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pullulan and pullulan hydrogels: their microstructure, engineering process and biocompatibility
Materials Science and Engineering C 58 (2016) 1046–1057
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Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
HLC/pullulan and pullulan hydrogels: their microstructure, engineering process and biocompatibility Xian Li a,b,1, Wenjiao Xue c,1, Yannan Liu b, Weina Li b, Daidi Fan b,⁎, Chenhui Zhu b, Yaoyu Wang a,⁎⁎ a b c
College of chemistry & materials science, Northwest University, Taibai North Road 229, Xi’an, Shaanxi 710069, China Shaanxi Key Laboratory of Degradable Biomedical Materials, Department of Chemical Engineering, Northwest University, Taibai North Road 229, Xi’an, Shaanxi 710069, China Shannxi provincial institute of microbiology, Xi’ an 710043, China
a r t i c l e
i n f o
Article history: Received 10 August 2015 Accepted 7 September 2015 Available online 11 September 2015 Keywords: HLC/pullulan hydrogels Pullulan hydrogels NaIO4 Subcutaneous Biodegradation
a b s t r a c t New locally injectable biomaterials that are suitable for use as soft tissue fillers are needed to address a significant unmet medical need. In this study, we used pullulan and human-like collagen (HLC) based hydrogels with various molecular weights (MWs) in combination therapy against tissue defects. Briefly, pullulan was crosslinked with NaIO4 to form a pullulan hydrogel and then may coupled with HLC using the reaction between the –NH2 end-group of HLC and the –CHO group present on the aldehyde pullulan to form the HLC/pullulan hydrogel, wherein the NaIO4 acted as the crosslinking and oxidizing agent. The good miscibility of pullulan and HLC in the hydrogels was confirmed via Fourier transform infrared spectroscopy, scanning electron microscopy, compression testing, enzyme degradation testing, cell adhesions, live/dead staining and subcutaneous filling assays. Here, pullulan hydrogels with various MWs were fabricated and physicochemically characterized. Limitations of the pullulan hydrogels included inflammation, poor mechanical strength, and degradation. By contrast, the properties of the HLC/pullulan hydrogels strongly enhanced. The efficacy of these hydrogels was evaluated both in vitro and in vivo. Our results indicate that HLC/pullulan hydrogels may have therapeutic value as efficient soft tissue fillers, with reduced inflammation, improved cell adhesion and delayed hydrogel degradation. © 2015 Elsevier B.V. All rights reserved.
1. Introduction Recently, microbial polysaccharides have received increasing interest because of their useful physiochemical properties and natural biodegradability. The repeating units of these exopolysaccharides are regular, branched or unbranched, and interconnected by glycosidic linkages. For example, due to its non-toxic, non-immunogenic, non-mutagenic and non-carcinogenic nature, pullulan, a non-ionic exopolysaccharide of fungal origin, has recently been evaluated for various biomedical applications, including tissue engineering, targeted drug/gene delivery, targeted drug therapy and wound healing [1–3]. Pullulan is one of the polymers obtained from the fermentation medium of black yeast (e.g., Auerobasidium pullulans). It comprises maltotriose units connected by α (1 → 4) glycosidic bond, whereas the consecutive maltotriose units are connected to each other by α (1 → 6) glycosidic bond [4–6]. It has a molecular weight (MW) range of 5,000–9,000,000 g/mol with a straight unbranched chain [7]. Pullulan has considerable mechanical strength and possesses numerous functional properties such as
⁎ Corresponding author. ⁎⁎ Correspondence to: Y. Wang, College of chemistry & materials science, Northwest University, Taibai North Road 229, Xi’an, Shaanxi 710069, China. E-mail addresses: [email protected] (D. Fan), [email protected] (Y. Wang). 1 These authors contributed equally to this work.
http://dx.doi.org/10.1016/j.msec.2015.09.039 0928-4931/© 2015 Elsevier B.V. All rights reserved.
adhesiveness, the ability to form a film and enzyme-mediated degradability [8]. However, a pullulan solution’s viscosity changes with its MW [9]. Therefore, it can be used as a low-viscosity filler in beverages and sauces, or it can be used as a high-viscosity filler in tissue engineering applications. Interestingly, the viscosity of a pullulan solution does not change with heat, changes in pH or in the presence of most metal ions, including sodium chloride [10]. Injectable hydrogels are promising scaffolds for cell encapsulation, tissue repair, tissue reconstruction and drug delivery. Novel polymeric materials such as polymer blends and composites have led to the advancement of technologies to develop injectable hydrogels [11]. Polymer blends are prepared by physically mixing two or more polymers, and they tend to have properties superior to any of the individual polymers [12]. Moreover, hydrogels, prepared as composites of polysaccharides and collagen with or without a crosslinking agent, have garnered great interest in tissue engineering applications. Numerous polysaccharides have been used for preparing hydrogels, including chitosan, pullulan, hyaluronic acid and carboxymethyl chitosan. Chitosan and hyaluronic acid have gained much popularity in the field of tissue engineering. Recently, pullulan, a new material and a non-ionic polysaccharide with unique characteristics (non-toxic, non-immunogenic, non-mutagenic and non-carcinogenic), has been explored for various biomedical applications [13]. Human-like collagen (HLC) has been used in hydrogels [14], vascular scaffolds [15] and artificial bones [16]
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due to its low rejection rate, good biocompatibility [17], biodegradability [18], workability and no risk of viral infection. Many papers have reported on using chitosan or chitosan derivatives and hyaluronic acid crosslinking with collagen or other crosslinking agents to form hydrogels [19]. Many articles have described the use of pullulan gels or nanoparticles for applications in gene-drug delivery and tissue repair. There are fewer articles reporting the formulation of pullulan-based copolymers. Gark et al. [20], Miyahara et al [21] and Rustad et al [22] reported that the pullulan and collagen hydrogels was used to force cell delivery in wounds, guide bone regeneration and as biomimetic hydrogel scaffold to improve the mesenchymal stem cell angiogenic capacity and stemness. Our report is mainly about different molecule weight of pullulan and collagen to fabricate injectable hydrogels by another method for skin reconstruction. Feng et al. reported the collagen and dialdehyde pullulan fabricate the complex scaffold materials [23]. Bruneel et al. reported that chemical modification of pullulan by periodate oxidation [24]. Based on the above studies, we choose the pullulan modification by periodate oxidation (NaIO4) and human-like collagen to fabricate hydrogels for skin reconstruction. Therefore, in this report, we investigated the use of pullulan and HLC with a crosslinking agent for the preparation of hydrogels. We used NaIO4 not only as an oxidant but also as a crosslinking agent in our hydrogel system. We prepared hydrogel systems with varying MWs (100,000–530,000 g/mol), with straight unbranched chains and with or without HLC. The HLC/pullulan hydrogels have biological effects that are different from pullulan hydrogels, and have excellent biological and mechanical properties.
HLC/pullulan hydrogels. Then, the 4% HLC solution was sterilized using a 0.22-μm sterilizing filter. The composite mixture was gently vortexes for 30 min at 4 °C to promote the homogeneous distribution of polymers within the hydrogel. The mixture was then mixing with 0.5 mL, 0.5 mol/L sterile NaIO4 and stirred for 20 min until dissolution was complete. The solutions were then placed in a water bath for gelling at 37 °C for 12 h. The pullulan and HLC/pullulan hydrogels were jellylike as shown in Figure 1 (A, B, C and D). The hydrogels were dialyzed in pyrogen-free water at room temperature for 12 h for hydrogel properties analysis and the in vitro biocompatibility analysis. Then, the hydrogels were extruded with a granulator to make them into small granules and were stored at 4 °C for in vivo biocompatibility analysis. All of the experiments were performed in a sterile environment. The hydrogels formed, purify and made into granules process is shown in Figure 1(F).
2. Materials and methods
The porosity of the pullulan and HLC/pullulan hydrogels was determined by the liquid displacement method. First, the volume and weight of the hydrogels were measured and recorded as V0 and W0, respectively. Second, the samples were immersed in dehydrated alcohol for 48 h until they were saturated, and then the hydrogels were weighed again (W1). Finally, the porosity was calculated according to the formula: P = (W1 − W0)/ρV0, where ρ represents the density of the dehydrated alcohol. The pore diameters of the hydrogel were measured using ImageJ software. Each pore of the sample was measured, and the average value of the measurements was calculated.
2.1. Materials HLC was expressed by E. coli by cloning a partial cDNA derived from human collagen mRNA [16]. HLC is a macromolecular watersoluble protein with a MW of 97,000 g/mol (China patent number: ZL01106757.8). The protein encoded by the partial cDNA is reference by Fan DD, et al [25]. The procedure to produce and purify HLC from the bacteria is reference by Ma et al [26] and Zhang HZ et al [27]. Pullulan (MW, 100,000 g/mol; 170,000 g/mol; 530,000 g/mol) was obtained from the Shaanxi Provincial Institute of Microbiology. NaIO4 was purchased from Shandong Frieda Biological Technology, Ltd. (China). The live/dead viability/cytotoxicity kits were purchased from Sigma-Aldrich. All of the other solvents and reagents were analytical grade. 2.2. Preparation of pullulan hydrogels and HLC/pullulan hydrogels Pullulan hydrogel and HLC/pullulan hydrogel compositions are shown in Table 1. One gram of pullulan powder was dissolved in pyrogen-free water with or without HLC to a total volume of 6 mL. HLC was mixed in at a MW of 100,000 g/mol, 170,000 g/mol or 530,000 g/mol pullulan (0.25 g HLC per 0.5 g pullulan) to prepare the
Table 1 The pullulan and HLC/pullulan hydrogels dompositions. Hydrogels
Sample
Pullulan (MW, g/mol)
HLC (%)
Pullulan (%)
NaIO4 (mol/L)
Pullulan hydrogels
Gel1 Gel2 Gel3 Gel4 Gel5 Gel6
100,000 170,000 530,000 100,000 170,000 530,000
0 0 0 4 4 4
8 8 8 8 8 8
0.5 0.5 0.5 0.5 0.5 0.5
HLC/pullulan hydrogels
2.3. FT-IR spectroscopic analysis The different components of the HLC/pullulan hydrogel systems were verified by FT-IR. Infrared spectra of the specimen powders, namely pullulan (MW of 100,000 g/mol, 170,000 g/mol and 530,000 g/mol) and NaIO4 as well as of the HLC and the six formulations of hydrogels were recorded using an FT-IR spectrophotometer (Thermo Fisher Scientific, Waltham, MA, USA). The samples were triturated with KBr at a ratio of 1:100 and pressed to form pelleted samples for FT-IR spectroscopic analysis at 500–4000 cm−1. 2.4. Porosity and pore diameter of the hydrogels
2.5. Microscopic investigations The pore morphology of the pullulan hydrogels and HLC/pullulan hydrogels was examined by scanning electron microscopy (SEM). The samples were carefully lyophilized to maintain their threedimensional porous structures. The lyophilized hydrogel samples were immersed in liquid nitrogen, and the vitrified samples were carefully cut with a cold knife. The cut samples were mounted, sputter-coated with gold, and observed under SEM using a Hitachi S-570 SEM microscope (Hitachi, Tokyo, Japan). 2.6. Swelling measurement Hydrogel water absorption capacity was calculated as a swelling ratio: Swelling ratio = (weight of wet sample − weight of dry sample)/weight of dry sample. For incubation studies, the hydrogels were incubated in deionized water or in PBS (pH = 7.4) at 4 °C or 37 °C. Excess liquid was gently shaken off, and the weights of the swollen gels were obtained. Hydrogel weights were measured at 24 h. Six samples were tested for each condition. 2.7. Compression experiment The hydrogels were tested using a gel strength instrument (Electronic Universal Testing Machine) as follows: first, the height of the compression plate was adjusted to 10 mm, and its final height was adjusted to 4 mm.
