of natural articular cartilage and swelling degree was greatly

depressed

by

using

these

strategies.

On

this

basis,

a

bilayer

hydrogel

scaffold based on the bionics principle for osteochondral regeneration was fabricated via chemical and physical crosslinking. Additionally, hydroxyapatite (HA) particles with two different sizes were introduced into the bilayer hydrogels, respectively: micro-HA in the top layer for promoting cartilage matrix deposition and HA nanocrystals in the bottom layer for enhancing compression modulus and osteogenesis. The osteochondral defect model of rabbits was used to evaluate the repair effect of the scaffolds with the bilayer structure and the results showed such as-synthesized scaffolds had a good osteochondral repair effect. KEYWORDS: OSTEOCHONDRAL; HYDROGEL; CARTILAGE TISSUE ENGINEERING; BIOMIMETIC; MECHANICAL PROPERTIES; SWELLING 1. INTRODUCTION Limited metabolic and biosynthetic activities for mature chondrocytes lead to low selfrepairing ability of articular cartilage. Overall, although current surgical techniques are reasonably effective in relieving symptoms, none of them achieve a complete and restorable solution1. Tissue engineering provides an alternative strategy for the repair of cartilage and osteochondral. Generally, articular cartilage has similar structure to hydrogel, hence, using hydrogel-based materials to repair

injured articular cartilage tissues has attracted increasing attention

2-5

.

Biopolymer hydrogels have low immuno response and good biocompatibility and biodegradability, and are thus promising biomaterials for osteochondral tissue engineering (OTE) and cartilage tissue engineering (CTE) applications6-9. And the migration of cells in many bio-hydrogels is more favorably than that in synthetic polymer hydrogels10-11. Long-term in vitro culture of chondrocytes showed that bio-hydrogels outperformed synthetic polymer hydrogels in


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the expression levels of collagen type II and aggrecan as well as in the deposition of glycosaminoglycan12. Although hydrogels based on biopolymers show the above advantages, the low mechanical properties often limit their applications in OTE13-14. Our recent work would be devoted to solving the poor loadbearing issue using a double-network (DN) hydrogel system which consists of two interpenetrating polymer networks. Such double-network can form an energy-dissipating structure15. Furthermore, DN hydrogels fabricated from many different polymer pairs can show much better mechanical properties than their individual components16. Poly (γ-glutamic acid)(PGA) is an unusual anionic polyelectrolyte that can be crosslinked by lysine in EDC/NHS method17. It was reported18 that PGA hydrogels crosslinked by lysine could form a porous structure with a pore size of about 200-300um, which is conducive to the growth of cells into the porous structure of such hydrogels. The PGA hydrogels have been successfully used for the wound healing applications 18-19. It is well known that glutamic acid accounts for the greatest quantitative (12.36%) of the amino acids in articular cartilage, thus, the use of glutamic acid will facilitate the mimicking of cartilage ECM20. However, PGA hydrogels can swell to a great extent in water owing to their high osmotic pressure21 and thus cannot be used as a stable scaffold in bone or cartilage tissue engineering, which severely restricts the applications of PGA hydrogels Alginate is a natural polysaccharide polymer, which can be physically crosslinked through divalent cations22 such as Ca2+. The resulting physical hydrogels have been widely used in cartilage tissue engineering research because of their good biocompatibility, low toxicity, and relatively low cost12,

23-25

.

However, low stability for the physically crosslinked alginate

hydrogels can lead to a rapid drop in the mechanical strengths of alginate hydrogels in the


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physiological environment, thus requiring other crosslinking mechanisms to further stiffen the network structures26. In this study, we would use covalently crosslinked PGA and ionically crosslinked alginate to form a DN structure while bacterial cellulose (BC) would be incorporated into the DN structure to prepare a high strength bio-hydrogel. Bacterial cellulose is a neutral polymer, consisting of hydrophilic ultrafine fibers27. Abundant hydroxyl groups in the fibers can form a large number of hydrogen bonds with the double network structure, as a result, the achieved strength would be reinforced by the hydrogen bonds which meanwhile would suppress swelling and increase Young's modulus of hydrogel，to ultimately allow these hydrogels to get closer to articular cartilage in chemical components 28-29, structure，water content30 and compression modulus31. Furthermore, to achieve an optimal repair effect for osteochondral defects, we would design a bilayer structure with two different hydrogel layers (the nonporous top layer and the porous bottom layer) to match the structures of osteochondral tissues. Additionally, hydroxyapatite (HA) particles with two different sizes would be pro-introduced into the bilayer hydrogels, respectively: micro-HA in the top layer for promoting cartilage matrix deposition32-35 and HA nanocrystals in the bottom layer for enhancing compression modulus and osteogenesis36-40. The repair effect of the bilayer structure would be evaluated by using the osteochondral defect model of rabbits. 2. Experimental section 2.1 Materials γ -Poly(glutamic acid) ( γ -PGA, MW= 1000 kDa) was obtained from Xi'an Bella Biotechnology Co., Ltd (Xi'an, China). Lysine, 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide


