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The functional response of alginate-gelatin-nanocrystalline cellulose injectable hydrogels toward delivery of cells and bioactive molecules
Acta Biomaterialia 36 (2016) 143–151
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Short communication
The functional response of alginate-gelatin-nanocrystalline cellulose injectable hydrogels toward delivery of cells and bioactive molecules K. Wang, K.C. Nune, R.D.K. Misra ⇑ Biomaterials and Macromolecular Science and Engineering Laboratory, Department of Metallurgical, Materials and Biomedical Engineering, University of Texas at El Paso, 500 W. University Avenue, El Paso, TX 79968, USA
a r t i c l e
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Article history: Received 1 December 2015 Received in revised form 22 January 2016 Accepted 9 March 2016 Available online 10 March 2016 Keywords: Hybrid hydrogel Alginate Nanocrystalline cellulose Biocompatibility Bone regeneration
a b s t r a c t Hybrid injectable hydrogels comprising of alginate, gelatin, and nanocrystalline cellulose (NCC) were conceived and processed through adaptation of interpenetrated network of alginate-gelatin, ionic crosslinking of alginate, and supramolecular interaction approach. The design of hybrid hydrogels was based on the hypothesis that it provides an environment that is favorable for cell proliferation, exchange of nutrients via porous structure, and are characterized by mechanical properties that closely resemble the native tissue. This aspect is important for the delivery of cells or biomolecules in bone tissue engineering. The hybrid hydrogels exhibited moderate swelling behavior on formation, and the porous structure of hydrogels as imaged via SEM was envisaged to facilitate easy migration of cells and rapid transportation of biomolecules. The hybrid hydrogels exhibited desired mechanical properties and were biocompatible as confirmed though MTT assay of fibroblasts. Interestingly, osteoblasts cultured within hydrogel using bone morphogenetic protein (BMP)-2 demonstrated the capability for encapsulation of cells and induced cell differentiation. The nanocrystalline cellulose significant impacted degradation and interaction between hydrogels and cells. Statement of Significance The study fundamentally explores a hypothesis driven novel hybrid hydrogel that provides an environment for favorable growth and proliferation of cells, exchange of nutrients and mechanical properties that closely match the native tissue. Ó 2016 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction Naturally derived polymers, including proteins and polysaccharides, have been extensively used as biomaterials. For instance, as hydrogels for wound dressing, tissue engineering scaffolds, and drug delivery vehicles [1–5]. Among these materials, alginate is a block copolymer with consecutive and alternately arranged b-Dmannuronic acid and a-L-guluronic acid residues and can be ionically crosslinked with divalent cations (e.g. Ca2+, Zn2+) [1,2]. The unique properties of alginate helps in the synthesis of standalone alginate hydrogel or in combination with other materials. However, because of the hydrophobic nature, alginate is not appropriate for cell adhesion unless incorporated with other hydrophilic materials [2,3]. The cell adherent components including collagen, gelatin, RGD (arginine-glycine-aspartate) or RGD derived peptide ⇑ Corresponding author. E-mail address: [email protected] (R.D.K. Misra). http://dx.doi.org/10.1016/j.actbio.2016.03.016 1742-7061/Ó 2016 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
sequences, and asialoglycoprotein receptors (ASGPR) can be integrated with alginate to enhance biological and physicochemical properties [1,4]. Although bioreagent grade alginate is perceived as biocompatible, a higher mannuronic acid content or impurities in alginate has potential to be immunogenic or induce foreign body reactions upon injection or implantation [2,3]. In recent years, the alginate-based hybrid hydrogels have received significant attention because of superior mechanical properties, chemical properties, and biological activity [5]. Essentially, two or more constituents are chemically bonded via two or more crosslinked networks, such that double or multiple network polymers can be formed to obtain superior mechanical properties. For instance, alginate-polyacrylamide double network hydrogels were characterized with 29 kPa Young’s modulus and over 1700% elongation [6]. Although the mechanical properties were superior to some natural polymers, standalone alginate was characterized by low tensile modulus and elongation [7]. Thus, the
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reinforcement of alginate-based hydrogel is important to render them suitable for load-bearing tissue regeneration applications. The nanocrystalline cellulose (NCC) is derived from plant or bacteria cellulose, and has good biocompatibility and is suitable for tissue integration [8,9]. NCC has been used in composite hydrogels of polyacrylamide, polyvinyl alcohol, and cyclodextrin to enhance mechanical properties [10–13]. NCC is not only a reinforcement material for hydrogels but also facilitates gelation in a supramolecular way through hydrophobic interaction and hydrogen bonds. Marine source NCC introduced oriented growth of myoblast in a radial pattern [14]. In another instance, alginate was combined with NCC to prepare a nanocomposite film for wound dressing [15]. Gelatin, a hydrolyzed product from collagen with high hydrophilicity, enhances cell adhesion and provides functional groups for chemical crosslinking. It can be blended with alginate to form spongy film, scaffolds and hydrogels, which provides a benign environment for cell adhesion and growth together with desirable mechanical properties [16,17]. The combination of alginate, gelatin and NCC is expected to improve mechanical and biological properties of NCC reinforced alginate hydrogel. In the study described here, novel hybrid hydrogels were synthesized involving chemical bonding of gelatin and alginate, ionic crosslinking of alginate with zinc ions, and supramolecular interaction with nanocrystalline cellulose. The interactions led to desired properties appropriate for regeneration of bone tissue, where the components of hydrogel precursors allow ease of injection, biocompatibility, effectiveness in nutrients and cell transportation, and cell adhesion and proliferation. BMP-2, a growth factor used extensively to enhance osteogenesis and osteoinduction enabled cell proliferation and differentiation of osteoblast within the hybrid hydrogels. The studied hybrid injectable hydrogels are potentially attractive as scaffold materials for bone regeneration. 2. Materials and methods 2.1. Materials Alginate (sodium salt from brown algae, 4-12 cP, bioreagent), zinc sulfate heptahydrate, N-hydroxy-succinimide (NHS) and 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) were obtained from Sigma–Aldrich USA. Gelatin (type A, from porcine skin, bioreagent) was obtained from MP Biomedicals USA. 2.2. Preparation of nanocrystalline cellulose NCC was prepared by acid hydrolysis of cellulose from filter paper (Whatman #1). After grinding, the cellulose powder was hydrolyzed, while stirring at 45 °C for 45 min using sulfuric acid (64 wt%) at 17.5 mL/g acid-to-pulp ratio [18]. To stop the hydrolysis reaction, deionized (DI) water (10-times the volume) was added. The suspension was centrifuged at 5000 rpm for 30 min and the acid residue was decanted. After washing the precipitate and dispersing in DI water, the suspension of NCC was dialyzed against DI water using a dialysis tube with 12,000 Da molecular weight cut-off (MWCO). The dialysis was continued until the pH of the NCC suspension was greater than 5, followed by dehydration of suspension in a desiccator. Finally, NCC was sonicated for 10 min using a QSonica 750 W probe sonicator to disperse in DI water. 2.3. Synthesis of hybrid injectable hydrogels NCC was dispersed in 2% alginate aqueous solution via sonication and then mixed with 2% gelatin aqueous solution in a 24-well tissue culture plate at varying volume ratio. 20 mg EDC and 10 mg NHS were added to each well and reacted at room
temperature for 24 h. Next, 1 mL 0.05 M ZnSO4 aqueous solution was added to each well for 30 min to complete ionic crosslinking, and the hydrogels were rinsed by 1X phosphate buffer saline (PBS). The details of ration of alginate, NCC, and gelatin are listed in Table 1. In the synthesis of hydrogel precursors the ratio of gelatin/alginate/EDC/NHS/NCC used was identical to that of groups 1 and 4. The reaction was carried out in a 100 mL round bottom flask and stirred at room temperature for 24 h. Next, the hydrogel precursor was withdrawn via 10 mL medical syringes with18-gauge needle.1 mL hydrogel precursor was injected into a covered plastic mold (15.6 mm diameter, 10 mm height, modified from 24-well plate) and then 0.5 mL 0.05 M ZnSO4 was injected to complete ionic crosslinking.
