Multifunctional Interpenetrating Polymer Network Hydrogels Based on ...

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Multifunctional Interpenetrating Polymer Network Hydrogels Based on Methacrylated Alginate for the Delivery of Small Molecule Drugs and Sustained Release of Protein Jun Zhao,† Xin Zhao,† Baolin Guo,*,† and Peter X. Ma*,†,‡,§,∥,⊥ †

Center for Biomedical Engineering and Regenerative Medicine, Frontier Institute of Science and Technology, Xi’an Jiaotong University, Xi’an 710049, China ‡ Department of Biomedical Engineering, §Department of Biologic and Materials Sciences, ∥Macromolecular Science and Engineering Center, and ⊥Department of Materials Science and Engineering, University of Michigan, Ann Arbor, Michigan 48109, United States S Supporting Information *

ABSTRACT: Multifunctional injectable thermo-/pH-responsive hydrogels as release systems for the oral delivery of small molecule drugs and the local delivery of protein are presented. The injectable interpenetrating polymer network (IPN) hydrogels based on poly(ethylene glycol) methacrylate, N-isopropylacrylamide, and methacrylated alginate were prepared by using ammonium persulfate (APS) and N,N,N′,N′-tetramethylethylenediamine (TEMED) as a redox initiator system at body temperature, and the obtained hydrogels overcame the instability of calcium cross-linked alginate hydrogels under physiological conditions. The hydrogels showed good mechanical strength by rheometer and exhibited temperature and pH sensitivity by a swelling test. Diclofenac sodium (DCS) as a model for small molecule water-soluble anti-inflammatory drugs and bovine serum albumin (BSA) as a model for protein drugs were encapsulated in situ in the hydrogel. The DCS and BSA release results indicated that these hydrogels, as carriers, have great potential for use in the oral delivery of small molecule drugs and for longterm localized protein release. Furthermore, the cytotoxicity of these hydrogels was studied via live/dead viability and alamarBlue assays using adipose tissue-derived mesenchymal stem cells.

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INTRODUCTION

bioactive macromolecules are conveniently mixed with the solution, and the mixed systems transfer into hydrogels once injected. Hydrogels then become an injectable site-specific drug depot for sustained release, which can reduce administration time and side effects.19 Furthermore, the pH responsiveness and/or thermoresponsiveness of the injectable materials can further tailor the release behavior of the injectable hydrogels.20−22 Because of their good biocompatibility and controllable biodegradability, natural materials have been widely used for injectable hydrogels for drug delivery.23,24 Sodium alginate, a naturally occurring nontoxic polysaccharide derived from brown seaweed, is one of the most versatile natural polymers

Drug delivery systems (DDS) and relevant release materials have attracted intense interest during recent decades.1−5 They can enhance therapeutic efficacy by providing a solubilizing environment where drugs, proteins, or other bioactive molecules can be entrapped and protected from degradation and from the harsh environment of the stomach.6,7 Because of their excellent biocompatibility and structural similarity to the highly hydrated structures of the human body and because of their tunable properties for controlling the release profile of the entrapped molecules, hydrogels are one of the most frequently used biomaterials in various drug delivery systems.8−12 In particular, intelligent injectable hydrogels have gained more and more attention in DDS and tissue engineering applications recently.13−18 Because the polymeric systems are flowable aqueous solutions before injection, the drugs, proteins, or other © 2014 American Chemical Society

Received: April 29, 2014 Revised: July 27, 2014 Published: August 7, 2014 3246

dx.doi.org/10.1021/bm5006257 | Biomacromolecules 2014, 15, 3246−3252

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Figure 1. Synthesis scheme of the injectable hydrogels and drug encapsulation.