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hydrochloride (EDC ﹒ HCl, ≥ 98.0%), and N-Hydroxysuccinimide (NHS, ≥ 98.0%) were purchased from Chengdu Dianchun Biotechnology Co., Ltd, (Chengdu, China). Sodium alginate (SA, MW=612kDa) ，BC suspension and CaCl2 were obtained from Shanghai Maichao Co., Ltd(Shanghai, China), Guilin Qihong Technology Co., Ltd(Guilin, China) and Kelong Chemical Reagent Co., Ltd (Chengdu, China), respectively. Other reagents used in this work were of analytical grade. 2.2 Preparation of hydrogels

Figure 1. Schematic of hydrogels formation and structures. (A) Schematic illustration of the preparation of BC-DN hydrogels. (B) Schematic depiction of the preparation of bilayered hydrogel scaffolds. (C) Schematic illustration of the structure of the bilayer hydrogel. (D) . SEM images of bilayer hydrogel scaffolds.


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PGA hydrogel was synthesized using an EDC/NHS mediated polymerization method similar to the method previously described18. PGA and lysine were dissolved in RO (Reverse osmosis) water to form a solution, and then EDC and NHS were mixed with the solution under vigorous stirring. After 5 minutes, a hydrogel was formed at room temperature. The BC-DN hydrogel was synthesized using a two-step method (Figure 1A). In the first step, PGA, lysine and SA were dissolved in BC suspension, and the mixture was stirred for 12h until the suspension solution became homogeneous and then EDC and NHS were mixed with the solution under vigorous stirring, a hydrogel was formed at room temperature after several minutes. In the second step, the hydrogel was immersed in the 1M CaCl2 solution for 48h until the swelling equilibrium was reached, obtaining a BC-DN hydrogel. Similarly, the DN hydrogel was synthesized using the similar method to the BC-DN hydrogel but without adding BC suspension. The used amount of precursors, EDC and NHS for the synthesis of different hydrogels is shown in Table S1.

2.3 Preparation of micro-HA and nano-HA The synthesis of micro-HA and nano-HA was based on the method reported previously41.The representative pictures are shown in the supporting information (Figure S1). 11.8 g of calcium nitrate tetrahydrate and 3.95 g of ammonium phosphate were dissolved in 100 ml and 60 ml distilled water, respectively, both solutions were then mixed under stirring and the pH value of the solutions was adjusted to 11 by adding ammonia. To obtain micro-HA and nano-HA, further hydrothermal reaction would be needed as described as follows.


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To achieve micro-HA, 9.3g of ethylenediaminetetraacetic acid disodium salt was added into the mixed solution, and then the hydrothermal reaction was conducted at 180℃for 10 hours. Subsequently, micro-HA was obtained after cleanout, desiccation and calcination. To synthesize nano-HA, 7.3g of ethylenediaminetetraacetic acid was added into the mixed solution, and then the hydrothermal reaction was carried out at 180℃for 6 hours. Subsequently, nano-HA was obtained after cleanout, desiccation and calcination.

2.4 Bilayer hydrogel scaffold composite fabrication The bilayer hydrogel composites were fabricated via a three-step crosslinking procedure (Figure 1B). The desired suspension (Table S2) for the chondrogenic (top) layer was poured into the bottom of a Teflon mold (100 mm diameter, 4 mm thickness). And then another composition for osteogenic was prepared, just by pouring the desired suspension into the partially filled Teflon molds to form the osteogenic (bottom) layer. The resulting bilayer hydrogels were immersed in the 1M CaCl2 solution to form the ion crosslinked network, and then immersed in RO water to remove the superfluous Ca2+. Notably, in order to get the compact chondrogenic layer and the porous osteogenic layer, the chemical crosslinking of the chondrogenic layer was carried out under an ice bath while the chemical crosslinking of the osteogenic layer was conducted using more catalyst (EDC/NHS) at room temperature.