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Figure 1. Visual inspection of hydrogels. (A) Gel1, Gel2, Gel3, (B) Gel4; (C) Gel5; (D) Gel6; (E)hydrogels particles; (F) the hydroges forming process.
The instrument disk was circular with a diameter of 10 mm, and the compression speed was adjusted to 2 mm/min. The gel was cut to a length of 10 mm and placed on the instrument disc. When the compression displacement reached 6 mm, the test was stopped. The compression modulus and the relationship between the compression load and compression displacement were calculated. The crosslinking density (ρ) [28] was determined from the compression modulus (G) and swelling ratio (Q) as follows: ρ = GQ1/3/RT, where G is the compression modulus, R is the gas constant (8.314 × 103 KPa · cm3 · mol−1 · K−1) and T is 310 K. 2.8. Quantification of crosslinking index The cross-linking index was measured using 2,4,6trinitrobenzenesulfonic acid (TNBS) [29] for the determination of free amine groups in the soluble and poorly soluble proteinaceous materials. Hydrogel mixtures containing HLC only and HLC/pullulan hydrogels (4 mg) were reacted with TNBS (0.01 M) in sodium hydrogen carbonate solution (pH 8.2, 4% w/v, 2 mL) for 2 h at 40 °C. Then, hydrochloric acid (6 M, 3 mL) was added to hydrolyze the water (5 mL) in the solution, and the absorbance was measured against a TNBS solution without gelatin that had been treated exactly the same as the other samples. Crosslinking index = (1 − absorbance of HLC/pullulan hydrogel/absorbance of HLC) × 100%.
2.10. Cell adhesion, proliferation and viability analysis Cell viability was evaluated by cell adhesion and live/dead staining was observed using a confocal microscopy and SEM, respectively. The live/dead viability/cytotoxicity kit for mammalian cells are using for live/dead staining as described previously [31]. Cryopreserved fibroblasts were thawed, centrifuged, and then cultured in DMEM containing 10% FBS. Confluent cells were passaged into multiple flasks for subsequent experiments. One day before culture, four wet hydrogel samples (10-mm diameter and 3-mm height) were sterilized at 121 °C for 5 min, rinsed with DMEM twice in a 48-well plate, and then equilibrated in the medium overnight until cell plating. For surface adhesion studies, logarithmic-phase fibroblasts (5 × 105 cells/mL) were seeded onto four hydrogels, and 1 mL of DMEM was added to each well in a 5% CO2, 37 °C incubator. Parts of the samples were fixed in glutaraldehyde at 72 h, and the cell morphology was observed by SEM after vacuum freeze drying. The remaining portions of the samples were stained by live/dead staining at 24, 48 and 72 h to evaluate cell death rate on the hydrogels, and the live/dead cells staining at 72 h were observed using a confocal microscopy. The death rate of cell was calculated by dead cell/live cell × 100% [14]. A control condition is conducted in normal fibroblast cultured by DMEM media, the test samples were cultured by the Gel1, Gel2, Gel3, Gel4, Gel5 and Gel6. Five parallel samples of each material were utilized.
2.9. In vitro degradation
2.11. The biocompatibility of hydrogel in vivo
Dry pullulan hydrogels (Gel1, Gel2 and Gel3) and HLC/pullulan hydrogels (Gel4, Gel5 and Gel6) were incubated with pullulanase (20 U/mL) in PBS and weighed every 30 min. Similar experiments were performed with collagenase I (20 U/mL) in PBS. Combination degradation studies using both pullulanase and collagenase I was conducted in PBS using the same concentrations as above [30]. The initial weight at time 0 was the dry weight of the hydrogels (W0) and the wet hydrogels after being degraded by enzymes were washed three times with water before vacuum drying (W1) and used for subsequent measurement. The weight loss (%) was calculated as follows: (W0 − W1)/ W0 × 100%. Experiments were performed six times for each condition at room temperature.
All experiments were performed in compliance with the relevant laws and institutional guidelines, and conducted with the approval of the Institute Animal Ethics Committee. This study was supported by Shaanxi Key Laboratory of Degradable Biomedical Materials and Northwest University. Wild-type adult mice (n = 18) were anesthetized with isoflurane inhalation. After cleansing the skin with 70% alcohol, 0.2 mL of pullulan hydrogels and HLC/pullulan hydrogels granules, obtained from the extruding granulator under ultraviolet (UV) light, were injected subcutaneously. The hydrogels were examined daily and harvested at 1 week. Digital photographs were taken to quantify the residual hydrogel size. Then, the hydrogels injected subcutaneously into wild-mice were taken out after 1 week and
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Figure 2. (a) FT-IR spectra of the pullulan hydrogels (a, b, c) and HLC/pullulan hydrogels (d, e, f). Small letters represent the following: a, Gel1; b, Gel2; c, Gel3; d, Gel4; e, Gel5; f, Gel6. (b) FT-IR spectra of pullulan (100,000, 170,000 and 530,000 g/mol), NaIO4, and HLC.
immediately fixed in 10% neutral buffered formalin. The tissue surrounding the hydrogels was tested by H&E staining and immunohistochemical analysis [19]. The TEM experimental process was reference by Xian Li et al [18].
2.12. Calculations and data analysis Data were collected in a Microsoft Excel 2000 database, and the results were presented as means and standard deviations using the Origins 7.0 software. Student's t-test was performed to determine the statistical significance between experimental groups. A value of p b 0.05 was considered statistically significant.
3. Results and discussion 3.1. FT-IR spectroscopic analysis The components of the hydrogel systems were assessed using Fourier transform infrared spectroscopy (FT-IR). The spectra obtained for pullulan, HLC, pullulan hydrogels, and HLC/pullulan hydrogels are displayed in Figure 2(a) and (b). The main characteristic absorption bands of the pullulan hydrogels at 1624 and 1444 cm−1 were caused by the C = O stretching. These results confirmed that the OH groups of pullulan were oxidized to C = O by NaIO4 in the pullulan-based hydrogels. Pullulan contains three different anhydroglucoside moieties in the repeating unit. Therefore, periodate oxidation of pullulan results
Figure 3. Synthesis of the pullulan and HLC/pullulan hydrogels. Schematic representation of the proposed mechanisms of the interactions between pullulan, HLC, and NaIO4.
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in different types of dialdehyde structures. The adjacent two hydroxyl groups of pullulan were periodate into dialdehyde or carboxide. The same results were reported by Bruneel et al [24] and Spatareanu A [32]. The HLC/pullulan hydrogels were identified in the respective spectra from the C = O asymmetric stretching band at 1644 cm−1. Although the HLC amide bands were partially masked by the adsorbed band at 1640 cm−1, their presence was denoted by the new band at 1515– 1540 cm−1 due to the absorption of the protonated amino groups. The characteristic peaks representing the stretching vibration of VC-N and the coupled vibrations of δNH and VC-N were observed at 1413 and 1515 cm−1, respectively and were may be consistent with the formation of -CONH at -C = O sites on pullulan coordinated with -NH2 sites on HLC in HLC/pullulan hydrogels. However, the adsorbed band at 1367 cm−1 may be represent ester group (COOR), which may be contribute the carboxyl group of HLC reaction with pullulan (Figure 3). These bands were absent in the pullulan and HLC. 3.2. The mechanism of hydrogel formation The HLC/pullulan hydrogels were synthesized by the oxidation and condensation of aldehydes and amines to form a Schiff base as shown in Figure 3. In the reaction between NaIO4 and the various MWs of pullulan, two adjacent hydroxyl groups of pullulan were oxidized to an aldehyde group in one step to form the pullulan hydrogels. After
pullulan powder resolved in deionized H2O, the pullulan particle contact the surface of water immediately, leading to the formation of water film enwrap the pullulan powder; then stirring, as need enough time for pullulan powder to saturate into water. The characteristic hydration properties of pullulan might be due to the flexibility of the pullulan skeleton [33]. On the other hand random coil proteins are able of forming coherent film due to the formation of extensive hydrogen bonds [34]. Therefore, HLC solution made it easier to dissolve the pullulan powder than deionized H2O. Various MWs of pullulan were cross-linked with HLC by adding NaIO4 to form the HLC/pullulan hydrogels in two steps. Crosslinking with HLC via NaIO4, caused the pullulan with unbranched chains in the hydrogel systems to increase in MW, improved its anti-degradation properties, and changed the spatial structure within the hydrogel system. This unique spatial structure may have influenced the properties, physiochemical characteristics and biocompatibility of the HLC/pullulan hydrogels. Therefore, developing HLC/pullulan hydrogels may produce ideal soft filler materials for tissue engineering applications. 3.3. Physical characteristics of the hydrogels To investigate the water retention properties of the hydrogels, swelling studies of the gels were performed with water and phosphatebuffered saline (PBS) at different temperatures as shown in Figure
Figure 4. Physical characterization and toxicity of different gel materials (Gel1, Gel2, Gel3, Gel4, Gel5, Gel6). Swelling ratios of gels in ddH2O (A) and PBS (B); elastic modulus (C), crosslinking density (D), porosity (E) and pore diameters (F) of gels; compression load of pullulan hydrogels (G) and HLC/pullulan hydrogels (H).