2.4. Characterization of hydrogel Hybrid hydrogels were immersed in 1X PBS at 37 °C and weighed periodically after removing excessive water from the hydrogel surface though filter paper. Then, hydrogels were frozen at 20 °C and lyophilized at 50 °C for 24 h, and weighed again. Swelling ratio of hydrogels was determined using Eq. (1):
Swelling Ratio ¼
Ws Wd 100% Wd
ð1Þ
where Ws is the weight of the swollen hydrogels and Wd is the weight of the hydrogels after freeze-drying. The mechanical properties of hybrid hydrogels were tested with a Dynamic Mechanical Analyzer (RSA3, TA Instrument). The compression test (compression modulus) was carried out unconfined. Six samples prepared in 24-well plate (18 mm diameter and 5 mm thickness) from each group were tested. Compression tests were carried out at room temperature at a speed of 0.1 mm/min and the compression modulus determined by using the stress value at 10% strain. Hydrogels were characterized using FT-IR spectrometer (JASCO FT/IR-4600). The freeze-dried hydrogel was mixed with KBr in the ratio of 1:100, and the pellet was pressed in 7 mm diameter mold. The wave number range was set between 4000 and 500 cm1 and the resolution was 4 cm1. The crystallinity of NCC and NCC-containing hydrogels was studied by X-ray diffraction (XRD) using a powder diffractometer (D8 Discover, Bruker) operating at 40 kV and 40 mA. The lyophilized NCC-free and NCC-containing hydrogels were stored in a desiccator and placed on a glass slide in the form of powder and exposed to Cu Ka X-ray. The freeze-dried hydrogels were observed by Hitachi S-4600 field emission scanning electron microscope (SEM) operated at 20 kV. Prior to examination, the hydrogels were sputter coated with gold. The average pore size of hydrogels was determined from micrographs using 10 randomly measured pores.
Table 1 The content of alginate, gelatin and nanocrystalline cellulose used for the synthesis of hydrogels, corresponding average pore diameter and compression modulus of hydrogels. Group
Alginate (mL)
Gelatin (mL)
NCC (mg)
Average pores (um)
Compression modulus (kPa)
1 2 3 4 5 6
1 1 0.5 1 1 1
1 0.5 1 1 1 1
0 0 0 5 10 20
35 ± 12 54 ± 9 48 ± 10 36 ± 9 47 ± 10 50 ± 11
50 ± 10 39 ± 17 37 ± 19 75 ± 15 84 ± 22 92 ± 28
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2.5. Cell culture
2.9. Fluorescence imaging and live/dead assay
NIH 3T3 mouse fibroblasts were cultured in alpha-minimum essential medium (a-MEM, Gibco) with 10% fetal bovine serum (FBS, Fisher Scientific, USA) and 1% antibiotic–antimycotic (10,000 unit/mL of penicillin, 10,000 lg/mL of streptomycin, Fisher Scientific, USA) with 4.5 g/L glucose in 75 mL cell culture flask (Corning, tissue cultured treated polystyrene (TCPS)). MC3T3-E1 mouse pre-osteoblasts were cultured in a-MEM with 10% FBS and 1% antibiotic–antimycotic in 75 mL cell culture flask. Both types of cells were cultured in an incubator at 5% CO2 and 37 °C. Culture medium was changed every 72 h. Cells were washed with 10 mL 1X PBS and detached with 1 mL trypsin–EDTA solution (0.25% trypsin, 0.02% EDTA, Fisher). The cell suspension was centrifuged at 2000 rpm for 5 min, and the supernatant discarded. Harvested cell pellet was dispersed by pipetting in 1 mL culture medium and counted using a hemocytometer.
The cell encapsulated hydrogels were rinsed with 1X PBS twice and incubated in 1 mL live-dead assay (Life Technologies, USA) at room temperature for 20 min using standard protocol. Live and dead cells were imaged with a Nikon eclipse Ni-U fluorescence microscope. The samples were sliced from and transferred to a clean glass slide and observed under 488 nm (green emission) exciting light to image viable cells and 570 nm (red emission) exciting light to image apoptotic cells.
2.10. Statistical analysis All quantitative data were expressed as the mean and standard deviation. Six samples were used for compression test, and three samples for degradation and cell viability assay. The data were analyzed using one-way analysis of variance (ANOVA), and a p-value less than 0.05 was considered significant.
2.6. Cytotoxicity Hybrid hydrogels (group 1 and group 4 in Table 1) were washed with 1 mL 1X PBS, followed by sterilizing under 254 nm UV light for 2 h prior to cell seeding. NIH-3T3 fibroblasts (10,000 cells/ cm2) were seeded on each hydrogel sample in a 24-well plate with 1 mL culture medium in each well, and incubated for 2, 4, and 7 days at 37 °C in a CO2 incubator. A minimum of three samples were studied and blank TCPS wells were used as controls. The cell viability was measured using MTT (3-(4,5-dimethylthiazol-2-yl)-2, 5-diphenyltetrazolium bromide) assay. The samples were incubated in 0.1 mL 0.5 mg/mL MTT assay and 1 mL fresh culture medium at 37 °C for 4 h. Next, the unreacted dye was withdrawn and 0.2 mL dimethyl sulfoxide was added to dissolve the intracellular purple formazan product into a colored solution. The solution was transferred to a 96-well plate and optical density (absorbance) of the solution was quantified by photospectrometry at 570 nm with a plate reader (Biotek, ELx800). The cell viability was quantified using Eq. (2):
Cell viability ¼ OD570nm ðSampleÞ=OD570nm ðcontrolÞ
ð2Þ
where OD is the optical density of absorbance.
2.7. Degradation Degradation of hybrid hydrogels was studied in vitro. Briefly, hybrid hydrogels at swelling equilibrium were weighed and placed in 10 mL 1X PBS with 10 unit/mL trypsin at 37 °C for 14 days. The hydrogel (group 1) immersed in 10 mL 1X PBS was used as control sample. The weight of hydrogels was measured periodically, and the final data presented as mean and standard derivation of three samples.
2.8. Cell encapsulation MC3T3-E1 osteoblast were detached and harvested as mentioned in section 2.5. The harvested cell pellets (15,000 cells) were mixed with 1.5 mL hydrogel precursor (group 1 and 4) in a 10 mL medical syringe and equally distributed to 3 wells in a 24well plate. 1 mL 0.05 M ZnSO4 solution was added to each well for ionic crosslinking, allowing 20 min for crosslinking. Finally, hydrogels were transferred to other wells that contained 1 mL medium with 50 ng BMP-2, and incubated for up to 7 days at 37 °C in a CO2 incubator (5% CO2 and 95% air). The culture medium was renewed every 72 h.