used in the medical and cosmetics fields.25,26 Alginate has been fabricated into large beads, microbeads, block gels, fibers, and in situ gelling systems.27,28 Because of the high number of carboxyl groups along the main chain of alginate, alginate exhibits pH sensitivity and gel formation ability in the presence of divalent cations such as Ca2+ or multivalent cations such as Fe3+ in aqueous medium.29−32 For example, when alginate and poly(Nisopropylacrylamide) solutions were mixed together and their droplets were cross-linked in CaCl2 solution, a semi-interpenetrating polymer network hydrogel beads was formed. Shi et al. showed that these beads had potential for use as a pH and temperature sustainable delivery system of indomethacin.33 In another example, microbeads based on alginate blended N,Ocarboxymethyl chitosan were prepared in Ca2+ solution, and the beads were used as a pH-sensitive system for the delivery of protein.34 Although ionically cross-linked alginate hydrogels are widely used as delivery carriers due to the mild gelling process, this process usually leads to poor control over the swelling and mechanical properties of the gels because of the rapid gelation in aqueous solution. Moreover, ionically cross-linked alginate hydrogels have exhibited limited long-term stability because of the loss of divalent cations under physiological conditions. Therefore, some photo-cross-linked alginate hydrogels have been reported.35−37 However, the photopolymerization process cannot always be carried out uniformly, especially when the penetration depth is quite limited and light distribution is inhomogeneous. Furthermore, the photoinitiator used in the systems is sometimes toxic and affects the drug quantification. A water-soluble redox initiator system composed of ammonium persulfate (APS) and N,N,N′,N′-tetramethylethylenediamine (TEMED) can overcome the limitation of photopolymerization to form injectable hydrogels.19,38−40 The aim of this work is to develop intelligent injectable hydrogels based on methacrylated alginate using APS/TEMED as the initiator system for the oral delivery of small molecule drugs and for localized, sustained protein release. The injectable hydrogels were synthesized by a copolymerization of methacrylated alginate, poly(ethylene glycol) methacrylate, and N-isopropylacrylamide under physiological conditions. Their chemical structures, rheological properties, pH/temperature-sensitive swelling properties, morphologies, cytotoxicities,

and release properties of a small molecule drug and a protein were evaluated. The results demonstrate that the injectable hydrogels are noncytotoxic and are promising candidates for the oral delivery of drugs and localized long-term protein release.

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EXPERIMENTAL SECTION

Materials. Alginic acid sodium salt (ALG), methacrylic anhydride (MA), N,N′-methylenebis(acrylamide) (BIS), N-isopropylacrylamide (NIPAm), and poly(ethylene glycol) methacrylate (PEGMA) with a molecular weight of 526 were obtained from Sigma. Ammonium persulfate (APS) and N,N,N′,N′-tetramethylethylenediamine (TEMED) from Sigma were used as initiator and catalyst, respectively. Bovine serum albumin (BSA) and diclofenac sodium (DCS), used as two kinds of model drugs, were from Sigma. Ethanol, NaOH, HCl, phosphate buffer saline (PBS), and deionized water are all analytical grade. Preparation of Methacrylated Alginate (ALGMA) Macromer. ALGMA was prepared through esterification of hydroxyl groups of alginate according to protocols previously described.41 Briefly, 1 g of alginate was dissolved in 100 mL of deionized water, and the pH of the solution was adjusted to 8 using a 5 mol/L NaOH solution. Twenty grams of methacrylic anhydride was then added slowly at 0−4 °C, and the pH was adjusted to 8 using a 5 mol/L NaOH solution every 30 min. The solution was then allowed to react for 24 h at 0−4 °C. The modified alginate was dialyzed in sterile water for 48 h to remove excess methacrylic anhydride, and stirring of the product was continued as a 5 mol/L NaOH solution was added dropwise until clarification was achieved. The final macromer was precipitated in cold ethanol and dried in a vacuum oven. Preparation of PEGMA-NIPAm Hydrogel and PEGMANIPAm-ALGMA Hydrogel. PEGMA-NIPAm (PN) hydrogel was synthesized by copolymerizing PEGMA and NIPAm using BIS as a cross-linker and APS and TEMED as an initiator system. Briefly, 50 mg of PEGMA and 50 mg of NIPAm were dissolved in PBS in a tube. BIS (3% w/w), APS (3% w/w), and 1.0 μL/g TEMED were added to the above solution under agitation. The mixture was put into a constant temperature oven at 37 °C. Hydrogels were obtained in a few minutes. The final product was called P50N50, and other ratios of PEGMA to NIPAm, such as 0:100, 30:70, 40:60, 60:40, 70:30, and 100:0, were also prepared using the same method, and they were named P0N100, P30N70, P40N60, P60N40, P70N30, and P100N0, respectively, according to their composition. 3247