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4(A) and (B). The swollen hydrogels retained their general shape and did not degrade after incubation in either deionized water or PBS. The swelling ratios for the pullulan hydrogels and the HLC/pullulan hydrogels incubated in deionized water and PBS at 4 °C and 37 °C were different, and we found that the swelling ratios for the pullulan hydrogel were higher than the ratios for the HLC/pullulan hydrogels. Hydrogels with the same pullulan to HLC ratio reached swelling equilibrium at 24 h independent of the crosslinking density or the osmolarity of the swelling medium. Hydrogels with different MWs of pullulan had different crosslinking densities. There was almost no hydrogel swelling during 12 ~ 48 h incubation in all conditions. The hydrogel reaching the swelling equilibrium was recorded at 24 h, indicating maximal water absorption after overnight incubation. In addition, when the swelling medium was changed from PBS to ddH2O, the gels swelled more than at equilibrium. The viscoelastic characteristics of the gels were assessed as shown in Figure 4(C). We found an increasing elastic modulus (3.8 KPa (Gel1) to 8.5 KPa (Gel3) and 10 KPa (Gel4) to 100 KPa (Gel6)) and decreasing swelling (22.6 ± 1 (Gel1) to 18.5 ± 0.9 (Gel3) and 17.8 ± 0.9 (Gel4) to 15.5 ± 0.7 (Gel6)) as the degree of cross-linking increased (Figure 4(D)). These findings are in agreement with the theoretically expected correlation between mechanical characteristics and cross-linking density [35]. The elastic modulus was directly proportional to the crosslinking density [36]. Despite the overall high water content of the materials and variability of the MW of pullulan from 100,000 to 530,000 g/mol, the meshwork stability was excellent. The higher MW of pullulan resulted in more stable molecular chains with more covalent cross-linking sites, thereby resulting in higher crosslinking densities in the pullulan hydrogels and HLC/pullulan hydrogels as shown in Figure 4(D). The relationship between compression load and the compression displacement of hydrogels is shown in Figure 4 (G, H). At the same compression displacement, hydrogels with different pullulan MWs showed different compression loads. As expected, the compression modulus
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increased as the compression load increased. The mechanical integrity of the hydrogels indicated that the compression modulus increased as the degree of cross-links increased, suggesting that the network became more resistant to change when a load was applied with more crosslinks. As expected, the HLC also affected the mechanical strength of the gels. The compression monotonically increased with the addition of HLC, presumably due to the increased cross-link density. The compression testing showed that the pullulan hydrogels were mechanically robust and showed varied loads from 1.3 KPa to 3 KPa depending on the NaIO4 crosslinking. Additionally, the addition of HLC led to a varied load from 2.5 KPa to 12 KPa. One striking observation was that pullulan and HLC appeared to affect the compression load similarly. Figure 4(E) shows effects of the MW of pullulan and the addition of HLC on particle size and the overall porosity of the prepared porous hydrogels. In all of the cases, the overall porosity increased linearly with decreasing MW of pullulan. The varying MW of pullulan had little impact on the increasing rate of overall porosity. However, the addition of HLC produced a more significant difference in overall porosity. For instance, the porosity of Gel1, Gel2, Gel3 was 80%, 79%, 71%, respectively, whereas the overall porosity of Gel4, Gel5, Gel6 was 69%, 65%, 60%, respectively. Clearly, the HLC created more opportunities for connections between pullulan powder particles, resulting in closer pore formations. The linear relationships implied that the overall porosity of hydrogels can be quantitatively adjusted by the addition of HLC and varying the MW of the pullulan powder particles. Figure 4(F) shows the pore size diameters of the various hydrogels. We found that the MW of pullulan influences the pore size and pore network of hydrogels. At the MW of 530,000 g/mol (Gel3, Gel6), the pore size diameters was higher than in the lower pullulan MW gels (Gel1, Gel4). From these studies, it can be concluded that the increase in porosity aids in higher MWs in gels. Based on the above analysis, our results indicate that pullulan and HLC had molecular interactions during the gelation process, and the chemical reaction was more obvious when the MW of pullulan was higher.
Figure 5. SEM micrographs × 250 of formed hydrogels A(Gel), B(Gel2), C(Gel3), D(Gel4), E(Gel5) and F(Gel6) at room temperature after vacuum freeze-drying for 48 h. Scale bar for all: 120 μm.
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3.4. Network structure
3.5. In vitro degradation by enzymes
Figure 5 shows representative scanning electron microscopy (SEM) images of the cross-section of the lyophilized pullulan hydrogels (Gel1, Gel2 and Gel3) and HLC/pullulan hydrogels (Gel4, Gel5 and Gel6). All the hydrogels were highly porous throughout the crosssection. Gel1 exhibited the largest pore size (average 52 μm) among all the hydrogel formulations. Figure 5 illustrates the differences in the cross-section structure of the hydrogels. The surface of pores of pullulan and HLC/pullulan hydrogels showed a rougher structure. However, the cross-section revealed relatively homogeneous and uniform structure throughout the section of the HLC/pullulan hydrogels. Due to the presence of HLC, which increased the relative cross-linking density of the hydrogel structure, Gel4, Gel5 and Gel6 appeared to have more compact porous structures with an average pore size range from 39 to 19 μm. Evidently, due to the higher cross-linking density, the porous sections of Gel6 were the thickest.
To investigate the degradation profiles of the hydrogels, we performed enzymatic degradation studies. Figure 6 shows that incubation of the pullulan hydrogels with collagenase I at 37 °C resulted in scaffold degradation after 50 h. Pullulanase incubation resulted in scaffold degradation after 30 h. Combination degradation experiments resulted in hydrogel dissolution after 40 h. However, HLC/pullulan hydrogels incubated in collagenase I, pullulanase and a combination of collagenase I/pullulanase did not degrade completely, and the weight loss range was from 50% to 80%. The proposed mechanism of enzymatic degradation (Figure 6(A)) was that the pullulanase could enhance the weight loss of pullulan hydrogels by acting as a surfactant to push the dispersed degradation products into water even though they could not dissolve in water. We also found that the presence of pullulanase could significantly accelerate the weight loss rates of all of the hydrogels, but it caused little difference in the decrease of MW and the change in
Figure 6. A schematic of the enzymatic degradation of the HLC/pullulan hydrogel in vitro. Pullulan degradation by pullulanase (A); HLC/pullulan hydrogel degradation by pullulanase (B), collagenase I (C) and pullulanase/collagenase I (D); (E) cross-linking index of hydrogels; weight loss of the hydrogels after degradation by pullulanase (F), collagenase I (G) and pullulanase/collagenase I (H).
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Figure 7. Effect of MWs of pullulan and addition of HLC on hydrogels cytocompatibility. A(Gel), B(Gel2), C(Gel3), D(Gel4), E(Gel5) and F(Gel6) SEM of the fibroblast adhesion to the hydrogels after 3 days. Scale bar for A, B, C, D, E, F: 100 μm and for a, b, c, d, e, f: 10 μm.
Figure 8. Evaluation of fibroblast live/dead staining. Fibroblast were cultured on the Gel1 (A), Gel2 (B), Gel3 (C), Gel4 (D), Gel5 (E) and Gel6 (F) and control group with live/dead staining at 3 days. Scale bar for all images: 200 μm. The death rates were shown in histogram (G and H) in 1, 2 and 3 days.
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hydrogel composition between the pullulan and HLC. Therefore, collagenase I was used for the degradation of the hydrogels. The activity site of collagenase I is the peptide hydroxyPro-Leu-Gly-Pro-Ala in the molecular chain of the collagen [18], and the activity site of pullulanase is the α-1, 6-glycosidic bond. The mechanism of degradation occurs as illustrated in Figure 6 (A, B, C, and D). The hydrogels comprised HLC and pullulan. After pullulanase degradation, the pullulan chains in the hydrogels were broken up into disaccharide polymerizate (Figure 6(B)). After collagenase I degradation, the hydroxyPro-Leu-Gly-ProAla peptides in the HLC chain in the hydrogels were cut, and the residual chains were combined with pullulan chains (Figure 6(C)). However, after treatment with a combination of collagenase I/pullulanase, the molecular chains were broken up, and the main backbone remained in the hydrogels (Figure 6(D)). During cross-linking, amine groups react with carbonyl groups. Thus, a higher cross-link index leads to lower free amine group content after cross-linking. Chemical cross-linking of HLC introduces covalent bonds into the network, thereby preventing dissolution of the gels. It was observed, however, that carboxylic acid-based cross-linking of
HLC was not in accordance with the number of free amine groups left after the cross-linking reaction. Thus, not all of the consumed amine groups were used for cross-linking. Therefore, 2, 4, 6 trinitrobenzene sulphonic acid (TNBS) [37] was used to measure the cross-linking index, which is the number of covalent bonds in the reaction between HLC and pullulan. Figure 6(E) shows that the cross-linking index are increased with the pullulan MW. The higher MW may accelerate the chemical cross-linking and then inhibit the degradation of the hydrogels. Therefore, the weight loss of Gel4, Gel5 and Gel6 was lower than the weight loss of Gel1, Gel2 and Gel3 (Figure 6 (F, G, and H)). 3.6. Cytotoxicity and cytocompatibility of hydrogels The compatibility of the hydrogels was evaluated by measuring the interaction of fibroblast attachment and hydrogels under in vitro conditions (Figure7). The fibroblast culture on the hydrogels at 3 days, the pullulan hydrogels and HLC/pullulan hydrogels caused different morphology and size of fibroblast at different MW and HLC addition. As expected, the number of the fibroblast increased substantially
Figure 9. 1 week after 0.2 mL of hydrogels (A-1) injected subcutaneously in wild-type mice (A-2), 1 week after hydrogels and tissue were taken out (A-3). H&E (A, B, C, D, E and F) and immunohistochemical staining with CD68 (a, b, c, d, e and f) in wild-type mice tissues for Gel1 (A, a), Gel2 (B, b), Gel3 (C, c), Gel4 (D, d), Gel5 (E, e), Gel6 (F, f) are injection at 1 week. The numbers represent the following: 1, chromatophore; 2, hydrogels; 3, adipocyte. Scale bar for A, B, C, D, E, F: 1 mm and for a, b, c, d, e, f: 200 μm.