3. Results 3.1. Hybrid injectable hydrogels The synthesis of hydrogel was carried out using EDC/NHS chemistry, zinc ion crosslinking of alginate and supramolecular interaction with NCC. A schematic of the reactions involved in the synthesis of hydrogel is presented in Fig. 1. The successful synthesis of hydrogels was confirmed by FT-IR (Fig. 2a). The presence of two adjacent absorption peaks in the region of 1595–1635 cm1 provided a clear evidence that gelatin and alginate were incorporated according to the previous literature [16], and peaks 1631 cm1 in NCC originated from OAH vibration. NCC spectrum exhibited strong absorption peak of CAO bond at 1151 cm1. The spectrum of hybrid hydrogel (group 4) was a combination of two components, where the broad peak from 2900 cm1 to 3600 cm1 illustrated strong OAH bond resulting from the interaction between alginate-gelatin backbone and NCC [21]. The crystallinity of NCC and air dried NCC-containing hydrogel was studied by XRD. In the NCC diffraction pattern, peaks at 2h = 16, 23, 33, 46 and 57° correspond to crystalline [1 1 0], [2 0 0], [1 1 1], [2 2 0] and [3 1 1] planes of cellulose consistent with literature [15,17]. In the XRD pattern of NCC-containing hydrogels (group 4), peaks at 2h = 16, 24 and 30° on a high baseline is an evidence of incorporation of nanocrystalline cellulose in the amorphous phase of alginate and gelatin. However, a few peaks were shifted and there was a decrease in the intensity of specific peaks compared to standalone NCC. A possible explanation is that hydrogen bonds, interactions between NCC, and the hydrogel matrix changed the oriented structure of the nanocrystals. It is also likely that the NCC induced less crystallinity within the hydrogel matrix such that the diffraction pattern was changed. The swelling behavior of injectable hydrogels at physiological pH is presented in Fig. 2c. Upon the formation of gelatin-alginate hydrogels, swelling ratio in the range of 900–1100% was obtained in group 1–6. However, incorporation of NCC (group 4–6) significantly changed the swelling behavior of hydrogels after the commencement of swelling test. The NCC-free hydrogels (groups 1– 3) indicated continuous increase in swelling ratio until 24 h, attaining 1300–1400% swelling ratio at the end of the test, while the change of NCC-containing hydrogels (groups 4–6) was not significant (1000–1100% final swelling ratio). NCC has high hydrophobicity and supramolecular interaction, which renders the hybrid hydrogels less likely to swell and reach equilibrium in a relatively short time.
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Fig. 1. Schematic of reactions and interactions involved in the synthesis of alginate-gelatin-NCC hydrogels. (a) Chemical crosslinking by EDC/NHS, (b) ionic crosslinking of guluronic acid on alginate, and (c) orientation of NCC by hydrogel bond and hydrophobic interaction.
The hybrid hydrogel prepared by injection in the mold are presented in Fig. 3(a) and (b). The alginate-gelatin appeared translucent without NCC (group 1), while the homogeneous dispersion of NCC rendered an opaque whitish appearance of hydrogels because of the dispersion of NCC (group 4). The incorporated NCC in hybrid hydrogels had dimensions in the range of 100–500 nm with an aspect ratio of 2–5 (Fig. 3c). Lyophilized hydrogels comprised of a number of pores as imaged via SEM, while NCC-free hydrogels had a smooth surface, and homogeneous distribution of nanoparticles was observed in NCC-containing hydrogels (Fig. 3d–g). The average dimension of pores in freeze-dried hydrogels was in the range of 35–55 lm (Table 1). Group 1 contained maximum water content (greater than 94% water) at equilibrium and smallest pore diameter among NCC-free hydrogels (Groups 1–3), which rendered a dense structure and enabled faster rate of chemical exchange with outer environment because of capillary effect
[20]. The size of pores is important from the viewpoint of migration of cells and transportation of biomolecules encapsulated in hydrogels [20]. Group 2 with higher content of alginate had higher hydrophobicity, while the mechanical properties were compromised in group 3 because of higher gelatin content. Furthermore, group 1 had highest compression modulus among groups 1–3, approaching 50 kPa. Thus, group 1 was most suitable for tissue regeneration in NCC-free hydrogels. Group 4 had swelling ratio and pore size similar to group 1, and incorporation of NCC increased compression modulus to 75 kPa. With increasing NCC content, there was significant decrease in swelling ratio and increase in pore size in groups 5 and 6 because of hydrophobicity of NCC, and compression modulus reached 100 kPa. Meanwhile, hybrid hydrogels containing high water-content provided an environment for cell growth, proliferation, and retention of water soluble molecules. From the above results, hybrid hydrogels (groups 1 and 4) were selected for further studies with cells.
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cells cultured for 4-days, indicated a significant increase in cell viability, whereas after 7-day culture, the rapid increase in cells was less evident. One possible reason is that the cell attachment on hydrogels took a few days and proliferation of cells was initiated after some degradation of hydrogels occurred because of release of gelatin. When fibroblast confluence on the surface of the hybrid hydrogel was high and a small fraction of gelatin component was degraded (Fig. 4b), the proliferation of cells was relatively slow. The incorporation of NCC within the hydrogel (group 4) inevitably compromised the cell viability of fibroblast compared to NCC-free samples such that the strong hydrophobicity of NCC reduced cell adhesion on the hydrogel surface. Although there was a significant increase in cell density compared to the control sample on day 4 and day 7 of culture, the proliferation of cells on hydrogels was slowed down. However, there was a continuous increase in cell density as evident on day 7, which suggested that NCC-containing hydrogels facilitated the proliferation of cells on a long term. Moreover, the proliferation rate estimated in terms of percentage proliferation of cells/day indicated a difference between NCC-free and NCC-containing hydrogels. The proliferation rate was 17% on NCC-free hydrogels and 8% on NCC-containing hydrogels. From the above results, we envisage that hybrid hydrogels, with or without NCC are essentially biocompatible. Furthermore, the adhesion of cells was homogeneous, as presented in Fig. 4b, indicating that NCC distribution on the hydrogel surface led to oriented growth of fibroblast. The degradation of hydrogels was also studied. In the presence of trypsin, the NCC-free hydrogels experienced 50% weight loss in 7 days. This demonstrates that majority of gelatin component in the hydrogel was susceptible to enzymatic degradation. However, NCC-containing hydrogels significantly improved resistance to biodegradation because of homogeneous dispersion and hydrophobicity of NCC, and only 20% weight loss was observed after 7 days. We propose that NCC-containing hybrid hydrogels appropriately meet the requirements for bone regeneration from the viewpoint of degradation. As regards, the behavior of encapsulated osteoblast in hydrogels after 7 days of culture, it can be said from Fig. 5 that the encapsulated osteoblast in NCC-free hydrogels had a relatively low number, and the evidence for cell proliferation and differentiation was not discernable. On the other hand, NCC-containing hydrogels, osteoblast proliferated and changed morphology as a result of initiation of differentiation, and a few dead cells were observed from Live/Dead assay. This behavior may be related to the gelatin component retained in the hydrogel. Hydrophobic alginate alone may not facilitate cell adhesion and proliferation, thus low proliferation of cells in NCC-free hydrogel was anticipated. With gelatin remaining in the NCC-containing hydrogels, osteoblast may adhere and proliferate. BMP-2 introduced differentiation of cells, and hydrogels had the ability for chemical transportation. Additionally, the reinforcement with NCC improved mechanical properties and created a favorable environment for growth of osteoblast.
Fig. 2. (a) FTIR spectra of hybrid hydrogels and NCC, (b) XRD pattern of NCC and NCC-containing hydrogel (group 4, air dried) and (c) swelling behavior of hybrid hydrogels at physiological pH.
