Performance optimization of injectable chitosan hydrogel by combining physical and chemical triple crosslinking structure Chen Chen, 1 Lei Wang, 2 Liandong Deng, 1 Renjie Hu, 2 Anjie Dong 1 1 School of Chemical Engineering and Technology, Tianjin University, Tianjin 300072, People’s Republic of China 2 Tianjin Institute of Medical and Pharmaceutical Science, Tianjin 300020, People’s Republic of China Received 9 March 2012; revised 4 June 2012; accepted 5 July 2012 Published online in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.a.34364 Abstract: To improve the biocompatibility and application properties of injectable chitosan hydrogel, an injectable triple crosslinking network hydrogel (CTGP) is prepared by physical interaction, Michael addition and disulfide bond formation based on thiolated chitosan (CS-TGA), b-glycerophosphate (b-GP) and poly(ethylene glycol) diacrylate (PEGDA) without the addition of cytotoxic crosslinkers and catalysts. Com- pared with the short gelation time of 2 min of CTG hydrogel (without PEGDA) at 37 C, CTGP hydrogel containing different molecular weight of PEGDA exhibits controllable gelation times from 1 to 22 min, which could meet the different demands in clinical application. Further, the compressive modulus is improved differently by introducing PEGDA into the system. The presence of PEGDA in CTGP hydrogel imparts better swelling property, and there is a sustained protein release from the hydrogel without any initial burst. In vitro cytotoxicity and hemolysis reveal that the gel is biocompatible. In vivo subdermal injection into mice models further confirms the non-cytotoxicity of the hydrogel and the hydrogel is highly resistant to degradation. V C 2012 Wiley Periodicals, Inc. J Biomed Mater Res Part A: 00A:000–000, 2012. Key Words: thiolated chitosan, b-glycerophosphate, poly(eth- ylene glycol) diacrylate, injectable triple crosslinking network hydrogel, biocompatibility How to cite this article: Chen C, Wang L, Deng L, Hu R, Dong A. 2012. Performance optimization of injectable chitosan hydrogel by combining physical and chemical triple crosslinking structure. J Biomed Mater Res Part A 2012:00A:000–000. INTRODUCTION Over the past decades, injectable thermo-sensitive hydro- gels, which are aqueous solutions at room temperature and can form a semi-solid depot after being injected into the desired tissue or organ, 1,2 have shown a great potential application in drug delivery systems, 3 tissue engineering, 4,5 and biomedical devices. 6,7 Injectable hydrogels have the advantage that some of the surgical implantations can be replaced by a simple minimally invasive injection procedure. Moreover, homogeneous encapsulation of cells and bioactive molecules like growth factors can be readily performed and the hydrogels can be formed in any desired shape in good alignment with the surrounding tissue. The typical obstacles for formulating injectable hydrogels for in vivo use include gelation occurring at a physiologically acceptable tempera- ture range within a reasonable time span and all compo- nents must be non-cytotoxic. Chitosan (CS), which is a natural cationic amino polysac- charide copolymer composed of glucosamine and N-acetyl- glucosamine, 8 has received a great deal of attention due to its well-documented biocompatibility, bioactivity, and low toxicity. 9–11 Injectable CS-based hydrogels have been pre- pared by either physical or chemical crosslinking methods. Combining CS with various polyols such as b-glycerophos- phate (b-GP) has been used to prepare physically crosslink- ing hydrogels with fast gelation property of a few minutes. 12–15 However, such physically crosslinking hydro- gels generally exhibit low stability, low mechanical strength, and fast degradation. The high concentrations of b-GP nec- essary for a fast gelation at body temperature can lead to potential toxicity. 15,16 Moreover, to get CS/b-GP hydrogel, CS has to be dissolved in acidic aqueous media. The poor solu- bility of CS in neutral solutions becomes a major obstacle which restricts its application. Thiolated CS, 17–19 a water soluble CS derivative, has been shown to form in situ hydrogel via forming disulfide bonds. Nevertheless, the relatively long gelation time in con- junction with the weak mechanical strength of the hydrogel formed render it less appealing for biomedical applica- tions. 20 Chemical crosslinking is an effective strategy to enhance the mechanical strength of hydrogels and the re- sistance to degradation, but usually potential cytotoxic cata- lysts, initiators, or crosslinkings are used. Herein, to improve the biocompatibility and application properties of injectable CS hydrogel, we combined physical and chemical crosslink- ing structure. The physical interaction between CS-TGA and Correspondence to: L. Deng; e-mail: [email protected]Contract grant sponsor: 863 Program; contract grant number: 2009AA03Z313 Contract grant sponsor: Tianjin Municipal Natural Science Foundation Key Project; contract grant number: 08JCZDJC17200 V C 2012 WILEY PERIODICALS, INC. 1