2.5 Characterization of the materials The chemical crosslinking of the samples were analyzed by a Fourier transform-infrared (FTIR) spectrometer (Spectrum One, Perkin Elmer, Norwalk, USA). The morphology of the


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hydrogels was observed using a scanning electron microscopy (SEM, Quanta 200, Philips, Netherlands). The tensile and compressive stress–strain measurement of the hydrogels was performed using an Instron-5967 universal testing system. Rheological measurements were carried out using a DHR-3 (TA Instruments, USA) rheometer. The swelling ratio of the gels was calculated in the RO water and PBS. The biocompatibility of hydrogels was evaluated via human umbilical vein endothelial cell. A detailed description about the characterization of the hydrogels is given in the supporting information. 2.6 Animal model A total of 30 female, skeletally mature (6 months old) New Zealand White rabbits were used in this study and randomly divided into two groups (n=15 for each group), in which single-layer pure hydrogel and bilayer composites hydrogel were used, respectively. All surgical procedures and subsequent animal care were approved by the Institutional Animal Care and Use Committee (IACUC) of Southwest Jiaotong University. A detailed description about animal experimental methods is indicated in the supporting information

2.7 Histological analysis Histological staining was performed according to the standard operating procedures from the manufacturers. The previously reported grading scale described by Kurtis et al42-43 was used to quantitatively evaluate the histological scores of cartilage repair in different groups. 2.8 Micro-CT analysis A viva micro-computed tomography (80 SCANCO MEDICAL) was used to evaluate the qualitative and quantitative bone regeneration level within the defects of the harvested samples at


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12 weeks post-operation. In order to actualize the best visualization of the mineralized tissues, the samples were scanned through a 360° rotation angle with a 30µm resolution and the used voltage of the X-ray source was 80 kV. The bone regeneration was measured as bone volume per tissue volume (BV/TV), bone area per tissue volume (BS/TV) and trabecular thickness (Tb. Th), respectively. 2.9 Statistical analysis All data were presented as mean ± standard deviation. Student’s t-test was used to determine the statistical significances of all data. If the p-value is lower than 0.05, the difference will be considered to be statistically significant. The displayed error bars represent the standard deviation (SD). 3 Result and discussion 3.1 FT-IR analysis of the reaction of lysine with PGA and hydrogen bonds of the hydrogels The FT-IR spectrum of PGA as revealed in Figure 2A showed the peaks at 1647 cm-1 , corresponding to the C=O stretching vibration (amide Ⅰ band), ascribed to saturated aliphatic carboxylic acid dimers, at 1565 cm-1 assigned to the C=O bond vibration (amide Ⅱ band), at 1260 cm-1 assigned to the C-N bond vibration (amide Ⅲ band) , and at 2170 cm-1 assigned to the NH3+ stretching vibrations. The FT-IR spectrum of lysine showed the peaks at 1580 cm-1 for amide Ⅰ due to the C=O stretching, at 1517 cm-1 for amine Ⅱ owing to the C–N stretching, at 2137 cm-1 assigned to the NH3+ stretching vibrations, and at 1350 cm-1 and 1425 cm-1 corresponding to the CH2 bond vibration, respectively. The FT-IR spectrum of the PGA-Ly hydrogel displayed the peaks at 1640 cm-1 assigned to amide Ⅰ band，at 1590 cm-1 for amide


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Ⅱ band and at 1255 cm-1 for amide Ⅲ band, respectively. It can be noted that the signiﬁcant enhancement of the characteristic peak at 2137 cm-1 assigned to the NH3+ stretching vibrations indicated that lysine had been grafted to PGA44. Also, the difference in the amide Ⅰ peak was associated with the weakened pick of the carboxyl group in the PGA-Ly hydrogel, which revealed the formation of new bonds17, 45. These results demonstrated that the PGA-Ly gels had been successfully obtained. The FT-IR spectra of the PGA-Ly, BC-PGA, and BC-DN gels are shown in Figure 2B. The OH stretching vibration frequency has been used to detect and determine the strength of hydrogen bonds for many years. The stronger hydrogen bond will lead to an increase in both the width and intensity of the absorption band from 3570 cm-1 to 3050 cm-1. Hence, this peak area can be used to semi-quantitatively analyze the strength of hydrogen bonds46-47. It can be seen that the peak area of the hydrogen bond of PGA-Ly gels was much smaller than that of BC-PGA gels and BCDN gels, indicating that the presence of BC caused the much stronger hydrogen bonding interactions between chains. However, it can be seen that the peak area of BC-DN gels decreased in the presence of SA-Ca2+ network，which is most likely due to the interaction between Ca2+ and COO- , which thus replaced partial hydrogen bonds.