3.2. Cell culture and degradation NIH 3T3 cell culture on group 1 and group 4 hydrogels were selected for evaluating the interaction between hydrogels and outer cells on injection and for determination of cytotoxicity of hybrid hydrogels. The cell viability of NIH 3T3 seeded on NCCfree hydrogel surface did not indicate any significant difference with TCPS control group after 2 days of culture (Fig. 4a). However,
4. Discussion The desirable hydrogels for tissue engineering are expected to provide favorable environment for cell proliferation, enable exchange of chemicals, and exhibit properties that closely resemble the native tissue. The extracellular matrix (ECM) of bone contains a large fraction of collagen, thus integration of collagen or proteins (e.g. gelatin) derived from collagen is critical for the effectiveness of hydrogels in bone regeneration [1,19]. The long-term stability of ionic crosslinked alginate hydrogel is low because of leaching of ions to the outer environment [2]. To resolve this issue, alginate-based hybrid hydrogels can be used in a divalent ion rich
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Fig. 3. The macroscopic view of (a) NCC-free hydrogel, (b) NCC-containing hydrogel. SEM micrographs of (c) NCC, (d) lyophilized NCC-free hydrogels (group 1), and NCCcontaining hydrogels (e) 5 mg NCC (group 4), (f) 10 mg NCC (group 5), and (g) 20 mg NCC (group 6).
environment, and the chemical crosslinking using non-toxic crosslinker EDC with NHS stabilizes the amide linkage against hydrolysis and improves hydrogel life. The hybrid hydrogels may combine the interpenetrated matrix of chemically crosslinked alginate and gelatin, and NCC is expected to distribute homogeneously and form oriented pattern throughout the network, involving hydrogen bonding and hydrophobic interaction [22–24]. Aforementioned interactions and chemical bonds are envisaged to significantly increase the mechanical properties of hydrogels that are most likely to meet the requirements for bone regeneration. Compared to the previous research on hybrid materials of alginate and gelatin that used physical blend or formation of
semi-interpenetrated networks, these type of materials have superior mechanical properties resulting from interpenetrated networks with reinforced nanomaterials and good stability against degradation [16,17]. Moreover, the alginate-gelatin precursor allows easy injection and avoids use of any organic solvent, (e.g. DMSO), hazardous chemicals (e.g. glutaraldehyde) and expensive crosslinking reagent (e.g. genipin) that were used in previous research for insitu gelation [16,17,25,26]. The hybrid hydrogels are suitable for easy injection, biocompatibility and cost-effectiveness. The Young’s modulus (10 kPa) and strength of standalone alginate gel does not match with the native bone ECM, considering that superior mechanical properties are required for bone
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Fig. 4. (a) Cell viability of NIH 3T3 fibroblasts on NCC-containing hydrogels (group 4) and NCC-free hydrogels as studied by MTT assay, (b) optical image of NIH 3T3 cells adhered on the surface of NCC-containing hydrogel, and (c) degradation profile of NCC-free and NCC-containing hydrogels.
regeneration [21]. The hybrid hydrogels had a compression modulus of 100 kPa, which is close to bone ECM (300 kPa) and other double-network synthesized hydrogels (100–500 kPa) [22,23]. Besides the improvement in mechanical properties, it is known that the zwitterion polymer network also provides a favorable environment to facilitate cell growth and proliferation [13]. The in-situ formation of hybrid hydrogels is convenient in a manner similar to gelation of other alginate-based hydrogels, and swelling of hydrogels can be controlled at a moderate level using NCC. It has been proven that the pore dimensions of hydrogels can be controlled by zeta-potential and ratio of components, and small pores may enable faster transportation of necessary bioactive molecules (e.g. BMP-2) and migration of cells within hydrogels [22]. Moreover, the orientation of NCC promotes growth of osteoblast within
hybrid hydrogels and allows osteoblast to form mimetics of native bone tissue [14]. Differentiation of osteoblast may also occur. Previous studies on injectable hydrogels obtained limited mineralization of bone [23], whereas alginate-gelatin hydrogels enable effective incorporation of Ca2+ and PO3 ions for mineralization. 4 All the above factors render an appropriate environment for osteogenic tissue [1]. Solubility considerations led to the use of CaCl2 for crosslinking of alginate-based injectable hydrogels [2]. However, CaCl2 is reported to be toxic if excessively present, and rapid reaction rate render poorly controlled gelation [2]. The use of zinc ions in alginate crosslinking are expected to avoid the disadvantages of CaCl2. Moreover, zinc ions provide antimicrobial properties to hydrogels, which minimizes and prevents possible infection during injection
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Fig. 5. Fluorescence micrographs (from left to right: dead assay, live assay, merged image) of osteoblast encapsulated in NCC-containing hydrogels (group 4) and (upper) NCC-free hybrid hydrogels (lower) after 7 days of culture.
[23]. Although calcium ions play an indispensable role in the mineralization of osteogenic tissue, the calcium-free hybrid injectable hydrogels enable exchange of ions (Ca2+ and PO3 4 ) with outer environment or provide solution to mineralization [24]. Incorporation of hydroxyapatite nanoparticles in hybrid hydrogels can be considered for enhancement of mineralization in future. The alginate-gelatin-NCC hydrogel precursor easily forms gel in-situ and acts as a delivery vehicle for cells. However, hybrid hydrogels cannot be compared with the native cortical or cancellous bone if mechanical properties are considered [22,24]. This limits the application of hydrogels as delivery vehicles and artificial ECM for regeneration of bone defects in a small region, rather than replacement of complete bone. Thus, if we inject the hydrogel into an implanted metallic or ceramic matrix, the matrix provides superior mechanical properties, whereas hydrogels regulate the behavior of encapsulated cells. The combination of hydrogel and biocompatible metallic or ceramic materials provide a solution for regeneration of hard tissue that involves large area of bone regeneration. These hybrid hydrogels are now subject of further studies in ECM protein secretion, gene expression, differentiation and mineralization of encapsulated stem cells, and studies with animal models. 5. Conclusions The synthesis of alginate-gelatin-NCC hybrid hydrogels involved EDC/NHS chemistry, ionic crosslinking of zinc ions and alginate, and supramolecular interaction with NCC, and optimization of constituents. Gelatin enabled effective cell adhesion, while alginate and zinc sulfate allowed controllable ionic crosslinking with minimal toxicity and NCC reinforced the mechanical properties. The hybrid hydrogels were characterized by desirable pore diameter and swelling behavior for cell growth and migration. The use of hydrophobic NCC increased crystallinity and stability against degradation. However, hydrophilicity was comprised if NCC was used excessively. The constituents of the hybrid hydrogels were biocompatible and had the capability to incorporate osteoblast and BMP-2. Homogeneous dispersion of NCC was envisaged to provide guidance to oriented cell growth on the surface or
within hydrogels. The studied hybrid hydrogels are potentially cost-effective for delivery of cells and growth factors for healing of bone defects. Further studies on stem cells differentiation and behavior in vivo are necessary to prove the promising future of the hybrid hydrogels as scaffold materials. Acknowledgement The authors thank the Department of Biosystem Engineering, Auburn University for help with the mechanical test. References [1] J.A. Rowley, G. Madlambayan, D.J. Mooney, Alginate hydrogels as synthetic extracellular matrix materials, Biomaterials 20 (1) (1999) 45–53. [2] K.Y. Lee, D.J. Mooney, Alginate: properties and biomedical applications, Prog. Polym. Sci. 37 (1) (2012) 106–126. [3] C. Azuma, K. Yasuda, Y. Tanabe, H. Taniguro, F. Kanaya, A. Nakayama, Y. Osada, Biodegradation of high-toughness double network hydrogels as potential materials for artificial cartilage, J. Biomed. Mater. Res. A 81 (2) (2007) 373–380. [4] D.J. Overstreet, D. Dutta, S.E. Stabenfeldt, B.L. Vernon, Injectable hydrogels, J. Polym. Sci. B: Polym. Phys. 50 (13) (2012) 881–903. [5] S. Van Vlierberghe, P. Dubruel, E. Schacht, Biopolymer-based hydrogels as scaffolds for tissue engineering applications: a review, Biomacromolecules 12 (5) (2011) 1387–1408. [6] J.Y. Sun, X. Zhao, W.R. Illeperuma, O. Chaudhuri, K.H. Oh, D.J. Mooney, Z. Suo, Highly stretchable and tough hydrogels, Nature 489 (7414) (2012) 133–136. [7] J. Li, W.R. Illeperuma, Z. Suo, J.J. Vlassak, Hybrid hydrogels with extremely high stiffness and toughness, ACS Macro Lett. 3 (6) (2014) 520–523. [8] W.K. Czaja, D.J. Young, M. Kawecki, R.M. Brown, The future prospects of microbial cellulose in biomedical applications, Biomacromolecules 8 (1) (2007) 1–12. [9] W. Bai, J. Holbery, K. Li, A technique for production of nanocrystalline cellulose with a narrow size distribution, Cellulose 16 (3) (2009) 455–465. [10] X. Yang, E. Bakaic, T. Hoare, E.D. Cranston, Injectable polysaccharide hydrogels reinforced with cellulose nanocrystals: morphology, rheology, degradation, and cytotoxicity, Biomacromolecules 14 (12) (2013) 4447–4455. [11] A. Jalal Uddin, J. Araki, Y. Gotoh, Toward ‘‘strong” green nanocomposites: polyvinyl alcohol reinforced with extremely oriented cellulose whiskers, Biomacromolecules 12 (3) (2011) 617–624. [12] L. Goetz, M. Foston, A.P. Mathew, K. Oksman, A.J. Ragauskas, Poly (methyl vinyl ether-co-maleic acid) Polyethylene glycol nanocomposites cross-linked in situ with cellulose nanowhiskers, Biomacromolecules 11 (10) (2010) 2660– 2666. [13] Z. Liu, P. Yao, Versatile injectable supramolecular hydrogels containing drug loaded micelles for delivery of various drugs, Polym. Chem. 5 (3) (2014) 1072– 1081.