Transcript
Performance optimization of injectable chitosan hydrogel bycombining physical and chemical triple crosslinking structure
Chen Chen,1 Lei Wang,2 Liandong Deng,1 Renjie Hu,2 Anjie Dong1
1School of Chemical Engineering and Technology, Tianjin University, Tianjin 300072, People’s Republic of China2Tianjin Institute of Medical and Pharmaceutical Science, Tianjin 300020, People’s Republic of China
Received 9 March 2012; revised 4 June 2012; accepted 5 July 2012
Published online in Wiley Online Library (wileyonlinelibrary.com). DOI: 10.1002/jbm.a.34364
Abstract: To improve the biocompatibility and application
properties of injectable chitosan hydrogel, an injectable triple
crosslinking network hydrogel (CTGP) is prepared by physical
interaction, Michael addition and disulfide bond formation
based on thiolated chitosan (CS-TGA), b-glycerophosphate
(b-GP) and poly(ethylene glycol) diacrylate (PEGDA) without
the addition of cytotoxic crosslinkers and catalysts. Com-
pared with the short gelation time of 2 min of CTG hydrogel
(without PEGDA) at 37�C, CTGP hydrogel containing different
molecular weight of PEGDA exhibits controllable gelation
times from 1 to 22 min, which could meet the different
demands in clinical application. Further, the compressive
modulus is improved differently by introducing PEGDA into
the system. The presence of PEGDA in CTGP hydrogel
imparts better swelling property, and there is a sustained
protein release from the hydrogel without any initial burst.
In vitro cytotoxicity and hemolysis reveal that the gel is
biocompatible. In vivo subdermal injection into mice models
further confirms the non-cytotoxicity of the hydrogel and the
hydrogel is highly resistant to degradation. VC 2012 Wiley
Periodicals, Inc. J Biomed Mater Res Part A: 00A:000–000, 2012.
How to cite this article: Chen C, Wang L, Deng L, Hu R, Dong A. 2012. Performance optimization of injectable chitosan hydrogel bycombining physical and chemical triple crosslinking structure. J Biomed Mater Res Part A 2012:00A:000–000.
INTRODUCTION
Over the past decades, injectable thermo-sensitive hydro-gels, which are aqueous solutions at room temperature andcan form a semi-solid depot after being injected into thedesired tissue or organ,1,2 have shown a great potentialapplication in drug delivery systems,3 tissue engineering,4,5
and biomedical devices.6,7 Injectable hydrogels have theadvantage that some of the surgical implantations can bereplaced by a simple minimally invasive injection procedure.Moreover, homogeneous encapsulation of cells and bioactivemolecules like growth factors can be readily performed andthe hydrogels can be formed in any desired shape in goodalignment with the surrounding tissue. The typical obstaclesfor formulating injectable hydrogels for in vivo use includegelation occurring at a physiologically acceptable tempera-ture range within a reasonable time span and all compo-nents must be non-cytotoxic.
Chitosan (CS), which is a natural cationic amino polysac-charide copolymer composed of glucosamine and N-acetyl-glucosamine,8 has received a great deal of attention due toits well-documented biocompatibility, bioactivity, and lowtoxicity.9–11 Injectable CS-based hydrogels have been pre-pared by either physical or chemical crosslinking methods.
Combining CS with various polyols such as b-glycerophos-phate (b-GP) has been used to prepare physically crosslink-ing hydrogels with fast gelation property of a fewminutes.12–15 However, such physically crosslinking hydro-gels generally exhibit low stability, low mechanical strength,and fast degradation. The high concentrations of b-GP nec-essary for a fast gelation at body temperature can lead topotential toxicity.15,16 Moreover, to get CS/b-GP hydrogel, CShas to be dissolved in acidic aqueous media. The poor solu-bility of CS in neutral solutions becomes a major obstaclewhich restricts its application.
