Covalently crosslinked chitosan hydrogel formed at neutral
pH and body temperature
Yi Hong,1 Zhengwei Mao,1 Hualin Wang,2 Changyou Gao,1 Jiacong Shen1
1
Key Laboratory of Macromolecule Synthesis and Functionalization of the Ministry of Education, Department
of Polymer Science and Engineering, Zhejiang University, Hangzhou 310027, People’s Republic of China
2
Department of Otorhinolaryngology, The Second Hospital Affiliated to Zhejiang University,
Hangzhou 310027, People’s Republic of China
Received 29 July 2005; accepted 21 March 2006
Published online 29 August 2006 in Wiley InterScience (www.interscience.wiley.com). DOI: 10.1002/jbm.a.30837
Abstract: Water-soluble chitosan having double bonds
(CS–MA–LA) was synthesized by sequential grafting of
methacrylic acid (MA) and lactic acid (LA) via the reaction
between amino groups and carboxyl groups under the
catalysis of carbodiimide. Its molecular structure was verified by FTIR and 1H NMR characterizations. Elemental
analysis measured grafting ratios of 19% and 10.33% for
MA and LA, respectively. CS–MA–LA was readily soluble
in pure water and did not precipitate till pH 9. Gelation of
the CS–MA–LA was realized by thermal treatment at body
temperature under the initiation of a redox system, ammonium persulfate (APS)/N, N,N0 ,N0 -tetramethylethylenediamine (TEMED). The gelation time could be mediated in
a wide range, e.g. from 6 to 20 min, by reaction tempera-
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HONG ET AL.
change in pH value, covalently or ionically crosslinking, chitosan hydrogel has been formed.13 However,
the acidic solubility and gelation methods employed
so far will surely limit the application of chitosan as an
injectable hydrogel for tissue regeneration in vivo. Up
to present, only two kinds of chitosan hydrogel systems have been developed as injectable scaffolds. For
example, glycerol-2-phosphate (b-GP) has been used
to adjust the chitosan solution from acidic to neutral,
thus chitosan hydrogel are formed at a temperature
close to 378C.2,14 By incorporation of viable chondrocytes into the hydrogel system, injection of the hydrogel into a mouse has formed proteoglycan-rich matrix
in vivo. Moreover, mesenchymal stem cells (MSCs)
have been mixed with an injectable thermosensitive
(water-soluble chitosan-g-poly(N-isopropylacrylamide))
hydrogel. In vivo results demonstrate that the MSCs
can be differentiated to chondrocytes, and cartilage is
formed after culturing for 14 weeks after the cell–
hydrogel complex is injected into the submucosal
layer of the bladder of rabbit.15
Therefore, modification of chitosan molecule to enhance its solubility at neutral pH and to develop a
friend gelation method is of both practical and technological significance. In the present study, a novel
chitosan gelation system is developed. The derivated
chitosan has good solubility at neutral pH, and can
be covalently thermocrosslinked to form hydrogel at
body temperature. For this purpose, methacrylic acid
(MA) and lactic acid (LA) are successively grafted
onto the chitosan molecules that endow the chitosan
with crosslinkable and water-soluble features, respectively. The chitosan hydrogel is then formed at
an elevated temperature under the initiation of a redox
system, ammonium persulfate (APS) and N,N,N0 ,
N0 -tetramethylethylenediamine (TEMED). Its cytotoxicity is further assessed by in vitro 3T3 fibroblast
culture.
EXPERIMENTAL SECTION
Materials
Chitosan (average MZ % 600,000) was obtained from Haidebei Marine Bioengineering Company, Ji’nan, China. MA
and APS were purified via distillation under reduced pressure and recrystallization, respectively. 1-Ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride (WSC) was
purchased from Sigma. LA and TEMED (>98%) were used as
received.
Synthesis of CS–MA–LA
Eight hundred milligrams of chitosan was dissolved in
100 mL water and 420 mL (0.48 mM) MA, to which 930 mg
Journal of Biomedical Materials Research Part A DOI 10.1002/jbm.a
(0.48 mM) WSC were added. The reaction lasted for 24 h
at room temperature under magnetic agitation. The pH
value of the mixture solution increased from 4 to *7 during this process owing to the alkaline nature of the resultant urea. In order to remove the unreacted MA and other
small molecular weight products, the resultant mixture
was sealed in a membrane with a cut off molecular weight
of 10,000 Da and dialyzed in a lager amount of triple-distilled water for 3 days. Finally, MA grafted chitosan (CS–
MA) was obtained by freeze-drying.