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Short communication
The functional response of alginate-gelatin-nanocrystalline cellulose injectable hydrogels toward delivery of cells and bioactive molecules K. Wang, K.C. Nune, R.D.K. Misra ⇑ Biomaterials and Macromolecular Science and Engineering Laboratory, Department of Metallurgical, Materials and Biomedical Engineering, University of Texas at El Paso, 500 W. University Avenue, El Paso, TX 79968, USA
a r t i c l e
i n f o
Article history: Received 1 December 2015 Received in revised form 22 January 2016 Accepted 9 March 2016 Available online 10 March 2016 Keywords: Hybrid hydrogel Alginate Nanocrystalline cellulose Biocompatibility Bone regeneration
a b s t r a c t Hybrid injectable hydrogels comprising of alginate, gelatin, and nanocrystalline cellulose (NCC) were conceived and processed through adaptation of interpenetrated network of alginate-gelatin, ionic crosslinking of alginate, and supramolecular interaction approach. The design of hybrid hydrogels was based on the hypothesis that it provides an environment that is favorable for cell proliferation, exchange of nutrients via porous structure, and are characterized by mechanical properties that closely resemble the native tissue. This aspect is important for the delivery of cells or biomolecules in bone tissue engineering. The hybrid hydrogels exhibited moderate swelling behavior on formation, and the porous structure of hydrogels as imaged via SEM was envisaged to facilitate easy migration of cells and rapid transportation of biomolecules. The hybrid hydrogels exhibited desired mechanical properties and were biocompatible as confirmed though MTT assay of fibroblasts. Interestingly, osteoblasts cultured within hydrogel using bone morphogenetic protein (BMP)-2 demonstrated the capability for encapsulation of cells and induced cell differentiation. The nanocrystalline cellulose significant impacted degradation and interaction between hydrogels and cells. Statement of Significance The study fundamentally explores a hypothesis driven novel hybrid hydrogel that provides an environment for favorable growth and proliferation of cells, exchange of nutrients and mechanical properties that closely match the native tissue. Ó 2016 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
1. Introduction Naturally derived polymers, including proteins and polysaccharides, have been extensively used as biomaterials. For instance, as hydrogels for wound dressing, tissue engineering scaffolds, and drug delivery vehicles [1–5]. Among these materials, alginate is a block copolymer with consecutive and alternately arranged b-Dmannuronic acid and a-L-guluronic acid residues and can be ionically crosslinked with divalent cations (e.g. Ca2+, Zn2+) [1,2]. The unique properties of alginate helps in the synthesis of standalone alginate hydrogel or in combination with other materials. However, because of the hydrophobic nature, alginate is not appropriate for cell adhesion unless incorporated with other hydrophilic materials [2,3]. The cell adherent components including collagen, gelatin, RGD (arginine-glycine-aspartate) or RGD derived peptide ⇑ Corresponding author. E-mail address: [email protected] (R.D.K. Misra). http://dx.doi.org/10.1016/j.actbio.2016.03.016 1742-7061/Ó 2016 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.
sequences, and asialoglycoprotein receptors (ASGPR) can be integrated with alginate to enhance biological and physicochemical properties [1,4]. Although bioreagent grade alginate is perceived as biocompatible, a higher mannuronic acid content or impurities in alginate has potential to be immunogenic or induce foreign body reactions upon injection or implantation [2,3]. In recent years, the alginate-based hybrid hydrogels have received significant attention because of superior mechanical properties, chemical properties, and biological activity [5]. Essentially, two or more constituents are chemically bonded via two or more crosslinked networks, such that double or multiple network polymers can be formed to obtain superior mechanical properties. For instance, alginate-polyacrylamide double network hydrogels were characterized with 29 kPa Young’s modulus and over 1700% elongation [6]. Although the mechanical properties were superior to some natural polymers, standalone alginate was characterized by low tensile modulus and elongation [7]. Thus, the
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reinforcement of alginate-based hydrogel is important to render them suitable for load-bearing tissue regeneration applications. The nanocrystalline cellulose (NCC) is derived from plant or bacteria cellulose, and has good biocompatibility and is suitable for tissue integration [8,9]. NCC has been used in composite hydrogels of polyacrylamide, polyvinyl alcohol, and cyclodextrin to enhance mechanical properties [10–13]. NCC is not only a reinforcement material for hydrogels but also facilitates gelation in a supramolecular way through hydrophobic interaction and hydrogen bonds. Marine source NCC introduced oriented growth of myoblast in a radial pattern [14]. In another instance, alginate was combined with NCC to prepare a nanocomposite film for wound dressing [15]. Gelatin, a hydrolyzed product from collagen with high hydrophilicity, enhances cell adhesion and provides functional groups for chemical crosslinking. It can be blended with alginate to form spongy film, scaffolds and hydrogels, which provides a benign environment for cell adhesion and growth together with desirable mechanical properties [16,17]. The combination of alginate, gelatin and NCC is expected to improve mechanical and biological properties of NCC reinforced alginate hydrogel. In the study described here, novel hybrid hydrogels were synthesized involving chemical bonding of gelatin and alginate, ionic crosslinking of alginate with zinc ions, and supramolecular interaction with nanocrystalline cellulose. The interactions led to desired properties appropriate for regeneration of bone tissue, where the components of hydrogel precursors allow ease of injection, biocompatibility, effectiveness in nutrients and cell transportation, and cell adhesion and proliferation. BMP-2, a growth factor used extensively to enhance osteogenesis and osteoinduction enabled cell proliferation and differentiation of osteoblast within the hybrid hydrogels. The studied hybrid injectable hydrogels are potentially attractive as scaffold materials for bone regeneration. 2. Materials and methods 2.1. Materials Alginate (sodium salt from brown algae, 4-12 cP, bioreagent), zinc sulfate heptahydrate, N-hydroxy-succinimide (NHS) and 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide (EDC) were obtained from Sigma–Aldrich USA. Gelatin (type A, from porcine skin, bioreagent) was obtained from MP Biomedicals USA. 2.2. Preparation of nanocrystalline cellulose NCC was prepared by acid hydrolysis of cellulose from filter paper (Whatman #1). After grinding, the cellulose powder was hydrolyzed, while stirring at 45 °C for 45 min using sulfuric acid (64 wt%) at 17.5 mL/g acid-to-pulp ratio [18]. To stop the hydrolysis reaction, deionized (DI) water (10-times the volume) was added. The suspension was centrifuged at 5000 rpm for 30 min and the acid residue was decanted. After washing the precipitate and dispersing in DI water, the suspension of NCC was dialyzed against DI water using a dialysis tube with 12,000 Da molecular weight cut-off (MWCO). The dialysis was continued until the pH of the NCC suspension was greater than 5, followed by dehydration of suspension in a desiccator. Finally, NCC was sonicated for 10 min using a QSonica 750 W probe sonicator to disperse in DI water. 2.3. Synthesis of hybrid injectable hydrogels NCC was dispersed in 2% alginate aqueous solution via sonication and then mixed with 2% gelatin aqueous solution in a 24-well tissue culture plate at varying volume ratio. 20 mg EDC and 10 mg NHS were added to each well and reacted at room
temperature for 24 h. Next, 1 mL 0.05 M ZnSO4 aqueous solution was added to each well for 30 min to complete ionic crosslinking, and the hydrogels were rinsed by 1X phosphate buffer saline (PBS). The details of ration of alginate, NCC, and gelatin are listed in Table 1. In the synthesis of hydrogel precursors the ratio of gelatin/alginate/EDC/NHS/NCC used was identical to that of groups 1 and 4. The reaction was carried out in a 100 mL round bottom flask and stirred at room temperature for 24 h. Next, the hydrogel precursor was withdrawn via 10 mL medical syringes with18-gauge needle.1 mL hydrogel precursor was injected into a covered plastic mold (15.6 mm diameter, 10 mm height, modified from 24-well plate) and then 0.5 mL 0.05 M ZnSO4 was injected to complete ionic crosslinking.