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Petri dishes, and the plates were washed extensively with DPBS to remove residual nonadherent red blood cells. The medium was changed every 3 days. To expand the cells, they were detached using 0.25% trypsin and passaged at 1:5 dilutions. Passage 4 (P4) cells were used for the viability and proliferation tests in this work. In Vitro Cell Compatibility of Hydrogel. To evaluate the cytotoxicity of the hydrogel, a direct contact test between the hydrogel and ADMSCs was used to measure the potentially toxic effect of the hydrogels on cells.42−44 The monomer solution was sterilized by filtration (0.22 μm filter, Millipore) and poured into a sterile Petri dish to form a hydrogel film with a 1.5 mm thickness at 37 °C in the incubator. The hydrogel film was then cut into 5 mm diameter disks and washed twice with cell culture medium for 30 min each at 37 °C under agitation. The complete growth medium was DMEM supplemented with 10% fetal bovine serum, 1.0 × 105 U/L penicillin, and 100 mg/L streptomycin. ADMSCs were seeded in a 96-well plate (Costar) at a density of 6000 cells/cm2. When cells had adhered to the plate (∼4 to 5 h), the hydrogel disks were placed into the wells. The culture medium was changed every 2 days. The ADMSCs proliferation and viability on hydrogel were evaluated by alamarBlue and live/dead viability/cytotoxicity assays, respectively. After being cultured for 1, 4, and 7 days, the hydrogel disks were removed, and 10 μL of the alamarBlue reagent was added into each well. The plate was incubated for 4 h at 37 °C in a humidified incubator containing 5% CO2. One hundred microliters of the medium in each well was removed into a 96-well black plate (Costar). Fluorescence was read using 530 nm as the excitation wavelength and 600 nm as the emission wavelength using a microplate reader (Molecular Devices). Cells seeded on a tissue culture polystyrene (TCPS) plate served as the positive control group. Tests were repeated four times for each group. The hydrogel and control groups were stained using a live/dead viability/cytotoxicity kit (Molecular Probes) for 45 min at 37 °C after being washed three times with PBS. Cell adhesion and proliferation were observed under an inverted fluorescence microscope (IX53, Olympus). Statistical Analysis. All experimental data from the studies were analyzed using Student’s t test. P < 0.05 was considered to be statistically significant. Results are expressed as the mean ± SD.