Half of the CS–MA (*400 mg) was dissolved in 50 mL
water containing 210 mL (0.2 mM) LA overnight, to which
460 mg WSC was added. The mixture was stirred for 24 h
at room temperature. Following the purification steps
described earlier, the water-soluble and cross-linkable chitosan (CS–MA–LA) was obtained. The yields of both CS–MA
and CS–MA–LA were over 90%.
Measurement of water solubility
The water solubility was measured according to literature.16 Briefly, chitosan, CS–MA, and CS–MA–LA were
dissolved separately in water (1 mg/mL) with a pH value
of 3. At this pH value, chitosan and its derivatives are
completely soluble. The pH value of each solution was
gradually raised by addition of 0.1 mol/L NaOH solution,
and the pH value at which precipitation occurred was
measured using a pH meter. The concentrations of the chitosan and its derivatives should not be diluted below 70%
of the original concentration (1 mg/mL) upon addition of
NaOH solution.
Gelation of CS–MA–LA
CS–MA–LA aqueous solution was gelated by radical polymerization under the initiation of a redox system including
oxidant APS and reducer TEMED. APS and TEMED were
previously made into 1M solutions, respectively. 1% CS–
MA–LA aqueous solution was sequentially mixed with APS
and TEMED solutions. The mole ratio of APS and TEMED
was kept same for all the experiments. Then this mixture was
injected into a mold by a syringe. The gelation was conducted in the mold at a given temperature. The gelation time
was recorded right from the mixing to the state that the mixture lost its flow ability. Three to five parallel experiments were
conducted and average data were reported as mean 6 standard
deviation. Initiator’s concentration and reaction temperature
were varied to evaluate the gelation time.
Swelling ratio of the hydrogel
For measuring the swelling ratio, chitosan hydrogels
were prepared by gelation of 1% CS–MA–LA solution and
different concentration of APS/TEMED for 24 h at 378C.
The chitosan hydrogels were balanced in PBS for 24 h at
378C and weighted (W1). Then the hydrogels were dehydrated under reduced pressure at 358C to constant weights
(W2). The swelling ratio of the hydrogel is defined as (W1 W2)/W2.
CHITOSAN HYDROGEL FORMED AT NEUTRAL pH AND BODY TEMPERATURE
Crosslinking yield
One percent CS–MA–LA solution was gelated at different concentration of APS/TEMED for 24 h at 378C. The
chitosan hydrogels were freeze-dried and weighted (W3).
The freeze-dried hydrogels were immersed in 3% acetic
acid solution for 24 h, followed by extensive washing with
deionized water. Then the hydrogels were freeze-dried again
to constant weights (W4). The crosslinking yield is defined as
(W4/W3) 100%.
Cytotoxicity test at different cell seeding density
and APS/TEMED concentration
The cytotoxicity of the chitosan hydrogel was assessed by
culture of 3T3 fibroblasts supplemented with the extractant
of the hydrogel. APS and TEMED were previously made
into 1M PBS solutions and then sterilized by membrane filtration with a pore size of 0.22 mm, respectively. CS–MA–LA
was sterilized under UV radiation for 3 h and then dissolved
in PBS. Hydrogels were fabricated from a 1% CS–MA–LA/
PBS solution and 5 or 10 mM APS/TEMED at 378C. The
hydrogels were treated with Dulbecco’s minimum essential
medium (DMEM) supplemented with 10% fetal bovine serum (FBS) at a ratio of 100 mg hydrogel/mL DMEM for 24 h
at 378C. The solutions extracted from the hydrogels gelated
at 5 mM and 10 mM APS/TEMED are designated as E5 and
E10, respectively. DMEM supplemented with 10% FBS was
used as a negative control.
Different numbers of 3T3 fibroblasts were seeded onto a
96-well culture plate. The cells were cultured in 200 mL
DMEM supplemented with 10% FBS at 378C in 5% carbon
dioxide atmosphere. After 12 h, DMEM was removed and
200 mL E5 was added. After incubation at 378C for 24 h, the
cytoviability was quantitatively measured by MTT assay.
The preserved ratio of cytoviability is defined as A2/A1
100%, where A1 and A2 represent the absorbance of the
negative control and the samples, respectively. All data
were averaged from 3 parallel experiments and expressed
as mean 6 standard deviation.