2.4. Characterization of hydrogel Hybrid hydrogels were immersed in 1X PBS at 37 °C and weighed periodically after removing excessive water from the hydrogel surface though filter paper. Then, hydrogels were frozen at 20 °C and lyophilized at 50 °C for 24 h, and weighed again. Swelling ratio of hydrogels was determined using Eq. (1):
Swelling Ratio ¼
Ws Wd 100% Wd
ð1Þ
where Ws is the weight of the swollen hydrogels and Wd is the weight of the hydrogels after freeze-drying. The mechanical properties of hybrid hydrogels were tested with a Dynamic Mechanical Analyzer (RSA3, TA Instrument). The compression test (compression modulus) was carried out unconfined. Six samples prepared in 24-well plate (18 mm diameter and 5 mm thickness) from each group were tested. Compression tests were carried out at room temperature at a speed of 0.1 mm/min and the compression modulus determined by using the stress value at 10% strain. Hydrogels were characterized using FT-IR spectrometer (JASCO FT/IR-4600). The freeze-dried hydrogel was mixed with KBr in the ratio of 1:100, and the pellet was pressed in 7 mm diameter mold. The wave number range was set between 4000 and 500 cm1 and the resolution was 4 cm1. The crystallinity of NCC and NCC-containing hydrogels was studied by X-ray diffraction (XRD) using a powder diffractometer (D8 Discover, Bruker) operating at 40 kV and 40 mA. The lyophilized NCC-free and NCC-containing hydrogels were stored in a desiccator and placed on a glass slide in the form of powder and exposed to Cu Ka X-ray. The freeze-dried hydrogels were observed by Hitachi S-4600 field emission scanning electron microscope (SEM) operated at 20 kV. Prior to examination, the hydrogels were sputter coated with gold. The average pore size of hydrogels was determined from micrographs using 10 randomly measured pores.
Table 1 The content of alginate, gelatin and nanocrystalline cellulose used for the synthesis of hydrogels, corresponding average pore diameter and compression modulus of hydrogels. Group
Alginate (mL)
Gelatin (mL)
NCC (mg)
Average pores (um)
Compression modulus (kPa)
1 2 3 4 5 6
1 1 0.5 1 1 1
1 0.5 1 1 1 1
0 0 0 5 10 20
35 ± 12 54 ± 9 48 ± 10 36 ± 9 47 ± 10 50 ± 11
50 ± 10 39 ± 17 37 ± 19 75 ± 15 84 ± 22 92 ± 28
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2.5. Cell culture
2.9. Fluorescence imaging and live/dead assay
NIH 3T3 mouse fibroblasts were cultured in alpha-minimum essential medium (a-MEM, Gibco) with 10% fetal bovine serum (FBS, Fisher Scientific, USA) and 1% antibiotic–antimycotic (10,000 unit/mL of penicillin, 10,000 lg/mL of streptomycin, Fisher Scientific, USA) with 4.5 g/L glucose in 75 mL cell culture flask (Corning, tissue cultured treated polystyrene (TCPS)). MC3T3-E1 mouse pre-osteoblasts were cultured in a-MEM with 10% FBS and 1% antibiotic–antimycotic in 75 mL cell culture flask. Both types of cells were cultured in an incubator at 5% CO2 and 37 °C. Culture medium was changed every 72 h. Cells were washed with 10 mL 1X PBS and detached with 1 mL trypsin–EDTA solution (0.25% trypsin, 0.02% EDTA, Fisher). The cell suspension was centrifuged at 2000 rpm for 5 min, and the supernatant discarded. Harvested cell pellet was dispersed by pipetting in 1 mL culture medium and counted using a hemocytometer.
The cell encapsulated hydrogels were rinsed with 1X PBS twice and incubated in 1 mL live-dead assay (Life Technologies, USA) at room temperature for 20 min using standard protocol. Live and dead cells were imaged with a Nikon eclipse Ni-U fluorescence microscope. The samples were sliced from and transferred to a clean glass slide and observed under 488 nm (green emission) exciting light to image viable cells and 570 nm (red emission) exciting light to image apoptotic cells.
2.10. Statistical analysis All quantitative data were expressed as the mean and standard deviation. Six samples were used for compression test, and three samples for degradation and cell viability assay. The data were analyzed using one-way analysis of variance (ANOVA), and a p-value less than 0.05 was considered significant.
2.6. Cytotoxicity Hybrid hydrogels (group 1 and group 4 in Table 1) were washed with 1 mL 1X PBS, followed by sterilizing under 254 nm UV light for 2 h prior to cell seeding. NIH-3T3 fibroblasts (10,000 cells/ cm2) were seeded on each hydrogel sample in a 24-well plate with 1 mL culture medium in each well, and incubated for 2, 4, and 7 days at 37 °C in a CO2 incubator. A minimum of three samples were studied and blank TCPS wells were used as controls. The cell viability was measured using MTT (3-(4,5-dimethylthiazol-2-yl)-2, 5-diphenyltetrazolium bromide) assay. The samples were incubated in 0.1 mL 0.5 mg/mL MTT assay and 1 mL fresh culture medium at 37 °C for 4 h. Next, the unreacted dye was withdrawn and 0.2 mL dimethyl sulfoxide was added to dissolve the intracellular purple formazan product into a colored solution. The solution was transferred to a 96-well plate and optical density (absorbance) of the solution was quantified by photospectrometry at 570 nm with a plate reader (Biotek, ELx800). The cell viability was quantified using Eq. (2):
Cell viability ¼ OD570nm ðSampleÞ=OD570nm ðcontrolÞ
ð2Þ
where OD is the optical density of absorbance.
