*2.6. Statistics*

All calculations are made using SigmaStat statistical software (Jandel Science, San Rafael, CA, USA) [10]. Statistical significance in the Student t-test corresponded to a confidence level of 95%. Data presented are mean ± SD from at least triplicate measurements. Di fferences were considered statistically significant at *p* < 0.05.

#### **3. Results and Discussion**

#### *3.1. Fabricating Fluidity HA Hydrogels by Adjusting Parameters of HA, and Reactions Conditions of BDDE Crosslinking Reactions*

HA hydrogels have high water contents because HA contents a large amounts of carboxyl groups and have been used for tissue repairs and in drug delivery systems [11]. However, without chemical modifications such as crosslinking of HA hydrogels, the gels would be easily disassembled in aqueous environment and then lost their mechanical properties which would usually deviate from those of human tissues including NP. Hydrogels for tissue repairs, including IVD repair, crosslinking HA to produce the HA network hydrogel has been widely prepared by oxidized HA or methacrylated HA, in order to improve and sustain the mechanical properties of HA [12–14]. However, it involves complex and delicate chemical reactions. Alternatively, HA network hydrogel was widely fabricated by crosslinking hydroxyl groups of *<sup>N</sup>*-acetyl-d-glucosamine (NAG) in HA using the epoxide groups of BDDE at a high pH condition (pH >11). In this study, HA network hydrogels were produced as the aforementioned method with modifying reaction parameters (Scheme 1). For the reactions in which HA was crosslinked by BDDE, several parameters such as the MW and concentrations of HA, the concentrations of BDDE, and the reaction time, would influence various properties of HA hydrogels [2,30]. For example, the uses of various concentrations of BDDE (0.01~20%) under alkaline conditions to crosslink various concentrations of HA with high molecular weight (2.65~10%, MW >10<sup>3</sup> kDa) to produce HA crosslinked hydrogels have been extensively investigated [30]. Here, a low concentration (~2.5%) and a low MW (~200 kDa) of HA was crosslinked using around 2.0% BDDE for 5–7 h to produce HA crosslinked hydrogels (Scheme 1), which had low G and G" (e.g., 0.24 ± 0.092 kPa and 0.09 ± 0.005 kPa, *n* = 3, respectively) with phase angle, δ, 21.4◦, indicating that viscoelastic HA crosslinked hydrogels were high fluidity. Hence, HA crosslinked hydrogels could be mixed well with varying amounts of SF for producing HS-IPN hydrogels.

#### Using HRP/H2O2 Reactions to Crosslink SF and Producing HS-IPN Hydrogels

Although the de-sericin process for SF polymers a ffects its molecular weight [31], the de-sericin procedures herein were well controlled which ensured the molecular weight of SF was approximately 185 kDa, as determined by SDS-PAGE (data not shown). Since high fluidity of HA crosslinked hydrogels could be well mixed with SF, using HRP/H2O2 enzymatic reactions to crosslink Tyr in SF within the hydrogels could homogenously take place to produce SF-IPN hydrogels and, consequently, produce HS-IPN hydrogels (Scheme 1). To further stabilize the HS-IPN hydrogels, 0.15 mM of PEI added into the hydrogels was further crosslinked by EDC/NHS for about 30 min. to produce amide bonds of carboxyl groups of HA with amine groups in PEI and in SF in the hydrogels, respectively. (Scheme 1).

According to the results of rheological study (in Section 3.3), the viscoelastic properties of HS7-IPN hydrogels were similar to matrix for human NP. The HS7-IPN hydrogels were chosen as a model hydrogel for inducing hBMSC to di fferentiate to NP cells or NP tissue engineering in this study.