Thiolated CS,17–19 a water soluble CS derivative, hasbeen shown to form in situ hydrogel via forming disulfidebonds. Nevertheless, the relatively long gelation time in con-junction with the weak mechanical strength of the hydrogelformed render it less appealing for biomedical applica-tions.20 Chemical crosslinking is an effective strategy toenhance the mechanical strength of hydrogels and the re-sistance to degradation, but usually potential cytotoxic cata-lysts, initiators, or crosslinkings are used. Herein, to improvethe biocompatibility and application properties of injectableCS hydrogel, we combined physical and chemical crosslink-ing structure. The physical interaction between CS-TGA and
Contract grant sponsor: 863 Program; contract grant number: 2009AA03Z313
Contract grant sponsor: Tianjin Municipal Natural Science Foundation Key Project; contract grant number: 08JCZDJC17200
VC 2012 WILEY PERIODICALS, INC. 1
low amount of b-GP ensured a controllable gelation. Thechemical crosslinking, formed by Michael addition betweenacrylate groups on poly(ethylene glycol) diacrylate (PEGDA)and thiol groups on CS-TGA, as well as disulfide bondbetween thiol groups occurring at physiological conditionswithout the need of toxic catalysts, enhanced the mechanicalstrength. The structural contribution to gelation, swelling,degradation, mechanical properties, and biocompatibilitywere studied in vitro and in vivo.
MATERIALS AND METHODS
MaterialsCS (Mn ¼ 50,000 g mol�1, degree of deacetylation ¼ 95%)was purchased from Aokang Bio-Technology (Shandong,China). CS-TGA (thiol groups immobilized on CS was 4.11mmol g�1) was synthesized by the covalent attachment of thi-oglycolic acid (TGA) to the amine groups of CS according tothe literature.17 PEGDA (Mn ¼ 258, 700, 2000 g mol�1) werepurchased from Sigma (USA). b-glycerophosphate disodiumsalt (b-GP) was obtained from Sigma (USA). Bovine serum al-bumin (BSA, Mn ¼ 66,000 g mol�1) was supplied by RocheCompany. Coomassie brilliant blue G250 was purchased fromBiosharp Company. All the reagents were of analytic grade.
Preparation of CTGP hydrogelCTGP hydrogel was prepared by the following steps. Briefly,CS-TGA and PEGDA were dissolved in b-GP aqueous solutionunder stirring. The homogeneous mixture obtained wouldform hydrogel by being heated in a water bath at 37�C. Thestandard recipe for preparing hydrogel is listed in Table I.In addition, CTG hydrogel was also prepared according tothe same method without adding PEGDA.
Characterization of hydrogelsInterior morphologies. Morphologies of hydrogels werecharacterized by utilizing scanning electron microscopy(SEM) after gelation. The hydrogel samples were quicklyfrozen in liquid nitrogen and further lyophilized with afreeze dryer system under vacuum at �50�C for at least48 h until all of the water was sublimed. The freeze-driedhydrogel was then carefully fractured and sputter-coatedwith gold. The interior morphology of hydrogel was visual-ized with SEM (S4800, Hitachi, Japan).
Rheological analysis. Rheological experiments were carriedout with a Stress Tech Fluids Rheometer (RheologicalInstruments AB, Sweden). A smooth parallel plate (20 mmdiameter, 0�) geometry was used for both the temperature-and the time-sweep rheological experiments. All experi-ments were carried out at a frequency of 1 Hz, a strain of1% and a gap size of 1 mm. The temperature-sweep rangedfrom 20�C to 50�C with a heating rate of 1�C/min and thetime-sweep was performed for 2 h at 37�C. To preventwater evaporation, a layer of oil was introduced around thesample in both time- and temperature-sweep experiments.21
Mechanical testing. The mixture solutions described abovewere injected into a 48-well culture plate and incubated at
37�C to obtain columned hydrogels (11-mm diameter, 8-mmheight). Compressive modulus of the hydrogel at room tem-perature was measured in the elastic region using materialstesting machines (350-2208, Testometric). Unconfined com-pression was applied at a constant rate of 1 mm/min up toa strain of 20%.22 The experiments were performed intriplicate.