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Figure 2. FT-IR analysis and dynamic rheology behaviors of the hydrogels. (A) FI-IR spectra of pure γ-PGA, lysine and PGA-Ly hydrogels. (B) FI-IR spectra of PGA gel, BC-PGA gel and BCDN gel, and semi-quantitative analysis for the strength of hydrogen bonds: the area can be used to characterize the strength of hydrogen bonds. (C, D, E, F) Frequency (ω) dependence of storage modulus G', loss modulus G" and loss factor tan δ of the PGA gel (C), BC-PGA gel (D), DN gel (E) and BC-DN gel (F).


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3.2 The rheological Studies of the hydrogels Dynamic mechanical study can also be used as a tool to clarify the viscoelasticity of the hydrogel system. At room temperature, the PGA SN gel exhibited viscoelasticity. As shown in Figure 2C, G' was much higher than G" at high frequencies, however, both tended to be the same at low frequencies, thereby showing strong frequency dependence. The loss tangent tan δ = G'/ G" is generally used as a standard measure of the extent of viscous dissipation in the material. The tan δ of SN gel was high (0.25-0.42) and showed obvious frequency dependence, as expected for the viscoelastic materials. Figures 2 E and F show that tan δ values of the PGA/SA DN and BC-DN gels were nearly constant within the whole frequency range, but lower than that of SN gel, indicating a viscoelastic decreasing behavior. On the other hand, it can be noted that the tan δ values of the PGA/SA DN gel and the BC-DN gel were at least an order of magnitude higher than that (~0.002-0.02) of the PAAm (polyacrylamide) hydrogels created by the conventional chemical crosslinking48. This suggests that both PGA/SA DN and the BC-DN gels have some viscous character, which is important for the explanation of the high extensibility and toughness as will be seen later. Correlation between G' and frequency is attributed to the fundamental polymer relaxation that reflects their conformational changes or their distinct physical state49. Figure 2F displays a nearly frequency independent response of G' in the range of 0.1 to 100 rad/s, which corresponds to the rubber plateau, meaning that we can use the theory of rubber elasticity to analyses the microstructure of the BC-DN gel. The G' can also be used for estimating the extent of hydrogel network formation. G' values of BC-PGA and BC-DN gels were higher than those of PGA SN and PGA/SA DN gels. It implies that the networks became much denser in the presence of BC,


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thus confirming the significant effect of incorporating BC into the hydrogel networks on the properties of the hydrogels. Dynamic mechanical analysis is also one of the most effective and sensitive methods to study the movement of polymer chains. Herein, we used frequency dependence of tan δ to study the effect of BC on the molecular chain motion of the hydrogels50. As shown in Figure 2E, multiple peaks were present for the tan δ of the PGA/SA DN gel, thus indicating that there existed multiple motor units on chains. For BC-PGA and BC-DN gels, much fewer peaks were observed compared with those of PGA SN and PGA/SA DN gels, further indicating that the motion of PGA/SA molecular chains could be limited by BC. These results were in good agreement with the FTIR results which showed stronger hydrogen bonds associated with BC. 3.3 The mechanical properties of the hydrogels


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Figure 3. The mechanical properties of the hydrogels. (A) The stress-strain curves of the SN and DN hydrogels under tensile test. (B) The elongation and strength at break of the SN and DN hydrogels. (C) The stress-strain curves of the SN and DN hydrogels under compression. (D) The