Each well of the 96-well culture plate was seeded with
1 105 3T3 fibroblasts. The cells were cultured in 200 mL
DMEM supplemented with 10% FBS at 378C in 5% carbon
dioxide atmosphere. After 12 h, DMEM was removed and
200 mL E5 or E10 was added. The cytoviability was quantitatively measured by MTT assay after incubation at 378C for
1d–4d. The morphology of the cells at 1d and 4d was
observed under confocal laser scanning microscopy (CLSM,
Bio-Rad Radiance 2100). The cells were incubated in 5 mg/
mL fluorescein diacetate (FDA)/PBS solution for 10 min. In
this process, FDA (no fluorescence) could penetrate through
cell membranes and was hydrolyzed into fluorescein by viable cells, which was then excited at 488 nm under CLSM.17
MTT assay
After the cells were cultured for a given time, 20 mL MTT
(3-(4,5-dimethyl) thiazol-2-yl-2,5-dimethyl tetrazolium bromide, 5 mg/mL) was added into each well. The cells were
continually cultured for another 4 h. During this period, via-
915
ble cells could reduce the MTT to formazan pigment, which
was dissolved by 200 mL dimethyl sulphoxide after removal
of the culture medium. The absorbance at 570 nm was
recorded under a microplate reader (Bio-Rad 550).
In vivo inflammatory reaction of the
chitosan hydrogel
A mixture of 1% CS–MA–LA/PBS solution and 5 mM APS/
TEMED solution was subcutaneously injected into white
mice, each with 0.5 mL liquid. After implantation for 1d, 3d,
and 10d, the mice were sacrificed and anatomized to investigate the inflammatory reaction. The skins and hydrogels
were harvested for histological evaluation. The sections of
the skins and the hydrogels were stained by H&E.
Characterizations
FTIR spectra were recorded on a BRUKER VECOTR22 spectrometer. 1H NMR spectrum of CS–MA–LA was recorded
on an ANAVCE DMX500 with D2O as solvent working at
500 MHz. Elemental analysis was performed on an elemental analyzer (Flash EA-1112).
Statistical analysis
Data were analyzed using ANOVA. Significance was
determined at a value of p < 0.05.
RESULTS AND DISCUSSION
Synthesis and characterization of CS–MA–LA
Chitosan can be regarded as a copolymer of N-acetylglucosamine and N-glucosamine units randomly distributed throughout the molecular chain. It is dissolved only in acidic solution for its strong intermolecular hydrogen bonding. It contains abundant amino
groups, through which both polymerizable (e.g. acrylate) and water-soluble groups can be conveniently introduced. In the present work, MA and LA are sequentially grafted onto the chitosan chains via the combination between the carboxyl groups and the amino
and/or hydroxyl groups to yield water soluble and
polymerizable CS–MA–LA under the catalysis of
carbodiimide (Scheme 1). Since MA and LA are both
weak acids and chitosan can be directly dissolved in
their solutions, the reaction is easily proceeded without involvement of other acid. The byproducts of
small molecular weight and unreacted monomers
are then removed by dialysis.18–21
FTIR and 1H NMR characterizations confirmed the
structure of chitosan and its derivatives (Fig. 1). The
IR spectrum of chitosan [Fig. 1(a)] illustrates peaks
assigned to the saccharide structure at 1153, 1081,
Journal of Biomedical Materials Research Part A DOI 10.1002/jbm.a
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HONG ET AL.
TABLE I
The grafting ratios and maximum soluble pHs
of chitosan, CS–MA and CS–MA–LA
Sample
C/N Ratio
Grafting
Ratio (%)
Maximum
Soluble pHb
Chitosana
CS–MA
CS–MA–LA
6.44
7.20
7.51
78.00
19.00
10.33
<5
<7
<9
a
The deacetylation degree of chitosan was calculated
according to the C/N ratio.
b
Data represent the highest pH value at which the precipitation occur.
Scheme 1. Synthesis route and molecular structure of
CS–MA–LA.
1029, and 898 cm1.22,23 The peaks at around 1653
and 1598 cm1 are assigned to amide I band and NH2
group, respectively. Accompanying with the weakening of absorbance at 1598 cm1, new peaks at 1626
and 1574 cm1 emerge in the IR spectrum of CS–MA
[Fig. 1(b)], which should be assigned to the C¼
¼C
double bond and the amide II band, respectively.