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when the hydrophilic –OH groups of pullulan were replaced with aldehyde groups. Owing to the aldehyde effect, the hydrophobic character of the pullulan hydrogels was considerably increased. The cause of cytotoxicity appeared to be the ability of the hydroxyl groups to adsorb to the cellular surface, thus altering their normal enzymatic activity. In addition, when the aldehyde groups were replaced with amide groups, the cell viability in the presence of HLC/pullulan hydrogels was higher than for the pullulan hydrogels because the hydrophilic nature of the HLC/ pullulan hydrogels increased significantly. For example, Gel6 have a large number of fibroblast attach. However, the difference of
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morphology of cell attach on all hydrogels is little. Therefore, live/dead staining is further to demonstrate the cell viability. As shown in Figure 8, the cells remained viable with few dead cells, the death rate is 15%, 14%, 11%, 9%, 8% and 7% in 1 day, respectively. Image analysis using live/dead staining showed a significantly higher number of death cells attached on the Gel1, Gel2 and Gel3 compared with the Gel4, Gel5 and Gel6. The Gel1, Gel2 and Gel3 exhibited more cytotoxicity, with many dead cells on the surfaces and a cell death rate of 45%, 41% and 36% after culture at 72 h. However, the Gel4, Gel5 and Gel6 showed lower cytotoxicity, with few dead cells on the surface and a cell death
Figure 10. TEM images at 5000× magnification of the epidermis tissue (A1, B1, C1, D1, E1 and F1) and dermis tissue (A2, B2, C2, D2, E2 and F2) surrounding the hydrogel. Gel1 (A1 and A2), Gel2 (B1 and B2), Gel3 (C1 and C2), Gel4 (D1 and D2), Gel5 (E1 and E2) and Gel6 (F1 and F2) are shown 1 week after injection. The numbers represent the following: 1, macrophages; 2, lipid droplets; 3, inflammatory cells; 4, collagen fibers. Scale bar for all: 10 μm.
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rate of 30%, 26% and 14% after culture at 72 h. Therefore, Compared with pullulan hydrogels, the HLC/pullulan hydrogels could be improved the fibroblasts attach more obveriously and inhibited the cell death.
plays an important role in their anti-inflammatory and biodegradability properties. 4. Conclusions
3.7. H&E and immunohistochemical staining of skin tissue 0.2 mL hydrogels (Figure 9 A-1) were subcutaneously injected into wild-type mice, a small protuberates on back (Figure 9 A-2). The hydrogels and mouse tissues were taken (Figure 9 A-3) and were then assessed by H&E staining and immunohistochemical staining with CD68 as shown in Figure 9. The positive staining mainly appeared on chromatophore, hydrogels and adipocyte. The staining extensiveness and density of Gel4, Gel5 and Gel6 (Figure 9 d, e and f) were milder than Gel1, Gel2 and Gel3 (Figure 9 a, b and c) after staining, respectively. According to these observations, the inflammatory response of mice was indeed affected by the addition of HLC to the hydrogels (Gel4, Gel5 and Gel6), indicating that the anti-inflammatory response offered to the HLC decreased the staining density. However, further H&E staining might reveal the relationship between degradation and inflammatory response. Numerous studies have highlighted a variety of hydrogels that have been found to be non-biodegradable. Our results regarding the inflammatory response and biodegradability of hydrogels agreed with the results of previous reports investigating the addition of HLC for the identification of injectable hydrogels. Our findings confirm that HLC generates anti-inflammatory responses and is anti-biodegradable. Moreover, we found that Gel4, Gel5 and Gel6 (Figure 9 D, E and F) presented in the form of a circle rather than a flat morphology in the hydrogel regions (Figure 9 A, B and C). Our findings therefore demonstrate that HLC can be a significant factor in determining the hydrogel’s inflammatory behavior and biodegradability. 3.8. Transmission electron microscopy (TEM) analysis To perform an in situ biocompatibility assay of the pullulan and HLC/ pullulan hydrogels injected under the dermis of mice, the cell morphology and quantity of site surrounding the epidermis tissue (A1, B1, C1, D1, E1 and F1) and dermis tissue (A2, B2, C2, D2, E2 and F2) after hydrogel injection were determined by TEM as shown in Figure 10. The different thicknesses and cell morphologies of the epidermis tissue were observed in Figure 10 (A1, B1, C1, D1, E1 and F1), the periostraca of Gel1, Gel4, Gel5 and Gel6 were thinner than the periostraca of Gel2 and Gel3, and the gaps between the periostraca were shown clearly. The corneous layer, comprising a small amount of granular layer cells in this case, and the granular layer of transparent cutin could be observed in the cytoplasm. The particle shape was irregular, and the electron density was higher. Spine cells in the granular layer showed that the cells were polygonal and small, and fewer basal cells below the base layer of the spine cells were observed. The dermis layer below the basement membrane had a large number of collagen fibers, and the collagen fibers were transverse and longitudinally staggered, with spindle-shaped fibroblasts. In Figure 10 (A2, B2, C2, D2, E2 and F2), the macrophages, lipid droplets, inflammatory and collagen fibers can be observed in the dermis layer. The macrophages and the inflammatory cells were fewer but normal as seen in Figure 10, whereas the quantity of the lipid droplets was higher in Gel1, Gel2 and Gel3 compared with Gel4, Gel5 and Gel6. This finding may be because the tissue surrounding the hydrogels adsorbed the hydrogels and produced the metabolites after hydrogel injection into the dermis layer. More lipid droplets caused more rapid degradation of the hydrogels by the tissue. There are several lines of evidence supporting the idea of an intricate network of relationships between the inflammatory responses and lipid biology [38], including the evidence of lipid drops (LDs) [39]. During clinical and experimental sepsis conditions, the accumulation of LDs and the association of proteins involved in the inflammatory response to LDs have been found in immune cells [40]. Therefore, the HLC added in hydrogel systems
Novel injectable hydrogels were prepared from various MWs of pullulan, HLC and NaIO4, and the relationships between chemical structures, mechanical properties, morphological characteristics, and the degradation (in vivo and in vitro) of the pullulan and HLC/pullulan hydrogels were investigated in this study. The pullulan hydrogels were formed due to NaIO4 oxidizing the hydroxyl groups in pullulan molecules into aldehyde groups, whereas the amide bond link formed between the amino group of HLC and the aldehyde groups in aqueous media after crosslinking with NaIO4. The acute toxicity experiments in vivo showed that the HLC/pullulan hydrogels were safe for use in facial plastic surgeries. The intention of fibroblasts on hydrogels was attach significantly after 3 days. Compared with pullulan hydrogels, the HLC/pullulan hydrogels could be improved the fibroblasts attach more obveriously and inhibited the cell death. These results contributed to the establishment of an approach to optimize hydrogel composition based on the interaction between pullulan of different MW. We aimed to improve the properties of polysaccharide-based hydrogels compared with other sources using HLC with specific binding of polysaccharides using a cross-linking agent to improve the HLC/polysaccharide hydrogel properties for use in tissue engineering. Ultimately, the good mechanical strength, controllable degradation time and tissue compatibility demonstrated by the HLC/pullulan hydrogels may present advantages for in vivo applications. Acknowledgments This study was financially supported by the National Natural Science Foundation of China (21476182); the National High Technology Research and Development Program of China (2014AA02108); the Scientific Research Program of Shannixi Provincial Department of Education, China (14JS102); and the Scientific and Technologic Research Program of Shaanxi Province Academy of Sciences, China (No. 2014 k-01). References [1] X.C. Yang, Y.L. Niu, N.N. Zhao, C. Mao, F.J. Xu, A biocleavable pullulan-based vector via ATRP for liver cell-targeting gene delivery, Biomaterials 35 (2014) 3873–3884. [2] G. Fundueanu, M. Constantin, I. Oanea, V. Harabagiu, P. Ascenzi, B.C. Simionescu, Entrapment and release of drugs by a strict "on-off" mechanism in pullulan microspheres with pendant thermosensitive groups, Biomaterials 31 (2010) 9544–9553. [3] M.R. Rekha, K. Pal, P. Bala, M. Shetty, I. Mittra, G.S. Bhuvaneshwar, C.P. Sharma, Pullulan-histone antibody nanoconjugates for the removal of chromatin fragments from systemic circulation, Biomaterials 34 (2013) 6328–6338. [4] M.R. Rekha, C.P. Sharma, Blood compatibility and in vitro transfection studies on cationically modified pullulan for liver cell targeted gene delivery, Biomaterials 30 (2009) 6655–6664. [5] K.R. Sugumaran, R.V. Sindhu, S. Sukanya, N. Aiswarya, V. Ponnusami, Statistical studies on high molecular weight pullulan production in solid state fermentation using jack fruit seed, Carbohydrate Polymers 98 (2013) 854–860. [6] H. Bender, J. Lechmann, K. Wallenfells, Pullulan in extracellular glucans von Pullularia pullulans, Biochemica Biophysica Acta 36 (1959) 309–316. [7] T. Kimoto, T. Shibuya, S. Shiobara, Safety studies of a novel starch, pullulan: Chronic toxicity in rats and bacterial mutagenecity, Food and Chemical Toxicology 35 (1997) 323–329. [8] K.I. Shingel, Current knowledge on biosynthesis, biological activity, and chemical modification of the exopolysaccharide, pullulan, Carbohydrate Research 339 (2004) 447–460. [9] Y. Tsijisaka, M. Mitsushashi, Pullulan in industrial gum: Polysaccharides and their derivatives, In: Whistler R, BeMiller JN, editors. San Diego: Academic; 1993. 447–460. [10] V.D. Prajapati, G.K. Jani, S.M. Khanda, Pullulan: An exopolysaccharide and its various applications, Carbohydrate Polymers 95 (2013) 540–549. [11] K. Cheek, Determination of polymer–polymer miscibility by viscometry, European Polymer Journal 26 (1990) 423. [12] Z. Sun, W. Wang, Z. Feng, Criterion of polymer–polymer miscibility determined by viscometry, European Polymer Journal 51 (1992) 1259–1261. [13] H. Yim, S.G. Yang, Y.S. Jeon, et al., The performance of gadolinium diethylene triamine pentaacetate-pullulan hepatocyte-specific T1 contrast agent for MRI, Biomaterials 32 (2011) 5187–5194.