The PEGMA-NIPAm-ALGMA (PNA) hydrogel was synthesized in a similar manner, and the synthesis scheme is shown in Figure 1. ALGMA (10, 20, or 30 wt %) was added to the PEGMA and NIPAm mixture, and the synthesis procedure was the same as that for the PN hydrogels. These hydrogels were named P50N50A10, P50N50A20, and P50N50A30, respectively. Fourier Transform Infrared (FT-IR) Spectra Characterization. FT-IR spectra of P50N50A20, P50N50, ALGMA, NIPAm, and PEGMA were recorded in the range of 4000−650 cm−1 using a Nicolet 6700 FT-IR spectrometer (Thermo Scientific Instrument). Rheological Characterization. The storage moduli and gelation time of these hydrogels were determined by a TA rheometer (DHR2). The transition behavior of the P50N50 polymer aqueous solution with different amounts of ALGMA was investigated by placing the mixture between 40 mm parallel plates with a 25 mm diameter and a gap of 1 mm. The experiment was performed using a time sweep test at 37 °C with a frequency of 1 Hz and 1% strain. Swelling Properties of Hydrogels. The swelling properties of PNA hydrogels were determined under different conditions. The four conditions included a pH 2.1 solution and a pH 7.4 solution either at body temperature (37 °C) or at room temperature (25 °C). The confirmed weight of hydrogel was placed into a 20 mL bottle under different conditions. Over a specified period of time, the hydrogels were taken out, and the weight was determined after sweeping excess water on the surface of the hydrogel. The swelling ratio (SR) of hydrogel was calculated as follows: SR = (Wt − Wo)/Wo × 100%, where Wt and Wo are the weight of the swollen hydrogel and the initial weight, respectively. For each sample, the experiments were repeated three times, and the final results were calculated by averaging the replicates. Morphology of Hydrogel. The morphology of PNA hydrogels after swelling in different pH (7.4 and 2.1) solutions at different temperatures (25 and 37 °C) was observed by scanning electron microscopy (FE-SEM, SU-8000, Hitachi, Japan). The hydrogels, after reaching swelling equilibrium, were freeze-dried for 48 h, and hydrogel surfaces were then coated with a gold layer before observation under SEM. In Vitro Drug Release Test. DCS and BSA were chosen as model drugs for the controlled release experiment. DCS (5.5 mg/mL) and BSA (40 mg/mL) were physically mixed with the monomer solution before forming the gel and stirred for 30 min. After the drugs were totally dissolved, APS/TEMED was added into the mixture, and the solution was placed in a 37 °C oven for 15 min to obtain the drugloaded hydrogels. The drug-loaded hydrogels were then placed into a flask containing 100 mL of PBS buffer solution (pH 7.4). At the same time, the drug-free hydrogels were placed under the same conditions as those of the control group. The flasks were put into a constant temperature shaking incubator at 37 °C at 100 rpm. At each predetermined time interval, 1 mL of the release media was taken out; meanwhile, 1 mL of fresh PBS was supplemented. The concentration of drugs in solution was quantified via an ultraviolet spectrophotometer (PerkinElmer Lambda 35) using a predetermined calibration. The UV−vis absorption of DCS and BSA was measured at 276 and 278 nm, respectively. Tests were repeated three times for each sample, and the average values are reported. Adipose-Derived Mesenchymal Stem Cells Isolation and Expansion. Rabbit adipose-derived mesenchymal stem cells (ADMSCs) were isolated from subcutaneous adipose tissue of a 2 week-old New Zealand rabbit. The guidelines of the National Institute for the Care and Use of Laboratory Animals were observed. The subcutaneous adipose tissue was dissected and washed extensively with Dulbecco’s phosphate buffered saline (DPBS, Gibco), and it was then digested with 0.1% type I collagenase (Gibco) at 37 °C for 45 min and centrifuged at 800 rpm for 10 min to obtain a cell pellet and remove the mature adipocytes. The pellet was resuspended in Dulbecco’s modified Eagle’s medium (DMEM) (Gibco) with 10% fetal bovine serum (Gibco), 1.0 × 105 U/L penicillin (Hyclone), and 100 mg/L streptomycin (Hyclone) and filtered through a 100 μm nylon mesh to remove cellular debris. The obtained cells were incubated at 37 °C in a humidified incubator containing 5% CO2 to allow them adhere on the