This result demonstrates that MA has been successfully grafted. After grafting LA [Fig. 1(c)], the peak
intensity at 1574 cm1 is further increased in the IR
spectrum of CS–MA–LA. Moreover, a new peak at
1734 cm1 assigned to ester bond appears. This would
mean that the carboxylic acid groups of LA react not
only with amino groups, but also with hydroxyl
groups of chitosan. That considerable amount of esters
is formed after LA grafting is understandable, since at
Figure 1. FTIR spectra of (a) chitosan, (b) CS–MA, (c) CS–
MA–LA and (d) chitosan hydrogel. [Color figure can be
viewed in the online issue, which is available at www.
interscience.wiley.com]
Journal of Biomedical Materials Research Part A DOI 10.1002/jbm.a
this moment the absolute amount of NH2 groups has
been largely decreased by reaction with MA. Nevertheless, the IR spectrum has undoubtedly evidenced
the introduction of LA.
Characterization of CS–MA–LA under 1H NMR
confirms also its molecular structure. Chemical shifts
belonging to the saccharide structure are assigned as
follows: 1H NMR (D2O) d ¼ 2.79(H2), d ¼ 3.43–
3.91(H3–H6), d ¼ 1.91(NCOCH3).24 Chemical shifts
at d ¼ 5.64 and d ¼ 5.42 are assigned to H2C¼
¼C
(a) of MA, respectively. Chemical shifts at d ¼ 1.84
and d ¼ 1.20 are assigned to methyl groups of MA
(c) and LA (b), respectively. Moreover, MA and LA
substitution degrees are 19% and 10.33% as detected
by elemental analysis, respectively (Table I). All these
results have confirmed that both MA and LA have
been grafted onto the chitosan chains.
Water solubility
It was shown in Table I that the highest soluble pH
increased from 5 to 9 when the chitosan was grafted
with MA and LA. When the acidic solution of chitosan
was neutralized with NaOH, it was sedimentated at
approximately pH 5, implying that chitosan is not
soluble in water. After grafting of MA, the precipitation pH value was raised to 7. However, when immersed under pure water, this MA grafted chitosan
could be only swollen but not dispersed. By contrast,
the CS–MA–LA could be completely dissolved in pure
water with low viscosity. This has endowed the polysaccharide with great opportunity as injectable scaffold in situ. The substitution of MA and LA can alter
the regular molecular structure of chitosan, and thus
weaken the intermolecular hydrogen bonding. Yet
MA only is not able to yield grafting product with sufficient water solubility, since the H-bonding cannot be
adequately screened. It was reported previously that
the chitosan can become water-soluble when the deacetylation degree is *50%.25 Yet this deacetylation
degree is hardly controllable. A further introduction
of LA may not only make the chitosan molecular
structure more irregular, but also can enhance the
CHITOSAN HYDROGEL FORMED AT NEUTRAL pH AND BODY TEMPERATURE
917
H-bonding between chitosan and water molecules because of the existence of pendent hydroxyl group in LA.
water-soluble chitosan has the potential capability to
be used as injectable and in situ gelable scaffold.
Hydrogel formation
Gelation time
Different from the before-mentioned gelation systems for chitosan,2,14,15 chitosan hydrogel described
here is formed via crosslinking reaction between the
double bonds. For this to occur, a redox system is used
so that polymerization can be performed at body temperature. It has been identified that the APS/TEMED
initiation system is water-soluble and cytocompatible
and thus is used to initiate the polymerization of
PPF.26–28 Figure 2 illustrates the macroscopic gelation
process. After addition of APS/TEMED and incubation at 378C for a few minutes, liquid CS–MA–LA
solution [Fig. 2(a)] was transferred into transparent
chitosan hydrogel [Fig. 2(b)] which could sustain its
macroscopic shape [Fig. 2(c)]. In the FTIR spectrum
of chitosan hydrogel [Fig. 1(d)], the absorbance at
1626 cm1 (C¼
¼C double bonds) has disappeared,
indicating the occurrence of polymerization. This has
preliminarily demonstrated that the as-synthesized
As an injectable biomaterial in clinical application
it is required that the polymer solution is stable at
room temperature for a relatively long period, and
forms hydrogel at body temperature (378C) rapidly.
Therefore, the gelation time was investigated by varying the initiator concentration and the incubation temperature (Fig. 3).
The gelation time decreased rapidly along with
the increase of initiator’s concentration as shown in
Figure 3(a). When the APS/TEMED concentration
was set at 2.5 mM, the gelation time was longer than
30 min at 378C (the data was not shown in [Fig. 3(a)]).
When the concentration was set at 5 and 10 mM,
formation of the hydrogel required only *5.5 and
*1.5 min, respectively. At still higher concentration, e.g. 15 mM, the gelation was completed within
*30 s. It is understandable that with higher concentration of initiators larger amount of free radicals will
Figure 2. (a) 1% CS–MA–LA water solution, (b) gelation of (a) in 5 mM APS/TEMED at 378C, and (c) shape persistent
behavior of (b). Images were taken with a digital camera. [Color figure can be viewed in the online issue, which is available at www.interscience.wiley.com]
Journal of Biomedical Materials Research Part A DOI 10.1002/jbm.a
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HONG ET AL.