#### *3.2. Characterizations of HS-IPN Hydrogels*

#### 3.2.1. ATR-FTIR Spectra of HA, SF, and the Model Hydrogels Consisted of Varying Strains of SF

The model hydrogels were characterized by spectra using an ATR-FTIR spectrophotometer (Figure 1). In the Figure 1, the transmission spectra for the peaks of carboxyl groups of H gels such as 1600 and 1402 cm<sup>−</sup><sup>1</sup> were about the same as those of HA polymers. Moreover, the peak of amide II of H gels was 1556 cm<sup>−</sup>1, which was similar to that of HA polymers. The transmission spectra for the peaks of amide I, II and III groups of S gels were 1640, 1511 and 1230 cm<sup>−</sup><sup>1</sup> that were about the same as SF polymers. The transmission spectra from HS-IPN gels produced from di fferent strains of SF (e.g., strain A, B and C) contained several characteristic peaks as those of H gels and S gels with minor variations. For instance, the peaks of amide II of HS-IPN gels shifted from 1556 cm<sup>−</sup><sup>1</sup> of H gels to 1528, 1522 and 1525 cm<sup>−</sup><sup>1</sup> for strain A, B and C, respectively (Figure 1). The presences of the peak of amide I of HS-IPN hydrogels only minor shifted from 1640 cm<sup>−</sup><sup>1</sup> to 1635 cm<sup>−</sup><sup>1</sup> of S gels, and there was no di fference among the model hydrogels consisted of varying strains of SF. Although the quantities of functional groups for varying weight ratios of HA/SF of the HS-IPN hydrogels were di fferent such as HS1-IPN and the model hydrogels, the ATR-FTIR spectra for those hydrogels were similar to those of the model hydrogels and not able to characterize their di fferences (Figure 1). Hence, other characterizations including rheological properties of the hydrogels needed to be carried out to determine the di fferences among the model hydrogels consisted of varying strains of SF.

**Figure 1.** ATR-FTIR transmission spectra of varying functional groups of SF, HA polymers, and H, S and HS7-IPN hydrogels. The characteristic peaks of H hydrogels such as carboxyl groups (COOH), and S hydrogels such as amide I, II and III, respectively, were found in HS-IPN hydrogels without varying by the SF of strains of A, B and C.

3.2.2. The Pore Structures of Scaffolds and Swelling Ratios for H, S and the Model Hydrogels

To examine the pore structures of scaffolds for H, S and the model hydrogels, SEM micrographs of the scaffolds were carried out and presented (Figure 2). The mean pore sizes in H and S scaffolds were approximately 49.36 ± 15.04 μm, and 13.40 ± 1.25 μm, respectively, (*n* = 3). The mean pore size of the scaffolds of S gels was significantly smaller (*p* < 0.01, *n* = 3) than that for the scaffold of H gels (Figure 2). Interestingly, the mean pore size of the scaffolds of the model hydrogels was around 38.96 ± 5.05 μm (*n* = 3) which was between those H and S scaffolds. (Figure 2). In addition, the pore size of the scaffolds of the model hydrogels was suitable for the proliferations of cells, including hBMSC [7].

The swelling ratios of H, S, H/PEI crosslinking and the model hydrogels in PBS were examined and shown (Figure 3). Since H gels were produced by highly hydrophilic polymers, the swelling ratios of the hydrogels fast increased within 12 h till 550% and then slowly increased up to ~700% at 48 h which was significantly higher than other hydrogels (*p* < 0.01, *n* = 3). Interestingly, after the first 6 h of fast swelling stage, the swelling ratios for the model and H/PEI crosslinked hydrogels were in plateau regions (e.g., 185.9 ± 24.4%, *n* = 3) until the end of study, 48 h. However, after swelling in the first 4 h, the swelling ratios for S hydrogels in PBS were about 24.9 ± 0.9% till 48 h, which were significantly lower than others (*p* < 0.01, *n* = 3) (Figure 3). The swelling ratios of the model hydrogels were very close to those of H/PEI crosslinked hydrogels, and the values were between those for H and S gels [32]. Notably, H/PEI crosslinked hydrogels highly reduced the swelling ratios in PBS compared with those for H gels, indicating that EDC/NHS reactions among HA and PEI in hydrogels would effectively crosslink carboxylic and amine groups to produce the amide bonds within the gels which resulted in significant decreases of the interactions among carboxylic groups of H gels and ambient H2O (Figure 3). Since EDC/NHS reactions were also carried out for preparing the model hydrogels as those for H/PEI crosslinked hydrogels, the swelling ratios for the model ones were similar to those of those for the later gels although the model hydrogels consisted of SF-IPN. In addition, the swelling ratios for other HS-IPN hydrogels (e.g., HS3-IPN hydrogels) would be similar to those for the model ones.