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compression modulus of the SN and DN hydrogels. (E) The stress-strain curves of the BC-DN hydrogels under compression. (F) The compression modulus of the BC-DN hydrogels. To quantify the tensile properties of SN and DN hydrogels in response to the ratio of SA, a series of tensile tests were performed. As shown in Figures 3B, the SN gel (PGA-Ly gel), which only contained chemical crosslinks, was brittle and weak, showing a fracture stress of 0.103 MPa and a fracture strain of 5.66 mm.mm−1. On the other hand, the DN gels with strong ionic crosslinks showed a much higher toughness. The fracture stress (0.83 MPa) of the DN2 gel was eight times higher than that of the SN gel while the elongation (10.46 mm.mm-1) of the DN2 gel was about two times higher than the SN gel. The enhanced toughness of the PGA/SA DN gels can be explained by the energy-dissipating mechanism14. There are a large number of G blocks on each alginate chain. Two adjacent G blocks can form ionic crosslink with Ca2+ between them. During the stretching process, the alginate network can unzip progressively as the stretch increases and thus the energy will be dissipated, as a result, the occurrence of the stress concentration in the PGA networks will be avoided and the PGA network can still remain intact. Moreover, it can be seen that the toughness of the DN gel was even more remarkable when compared with their parents, the stress and stretch of which at rupture were only 3.7 kPa and 1.2 mm.mm-1 for the alginate SN gel14 and 0.103 MPa and 5.66 mm.mm−1 for the PGA SN gel, respectively. Therefore, such tensile properties for the DN gels should be attributed to a synergistic effect of the binary structure rather than a linear combination of two component networks. As observed, the elastic modulus of DN hydrogels increased gradually with raising the SA content while the elongation of the DN gels exhibited a trend of first increase and then decrease. When the SA content in the DN gels was at a high level, a large number of ionic crosslinks between Ca2+ and G Blocks could be formed so that the DN gels became stiff, leading


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to the fracture of the DN hydrogels at a low strain and the obvious decrease of the tensile property. Compressive properties are important features of hydrogels for performance evaluation, in which the uniaxial compressive strength of the DN and BC-DN gels was comprehensively studied. The compression modulus was measured based on the slope of the linear portion of the resultant stress-strain curve (strain 0-22%). The compressive property of the SN and DN gels as a function of different SA contents was compared in Figures 3C-D. The compression modulus of the DN3 hydrogel reached 0.168 MPa, a 5-fold increase over that of the pure PGA SN hydrogel (0.034 MPa). The increase of the compression modulus might be due to the constraining effect of the tightly cross-linked SA-Ca2+ network on the softly cross-linked PGA network. Intertwist of the two types of cross-linked polymer chains causes an interpenetrated network (IPN), resulting in a denser network structure for the DN hydrogels. Basically, the compression modulus increased with raising the SA content, indicating the compression modulus of the DN hydrogels can be adjusted just by increasing or decreasing the content of SA. The compressive property of the BC-DN gel samples as a function of different BC concentrations was compared in Figure 3E-F. The compression modulus of the hydrogels increased with increasing the content of BC, and the compression modulus of the BC-DN3 hydrogel reached 0.322 MPa, which is comparable with that of natural articular cartilage tissue (0.31MPa)31. This result can be explained by the illustration of BC-DN hydrogel (Figure 1A). Hydrophilic BC is a neutral polymer with abundant hydroxyl groups and hence BC can form a large number of hydrogen bonds with PGA and SA, which has been confirmed by the FTIR results. The chains of both PGA and SA networks were tied together by BC to form chain


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bundles, rigidifying the molecular chain and thus leading to the improvement of the mechanical strength of the BC-DN hydrogels. 3.4 Morphology of bilayer hydrogel The bilayer structure was confirmed by the SEM images as shown in Figure 1D. The bottom layer exhibited a porous structure with a pore size range of 100 to 500um while the top layer showed a compact structure (less porous). The bilayer hydrogel exhibited a perfect integrated structure without any boundary and fissure. 3.4 Swelling behaviors of the hydrogels Swelling was investigated for different hydrogels. The PGA SN hydrogel displayed extraordinary swelling properties in water or in PBS. As shown in Figure 4C- D, the swelling ratio of PGA SN hydrogel could reach 29.5 in PBS and over 150 in RO water. In comparison, DN gels exhibited a much lower swelling ratio, whose value was in the range of 70 to 15.5 in RO water and 16.5 to 9.2 in PBS, respectively, and decreased with increasing the SA content. Figures 4 E and F show that swelling ratio was further reduced when BC was used. The swelling ratio decreased from 15 to 7.5 in RO water, from 11.6 to 7.3 in PBS, respectively, with increasing the BC content from 0.25% to 0.75%. A low swelling ratio will be beneficial for the scaffold to be fixed on the defect area. The different swelling behaviors of the hydrogels can be explained by Figure 4A. The equilibrium swelling degree of a gel is determined by the elastic tension and the osmotic between the polymer networks and the external solution51-52. The osmotic promotes swelling while the elastic tension suppresses swelling. It is clear that when osmotic remains the same, the equilibrium swelling degree of a gel is determined mainly by the elastic tension. PGA-Ly


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networks, carrying a lot of hydrophilic groups and ionic functional groups, can swell in water and can hold a large amount of water while maintaining the physical dimension structure. It is noteworthy that the SA-Ca hydrogel does not exhibit regular volumetric swelling owing to the ion-responsive shrinking behavior53, which thus suppresses the swelling of the DN gels. In BCDN hydrogel networks, BC can be connected with PGA polymer chains through hydrogen bonding, which limits the chain motor or stretching of PGA networks, further suppressing the swelling of the BC-DN gels.