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Contents lists available at ScienceDirect
Materials Science and Engineering C journal homepage: www.elsevier.com/locate/msec
HLC/pullulan and pullulan hydrogels: their microstructure, engineering process and biocompatibility Xian Li a,b,1, Wenjiao Xue c,1, Yannan Liu b, Weina Li b, Daidi Fan b,⁎, Chenhui Zhu b, Yaoyu Wang a,⁎⁎ a b c
College of chemistry & materials science, Northwest University, Taibai North Road 229, Xi’an, Shaanxi 710069, China Shaanxi Key Laboratory of Degradable Biomedical Materials, Department of Chemical Engineering, Northwest University, Taibai North Road 229, Xi’an, Shaanxi 710069, China Shannxi provincial institute of microbiology, Xi’ an 710043, China
a r t i c l e
i n f o
Article history: Received 10 August 2015 Accepted 7 September 2015 Available online 11 September 2015 Keywords: HLC/pullulan hydrogels Pullulan hydrogels NaIO4 Subcutaneous Biodegradation
a b s t r a c t New locally injectable biomaterials that are suitable for use as soft tissue fillers are needed to address a significant unmet medical need. In this study, we used pullulan and human-like collagen (HLC) based hydrogels with various molecular weights (MWs) in combination therapy against tissue defects. Briefly, pullulan was crosslinked with NaIO4 to form a pullulan hydrogel and then may coupled with HLC using the reaction between the –NH2 end-group of HLC and the –CHO group present on the aldehyde pullulan to form the HLC/pullulan hydrogel, wherein the NaIO4 acted as the crosslinking and oxidizing agent. The good miscibility of pullulan and HLC in the hydrogels was confirmed via Fourier transform infrared spectroscopy, scanning electron microscopy, compression testing, enzyme degradation testing, cell adhesions, live/dead staining and subcutaneous filling assays. Here, pullulan hydrogels with various MWs were fabricated and physicochemically characterized. Limitations of the pullulan hydrogels included inflammation, poor mechanical strength, and degradation. By contrast, the properties of the HLC/pullulan hydrogels strongly enhanced. The efficacy of these hydrogels was evaluated both in vitro and in vivo. Our results indicate that HLC/pullulan hydrogels may have therapeutic value as efficient soft tissue fillers, with reduced inflammation, improved cell adhesion and delayed hydrogel degradation. © 2015 Elsevier B.V. All rights reserved.
1. Introduction Recently, microbial polysaccharides have received increasing interest because of their useful physiochemical properties and natural biodegradability. The repeating units of these exopolysaccharides are regular, branched or unbranched, and interconnected by glycosidic linkages. For example, due to its non-toxic, non-immunogenic, non-mutagenic and non-carcinogenic nature, pullulan, a non-ionic exopolysaccharide of fungal origin, has recently been evaluated for various biomedical applications, including tissue engineering, targeted drug/gene delivery, targeted drug therapy and wound healing [1–3]. Pullulan is one of the polymers obtained from the fermentation medium of black yeast (e.g., Auerobasidium pullulans). It comprises maltotriose units connected by α (1 → 4) glycosidic bond, whereas the consecutive maltotriose units are connected to each other by α (1 → 6) glycosidic bond [4–6]. It has a molecular weight (MW) range of 5,000–9,000,000 g/mol with a straight unbranched chain [7]. Pullulan has considerable mechanical strength and possesses numerous functional properties such as
⁎ Corresponding author. ⁎⁎ Correspondence to: Y. Wang, College of chemistry & materials science, Northwest University, Taibai North Road 229, Xi’an, Shaanxi 710069, China. E-mail addresses: [email protected] (D. Fan), [email protected] (Y. Wang). 1 These authors contributed equally to this work.
http://dx.doi.org/10.1016/j.msec.2015.09.039 0928-4931/© 2015 Elsevier B.V. All rights reserved.
adhesiveness, the ability to form a film and enzyme-mediated degradability [8]. However, a pullulan solution’s viscosity changes with its MW [9]. Therefore, it can be used as a low-viscosity filler in beverages and sauces, or it can be used as a high-viscosity filler in tissue engineering applications. Interestingly, the viscosity of a pullulan solution does not change with heat, changes in pH or in the presence of most metal ions, including sodium chloride [10]. Injectable hydrogels are promising scaffolds for cell encapsulation, tissue repair, tissue reconstruction and drug delivery. Novel polymeric materials such as polymer blends and composites have led to the advancement of technologies to develop injectable hydrogels [11]. Polymer blends are prepared by physically mixing two or more polymers, and they tend to have properties superior to any of the individual polymers [12]. Moreover, hydrogels, prepared as composites of polysaccharides and collagen with or without a crosslinking agent, have garnered great interest in tissue engineering applications. Numerous polysaccharides have been used for preparing hydrogels, including chitosan, pullulan, hyaluronic acid and carboxymethyl chitosan. Chitosan and hyaluronic acid have gained much popularity in the field of tissue engineering. Recently, pullulan, a new material and a non-ionic polysaccharide with unique characteristics (non-toxic, non-immunogenic, non-mutagenic and non-carcinogenic), has been explored for various biomedical applications [13]. Human-like collagen (HLC) has been used in hydrogels [14], vascular scaffolds [15] and artificial bones [16]
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due to its low rejection rate, good biocompatibility [17], biodegradability [18], workability and no risk of viral infection. Many papers have reported on using chitosan or chitosan derivatives and hyaluronic acid crosslinking with collagen or other crosslinking agents to form hydrogels [19]. Many articles have described the use of pullulan gels or nanoparticles for applications in gene-drug delivery and tissue repair. There are fewer articles reporting the formulation of pullulan-based copolymers. Gark et al. [20], Miyahara et al [21] and Rustad et al [22] reported that the pullulan and collagen hydrogels was used to force cell delivery in wounds, guide bone regeneration and as biomimetic hydrogel scaffold to improve the mesenchymal stem cell angiogenic capacity and stemness. Our report is mainly about different molecule weight of pullulan and collagen to fabricate injectable hydrogels by another method for skin reconstruction. Feng et al. reported the collagen and dialdehyde pullulan fabricate the complex scaffold materials [23]. Bruneel et al. reported that chemical modification of pullulan by periodate oxidation [24]. Based on the above studies, we choose the pullulan modification by periodate oxidation (NaIO4) and human-like collagen to fabricate hydrogels for skin reconstruction. Therefore, in this report, we investigated the use of pullulan and HLC with a crosslinking agent for the preparation of hydrogels. We used NaIO4 not only as an oxidant but also as a crosslinking agent in our hydrogel system. We prepared hydrogel systems with varying MWs (100,000–530,000 g/mol), with straight unbranched chains and with or without HLC. The HLC/pullulan hydrogels have biological effects that are different from pullulan hydrogels, and have excellent biological and mechanical properties.
HLC/pullulan hydrogels. Then, the 4% HLC solution was sterilized using a 0.22-μm sterilizing filter. The composite mixture was gently vortexes for 30 min at 4 °C to promote the homogeneous distribution of polymers within the hydrogel. The mixture was then mixing with 0.5 mL, 0.5 mol/L sterile NaIO4 and stirred for 20 min until dissolution was complete. The solutions were then placed in a water bath for gelling at 37 °C for 12 h. The pullulan and HLC/pullulan hydrogels were jellylike as shown in Figure 1 (A, B, C and D). The hydrogels were dialyzed in pyrogen-free water at room temperature for 12 h for hydrogel properties analysis and the in vitro biocompatibility analysis. Then, the hydrogels were extruded with a granulator to make them into small granules and were stored at 4 °C for in vivo biocompatibility analysis. All of the experiments were performed in a sterile environment. The hydrogels formed, purify and made into granules process is shown in Figure 1(F).
2. Materials and methods
The porosity of the pullulan and HLC/pullulan hydrogels was determined by the liquid displacement method. First, the volume and weight of the hydrogels were measured and recorded as V0 and W0, respectively. Second, the samples were immersed in dehydrated alcohol for 48 h until they were saturated, and then the hydrogels were weighed again (W1). Finally, the porosity was calculated according to the formula: P = (W1 − W0)/ρV0, where ρ represents the density of the dehydrated alcohol. The pore diameters of the hydrogel were measured using ImageJ software. Each pore of the sample was measured, and the average value of the measurements was calculated.
2.1. Materials HLC was expressed by E. coli by cloning a partial cDNA derived from human collagen mRNA [16]. HLC is a macromolecular watersoluble protein with a MW of 97,000 g/mol (China patent number: ZL01106757.8). The protein encoded by the partial cDNA is reference by Fan DD, et al [25]. The procedure to produce and purify HLC from the bacteria is reference by Ma et al [26] and Zhang HZ et al [27]. Pullulan (MW, 100,000 g/mol; 170,000 g/mol; 530,000 g/mol) was obtained from the Shaanxi Provincial Institute of Microbiology. NaIO4 was purchased from Shandong Frieda Biological Technology, Ltd. (China). The live/dead viability/cytotoxicity kits were purchased from Sigma-Aldrich. All of the other solvents and reagents were analytical grade. 2.2. Preparation of pullulan hydrogels and HLC/pullulan hydrogels Pullulan hydrogel and HLC/pullulan hydrogel compositions are shown in Table 1. One gram of pullulan powder was dissolved in pyrogen-free water with or without HLC to a total volume of 6 mL. HLC was mixed in at a MW of 100,000 g/mol, 170,000 g/mol or 530,000 g/mol pullulan (0.25 g HLC per 0.5 g pullulan) to prepare the
Table 1 The pullulan and HLC/pullulan hydrogels dompositions. Hydrogels
Sample
Pullulan (MW, g/mol)
HLC (%)
Pullulan (%)
NaIO4 (mol/L)
Pullulan hydrogels
Gel1 Gel2 Gel3 Gel4 Gel5 Gel6
100,000 170,000 530,000 100,000 170,000 530,000
0 0 0 4 4 4
8 8 8 8 8 8
0.5 0.5 0.5 0.5 0.5 0.5
HLC/pullulan hydrogels
2.3. FT-IR spectroscopic analysis The different components of the HLC/pullulan hydrogel systems were verified by FT-IR. Infrared spectra of the specimen powders, namely pullulan (MW of 100,000 g/mol, 170,000 g/mol and 530,000 g/mol) and NaIO4 as well as of the HLC and the six formulations of hydrogels were recorded using an FT-IR spectrophotometer (Thermo Fisher Scientific, Waltham, MA, USA). The samples were triturated with KBr at a ratio of 1:100 and pressed to form pelleted samples for FT-IR spectroscopic analysis at 500–4000 cm−1. 2.4. Porosity and pore diameter of the hydrogels
2.5. Microscopic investigations The pore morphology of the pullulan hydrogels and HLC/pullulan hydrogels was examined by scanning electron microscopy (SEM). The samples were carefully lyophilized to maintain their threedimensional porous structures. The lyophilized hydrogel samples were immersed in liquid nitrogen, and the vitrified samples were carefully cut with a cold knife. The cut samples were mounted, sputter-coated with gold, and observed under SEM using a Hitachi S-570 SEM microscope (Hitachi, Tokyo, Japan). 2.6. Swelling measurement Hydrogel water absorption capacity was calculated as a swelling ratio: Swelling ratio = (weight of wet sample − weight of dry sample)/weight of dry sample. For incubation studies, the hydrogels were incubated in deionized water or in PBS (pH = 7.4) at 4 °C or 37 °C. Excess liquid was gently shaken off, and the weights of the swollen gels were obtained. Hydrogel weights were measured at 24 h. Six samples were tested for each condition. 2.7. Compression experiment The hydrogels were tested using a gel strength instrument (Electronic Universal Testing Machine) as follows: first, the height of the compression plate was adjusted to 10 mm, and its final height was adjusted to 4 mm.