In vitro swelling and degradation. 0.5 mL CTGP700-4solution was placed in a test tube, incubated at 37�C toform a hydrogel and accurately weighed (Wi0). Five millili-ters of phosphate buffer saline (PBS, pH 7.4) was appliedon top of the hydrogel. The tube was set on a constant tem-perature shaker (SHZ-88, Jiangsu, China) at 37�C and70 rpm. A 5 mL aliquot of PBS was replaced by fresh PBSat regular time intervals, and the hydrogels were weighed(Wt0). The swelling ratio of the hydrogel was calculatedfrom the equation Wt0/Wi0 � 100%. In the degradationexperiment, the initial and following hydrogels were lyophi-lized by freeze drying and weighed (Wi, Wt). The degrada-tion of the hydrogel was evaluated by the weight lossaccording to the equation (Wi � Wt)/Wi � 100%. Theexperiments were performed in triplicate. The freeze-driedhydrogels were characterized by SEM.
Incorporation of BSA and In vitro release study. 0.5 mLhydrogel mixture with a defined amount of BSA (0.5, 1, 2wt %) was placed into a test tube and incubated at 37�C toform a hydrogel, and a 5 mL of PBS (pH 7.4) was thenadded into the test tube. In vitro release was operatedunder shaking (70 rpm) at 37�C. At predetermined inter-vals, a 2.5 mL aliquot of PBS was withdrawn for measuringthe amount of BSA released from hydrogel and subse-quently an equal volume of fresh PBS was supplemented tokeep the volume of the release medium constant. The con-centration of BSA in the release medium was determined byBradford assay.
In vitro cell proliferation assay. To investigate the toxiceffect of PEGDA700 on cell culture and cell proliferation, theviability of mouse fibroblast cells (L929) was investigated.The cells were cultured on a 96-well tissue culture plates(100 lL/well; Cellstar) at 1 � 105 cells/mL in DMEM
a CTGPa-b: a is representative for the molecular weight of PEGDA
and b stands for the different hydrogel series with different concentra-
tions of PEGDA.b PEGDA700 is representative for poly(ethylene glycol) diacrylate,
Mn ¼ 700 g mol�1.
2 CHEN ET AL. PERFORMANCE OPTIMIZATION OF INJECTABLE CHITOSAN HYDROGEL
(Dulbecco’s modified Eagle’s medium, Sigma-Aldrich) con-taining 10% fetal bovine serum (FBS, Biochrom Ag, Ger-many), and then incubated at 37�C in 5% CO2 for 24 h.PEGDA700 was added to cell culture media, and the finalconcentrations of PEGDA700 in the media were 8, 16, 32mmol L�1. 100 lL of each media was separately added intothe individual well to replace the original medium, in whichL929 cells had been seeded. The cell viability cultured withDMEM medium was used as control. The cell viability wasevaluated every 24 h. The absorbance at 570 nm which rep-resented the cell viability was measured with a WellscanMK-3 Microplate Spectrophotometer (Labsystems Dragon,Finland).
To investigate the biological effect of CTGP700-4 hydro-gel extract on cells, viability of mouse fibroblast cells(L929) was also investigated. CTGP700-4 hydrogel was pre-pared with 8 mmol L�1 PEGDA700. The hydrogel extractwas prepared by immersing 0.1 g hydrogel into 1 mLDMEM medium for 72 h. The same experiment was per-formed as described above, except for replacing the originalmedium with 100 lL extract.
Blood compatibility evaluation. The blood compatibility ofCTGP700-4 hydrogel extract was evaluated by hemolysisexperiment.23,24 New Zealand white rabbit blood (2 mL)anticoagulated with 2% potassium oxalate (0.1 mL) was fur-ther diluted by adding 4.5 mL physiological saline. Thediluted blood (0.2 mL) was added to 8 mL hydrogel extracts(prepared in the same method as above) in test tube. Posi-tive controls (100% hemolysis induced by replacing thehydrogel extract with 8 mL of distilled water) and negativecontrols (0% hemolysis, 8 mL physiological saline with nohydrogel extract) were also set up. All the test tubes con-taining the hydrogel extracts and control solutions wereincubated at 37�C for 1 h. After the incubation, the tubeswere centrifuged at 2000 rpm for 5 min at room tempera-ture. The supernatant solutions were removed into new testtubes for photographing. The absorbance of the supernatantsolution was measured with UV–Vis spectrophotometer at545 nm. The hemolysis was compared with the control sam-ples. Each experiment was performed in triplicate.
Histopathological analysis. Kunming mice (20–30 g, 8weeks) were used to examine the biocompatibility of
CTGP700-4 hydrogel. After being sterilized, 0.4 mLCTGP700-4 solution was subcutaneously injected in theback of the mice. The mice injected with 0.4 mL of b-GP(0.1 mol L�1) solution were used as the control. The weightof each mouse was taken daily before their scheduled sacri-fice on days 1, 7, 21, and 35 following injection. The skinwas cut and the injection site was exposed.