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RESULTS AND DISCUSSION Synthesis of the PNA Hydrogel. The PNA hydrogel synthesis scheme is shown in Figure 1. The free radical copolymerization among the double bonds from PEGMA, NIPAm, and ALGMA enable the formation of the interpenetrating polymer network among the different components (Figure 1), which could improve the mechanical properties of the hydrogels. The polymerization system was initiated by a redox initiator system (APS/TEMED), and the hydrogels were formed in situ in the injection area and could fill any irregular shape.19,38−40 This fabrication process also allows for bioactive molecules to be easily and uniformly loaded in the hydrogel in situ. Furthermore, the polymerization and drug loading process can be applied under physiological conditions, which can protect the model molecules, such as drugs and proteins, encapsulated in the system. The mechanical properties of the hydrogels are usually quite low, and the IPN structure would enhance the strength of the hydrogels.45,46 We optimized the mechanical properties of the hydrogels with different compositions using a rheometer. We first prepared the PEGMA-NIPAm (PN) hydrogels with different compositions, and we found that the pure PNIPAm (P0N100) hydrogel was quite soft and showed the lowest storage modulus of 179 Pa. After co-polymerized with PEGMA, the modulus of the PN hydrogel increased dramatically. P50N50 and P70N30 hydrogels showed higher modulus values compared with that of other compositions (Table 1 in the Supporting Information). Furthermore, the P50N50 hydrogel 3248


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PEGMA (Figure S3), the peak at 1725 cm−1 corresponds to the ester group, the band at 1633 cm−1 corresponds to the stretching vibration of carboxylic groups, and the band at 1025 cm−1 is due to the C−O stretching of alcoholic groups. The absorption peak of the double bonds around 1622 cm−1 for NIPAm and PEGMA completely disappeared in the FT-IR spectra of the PNA and P50N50 hydrogels (Figure S3), indicating that cross-linking took place during the reaction. The characteristic bands of alginate at 3340 cm−1 (Figure S3 in the Supporting Information) corresponding to the stretching vibration of hydroxyl groups shifted to 3275 cm−1 (Figure 3), indicating that the hydrogen bonds were formed among alginate, PEGMA, and PNIPAm, which should enhance the mechanical strength of the hydrogels. Swelling Properties of the Hydrogels. Figure 4 shows the swelling properties of the P50N50A20 hydrogel under

showed a much more obvious thermoresponsive property compared with that of the P70N30 hydrogel (Figure S1 in the Supporting Information). Therefore, the P50N50 hydrogel with a modulus of 4132 Pa was chosen for further research. In the next step, we tested the rheological properties of the PNA hydrogels, and the results are shown in Figure 2. The storage

Figure 2. Rheological behavior of the PN and PNA hydrogels.

modulus increased as the ALGMA content in the hydrogels increased. In particular, the modulus of the P50N50A20 and P50N50A30 hydrogels reached 12 800 and 13 750 Pa, respectively, and increased to 3.1−3.3 times that of the P50N50 hydrogel. This could be attributed to the formation of the IPN network among the different components and the hydrogen bonds among the alginate, PEGMA, and NIPAm. However, the P50N50A30 hydrogel is slightly brittle, and its gelation time is too short (16 s) for practical application. Therefore, we chose the P50N50A20 hydrogel with a gelation time of about 1 min (tested by rheometer) for the following swelling and drug release tests in this work. FT-IR spectra were used to characterize the chemical structure of the P50N50A20 hydrogel, as shown in Figure 3 and Figure S3 in the Supporting Information. Compared with that of NIPAm (Figure S3), the absorption peaks around 1364 cm−1 attributed to the methyl group −CH(CH3)2 and characteristic peaks at 1537 cm−1 of amide II mainly due to the C−N−H bending vibration of PNIPAm are present in the spectra of P50N50A20. Compared with the FT-IR spectra of

Figure 4. Swelling ratio of the P50N50A20 hydrogel under different conditions.