**Figure 2.** SEM micrographs of H, S, and HS-IPN hydrogels. The pore size of the model hydrogels was 38.96 ± 5.05 μm, (*n* = 3), in between those of H and S hydrogels.

**Figure 3.** The swelling ratios for H, S, H/PEI crosslinked and the model hydrogels in PBS were shown. The swelling ratios for the H gels were significantly higher than others (*p* < 0.01, *n* = 3), while those for S gels were significantly lower than others (*p* < 0.01, *n* = 3). The swelling ratios of the model hydrogels in PBS were similar to those for crosslinking H/PEI crosslinked hydrogels between those of H and S hydrogels (Data are mean ± SD, *n* = 3).

Notably, the results or data of HS-IPN gels at slow swelling ratio stages were similar to human NP, revealing that they may be suitable for use in TE of NP [7,28]. Moreover, the stability of the model hydrogels in PBS solution was examined by observation and measuring weight loss of the hydrogels after they were immersed into PBS solution for three weeks. The morphology of the model hydrogels was intact without disintegration by observation while approximately 4% of weight loss was found compared to that of the original hydrogels to be immersed in PBS. In addition, the samples became

more fluidity towards the end of the four weeks. Notably, the results of HS-IPN gels at slow swelling ratio stages were similar to human NP, revealing that they may be suitable for use in TE of NP [28].

#### *3.3. Varying Strains of SF Influenced the Rheological Properties of HS-IPN Hydrogels*

Viscoelastic flow parameters are important mechanical properties to characterize hydrogels for applications of bio-fluids such as lubricant in joints or NP. To determine, G, G" and δ values of HS-IPN hydrogels, they are determined using a parallel-plate rheometer operated in an oscillatory mode at a fixed strain (0.01 rad) with various angular frequencies (1–100 rad/s) as used by other studies [11,12,28]. G of the model hydrogels consisted of SF of strain A was 4.09 ± 0.32 kPa (*n* = 3, *p* < 0.01) which was significantly higher than those consisted of SF of strains B and C at a fixed angular frequency (Table 1). Since strain A of SF consists of more amount of Tyr than those in strain B and C, it could be assumed that more di-tyrosine bones were formed to crosslink peptides of strain A of SF to produce SF-IPN hydrogels than those in B and C ones. Hence, the rheological properties of the HS-IPN hydrogels were influenced by the strain of SF. The results rheological properties influenced by strains of SF, presented in Table 1, were qualitatively consistent with those UV-excited fluorescent intensities presented in Table 2. According to amino acid analysis for three strains of SF in this study, the total number of amino acids were about 5500, including about 5.1% of Tyr in strain A, while about 4.6% of tyrosine in strains B and C. The Tyr contents for strains of B and C were similar to other reports [27]. Notably, the formations of varying amounts of dityrosine bonds in SF-IPN hydrogels, crosslinked by HRP/H2O2 would emit varying intensities of blue fluorescence when the hydrogels were irradiated by UV [29]. The intensity of emitted blue fluorescence of SF-IPN hydrogels for strain A was significantly higher (*n* = 4, *p* < 0.01) than those of strains of B and C (Table 2), which was consistent with that Tyr contents in strain A (e.g., 277 ± 11, *n* = 4) is significantly higher than those of strain B and C (e.g., 255 ± 2, *n* = 4 for B), respectively. Therefore, the influences of strains of SF on the rheological properties of HS-IPN hydrogels possibly resulted from the tyrosine contents in each strain of SF.