Immediately after necropsy, skin tissues from the injec-tion sites were retrieved and fixed in Bouin’s solution for 7days, and then washed by flowing water for 24 h. Afterbeing bisected into pieces with a thickness of 2 mm, all tis-sues were initially dehydrated in a graded series of alcoholand then embedded in paraffin. The transverse sections (4–5 mm thick) were prepared using rotatory microtome andstained with hematoxylin–eosin dye. The histologicalchanges, such as acute-chronic inflammatory symptoms,fibroblastic proliferation, foreign material deposits, and anyother inflammation symptoms were evaluated through ob-servation under a light microscope.
All animals were housed individually in plastic cages ina controlled environment with free access to food andwater. All animal experiments were performed in accord-ance with the People’s Republic of China national standard(GB/T 16886.6-1997).
RESULTS AND DISCUSSION
Hydrogel formationThe feasibility of developing a thermo-gelling system basedon CTGP is demonstrated in Figure 1. CTGP700-4 aqueoussolution presented a flowing state at room temperature[Fig. 1(a)]. After being heated at 37�C for several minutes,the solution transformed to non-flowing hydrogel. Thehydrogel taken out of the test tube was regularly shapedwith a cylindrical appearance [Fig. 1(b)]. Compared withCTG hydrogel [Fig. 1(c)], CTGP700-4 hydrogel [Fig. 1(b)]appeared to be stronger and more transparent.
The concept of formation of the hydrogel is depicted inScheme 1. CS-TGA, a water soluble CS derivative formed byattaching TGA to CS, has both –NH2 and –SH groups. Withthe presence of b-GP, there is a physical interaction betweenb-GP and –NH2 groups on CS-TGA responded to tempera-ture,25 so that the system displays a fast gelation accordingto the temperature variation by forming physical crosslink-ing network firstly. PEGDA served as the crosslinker is
FIGURE 1. CTGP700-4 aqueous solution at room temperature (a), CTGP700-4 hydrogel (b), and CTG hydrogel (c) taken out of the test tube after
gelation at 37�C. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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capable of reacting with the free –SH groups on CS-TGA.The relatively fast formation of Michael addition results increating one chemical crosslinking network. This is followedby a slower interaction of disulfide bonds betweenunreacted –SH groups leading to the establishment of theother chemical crosslinking network. In addition, glycerol
groups of b-GP can promote the protective hydration of CS-TGA chains which makes the chains stretch freely andapproach each other easily so that it can help the formationof chemical crosslinking networks.26,27 By taking advantageof the disparity in their reaction times, the triple crosslink-ing network hydrogel was formed via a one-step process
SCHEME 1. Schematic representation of the theoretical CTGP hydrogel formation.
FIGURE 2. Interior morphologies of CTG hydrogel (a), CTGP258-4 hydrogel (b), CTGP700-4 hydrogel (c), and CTGP2000-2 hydrogel (d) (500�).
4 CHEN ET AL. PERFORMANCE OPTIMIZATION OF INJECTABLE CHITOSAN HYDROGEL
that simply blending the three components, without anycomplicated operations.
Morphologies of hydrogelsThe interior morphologies of the freeze-dried CTGP hydrogelswith different molecular weight of PEGDA and CTG hydrogelwere observed by SEM. As shown in Figure 2, all hydrogelshad a highly macroporous, sponge-like structure, and the av-erage pore size was 20–60 lm. CTGP258-4 hydrogel devel-oped a more crosslinked triple crosslinking network structure
than CTG hydrogel. However, with the increment of molecularweight of PEGDA, the pore size became larger. These were at-tributable to the molecular weight of PEGDA and the cross-linking densities formed in the hydrogels. As the molecularweight of PEGDA became higher, the solubility was better andit stretched sufficiently which created a larger pore size net-work. In addition, CTGP2000-2 hydrogel had the largest poresize also due to the lower crosslinking density. Herein, the po-rous structures of the hydrogels show potentials as carriersfor macromolecular compounds delivery system.