different conditions. All of the samples, tested at different temperatures and pH, reach swelling equilibrium before 9 h. Obviously, the swelling ratios of the hydrogel in a pH 2.1 solution are much lower than that in a pH 7.4 solution. This phenomenon can be attributed to the fact that the carboxylate groups on the alginate chain are transformed into −COOH under low pH conditions (pH 2.1) because the pKa of alginate is about 3.2. The hydrogen bonds among the hydroxyl groups (−OH) and −COOH groups in alginate and the −CONH− groups in PNIPAm and −C−O− and −OH groups in PEGMA restrict the swelling of the hydrogel. However, the carboxyl group in alginate changes into −COO− groups under higher pH (pH 7.4), and electrostatic repulsions between them lead to the swelling of the hydrogels. The temperature of the solution also affects the swelling ratio of the hydrogel. The swelling ratio of the hydrogel at 25 °C is much higher than it is at body temperature (37 °C). This is because PNIPAm is a temperature-sensitive polymer with a lower critical solution temperature at 32 °C, which means that PNIPAm is hydrophilic below LCST and is hydrophobic above its LCST.47 Therefore, the hydrogels at 25 °C swell much more than they do at 37 °C. The pH- and temperature-responsive nature of the hydrogels makes them quite attractive for use as a controlled drug delivery system. Morphology of the Hydrogels. The surface morphology of the hydrogels after swelling was observed by SEM (Figure S2 in the Supporting Information). Generally, all of the hydrogels showed a porous structure after swelling. The hydrogels in a

Figure 3. FT-IR spectra of the P50N50A20 hydrogel. 3249


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pH 7.4 solution at 25 °C showed the biggest cavities on the surface due to their having the highest swelling ratio among the samples (Figure S2a in the Supporting Information). However, the hydrogels in a pH 2.1 solution at 37 °C exhibited a condensed surface with the smallest pores due to their having the lowest swelling under these conditions. Release Properties of the Hydrogels. Because of its short half-life in plasma (1 to 2 h) and associated side effects, DCS is regarded as an ideal model molecule for use in a release system.48 The release profiles of DCS from the P50N50A20 hydrogel under different conditions are shown in Figure 5. At

would be beneficial for the practical application of the materials in an oral drug delivery system. Drugs administered orally usually pass through various segments of the gastrointestinal tract. In order to show the potential application of the hydrogels for use in oral delivery systems, we simulated the drug release behavior under gastrointestinal tract conditions. The drug-encapsulated PNA hydrogel was first immersed in a pH 2.1 solution (stomach conditions) for 3 h and was then transferred into a pH 7.4 solution (intestinal conditions) for 19 h, and the release results are shown in Figure 6. Less than 10% of the DCS in the

Figure 5. DCS release from the hydrogels at different temperatures and pH.

Figure 6. Release behavior of DCS from the P50N50A20 hydrogel in pH 2.1 buffer solution for 3 h followed by shifting it to a pH 7.4 solution for 19 h at 37 °C.

37 °C, the higher temperature accelerated the release of DCS from the hydrogel both in pH 2.1 and 7.4 solutions. This can be explained by the fact that PNIPAm was in a shrinking state at 37 °C above its LCST and because the effective cross-linking density of the PNA hydrogel network was reduced by the precipitation of PNIPAm, which would enhance the diffusion rate of the DCS out of the hydrogel network. On the other hand, the hydrogen bonds between the drug molecule DCS and PNA hydrogels were weakened at 37 °C, and the diffusion speed is higher at 37 °C than at 25 °C for a small molecule drug. All of these factors are beneficial for the improved release of DCS from the hydrogels at 37 °C than that at 25 °C. Therefore, the release of the small molecule drug DCS from the hydrogel was faster at 37 °C than at 25 °C. From Figure 5, we can also see that the pH of the solution has a significant effect on drug release. DCS release from the hydrogel in a pH 7.4 solution is 31−40% after 1 h, 68−78% after 5 h, and 86−95% after 12 h. However, a very slow release of DCS was observed at pH 2.1. The percent of DCS released at 1, 5, and 12 h is 5−6, 10−13, and 15−19%, respectively. This can be mainly attributed to the much lower swelling ratio of the hydrogel at pH 2.1 than that at pH 7.4. Furthermore, the interactions between DCS and the hydrogel also contribute to the slow release at pH 2.1. The −COOH group in DCS is in the −COO− state in an alkaline solution, so the water solubility of DCS greatly increases. Furthermore, the electrostatic repulsions between the −COO− group in DCS and −COO− in alginate greatly accelerate the release of DCS. However, in a pH 2.1 solution, the carboxyl group in DCS is in the −COOH state, and the hydrogen bonds formed between the −COOH from DCS and −OH, −NHCO, and −COOH in the network under acidic conditions further hinder the release of DCS from the hydrogel. The pH-sensitive release behavior of the hydrogel