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Figure 1. Visual inspection of hydrogels. (A) Gel1, Gel2, Gel3, (B) Gel4; (C) Gel5; (D) Gel6; (E)hydrogels particles; (F) the hydroges forming process.
The instrument disk was circular with a diameter of 10 mm, and the compression speed was adjusted to 2 mm/min. The gel was cut to a length of 10 mm and placed on the instrument disc. When the compression displacement reached 6 mm, the test was stopped. The compression modulus and the relationship between the compression load and compression displacement were calculated. The crosslinking density (ρ) [28] was determined from the compression modulus (G) and swelling ratio (Q) as follows: ρ = GQ1/3/RT, where G is the compression modulus, R is the gas constant (8.314 × 103 KPa · cm3 · mol−1 · K−1) and T is 310 K. 2.8. Quantification of crosslinking index The cross-linking index was measured using 2,4,6trinitrobenzenesulfonic acid (TNBS) [29] for the determination of free amine groups in the soluble and poorly soluble proteinaceous materials. Hydrogel mixtures containing HLC only and HLC/pullulan hydrogels (4 mg) were reacted with TNBS (0.01 M) in sodium hydrogen carbonate solution (pH 8.2, 4% w/v, 2 mL) for 2 h at 40 °C. Then, hydrochloric acid (6 M, 3 mL) was added to hydrolyze the water (5 mL) in the solution, and the absorbance was measured against a TNBS solution without gelatin that had been treated exactly the same as the other samples. Crosslinking index = (1 − absorbance of HLC/pullulan hydrogel/absorbance of HLC) × 100%.
2.10. Cell adhesion, proliferation and viability analysis Cell viability was evaluated by cell adhesion and live/dead staining was observed using a confocal microscopy and SEM, respectively. The live/dead viability/cytotoxicity kit for mammalian cells are using for live/dead staining as described previously [31]. Cryopreserved fibroblasts were thawed, centrifuged, and then cultured in DMEM containing 10% FBS. Confluent cells were passaged into multiple flasks for subsequent experiments. One day before culture, four wet hydrogel samples (10-mm diameter and 3-mm height) were sterilized at 121 °C for 5 min, rinsed with DMEM twice in a 48-well plate, and then equilibrated in the medium overnight until cell plating. For surface adhesion studies, logarithmic-phase fibroblasts (5 × 105 cells/mL) were seeded onto four hydrogels, and 1 mL of DMEM was added to each well in a 5% CO2, 37 °C incubator. Parts of the samples were fixed in glutaraldehyde at 72 h, and the cell morphology was observed by SEM after vacuum freeze drying. The remaining portions of the samples were stained by live/dead staining at 24, 48 and 72 h to evaluate cell death rate on the hydrogels, and the live/dead cells staining at 72 h were observed using a confocal microscopy. The death rate of cell was calculated by dead cell/live cell × 100% [14]. A control condition is conducted in normal fibroblast cultured by DMEM media, the test samples were cultured by the Gel1, Gel2, Gel3, Gel4, Gel5 and Gel6. Five parallel samples of each material were utilized.
2.9. In vitro degradation
2.11. The biocompatibility of hydrogel in vivo
Dry pullulan hydrogels (Gel1, Gel2 and Gel3) and HLC/pullulan hydrogels (Gel4, Gel5 and Gel6) were incubated with pullulanase (20 U/mL) in PBS and weighed every 30 min. Similar experiments were performed with collagenase I (20 U/mL) in PBS. Combination degradation studies using both pullulanase and collagenase I was conducted in PBS using the same concentrations as above [30]. The initial weight at time 0 was the dry weight of the hydrogels (W0) and the wet hydrogels after being degraded by enzymes were washed three times with water before vacuum drying (W1) and used for subsequent measurement. The weight loss (%) was calculated as follows: (W0 − W1)/ W0 × 100%. Experiments were performed six times for each condition at room temperature.
All experiments were performed in compliance with the relevant laws and institutional guidelines, and conducted with the approval of the Institute Animal Ethics Committee. This study was supported by Shaanxi Key Laboratory of Degradable Biomedical Materials and Northwest University. Wild-type adult mice (n = 18) were anesthetized with isoflurane inhalation. After cleansing the skin with 70% alcohol, 0.2 mL of pullulan hydrogels and HLC/pullulan hydrogels granules, obtained from the extruding granulator under ultraviolet (UV) light, were injected subcutaneously. The hydrogels were examined daily and harvested at 1 week. Digital photographs were taken to quantify the residual hydrogel size. Then, the hydrogels injected subcutaneously into wild-mice were taken out after 1 week and
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Figure 2. (a) FT-IR spectra of the pullulan hydrogels (a, b, c) and HLC/pullulan hydrogels (d, e, f). Small letters represent the following: a, Gel1; b, Gel2; c, Gel3; d, Gel4; e, Gel5; f, Gel6. (b) FT-IR spectra of pullulan (100,000, 170,000 and 530,000 g/mol), NaIO4, and HLC.
immediately fixed in 10% neutral buffered formalin. The tissue surrounding the hydrogels was tested by H&E staining and immunohistochemical analysis [19]. The TEM experimental process was reference by Xian Li et al [18].
2.12. Calculations and data analysis Data were collected in a Microsoft Excel 2000 database, and the results were presented as means and standard deviations using the Origins 7.0 software. Student's t-test was performed to determine the statistical significance between experimental groups. A value of p b 0.05 was considered statistically significant.
3. Results and discussion 3.1. FT-IR spectroscopic analysis The components of the hydrogel systems were assessed using Fourier transform infrared spectroscopy (FT-IR). The spectra obtained for pullulan, HLC, pullulan hydrogels, and HLC/pullulan hydrogels are displayed in Figure 2(a) and (b). The main characteristic absorption bands of the pullulan hydrogels at 1624 and 1444 cm−1 were caused by the C = O stretching. These results confirmed that the OH groups of pullulan were oxidized to C = O by NaIO4 in the pullulan-based hydrogels. Pullulan contains three different anhydroglucoside moieties in the repeating unit. Therefore, periodate oxidation of pullulan results
Figure 3. Synthesis of the pullulan and HLC/pullulan hydrogels. Schematic representation of the proposed mechanisms of the interactions between pullulan, HLC, and NaIO4.
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in different types of dialdehyde structures. The adjacent two hydroxyl groups of pullulan were periodate into dialdehyde or carboxide. The same results were reported by Bruneel et al [24] and Spatareanu A [32]. The HLC/pullulan hydrogels were identified in the respective spectra from the C = O asymmetric stretching band at 1644 cm−1. Although the HLC amide bands were partially masked by the adsorbed band at 1640 cm−1, their presence was denoted by the new band at 1515– 1540 cm−1 due to the absorption of the protonated amino groups. The characteristic peaks representing the stretching vibration of VC-N and the coupled vibrations of δNH and VC-N were observed at 1413 and 1515 cm−1, respectively and were may be consistent with the formation of -CONH at -C = O sites on pullulan coordinated with -NH2 sites on HLC in HLC/pullulan hydrogels. However, the adsorbed band at 1367 cm−1 may be represent ester group (COOR), which may be contribute the carboxyl group of HLC reaction with pullulan (Figure 3). These bands were absent in the pullulan and HLC. 3.2. The mechanism of hydrogel formation The HLC/pullulan hydrogels were synthesized by the oxidation and condensation of aldehydes and amines to form a Schiff base as shown in Figure 3. In the reaction between NaIO4 and the various MWs of pullulan, two adjacent hydroxyl groups of pullulan were oxidized to an aldehyde group in one step to form the pullulan hydrogels. After
pullulan powder resolved in deionized H2O, the pullulan particle contact the surface of water immediately, leading to the formation of water film enwrap the pullulan powder; then stirring, as need enough time for pullulan powder to saturate into water. The characteristic hydration properties of pullulan might be due to the flexibility of the pullulan skeleton [33]. On the other hand random coil proteins are able of forming coherent film due to the formation of extensive hydrogen bonds [34]. Therefore, HLC solution made it easier to dissolve the pullulan powder than deionized H2O. Various MWs of pullulan were cross-linked with HLC by adding NaIO4 to form the HLC/pullulan hydrogels in two steps. Crosslinking with HLC via NaIO4, caused the pullulan with unbranched chains in the hydrogel systems to increase in MW, improved its anti-degradation properties, and changed the spatial structure within the hydrogel system. This unique spatial structure may have influenced the properties, physiochemical characteristics and biocompatibility of the HLC/pullulan hydrogels. Therefore, developing HLC/pullulan hydrogels may produce ideal soft filler materials for tissue engineering applications. 3.3. Physical characteristics of the hydrogels To investigate the water retention properties of the hydrogels, swelling studies of the gels were performed with water and phosphatebuffered saline (PBS) at different temperatures as shown in Figure
Figure 4. Physical characterization and toxicity of different gel materials (Gel1, Gel2, Gel3, Gel4, Gel5, Gel6). Swelling ratios of gels in ddH2O (A) and PBS (B); elastic modulus (C), crosslinking density (D), porosity (E) and pore diameters (F) of gels; compression load of pullulan hydrogels (G) and HLC/pullulan hydrogels (H).