2.7. Degradation Degradation of hybrid hydrogels was studied in vitro. Briefly, hybrid hydrogels at swelling equilibrium were weighed and placed in 10 mL 1X PBS with 10 unit/mL trypsin at 37 °C for 14 days. The hydrogel (group 1) immersed in 10 mL 1X PBS was used as control sample. The weight of hydrogels was measured periodically, and the final data presented as mean and standard derivation of three samples.
2.8. Cell encapsulation MC3T3-E1 osteoblast were detached and harvested as mentioned in section 2.5. The harvested cell pellets (15,000 cells) were mixed with 1.5 mL hydrogel precursor (group 1 and 4) in a 10 mL medical syringe and equally distributed to 3 wells in a 24well plate. 1 mL 0.05 M ZnSO4 solution was added to each well for ionic crosslinking, allowing 20 min for crosslinking. Finally, hydrogels were transferred to other wells that contained 1 mL medium with 50 ng BMP-2, and incubated for up to 7 days at 37 °C in a CO2 incubator (5% CO2 and 95% air). The culture medium was renewed every 72 h.
3. Results 3.1. Hybrid injectable hydrogels The synthesis of hydrogel was carried out using EDC/NHS chemistry, zinc ion crosslinking of alginate and supramolecular interaction with NCC. A schematic of the reactions involved in the synthesis of hydrogel is presented in Fig. 1. The successful synthesis of hydrogels was confirmed by FT-IR (Fig. 2a). The presence of two adjacent absorption peaks in the region of 1595–1635 cm1 provided a clear evidence that gelatin and alginate were incorporated according to the previous literature [16], and peaks 1631 cm1 in NCC originated from OAH vibration. NCC spectrum exhibited strong absorption peak of CAO bond at 1151 cm1. The spectrum of hybrid hydrogel (group 4) was a combination of two components, where the broad peak from 2900 cm1 to 3600 cm1 illustrated strong OAH bond resulting from the interaction between alginate-gelatin backbone and NCC [21]. The crystallinity of NCC and air dried NCC-containing hydrogel was studied by XRD. In the NCC diffraction pattern, peaks at 2h = 16, 23, 33, 46 and 57° correspond to crystalline [1 1 0], [2 0 0], [1 1 1], [2 2 0] and [3 1 1] planes of cellulose consistent with literature [15,17]. In the XRD pattern of NCC-containing hydrogels (group 4), peaks at 2h = 16, 24 and 30° on a high baseline is an evidence of incorporation of nanocrystalline cellulose in the amorphous phase of alginate and gelatin. However, a few peaks were shifted and there was a decrease in the intensity of specific peaks compared to standalone NCC. A possible explanation is that hydrogen bonds, interactions between NCC, and the hydrogel matrix changed the oriented structure of the nanocrystals. It is also likely that the NCC induced less crystallinity within the hydrogel matrix such that the diffraction pattern was changed. The swelling behavior of injectable hydrogels at physiological pH is presented in Fig. 2c. Upon the formation of gelatin-alginate hydrogels, swelling ratio in the range of 900–1100% was obtained in group 1–6. However, incorporation of NCC (group 4–6) significantly changed the swelling behavior of hydrogels after the commencement of swelling test. The NCC-free hydrogels (groups 1– 3) indicated continuous increase in swelling ratio until 24 h, attaining 1300–1400% swelling ratio at the end of the test, while the change of NCC-containing hydrogels (groups 4–6) was not significant (1000–1100% final swelling ratio). NCC has high hydrophobicity and supramolecular interaction, which renders the hybrid hydrogels less likely to swell and reach equilibrium in a relatively short time.
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Fig. 1. Schematic of reactions and interactions involved in the synthesis of alginate-gelatin-NCC hydrogels. (a) Chemical crosslinking by EDC/NHS, (b) ionic crosslinking of guluronic acid on alginate, and (c) orientation of NCC by hydrogel bond and hydrophobic interaction.
The hybrid hydrogel prepared by injection in the mold are presented in Fig. 3(a) and (b). The alginate-gelatin appeared translucent without NCC (group 1), while the homogeneous dispersion of NCC rendered an opaque whitish appearance of hydrogels because of the dispersion of NCC (group 4). The incorporated NCC in hybrid hydrogels had dimensions in the range of 100–500 nm with an aspect ratio of 2–5 (Fig. 3c). Lyophilized hydrogels comprised of a number of pores as imaged via SEM, while NCC-free hydrogels had a smooth surface, and homogeneous distribution of nanoparticles was observed in NCC-containing hydrogels (Fig. 3d–g). The average dimension of pores in freeze-dried hydrogels was in the range of 35–55 lm (Table 1). Group 1 contained maximum water content (greater than 94% water) at equilibrium and smallest pore diameter among NCC-free hydrogels (Groups 1–3), which rendered a dense structure and enabled faster rate of chemical exchange with outer environment because of capillary effect
[20]. The size of pores is important from the viewpoint of migration of cells and transportation of biomolecules encapsulated in hydrogels [20]. Group 2 with higher content of alginate had higher hydrophobicity, while the mechanical properties were compromised in group 3 because of higher gelatin content. Furthermore, group 1 had highest compression modulus among groups 1–3, approaching 50 kPa. Thus, group 1 was most suitable for tissue regeneration in NCC-free hydrogels. Group 4 had swelling ratio and pore size similar to group 1, and incorporation of NCC increased compression modulus to 75 kPa. With increasing NCC content, there was significant decrease in swelling ratio and increase in pore size in groups 5 and 6 because of hydrophobicity of NCC, and compression modulus reached 100 kPa. Meanwhile, hybrid hydrogels containing high water-content provided an environment for cell growth, proliferation, and retention of water soluble molecules. From the above results, hybrid hydrogels (groups 1 and 4) were selected for further studies with cells.
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cells cultured for 4-days, indicated a significant increase in cell viability, whereas after 7-day culture, the rapid increase in cells was less evident. One possible reason is that the cell attachment on hydrogels took a few days and proliferation of cells was initiated after some degradation of hydrogels occurred because of release of gelatin. When fibroblast confluence on the surface of the hybrid hydrogel was high and a small fraction of gelatin component was degraded (Fig. 4b), the proliferation of cells was relatively slow. The incorporation of NCC within the hydrogel (group 4) inevitably compromised the cell viability of fibroblast compared to NCC-free samples such that the strong hydrophobicity of NCC reduced cell adhesion on the hydrogel surface. Although there was a significant increase in cell density compared to the control sample on day 4 and day 7 of culture, the proliferation of cells on hydrogels was slowed down. However, there was a continuous increase in cell density as evident on day 7, which suggested that NCC-containing hydrogels facilitated the proliferation of cells on a long term. Moreover, the proliferation rate estimated in terms of percentage proliferation of cells/day indicated a difference between NCC-free and NCC-containing hydrogels. The proliferation rate was 17% on NCC-free hydrogels and 8% on NCC-containing hydrogels. From the above results, we envisage that hybrid hydrogels, with or without NCC are essentially biocompatible. Furthermore, the adhesion of cells was homogeneous, as presented in Fig. 4b, indicating that NCC distribution on the hydrogel surface led to oriented growth of fibroblast. The degradation of hydrogels was also studied. In the presence of trypsin, the NCC-free hydrogels experienced 50% weight loss in 7 days. This demonstrates that majority of gelatin component in the hydrogel was susceptible to enzymatic degradation. However, NCC-containing hydrogels significantly improved resistance to biodegradation because of homogeneous dispersion and hydrophobicity of NCC, and only 20% weight loss was observed after 7 days. We propose that NCC-containing hybrid hydrogels appropriately meet the requirements for bone regeneration from the viewpoint of degradation. As regards, the behavior of encapsulated osteoblast in hydrogels after 7 days of culture, it can be said from Fig. 5 that the encapsulated osteoblast in NCC-free hydrogels had a relatively low number, and the evidence for cell proliferation and differentiation was not discernable. On the other hand, NCC-containing hydrogels, osteoblast proliferated and changed morphology as a result of initiation of differentiation, and a few dead cells were observed from Live/Dead assay. This behavior may be related to the gelatin component retained in the hydrogel. Hydrophobic alginate alone may not facilitate cell adhesion and proliferation, thus low proliferation of cells in NCC-free hydrogel was anticipated. With gelatin remaining in the NCC-containing hydrogels, osteoblast may adhere and proliferate. BMP-2 introduced differentiation of cells, and hydrogels had the ability for chemical transportation. Additionally, the reinforcement with NCC improved mechanical properties and created a favorable environment for growth of osteoblast.