hydrogel matrix was released in the first 3 h in the acidic solution, and this result agrees well with the release behavior presented in Figure 5. This indicated that the hydrogel has a protective effect for drugs administered orally, because most drugs remain in the hydrogel matrix when they pass through the low-pH environment of the stomach. When the hydrogel was moved into the alkaline solution (pH 7.4), an obvious release of DCS from the hydrogel was observed, and more than 68% of the DCS in the matrix was released 11 h after the hydrogel was transferred to the alkaline pH solution. After that, the DCS release continued, but the release rate became slower. Interestingly, a linear-like release profile of DCS from the hydrogels was observed in the pH 7.4 solution from 1 to 11 h (Figure 6). These results are very important because they indicate that it is easy to control and predict pharmacokinetics if these hydrogels are used in therapeutics, potentially achieving a high local bioactivity and low side effects of anti-inflammation drugs used for the treatment of disease. In order to demonstrate the potential application of the P50N50A20 hydrogel as an injectable delivery depot for localized protein release, we chose BSA as a model protein drug to evaluate the release properties of the PNA hydrogel in this study, and the results of BSA release from the hydrogel under different conditions are plotted in Figure 7. Under acidic conditions (pH 2.1), the release percentage of BSA from the matrix is less than 6% at both 25 and 37 °C for the first 9 h, and the release profile then reached equilibrium. This is probably because the small pores, resulting from the low swelling ratio of the hydrogels in a pH 2.1 solution, hinder the free diffusion of the high molecular weight BSA molecules. Furthermore, the attractions between carboxyl groups (−COOH) on the alginate chain and the amine group (−NH2) of BSA as well as the hydrogen bonds between the −COOH from alginate and 3250


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Figure 7. BSA release from the P50N50A20 hydrogel under different conditions.

Figure 8. Fluorescent images of ADMSCs cultured in the presence of the P50N50A20 hydrogel and control group (TCPS) by live/dead staining. (a−c) Days 1, 4, and 7 for the PNA hydrogel, respectively; (d−f) days 1, 4, and 7 for the control group (TCPS), respectively. Scale bar: 200 μm.

−COOH from BSA all restrict BSA release from the hydrogel. All of these reasons likely contribute to the low amount of BSA released in the acidic solution. In the pH 7.4 solution, the hydrogel showed a slight burst release with a release percentage of BSA of about 10−13% in the first hour, mainly attributed to some BSA molecules attached on the surface of the hydrogel. After 9 h, about 22% of BSA was released from the hydrogel, and after that, the BSA released gradually from the matrix. The release rate of BSA from the hydrogel at 25 °C was faster than that at 37 °C, and the release percentage was 85 and 66% after 13 days, respectively. This can be explained by the fact that the swelling ratio of hydrogel at 25 °C was higher than that at 37 °C, which is beneficial for the diffusion of BSA molecules, because when the hydrogels reached swelling equilibrium, the BSA release rate and percentage of release mainly relied on free diffusion which is dependent on the swelling ratio and pore size of the hydrogel. The hydrogel continued to release BSA for 13 days, which is much longer than the ionically cross-linked alginate hydrogels,34,49,50 indicating a much higher stability of the PNA hydrogels under physiological conditions. These results indicate that the injectable smart hydrogel has great potential application for use in the localized, sustained release of proteins. In Vitro Cell Compatibility of the Hydrogel. The cytotoxicity of the hydrogels was evaluated by performing live/dead viability/cytotoxicity and alamarBlue assays. ADMSCs (P4) were seeded in the wells at the same initial density, with cells contacting with the hydrogel. The cell viability and proliferation were examined on days 1, 4, and 7 (Figure 8). After incubating for 1 day, the live/dead assay result, in which the green color stands for live cells (stained with Calcein AM) and the red color stands for dead cells (stained with EthD-1), showed that both the hydrogel and control groups had few dead cells, with the cells showing a spindle-like morphology. When incubated for 4 days, both groups showed about 90% confluence. After continuing to incubate to the seventh day, both groups were 100% confluent, and there were a few dead cells in both groups. These results revealed that contacting the cells with the hydrogel did not affect the proliferation and morphology of ADMSCs, indicating that the hydrogels are noncytotoxic. The viability of ADMSCs was further quantitatively determined with alamarBlue assay, and the results are shown in Figure 9. After the first day, the cell proliferation from the hydrogel group was 97% of that of control group (TCPS), which was not statistically different (p > 0.05). This result