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Figure 4. Swelling behaviors of the hydrogels. (A) Illustration of the swelling–deswelling network structure SN, DN and BC-DN hydrogels. (B) Images of the hydrogels before and after swelling equilibrium. (C-D) The swelling kinetics of SN and DN gels in RO Water and PBS buffer solution, respectively. (E-F) The swelling kinetics of BC-DN hydrogels in RO water and PBS buffer solution, respectively.


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3.6 Cell compatibility of the hydrogels

Figure 5. Cell compatibility of the hydrogels. (A)Cell proliferation of the HUVEC cells on hydrogels evaluated using an MTT assay. (B)The fluorescence images of the HUVEC cells after 7 days of culture under AO/PI staining. The MTT assay was used to measure the cell proliferation in our study. As shown in Figure 5A, the number of cells increased significantly in all the groups from day 1 to day 7 and a similar


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increasing tendency was observed in all the groups over time. The OD value in each experimental group was not statistically different from the control group at days 1, 3, 5 and 7, indicating that the as-synthesized hydrogels were biocompatible. Moreover, the percentage of viable cells was characterized through LIVE/ DEAD fluorescence staining by using acridine orange (AO) and propidium iodide (PI). AO and PI can stimulate green and red fluorescence under a fluorescence microscope, respectively, and are used to label viable and dead cells, respectively. From Figure 5B, it could be seen that there were few dead cells in all the groups, further demonstrating that the as-synthesized hydrogels were biocompatible. 3.7 Gross observation for the repaired knees Postoperatively, the wound healed well and no tissue fluid exuded in the defect area, and additionally, no inflammation and swelling was found in the joint capsule. The general evaluation of knee repair at different time points for each group was presented in Figure 6. At 4 weeks, the defect in the single layer group was filled with granulation tissue while the defect in the bilayer group has been filled with fibrous or immature repair tissue. However, the surface of the newly-formed tissue at the defect site in both groups was rough. At 8 weeks, the boundaries between the implants and the surrounding tissue could still be seen for the single layer group while they almost disappeared for the bilayer group. At 12 weeks, the boundaries between the implants and the surrounding tissue for both groups have disappeared and the surface of the regenerated tissue at the defect site has become smooth. Obviously, the color of the newlyformed chondral tissue in the bilayer group was closer to the adjacent host chondral tissue. These results demonstrated that the defects treated with both scaffolds have been well repaired and the new cartilage had smooth surface and well integrated with the adjacent host cartilage. Comparatively, the bilayer group showed a better repair ability.


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Figure 6. Macroscopic images of knees repaired at different time points. (Single layer group and Bilayer group) 3.8 Micro-CT observations and quantification As shown in Figure 7A, the repair degrees of subchondral bone defects for both groups were different after operation for 12 weeks. From the results of 2D feature of the articular joint, it can be seen that a large area without regenerated subchondral bone still existed at the defects site in the single layer group while the subchondral bone reconstruction has been almost completed at the defects in the bilayer group. From the results of 3D reconstruction of the articular joint, it can be seen that a relatively large hollow still remained at the defects in the single layer group while there was only a tiny hole in the central area of the defect in the bilayer group.


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Furthermore, the quantitative analysis of bone histomorphometry parameters for the new subchondral bone from the interested region was performed. As shown in Figure 7B, it can be clearly seen that the average trabecular bone thickness, the ratio of the relative bone volume to the whole tissue volume and the ratio of bone surface to bone volume in the bilayer group were 0.17±0.01mm, 26±0.9% and 13±0.4 mm-1, respectively, which were much higher than these in the single layer group (0.11±0.01mm, 18±1.1% and 9±0.3 mm-1, respectively), showing the significantly positive effect of the bilayer integrated structure on promoting subchondral bone repair.

Figure 7. Micro-CT analysis of the articular joint. (A) 3D reconstruction and 2D feature of the articular joint after operation for 12 weeks (the osteochondral defect are marked using red circles and squares). (B) Quantification of the articular joint after operation for 12 weeks (*P

Biomimetic Bacterial Cellulose-Enhanced ... - ACS Publications

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