FIGURE 3. Rheological analysis of hydrogels. (a) Storage modulus (G0) and loss modulus (G00) as a function of temperature; (b) Storage modulus
(G0) and loss modulus (G00) at 37�C as a function of time; (c) Viscosity at 37�C as a function of time.ED1
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Rheological analysisThe characteristics of injectable hydrogel systems were epito-mized by their rheological properties. Figure 3(a) depicts thetemperature-sweep analysis of the representative CTGP258-4,CTGP700-4, CTGP2000-2, and CTG mixture solutions. Figure3(b,c) shows storage modulus (G0), loss modulus (G00), and vis-cosity of the mixture solutions as a function of time at 37�C.Both moduli elevated as gelation proceeded, but the builduprate of G0 was higher than that of G00 because of the interactionbetween the components and the formation of crosslinking.The different elevation rates of G0 and G00 led to a crossover(t ¼ tgel, the gel point), signifying the transition of theprecursor from a liquid phase to a solid phase.28 Both modulicontinued to increase and eventually plateaued, implying theformation of a well-developed three-dimensional network.
The gelation temperatures of CTGP258-4, CTGP700-4,CTGP2000-2, and CTG hydrogels were 48, 46, 23, and 30�C[Fig. 3(a)], and the gelation times of them in sequence were22, 14, 1, and 2 min [Fig. 3(b)]. During the same time pe-riod at a constant temperature of 37�C, G0 of CTG systemplateau, but G0 of CTGP system continued to increase due tothe formation of chemical crosslinking which made the sys-tem obtain higher elasticity [Fig. 3(b)]. Similar transitionswere also observed in viscosity [Fig. 3(c)]. Apparently, CTGhydrogel formed by physical crosslinking and disulfidebonding had a fast gelation but lower mechanical strength.CTGP hydrogel with triple crosslinking network establishedfirst via physical crosslinking, followed by Michael additionand disulfide bond displayed higher mechanical strength.The gelation behaviors of CTGP system were not identicaldue to the different molecular weight of PEGDA. PEGDAwith lower molecular weight was more hydrophobic, whichhindered the formation of crosslinking between doublebond and thiol group. In addition, the gelation propertieswere also concerned with the viscosity. The gelation timebecame longer with lower viscosity [Fig. 3(b,c)]. Neverthe-less, the gelation was still controllable in 22 min at thepresence of the b-GP, and the various gelation times pro-vided wider uses in clinical application.
Mechanical testingCompressive modulus which is an important mechanicalcharacteristic of hydrogel was studied by unconfined com-pression testing as shown in Figure 4. Compared with CTGhydrogel, the compressive modulus of CTGP hydrogels washigher because of Michael addition formation in concertwith the disulfide bond. With increasing the ratio of doublebond/thiol group, the compressive modulus was elevatedattributing to the increase of crosslinking density. Consistentwith the results of rheological testing, the compressive mod-ulus of CTGP hydrogel was also in the order of CTGP700 >
CTGP258 > CTGP2000.In summary, the results of mechanical testing strongly
agreed with those obtained from the rheological test, dem-onstrating that the mechanical property of CTG hydrogelcould be improved by introducing PEGDA. CTGP700-4hydrogel was chosen in the following experiments due to itsproper gelation time and better mechanical property com-pared with other CTGP hydrogels.
In vitro swelling and degradationThe swelling and degradation profiles of CTGP700-4 andCTG hydrogels are shown in Figure 5. During the initialtime period, the continuous swelling of hydrogel was attrib-uted to increased uptake of medium by the hydrogel. How-ever, the degraded fragments became soluble and werereleased from hydrogel over time. When the increment inswelling was overtaken by the weight loss of hydrogel, adecrement in the swelling ratio of hydrogel took place. Com-pared with CTG hydrogel, the swelling ratio of CTGP700-4hydrogel was higher. Hydrogel containing PEGDA couldretain more water. The degradation rate of CTGP700-4hydrogel was also higher. The disparity of the two hydrogelsin degradation was caused by the hydrolysis of the esterbond. There are two kinds of degradation in CTGP hydrogel:one is the degradation through hydrolysis of the ester bondbetween the thioether and PEG, as the hydrogels via Michaeladdition;29–32 the other is the biodegradation of CS. Therewas little hydrolysis of CS when the glycosidic bonds of theCS backbone were exposed to the environment containingno hydrolases.33,34 In addition, CTGP700-4 hydrogel sus-tained a highly macroporous, sponge-like structure, and theaverage pore size was increased gradually due to the degra-dation [Fig. 4(c–e)].