Figure 9. Proliferation of ADMSCs cultured in the presence of the P50N50A20 hydrogel, expressed as the percent of fluorescence intensity of the TCPS group on the first day. Mean for n = 4 ± SD. Error bar indicates standard deviation. *P < 0.05.

revealed that the hydrogel had no side effect on cell adhesion. When incubated for 4 days, the hydrogel group had a cell number of 85% of that of control group (p < 0.05), probably because the ADMSCs needed to adapt to the new environment, which might affect their proliferation rate. Nevertheless, cells in the hydrogel group demonstrated obvious cell proliferation (p < 0.05) after incubating them from day 1 to 4. Continuing to incubate cells from the fourth day to the seventh day, the hydrogel group showed continuous cell proliferation (p < 0.05). The cell number in the hydrogel group underwent a fast increase, reaching 97% of that of control group (p > 0.05) at the seventh day, and the few dead cells present in both groups might be due to normal cellular metabolism and apoptosis. The above results revealed that the hydrogel is noncytotoxic. Thus, the hydrogel has good biocompatibility and is a promising candidate for use in the controlled release of drugs.

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CONCLUSIONS Injectable, biocompatible, and robust IPN hydrogels based on methacrylated alginate were successfully synthesized, and their application as carriers for the oral delivery of drugs and local sustained release of proteins was demonstrated. The injectable hydrogels composed of poly(ethylene glycol) methacrylate, N3251


Biomacromolecules

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isopropylacrylamide, and methacrylated alginate were obtained using APS/TEMED as an initiator system under physiological conditions. Diclofenac sodium (DCS) and bovine serum albumin (BSA) were used as model molecules and were conveniently encapsulated in situ in the hydrogel. DCS release from hydrogel at pH 2.1 is quite low, and it increases dramatically at pH 7.4, indicating that the hydrogels have potential application for the oral delivery of molecules. BSA release from the hydrogel at pH 7.4 at 37 °C lasted for 13 days, demonstrating that the matrix is a good candidate for sustained protein release. The viability and proliferation of cells in the presence of the hydrogels indicate that these hydrogels are noncytotoxic. All of these results suggest that the nontoxic, injectable hydrogels are excellent candidates for use as polymeric carriers for the oral delivery of drugs and for localized long-term protein release.

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ASSOCIATED CONTENT

S Supporting Information *

Rheological properties and swelling properties of PN hydrogels; SEM images of P50N50A20 hydrogels; and FT-IR spectra of PEGMA, NIPAm, ALGMA, and P50N50 hydrogels. This material is available free of charge via the Internet at http:// pubs.acs.org.

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AUTHOR INFORMATION

Corresponding Authors

*(B.G.) Tel.:+86-29-83395361. Fax: +86-29-83395131. E-mail: [email protected]. *(P.X.M.) E-mail: [email protected]. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This project was supported by the National Natural Science Foundation of China (grant no. 21304073) and Xi’an Jiaotong University.

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