Fig. 2. (a) FTIR spectra of hybrid hydrogels and NCC, (b) XRD pattern of NCC and NCC-containing hydrogel (group 4, air dried) and (c) swelling behavior of hybrid hydrogels at physiological pH.
3.2. Cell culture and degradation NIH 3T3 cell culture on group 1 and group 4 hydrogels were selected for evaluating the interaction between hydrogels and outer cells on injection and for determination of cytotoxicity of hybrid hydrogels. The cell viability of NIH 3T3 seeded on NCCfree hydrogel surface did not indicate any significant difference with TCPS control group after 2 days of culture (Fig. 4a). However,
4. Discussion The desirable hydrogels for tissue engineering are expected to provide favorable environment for cell proliferation, enable exchange of chemicals, and exhibit properties that closely resemble the native tissue. The extracellular matrix (ECM) of bone contains a large fraction of collagen, thus integration of collagen or proteins (e.g. gelatin) derived from collagen is critical for the effectiveness of hydrogels in bone regeneration [1,19]. The long-term stability of ionic crosslinked alginate hydrogel is low because of leaching of ions to the outer environment [2]. To resolve this issue, alginate-based hybrid hydrogels can be used in a divalent ion rich
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Fig. 3. The macroscopic view of (a) NCC-free hydrogel, (b) NCC-containing hydrogel. SEM micrographs of (c) NCC, (d) lyophilized NCC-free hydrogels (group 1), and NCCcontaining hydrogels (e) 5 mg NCC (group 4), (f) 10 mg NCC (group 5), and (g) 20 mg NCC (group 6).
environment, and the chemical crosslinking using non-toxic crosslinker EDC with NHS stabilizes the amide linkage against hydrolysis and improves hydrogel life. The hybrid hydrogels may combine the interpenetrated matrix of chemically crosslinked alginate and gelatin, and NCC is expected to distribute homogeneously and form oriented pattern throughout the network, involving hydrogen bonding and hydrophobic interaction [22–24]. Aforementioned interactions and chemical bonds are envisaged to significantly increase the mechanical properties of hydrogels that are most likely to meet the requirements for bone regeneration. Compared to the previous research on hybrid materials of alginate and gelatin that used physical blend or formation of
semi-interpenetrated networks, these type of materials have superior mechanical properties resulting from interpenetrated networks with reinforced nanomaterials and good stability against degradation [16,17]. Moreover, the alginate-gelatin precursor allows easy injection and avoids use of any organic solvent, (e.g. DMSO), hazardous chemicals (e.g. glutaraldehyde) and expensive crosslinking reagent (e.g. genipin) that were used in previous research for insitu gelation [16,17,25,26]. The hybrid hydrogels are suitable for easy injection, biocompatibility and cost-effectiveness. The Young’s modulus (10 kPa) and strength of standalone alginate gel does not match with the native bone ECM, considering that superior mechanical properties are required for bone
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Fig. 4. (a) Cell viability of NIH 3T3 fibroblasts on NCC-containing hydrogels (group 4) and NCC-free hydrogels as studied by MTT assay, (b) optical image of NIH 3T3 cells adhered on the surface of NCC-containing hydrogel, and (c) degradation profile of NCC-free and NCC-containing hydrogels.
regeneration [21]. The hybrid hydrogels had a compression modulus of 100 kPa, which is close to bone ECM (300 kPa) and other double-network synthesized hydrogels (100–500 kPa) [22,23]. Besides the improvement in mechanical properties, it is known that the zwitterion polymer network also provides a favorable environment to facilitate cell growth and proliferation [13]. The in-situ formation of hybrid hydrogels is convenient in a manner similar to gelation of other alginate-based hydrogels, and swelling of hydrogels can be controlled at a moderate level using NCC. It has been proven that the pore dimensions of hydrogels can be controlled by zeta-potential and ratio of components, and small pores may enable faster transportation of necessary bioactive molecules (e.g. BMP-2) and migration of cells within hydrogels [22]. Moreover, the orientation of NCC promotes growth of osteoblast within
hybrid hydrogels and allows osteoblast to form mimetics of native bone tissue [14]. Differentiation of osteoblast may also occur. Previous studies on injectable hydrogels obtained limited mineralization of bone [23], whereas alginate-gelatin hydrogels enable effective incorporation of Ca2+ and PO3 ions for mineralization. 4 All the above factors render an appropriate environment for osteogenic tissue [1]. Solubility considerations led to the use of CaCl2 for crosslinking of alginate-based injectable hydrogels [2]. However, CaCl2 is reported to be toxic if excessively present, and rapid reaction rate render poorly controlled gelation [2]. The use of zinc ions in alginate crosslinking are expected to avoid the disadvantages of CaCl2. Moreover, zinc ions provide antimicrobial properties to hydrogels, which minimizes and prevents possible infection during injection
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Fig. 5. Fluorescence micrographs (from left to right: dead assay, live assay, merged image) of osteoblast encapsulated in NCC-containing hydrogels (group 4) and (upper) NCC-free hybrid hydrogels (lower) after 7 days of culture.
[23]. Although calcium ions play an indispensable role in the mineralization of osteogenic tissue, the calcium-free hybrid injectable hydrogels enable exchange of ions (Ca2+ and PO3 4 ) with outer environment or provide solution to mineralization [24]. Incorporation of hydroxyapatite nanoparticles in hybrid hydrogels can be considered for enhancement of mineralization in future. The alginate-gelatin-NCC hydrogel precursor easily forms gel in-situ and acts as a delivery vehicle for cells. However, hybrid hydrogels cannot be compared with the native cortical or cancellous bone if mechanical properties are considered [22,24]. This limits the application of hydrogels as delivery vehicles and artificial ECM for regeneration of bone defects in a small region, rather than replacement of complete bone. Thus, if we inject the hydrogel into an implanted metallic or ceramic matrix, the matrix provides superior mechanical properties, whereas hydrogels regulate the behavior of encapsulated cells. The combination of hydrogel and biocompatible metallic or ceramic materials provide a solution for regeneration of hard tissue that involves large area of bone regeneration. These hybrid hydrogels are now subject of further studies in ECM protein secretion, gene expression, differentiation and mineralization of encapsulated stem cells, and studies with animal models. 5. Conclusions The synthesis of alginate-gelatin-NCC hybrid hydrogels involved EDC/NHS chemistry, ionic crosslinking of zinc ions and alginate, and supramolecular interaction with NCC, and optimization of constituents. Gelatin enabled effective cell adhesion, while alginate and zinc sulfate allowed controllable ionic crosslinking with minimal toxicity and NCC reinforced the mechanical properties. The hybrid hydrogels were characterized by desirable pore diameter and swelling behavior for cell growth and migration. The use of hydrophobic NCC increased crystallinity and stability against degradation. However, hydrophilicity was comprised if NCC was used excessively. The constituents of the hybrid hydrogels were biocompatible and had the capability to incorporate osteoblast and BMP-2. Homogeneous dispersion of NCC was envisaged to provide guidance to oriented cell growth on the surface or
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