**Table 1.** The viscoelastic parameters of the model hydrogels fabricated from di fferent strains of SF measured at 0.01 rad and 10 rad/s (*n* = 3). The parameters of strain A such as G and |G\*|, complex shear modulus, were significantly higher than strain B and C.


**Table 2.** The intensity of blue fluorescence of SF hydrogels excited by UV irradiation, in terms of OD values, for di fferent strains of SF at 1.62 mM H2O2 [29]. (*n* = 4).


\*\* *p* < 0.01. Influence of the Weight Ratios of SF to HA in HS-IPN Hydrogels on Rheological Properties of Hydrogels

Other than the SF strains, the weight ratios of SF to HA in producing the HS-IPN hydrogels might also influence the viscoelastic properties of the hydrogels. For examining this factor, at a fixed HA content (4%), varying SF concentrations (in wt.%) in producing HS1-IPN to HS7-IPN hydrogels (or the model gels) were carried out at the aforementioned oscillatory conditions (Figure 4A). G values increased with increasing the concentrations of SF in the HS-IPN hydrogels. Therefore, the model hydrogels had the highest G and G" values (*p* < 0.001, *n* = 3) among the produced hydrogels herein tested at varying angular frequencies. For example, the G values of the model hydrogels (e.g., 4.09 kPa at 10 rad/s) were about 2.6 times higher than those of HS1-IPN gels.

**Figure 4.** (**A**) The rheological parameters (e.g., G and G") for all tested HS-IPN hydrogels increased with increasing the weight ratios of SF/HA in producing the hydrogels versus the varying angular frequencies, (**B**) increasing weight ratios of SF/HA in the HS-IPN hydrogels increased the δ values from about 4.2 (e.g., HS1-IPN) hydrogels to 8.0◦ (e.g., HS7-IPN) hydrogels vs. with varying angular frequencies, respectively (Data are mean ± SD, *n* = 3).

The phase angles, δ values, for all HS-IPN hydrogels were shown (Figure 4B) which increased from around 4.2◦ to 8.0◦ at 10 rad/s, respectively, (*n* = 3). The results of δ values indicated that the model hydrogels were less viscoelastic solid than others (e.g., HS1-IPN), respectively. In comparison, the rheological properties for HA crosslinked hydrogels, produced by the same protocols as those for HS-IPN hydrogels, were carried out and had low G (e.g., 0.24 ± 0.092 kPa, *n* = 3) with phase angle of 21.4◦. The results indicated that viscoelastic properties for HA crosslinked hydrogels were high fluidity with a very low elastic modulus. Hence, the results of rheological properties for HA crosslinked hydrogels were not suitable for TE of NP. Notably, the G and δ values at 10 rad/s for the HS7-IPN hydrogels produced herein were similar to those reported for native NP (5.0~10.3 kPa and 2.5◦~35◦, respectively) [7,18,28], and therefore, the hydrogels were selected as the model hydrogels for further this investigation.

The G for IPN hydrogels produced herein were similar to those of laminin-111-PEG hydrogels reported for the matrix of NP [24]. Interestingly, the rheological modulus for G or |G\*|of HS1-IPN hydrogels were fitted the requirements of hydrogels for cardiac repairs. According to the tyrosine contents of varying strains of SF were different (Table 2), and the strains of SF influenced the viscoelastic modules of HS-IPN hydrogels (Table 1). The contents of tyrosine of SF were one of an important factor on determining those modules of HS-IPN hydrogels although the phase angles of the hydrogels might not be the case as those modules (Table 1). Moreover, the model hydrogels contained more concentrations of SF and amounts of di-tyrosine bonds in qualitative in the hydrogels than other HS-IPN hydrogels that resulted in increasing their viscoelastic modulus (Figure 4A and Table 1). However, the amounts of di-tyrosine bonds in each HS-IPN hydrogel were not able to be determined quantitatively. Although, the bonds could be semi-quantitatively evaluated using the intensity of UV-excited blue fluorescence [29].