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4(A) and (B). The swollen hydrogels retained their general shape and did not degrade after incubation in either deionized water or PBS. The swelling ratios for the pullulan hydrogels and the HLC/pullulan hydrogels incubated in deionized water and PBS at 4 °C and 37 °C were different, and we found that the swelling ratios for the pullulan hydrogel were higher than the ratios for the HLC/pullulan hydrogels. Hydrogels with the same pullulan to HLC ratio reached swelling equilibrium at 24 h independent of the crosslinking density or the osmolarity of the swelling medium. Hydrogels with different MWs of pullulan had different crosslinking densities. There was almost no hydrogel swelling during 12 ~ 48 h incubation in all conditions. The hydrogel reaching the swelling equilibrium was recorded at 24 h, indicating maximal water absorption after overnight incubation. In addition, when the swelling medium was changed from PBS to ddH2O, the gels swelled more than at equilibrium. The viscoelastic characteristics of the gels were assessed as shown in Figure 4(C). We found an increasing elastic modulus (3.8 KPa (Gel1) to 8.5 KPa (Gel3) and 10 KPa (Gel4) to 100 KPa (Gel6)) and decreasing swelling (22.6 ± 1 (Gel1) to 18.5 ± 0.9 (Gel3) and 17.8 ± 0.9 (Gel4) to 15.5 ± 0.7 (Gel6)) as the degree of cross-linking increased (Figure 4(D)). These findings are in agreement with the theoretically expected correlation between mechanical characteristics and cross-linking density [35]. The elastic modulus was directly proportional to the crosslinking density [36]. Despite the overall high water content of the materials and variability of the MW of pullulan from 100,000 to 530,000 g/mol, the meshwork stability was excellent. The higher MW of pullulan resulted in more stable molecular chains with more covalent cross-linking sites, thereby resulting in higher crosslinking densities in the pullulan hydrogels and HLC/pullulan hydrogels as shown in Figure 4(D). The relationship between compression load and the compression displacement of hydrogels is shown in Figure 4 (G, H). At the same compression displacement, hydrogels with different pullulan MWs showed different compression loads. As expected, the compression modulus
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increased as the compression load increased. The mechanical integrity of the hydrogels indicated that the compression modulus increased as the degree of cross-links increased, suggesting that the network became more resistant to change when a load was applied with more crosslinks. As expected, the HLC also affected the mechanical strength of the gels. The compression monotonically increased with the addition of HLC, presumably due to the increased cross-link density. The compression testing showed that the pullulan hydrogels were mechanically robust and showed varied loads from 1.3 KPa to 3 KPa depending on the NaIO4 crosslinking. Additionally, the addition of HLC led to a varied load from 2.5 KPa to 12 KPa. One striking observation was that pullulan and HLC appeared to affect the compression load similarly. Figure 4(E) shows effects of the MW of pullulan and the addition of HLC on particle size and the overall porosity of the prepared porous hydrogels. In all of the cases, the overall porosity increased linearly with decreasing MW of pullulan. The varying MW of pullulan had little impact on the increasing rate of overall porosity. However, the addition of HLC produced a more significant difference in overall porosity. For instance, the porosity of Gel1, Gel2, Gel3 was 80%, 79%, 71%, respectively, whereas the overall porosity of Gel4, Gel5, Gel6 was 69%, 65%, 60%, respectively. Clearly, the HLC created more opportunities for connections between pullulan powder particles, resulting in closer pore formations. The linear relationships implied that the overall porosity of hydrogels can be quantitatively adjusted by the addition of HLC and varying the MW of the pullulan powder particles. Figure 4(F) shows the pore size diameters of the various hydrogels. We found that the MW of pullulan influences the pore size and pore network of hydrogels. At the MW of 530,000 g/mol (Gel3, Gel6), the pore size diameters was higher than in the lower pullulan MW gels (Gel1, Gel4). From these studies, it can be concluded that the increase in porosity aids in higher MWs in gels. Based on the above analysis, our results indicate that pullulan and HLC had molecular interactions during the gelation process, and the chemical reaction was more obvious when the MW of pullulan was higher.
Figure 5. SEM micrographs × 250 of formed hydrogels A(Gel), B(Gel2), C(Gel3), D(Gel4), E(Gel5) and F(Gel6) at room temperature after vacuum freeze-drying for 48 h. Scale bar for all: 120 μm.
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3.4. Network structure
3.5. In vitro degradation by enzymes
Figure 5 shows representative scanning electron microscopy (SEM) images of the cross-section of the lyophilized pullulan hydrogels (Gel1, Gel2 and Gel3) and HLC/pullulan hydrogels (Gel4, Gel5 and Gel6). All the hydrogels were highly porous throughout the crosssection. Gel1 exhibited the largest pore size (average 52 μm) among all the hydrogel formulations. Figure 5 illustrates the differences in the cross-section structure of the hydrogels. The surface of pores of pullulan and HLC/pullulan hydrogels showed a rougher structure. However, the cross-section revealed relatively homogeneous and uniform structure throughout the section of the HLC/pullulan hydrogels. Due to the presence of HLC, which increased the relative cross-linking density of the hydrogel structure, Gel4, Gel5 and Gel6 appeared to have more compact porous structures with an average pore size range from 39 to 19 μm. Evidently, due to the higher cross-linking density, the porous sections of Gel6 were the thickest.
To investigate the degradation profiles of the hydrogels, we performed enzymatic degradation studies. Figure 6 shows that incubation of the pullulan hydrogels with collagenase I at 37 °C resulted in scaffold degradation after 50 h. Pullulanase incubation resulted in scaffold degradation after 30 h. Combination degradation experiments resulted in hydrogel dissolution after 40 h. However, HLC/pullulan hydrogels incubated in collagenase I, pullulanase and a combination of collagenase I/pullulanase did not degrade completely, and the weight loss range was from 50% to 80%. The proposed mechanism of enzymatic degradation (Figure 6(A)) was that the pullulanase could enhance the weight loss of pullulan hydrogels by acting as a surfactant to push the dispersed degradation products into water even though they could not dissolve in water. We also found that the presence of pullulanase could significantly accelerate the weight loss rates of all of the hydrogels, but it caused little difference in the decrease of MW and the change in
Figure 6. A schematic of the enzymatic degradation of the HLC/pullulan hydrogel in vitro. Pullulan degradation by pullulanase (A); HLC/pullulan hydrogel degradation by pullulanase (B), collagenase I (C) and pullulanase/collagenase I (D); (E) cross-linking index of hydrogels; weight loss of the hydrogels after degradation by pullulanase (F), collagenase I (G) and pullulanase/collagenase I (H).
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Figure 7. Effect of MWs of pullulan and addition of HLC on hydrogels cytocompatibility. A(Gel), B(Gel2), C(Gel3), D(Gel4), E(Gel5) and F(Gel6) SEM of the fibroblast adhesion to the hydrogels after 3 days. Scale bar for A, B, C, D, E, F: 100 μm and for a, b, c, d, e, f: 10 μm.
Figure 8. Evaluation of fibroblast live/dead staining. Fibroblast were cultured on the Gel1 (A), Gel2 (B), Gel3 (C), Gel4 (D), Gel5 (E) and Gel6 (F) and control group with live/dead staining at 3 days. Scale bar for all images: 200 μm. The death rates were shown in histogram (G and H) in 1, 2 and 3 days.
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hydrogel composition between the pullulan and HLC. Therefore, collagenase I was used for the degradation of the hydrogels. The activity site of collagenase I is the peptide hydroxyPro-Leu-Gly-Pro-Ala in the molecular chain of the collagen [18], and the activity site of pullulanase is the α-1, 6-glycosidic bond. The mechanism of degradation occurs as illustrated in Figure 6 (A, B, C, and D). The hydrogels comprised HLC and pullulan. After pullulanase degradation, the pullulan chains in the hydrogels were broken up into disaccharide polymerizate (Figure 6(B)). After collagenase I degradation, the hydroxyPro-Leu-Gly-ProAla peptides in the HLC chain in the hydrogels were cut, and the residual chains were combined with pullulan chains (Figure 6(C)). However, after treatment with a combination of collagenase I/pullulanase, the molecular chains were broken up, and the main backbone remained in the hydrogels (Figure 6(D)). During cross-linking, amine groups react with carbonyl groups. Thus, a higher cross-link index leads to lower free amine group content after cross-linking. Chemical cross-linking of HLC introduces covalent bonds into the network, thereby preventing dissolution of the gels. It was observed, however, that carboxylic acid-based cross-linking of
HLC was not in accordance with the number of free amine groups left after the cross-linking reaction. Thus, not all of the consumed amine groups were used for cross-linking. Therefore, 2, 4, 6 trinitrobenzene sulphonic acid (TNBS) [37] was used to measure the cross-linking index, which is the number of covalent bonds in the reaction between HLC and pullulan. Figure 6(E) shows that the cross-linking index are increased with the pullulan MW. The higher MW may accelerate the chemical cross-linking and then inhibit the degradation of the hydrogels. Therefore, the weight loss of Gel4, Gel5 and Gel6 was lower than the weight loss of Gel1, Gel2 and Gel3 (Figure 6 (F, G, and H)). 3.6. Cytotoxicity and cytocompatibility of hydrogels The compatibility of the hydrogels was evaluated by measuring the interaction of fibroblast attachment and hydrogels under in vitro conditions (Figure7). The fibroblast culture on the hydrogels at 3 days, the pullulan hydrogels and HLC/pullulan hydrogels caused different morphology and size of fibroblast at different MW and HLC addition. As expected, the number of the fibroblast increased substantially
Figure 9. 1 week after 0.2 mL of hydrogels (A-1) injected subcutaneously in wild-type mice (A-2), 1 week after hydrogels and tissue were taken out (A-3). H&E (A, B, C, D, E and F) and immunohistochemical staining with CD68 (a, b, c, d, e and f) in wild-type mice tissues for Gel1 (A, a), Gel2 (B, b), Gel3 (C, c), Gel4 (D, d), Gel5 (E, e), Gel6 (F, f) are injection at 1 week. The numbers represent the following: 1, chromatophore; 2, hydrogels; 3, adipocyte. Scale bar for A, B, C, D, E, F: 1 mm and for a, b, c, d, e, f: 200 μm.