In vitro BSA release behaviorsIn vitro release behaviors of protein from CTGP and CTGhydrogels were investigated using BSA as a model protein.As shown in Figure 6(a), the initial burst releases of BSAfrom the hydrogels were not observed and the release rateswere sustained. The linear correlation coefficients of theprofiles shown in insert Figure 6(a) were 0.988, 0.992, and0.996, and the releases scaled almost linearly with thesquare root of time, indicating that the releases were closeto first-order kinetics.32 However, the imperfect shape of theprofiles might result from diffusion mechanisms, in whichdiffusion of BSA from the bottom layer might take longerthan that from the upper layer. It could be noticed that the
FIGURE 4. Compressive modulus of hydrogels at room temperature
(n ¼ 3).
6 CHEN ET AL. PERFORMANCE OPTIMIZATION OF INJECTABLE CHITOSAN HYDROGEL
release rate from CTGP hydrogel was slower than CTGhydrogel. By introducing PEGDA to the system, the triplecrosslinking network made the hydrogel more compact thandouble network.
The effect of BSA-loaded amount in the hydrogel on therelease profile was also investigated at 37�C with CTGP700-4 hydrogel. As shown in Figure 6(b), the cumulative releasewas decreased with increasing BSA-loaded amount. The lin-ear correlation coefficients of the profiles shown in insetFigure 6(b) were 0.992, 0.996, and 0.992, thus the releaseswere also close to first-order kinetics. At a lower BSA-loaded amount, the release mainly depended on diffusion-controlled release since BSA was free and the mesh size ofhydrogel network was larger than the hydrodynamic size ofBSA.35,36 However, with the increment of BSA-loaded
amount, the cumulative release declined. The aggregationand subsequent precipitation of BSA might have occurred inhigher BSA concentrations, which led to decreased releases.
It was observed from Figure 6 that BSA was not com-pletely released. Eighty-eight percent of entrapped BSA wasreleased at most until reaching plateau because some of thefree thiol groups on BSA could have reacted with PEGDA orCS-TGA, which restricted the release.
In vitro cell proliferation assay of PEGDA700and CTGP700-4 hydrogelTo evaluate the effect of PEGDA700 on cell viability, L929were cultured with media containing certain concentrationsof PEGDA700 for 72 h. As shown in Figure 7(a), after 24 h,cell viability measured by MTT assay at 16 mmol L�1 and
FIGURE 6. In vitro release profiles of BSA from hydrogels at 37�C. (a) BSA (1 wt %) release profiles from CTG and CTGP hydrogels. The inset
shows the cumulative release as a function of the square root of time (linear correlation coefficients are 0.988, 0.992, and 0.996) (n ¼ 3); (b) The
effect of BSA-loaded amount on BSA release from CTGP700-4 hydrogel. The inset shows the cumulative release as a function of the square root
of time (linear correlation coefficients are 0.992, 0.996, and 0.992) (n ¼ 3).
FIGURE 5. Swelling and degradation profiles of hydrogels exposed to PBS (pH ¼ 7.4) at 37�C (n ¼ 3). (a) Swelling profile of CTGP700-4 and
CTG hydrogels against time. (b) Weight loss profile of CTGP700-4 and CTG hydrogels against degradation time. (c–e) The interior morphology
of CTGP700-4 hydrogel after degradation for 3, 18, and 30 days, respectively (300�).
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32 mmol L�1 PEGDA700 was only 77% and 55%, respec-tively, and further diminished over time. Cell viability was88% in culture media containing 8 mmol L�1 of PEGDA700after 24 h and the cells proliferated after incubating for 72h, indicating no cytotoxic effect on cell survival and growth.The result suggested that cell viability was affected by
PEGDA700 in a dose-dependent manner and there was aninverse relationship between PEGDA700 and cell viability.
The effect of CTGP700-4 hydrogel containing 8 mmolL�1 PEGDA700 on cell viability was also evaluated by inves-tigating the cytotoxicity of CTGP700-4 hydrogel extracttoward L929 via MTT assay. Figure 7(b) shows the cell
FIGURE 8. Subdermal injection of CTGP700-4 hydrogel. (a) The apparent morphology of the hydrogel 35 days after injection; (b) The interior
morphology of the hydrogel 35 days after injection (450�). (c–f) Representative histopathological changes noted at the injection sites: (c)
CTGP700-4 hydrogel, day 1, (100�); (d) CTGP700-4 hydrogel, day 7, (100�); (e) CTGP700-4 hydrogel, day 21, (100�); (f) CTGP700-4 hydrogel, day
35, (100�) (M: Macrophage; L: Lymphocyte; FC: Fibroblast cells; FT: Fibrous tissues; H: Hydrogel). [Color figure can be viewed in the online
issue, which is available at wileyonlinelibrary.com.]