Although using NHS/EDC to crosslink HA/SF to facilely prepare HA/SF hydrogels was recently reported [23], it was only one step of our processes to prepare hydrogels. However, they did not perform the rheological study for the aforementioned hydrogels [23]. Recently, using sonication and UV photo-polymerization to prepare SF/methacrylated HA or to produce SF-based IPN hydrogels has been reported by Xiao et al. [33], respectively. However, the rheological properties of the hydrogels were not determined. Interestingly, using HRP/H2O2 to crosslink tyramine-substitute HA and 2% of SF to produce hydrogels with varying HA contents has been reported by Raia et al. [24]. HA and SF-IPN hydrogels produced in their study are configured by di-tyramine bonds in HA crosslinked network, di-tyrosine bonds in SF crosslinked network and tyramine-tyrosine bonds in HA/SF-IPN, which bonding structures of their hydrogels were distinct from those in HS-IPN hydrogels produced in this study. Hence, the rheological properties of our hydrogels were different from theirs [24]. Moreover, the tyramine-substitute HA needed to be complexly and chemically synthesized for the research which was not a commercially available biomaterial, while the biomaterials were generally commercially available.

#### *3.4. Confined Compressive Modules of the Model Hydrogels*

The confined compressive stress of the model hydrogels was conducted on the hydrogels under 5% strain with value of 0.109 ± 0.011 MPa (*n* = 3) which was slightly lower than the for human NP (e.g., around 0.5–1.5 MPa) [7]. Notably, the compressive modulus for the model hydrogels was 2.29 ± 0.05 MPa (*n* = 3) was similar to that of human NP [7]. Although using NHS/EDC to crosslink HA/SF to facilely prepare HA/SF hydrogels was recently reported by Yang et al. [23] that was generally simple to produce the hydrogels than those produced by the procedures for this study, the confined compressive stress of their products was about 12 kPa at 30% of strain, which was about 25 times less than the stress of our model hydrogels (e.g., 314 ± 2.8 kPa, at 15% strain, *n* = 3). Hence, the compressive stress for their hydrogels would not fit the need of human NP.

Recently, Xiao et. al. [34] reported that using sonication and further UV photo-polymerization to prepare SF-IPN and methacrylated HA to produce network SF/HA hydrogels, which had a stiff but brittle SF structure while SF-IPN hydrogels produced herein (Scheme 1) were non-brittle. However, the confined compression modules of their hydrogels were not examined [23,31] as did in this study.

#### *3.5. Cytotoxicity Examinations for the Model Hydrogels*

In-vitro cytotoxicity for the model hydrogels were performed according to the requirements of ISO10993-5. Briefly, after L929 cells were cultured with vary concentrations of extraction mediums, for 24 h, and the MTT assay to the cells were carried out to examine the cell viabilities [26]. According to the MTT assay shown in Figure 5, the L929 viability of the group with a dilution ratio of 0.5 exceeded that of the group of extraction medium only which met the requirements of ISO10993-5, revealing that the model hydrogel was biocompatible and suitable for use in TE. The results of the biocompatibility of the hydrogels were similar to other SF-based biomaterials [8,31].