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when the hydrophilic –OH groups of pullulan were replaced with aldehyde groups. Owing to the aldehyde effect, the hydrophobic character of the pullulan hydrogels was considerably increased. The cause of cytotoxicity appeared to be the ability of the hydroxyl groups to adsorb to the cellular surface, thus altering their normal enzymatic activity. In addition, when the aldehyde groups were replaced with amide groups, the cell viability in the presence of HLC/pullulan hydrogels was higher than for the pullulan hydrogels because the hydrophilic nature of the HLC/ pullulan hydrogels increased significantly. For example, Gel6 have a large number of fibroblast attach. However, the difference of
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morphology of cell attach on all hydrogels is little. Therefore, live/dead staining is further to demonstrate the cell viability. As shown in Figure 8, the cells remained viable with few dead cells, the death rate is 15%, 14%, 11%, 9%, 8% and 7% in 1 day, respectively. Image analysis using live/dead staining showed a significantly higher number of death cells attached on the Gel1, Gel2 and Gel3 compared with the Gel4, Gel5 and Gel6. The Gel1, Gel2 and Gel3 exhibited more cytotoxicity, with many dead cells on the surfaces and a cell death rate of 45%, 41% and 36% after culture at 72 h. However, the Gel4, Gel5 and Gel6 showed lower cytotoxicity, with few dead cells on the surface and a cell death
Figure 10. TEM images at 5000× magnification of the epidermis tissue (A1, B1, C1, D1, E1 and F1) and dermis tissue (A2, B2, C2, D2, E2 and F2) surrounding the hydrogel. Gel1 (A1 and A2), Gel2 (B1 and B2), Gel3 (C1 and C2), Gel4 (D1 and D2), Gel5 (E1 and E2) and Gel6 (F1 and F2) are shown 1 week after injection. The numbers represent the following: 1, macrophages; 2, lipid droplets; 3, inflammatory cells; 4, collagen fibers. Scale bar for all: 10 μm.
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rate of 30%, 26% and 14% after culture at 72 h. Therefore, Compared with pullulan hydrogels, the HLC/pullulan hydrogels could be improved the fibroblasts attach more obveriously and inhibited the cell death.
plays an important role in their anti-inflammatory and biodegradability properties. 4. Conclusions
3.7. H&E and immunohistochemical staining of skin tissue 0.2 mL hydrogels (Figure 9 A-1) were subcutaneously injected into wild-type mice, a small protuberates on back (Figure 9 A-2). The hydrogels and mouse tissues were taken (Figure 9 A-3) and were then assessed by H&E staining and immunohistochemical staining with CD68 as shown in Figure 9. The positive staining mainly appeared on chromatophore, hydrogels and adipocyte. The staining extensiveness and density of Gel4, Gel5 and Gel6 (Figure 9 d, e and f) were milder than Gel1, Gel2 and Gel3 (Figure 9 a, b and c) after staining, respectively. According to these observations, the inflammatory response of mice was indeed affected by the addition of HLC to the hydrogels (Gel4, Gel5 and Gel6), indicating that the anti-inflammatory response offered to the HLC decreased the staining density. However, further H&E staining might reveal the relationship between degradation and inflammatory response. Numerous studies have highlighted a variety of hydrogels that have been found to be non-biodegradable. Our results regarding the inflammatory response and biodegradability of hydrogels agreed with the results of previous reports investigating the addition of HLC for the identification of injectable hydrogels. Our findings confirm that HLC generates anti-inflammatory responses and is anti-biodegradable. Moreover, we found that Gel4, Gel5 and Gel6 (Figure 9 D, E and F) presented in the form of a circle rather than a flat morphology in the hydrogel regions (Figure 9 A, B and C). Our findings therefore demonstrate that HLC can be a significant factor in determining the hydrogel’s inflammatory behavior and biodegradability. 3.8. Transmission electron microscopy (TEM) analysis To perform an in situ biocompatibility assay of the pullulan and HLC/ pullulan hydrogels injected under the dermis of mice, the cell morphology and quantity of site surrounding the epidermis tissue (A1, B1, C1, D1, E1 and F1) and dermis tissue (A2, B2, C2, D2, E2 and F2) after hydrogel injection were determined by TEM as shown in Figure 10. The different thicknesses and cell morphologies of the epidermis tissue were observed in Figure 10 (A1, B1, C1, D1, E1 and F1), the periostraca of Gel1, Gel4, Gel5 and Gel6 were thinner than the periostraca of Gel2 and Gel3, and the gaps between the periostraca were shown clearly. The corneous layer, comprising a small amount of granular layer cells in this case, and the granular layer of transparent cutin could be observed in the cytoplasm. The particle shape was irregular, and the electron density was higher. Spine cells in the granular layer showed that the cells were polygonal and small, and fewer basal cells below the base layer of the spine cells were observed. The dermis layer below the basement membrane had a large number of collagen fibers, and the collagen fibers were transverse and longitudinally staggered, with spindle-shaped fibroblasts. In Figure 10 (A2, B2, C2, D2, E2 and F2), the macrophages, lipid droplets, inflammatory and collagen fibers can be observed in the dermis layer. The macrophages and the inflammatory cells were fewer but normal as seen in Figure 10, whereas the quantity of the lipid droplets was higher in Gel1, Gel2 and Gel3 compared with Gel4, Gel5 and Gel6. This finding may be because the tissue surrounding the hydrogels adsorbed the hydrogels and produced the metabolites after hydrogel injection into the dermis layer. More lipid droplets caused more rapid degradation of the hydrogels by the tissue. There are several lines of evidence supporting the idea of an intricate network of relationships between the inflammatory responses and lipid biology [38], including the evidence of lipid drops (LDs) [39]. During clinical and experimental sepsis conditions, the accumulation of LDs and the association of proteins involved in the inflammatory response to LDs have been found in immune cells [40]. Therefore, the HLC added in hydrogel systems
Novel injectable hydrogels were prepared from various MWs of pullulan, HLC and NaIO4, and the relationships between chemical structures, mechanical properties, morphological characteristics, and the degradation (in vivo and in vitro) of the pullulan and HLC/pullulan hydrogels were investigated in this study. The pullulan hydrogels were formed due to NaIO4 oxidizing the hydroxyl groups in pullulan molecules into aldehyde groups, whereas the amide bond link formed between the amino group of HLC and the aldehyde groups in aqueous media after crosslinking with NaIO4. The acute toxicity experiments in vivo showed that the HLC/pullulan hydrogels were safe for use in facial plastic surgeries. The intention of fibroblasts on hydrogels was attach significantly after 3 days. Compared with pullulan hydrogels, the HLC/pullulan hydrogels could be improved the fibroblasts attach more obveriously and inhibited the cell death. These results contributed to the establishment of an approach to optimize hydrogel composition based on the interaction between pullulan of different MW. We aimed to improve the properties of polysaccharide-based hydrogels compared with other sources using HLC with specific binding of polysaccharides using a cross-linking agent to improve the HLC/polysaccharide hydrogel properties for use in tissue engineering. Ultimately, the good mechanical strength, controllable degradation time and tissue compatibility demonstrated by the HLC/pullulan hydrogels may present advantages for in vivo applications. Acknowledgments This study was financially supported by the National Natural Science Foundation of China (21476182); the National High Technology Research and Development Program of China (2014AA02108); the Scientific Research Program of Shannixi Provincial Department of Education, China (14JS102); and the Scientific and Technologic Research Program of Shaanxi Province Academy of Sciences, China (No. 2014 k-01). References [1] X.C. Yang, Y.L. Niu, N.N. Zhao, C. Mao, F.J. Xu, A biocleavable pullulan-based vector via ATRP for liver cell-targeting gene delivery, Biomaterials 35 (2014) 3873–3884. [2] G. Fundueanu, M. Constantin, I. Oanea, V. Harabagiu, P. Ascenzi, B.C. Simionescu, Entrapment and release of drugs by a strict "on-off" mechanism in pullulan microspheres with pendant thermosensitive groups, Biomaterials 31 (2010) 9544–9553. [3] M.R. Rekha, K. Pal, P. Bala, M. Shetty, I. Mittra, G.S. Bhuvaneshwar, C.P. Sharma, Pullulan-histone antibody nanoconjugates for the removal of chromatin fragments from systemic circulation, Biomaterials 34 (2013) 6328–6338. [4] M.R. Rekha, C.P. Sharma, Blood compatibility and in vitro transfection studies on cationically modified pullulan for liver cell targeted gene delivery, Biomaterials 30 (2009) 6655–6664. [5] K.R. Sugumaran, R.V. Sindhu, S. Sukanya, N. Aiswarya, V. Ponnusami, Statistical studies on high molecular weight pullulan production in solid state fermentation using jack fruit seed, Carbohydrate Polymers 98 (2013) 854–860. [6] H. Bender, J. Lechmann, K. Wallenfells, Pullulan in extracellular glucans von Pullularia pullulans, Biochemica Biophysica Acta 36 (1959) 309–316. [7] T. Kimoto, T. Shibuya, S. Shiobara, Safety studies of a novel starch, pullulan: Chronic toxicity in rats and bacterial mutagenecity, Food and Chemical Toxicology 35 (1997) 323–329. [8] K.I. Shingel, Current knowledge on biosynthesis, biological activity, and chemical modification of the exopolysaccharide, pullulan, Carbohydrate Research 339 (2004) 447–460. [9] Y. Tsijisaka, M. Mitsushashi, Pullulan in industrial gum: Polysaccharides and their derivatives, In: Whistler R, BeMiller JN, editors. San Diego: Academic; 1993. 447–460. [10] V.D. Prajapati, G.K. Jani, S.M. Khanda, Pullulan: An exopolysaccharide and its various applications, Carbohydrate Polymers 95 (2013) 540–549. [11] K. Cheek, Determination of polymer–polymer miscibility by viscometry, European Polymer Journal 26 (1990) 423. [12] Z. Sun, W. Wang, Z. Feng, Criterion of polymer–polymer miscibility determined by viscometry, European Polymer Journal 51 (1992) 1259–1261. [13] H. Yim, S.G. Yang, Y.S. Jeon, et al., The performance of gadolinium diethylene triamine pentaacetate-pullulan hepatocyte-specific T1 contrast agent for MRI, Biomaterials 32 (2011) 5187–5194.
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