FIGURE 7. In vitro cytotoxicity evaluation of PEGDA700 (a) and CTGP700-4 hydrogel extracts (without dilution) (b) toward L929 (n ¼ 5).
8 CHEN ET AL. PERFORMANCE OPTIMIZATION OF INJECTABLE CHITOSAN HYDROGEL
viability cultured with CTGP700-4 hydrogel extract. The via-bility of L929 was greater than 98% after 24 h and therewas an increase in cell viability after a week. According tothe relationship of relative growth rate and cytotoxicitygrade in United States Pharmacopeia (USP), there was nocytotoxicity of CTGP700-4 hydrogel extract toward L929.The result indicated that CTGP700-4 hydrogel extract exhib-ited a good cellular compatibility.
Blood compatibilityThe blood compatibility of CTGP700-4 hydrogel extract wasevaluated by hemolysis assay. All supernatants removedfrom CTGP700-4 hydrogel extract groups were as clear andcolorless as negative control groups. In contrast, the positivecontrol groups produced homogenous red samples, indicat-ing cell membrane rupture had occurred. The absorbance ofthe supernatant solutions was analyzed in one-way ANOVA.There were no significant differences (0.074, p>0.05)between CTGP700-4 hydrogel extract groups and negativecontrols groups. Significant differences (0.0001, p <0.01)were existed between hydrogel extracts samples and posi-tive controls. The hemolysis assay results demonstrated thatCTGP700-4 hydrogel extract was non-hemolytic in natureand had good blood compatibility.
In vivo biocompatibilityAll mice were healthy throughout the experiments, asjudged from the body weight experiments and pathologicalchanges at the injection site. After the mice were sacrificedat predetermined time points, the apparent morphology ofthe hydrogel was observed [Fig. 8(a)]. CTGP700-4 hydrogelwas found to maintain its integrity after 35 days. The inte-rior morphology of hydrogel was observed by SEM. Asshown in Figure 8(b), the hydrogel sustained a highly mac-roporous structure.
As shown in Figure 8(c–f), the tissue samples takenfrom the injection sites at different time were analyzed. Oneday after injection the inflammation was minimal with a fewinflammatory cells appearing, which mainly were neutro-cytes and lymphocytes. At 7th day, a layer of fibrous tissueappeared under superficial layer of the muscularis. A fewlymphocytes could still be found, but there were more mac-rophages occurred. After 21 days, the layer of fibrous tissueand a few lymphocytes still existed. Among the tissue, smallhydrogel pieces could be seen. After 35 days, the inflamma-tion was not obvious and fibrous tissue became thinner astime went by.37 Overall, no significant inflammation wascaused by the injection of CTGP700-4 hydrogel. The tissuereaction to CTGP700-4 hydrogel implant is acceptable andCTGP700-4 hydrogel is biocompatible for use in therapeuticagent delivery systems in vivo and tissue engineering.
CONCLUSION
CTGP hydrogel, composed of CS-TGA, b-GP, and PEGDA, hasbeen formulated without any potentially cytotoxic cross-linkers and catalysts. CTGP hydrogel created by physicalinteraction, Michael addition, and disulfide bond combinesthe advantages of physical and chemical crosslinking. CTGP
hydrogel undergoes a controllable sol-gel transition at bodytemperature with a low amount of b-GP which decreasesthe cytotoxicity. The results of mechanical test stronglyagree with the rheological analysis, CTGP hydrogel ismechanically strong due to the triple crosslinking networkstructure. The swelling ratio of CTGP hydrogel is higherthan CTG hydrogel due to the presence of PEGDA. There isa sustained release of BSA from CTGP hydrogel with noinitial burst in vitro. In vitro cytotoxicity, hemolysis, andin vivo subdermal injection experiments verify that CTGPhydrogel is biocompatible. Therefore, this injectable triplecrosslinking network hydrogel is a promising candidate forinjectable protein delivery system and clinically relatedapplications, where the biocompatibility, fast gelation, me-chanical strength, and durability are the essentialrequirements.
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