#### *3.6. Di*ff*erentiations of hBMSC to NP Cells in the Model Hydrogels*

To evaluate the di fferentiation of hBMSCs to NP cells, they were induced by TGF-β3 in the model hydrogels. The morphology of hBMSC, and the accumulations of NP-related ECM deposits in the hydrogels were examined using vimentin stain and immuno-histochemical (IHC) analysis, such as glycosaminoglycan (GAG) and collagen type II stain, respectively. The cell cultivated in cultural wells was also stained as a control group. Figure 6A exhibited that the deposition of glycosaminoglycan (GAG, in blue), one of the main ECMs in the NP, in hBMSC-laden model hydrogels was much more than that in the control [34]. Collagen type II, an important component of ECM in NP tissue, forms a fibrillar network that traps proteoglycan and resists swelling [11,12,20]. The deposition of GAG (in blue) and accumulation of collagen type II (in brown) in the model hydrogels containing di fferentiated hBMSC increased with increasing the cultural period (Figure 6A,B, respectively). The depositions of the GAG and collagen type II were broad distribution in the model hydrogels, revealing that the hydrogels herein were suitable to the di fferentiations of hBMSCs to NP cells. Enhancing expressions in GAG and collagen type II of the di fferentiations of hBMSCs to NP cells in the model hydrogels compared to those expressions in the control group revealed that they promoted the di fferentiations of hBMSC (Figure 6A,B, respectively). According to those results, the mechanical properties of the model hydrogels compared with hardness matrix (i.e., cultural wells) were suitable to the di fferentiations of hBMSCs to NP cells.

**Figure 5.** L929 Cell viability incubated by the extraction mediums (EM), taken from supernatants which was incubated with the model hydrogels for 24 h, at different dilute ratios according to the guidelines of ISO 10993-5. The relative L929 viability (%) of the group with a dilution ratio of 0.5 exceeded that of the group of EM only. The results of MTT assay for L929 Cell viability showed that the model hydrogels were biocompatible biomaterials (data are mean ±SD, *n* = 3).

(**A**) 

**Figure 6.** Immuno-histochemical (IHC) analysis of hBMSCs in the model hydrogels after hBMSC were induced by TGF-β3 to differentiate them to NP cells for 7 d and 14 d, respectively. Deposition of (**A**) GAG (in blue) and, (**B**) Collagen type II (in brown) in the model hydrogels were shown which were significantly higher than those of cultivated in cultural wells (e.g., control group).

(**B**)

Aggrecan (AGN), collagen type II (Col II) and collagen type I (Col II) were selected as test genes for examining the differentiations of hBMSCs in the model hydrogels [11,12,21]. The gene expressions of chondrogenesis of hBMSCs, cultured in the model hydrogels, revealed significant up-regulation of both AGN and Col II but significant down-regulation of Col I, compared to those of hBMSCs cultivated in cultural wells (e.g., the control group) for 7 d (Figure 7A–C). The results gene expressions of differentiations of hBMSCs to NP cells presented herein were similar to those of other studies [11–13] although the compositions of their hydrogels were different from this study. AGN and Col II gene expressions increased with the 7 d of differentiations of hBMSCs to NP cells in the model hydrogels that were consistent with the increasing depositions of GAG and collagen type II of IHC stains (Figure 6A,B, respectively). However, the gene expressions of the differentiations of hBMSCs to NP cells in the model hydrogels (Figure 7A–C, respectively) were not further increased in the period of 7 d to 14 d. The possible mechanisms such like reducing the activity of CD44 of HA for the differences of results the differentiations of hBMSCs to NP cells in Figures 6 and 7 at 14 d need to be further studied. Although adjust mechanical properties of HA-based or methacrylated HA-based containing SF hydrogels using varying techniques for other TE studies have been reported by others [11,23,24]. This study reported an examination of the differentiations of hBMSCs to NP cells in HS-IPN hydrogels. Nevertheless, the model hydrogels promoted the differentiations of hBMSCs to NP cells for tissue engineering of NP.

60

**Figure 7.** Real-time PCR analysis for the relative quantities of NP-specific gene expressions of inducing differentiation of hBMSC to NP cells on the model hydrogels for 7 d and 14 d. Gene expressions of AGN (**A**), COL II (**B**) and COL I (**C**) for the control group (e.g., ctrl) were the data for 7 d incubated at cultural wells. (\* *p* < 0.05, \*\* *p* < 0.01, \*\*\**p* < 0.001; data presented are mean ±SD, *n* = 3).