Figure 3. Gelation time of 1 wt % CS–MA–LA solution as a function of (a) APS/TEMED concentration at 378C, and (b)
incubation temperature with APS/TEMED concentration of 5 mM.
be created. Consequently the crosslinking polymerization can proceed in a relatively fast rate.
APS/TEMED initiating system is sensitive to temperature.27 Higher temperature can accelerate the generation and diffusion of free radicals, and also improve the motion ability of macromolecular chains. As
shown in Figure 3(b), at an APS/TEMED concentration of 5 mM the gelation time decreased from *20 to
*5.5 min when the temperature was raised from 25
to 378C. As an injectable hydrogel, an appropriate
gelation time is important. The property of the chitosan solution, i.e. rather stable at room temperature
while can be gelated at body temperature, is very promising for clinical application. The longer gelation time
at room temperature benefits for operation, while the
shorter gelation time at body temperature can prevent
from the liquid diffusion and favor the shape persistence. It is worth noting that no apparent temperature
increase was measured during the gelation process.
Polymerization of the chitosan macromonomers is
largely dependent on the encountering probability of
C¼
¼C double bonds. Crosslinking can take place only if
a macromolecular radical is close enough to another
C¼
¼C bond. The macromolecular chains have very low
moving ability and are confined within a limited spatial volume. Higher initiator’s concentration will create
more macromolecular radicals at a definite volume.
As a result, there will be a higher chance for the macromolecular radicals to react with other C¼
¼C bonds.
Hence, a higher crosslinking yield can be produced.
Cytotoxicity
To assess the toxicity of the chitosan hydrogel, 3T3
fibroblasts were cultured in medium of extractant.
First, the cytotoxicity under different cell seeding
Swelling behavior of the chitosan hydrogels
Figure 4 presents the swelling behavior of the chitosan hydrogels formed at 378C with different APS/
TEMED concentration. At a lower concentration of
APS/TEMED (<7.5 mM), the balanced swelling ratio
of the hydrogels decreased initially as a function of the
initiator’s concentration (p < 0.05). When the initiator’s
concentration was higher than 7.5 mM, the swelling
ratio was kept at a relatively low level without significant difference (p > 0.05). This feature matches inversely with the crosslinking yield of the hydrogels, e.g.
with higher crosslinking yield the swelling ratio is
lower. When the crosslinking density is higher, the
swelling of the hydrogel is largely restricted. Moreover, it is not strange that the crosslinking yield shows
a positive correlation with the initiator’s concentration.
Journal of Biomedical Materials Research Part A DOI 10.1002/jbm.a
Figure 4. Swelling ratio and crosslinking yield of chitosan
hydrogel as a function of APS/TEMED concentration. The
hydrogels were immersed under PBS at 378C. [Color figure
can be viewed in the online issue, which is available at www.
interscience.wiley.com]
CHITOSAN HYDROGEL FORMED AT NEUTRAL pH AND BODY TEMPERATURE
number was investigated (Fig. 5). After cultured for
24 h, the cytoviability measured by MTT assay in E5
medium and in DMEM control was compared as a
function of cell seeding number. The overall cell viability in both control and E5 medium increased along
with the cell seeding number till 5 104 (p < 0.05).
With still higher cell number, no significant difference
was found (p > 0.05). Compared with the negative
control, the cytoviability of E5 was significantly low
(p < 0.05) when the cell number was smaller than 2 104. Above this seeding number, no significant difference was detected. This alteration tendency is more
clearly illustrated by the preserved viability ratio as
shown in the inset of Figure 5. Hence, one can conclude
that the extractant from the hydrogel has negative effect
on the viability of 3T3 cells when the cell number is
small, i.e. some degree of cytotoxicity which should be
mainly attributed to the initiators. However, with
enough number of cells, the cytotoxicity introduced by
the initiators is very minimal and neglectable.
To further identify the cytotoxicity of the hydrogel,
cell culture with an initial seeding number of 1 105
in mediums of negative control, E5 and E10 was performed. As shown in Figure 6, the cytoviability of both
the E5 and the control was increased as a function of
culture time, indicating that the cells in these culture
mediums can normally proliferate. By contrast, the
cytoviability of sample E10 was steadily decreased as
a function of the culture time, implying that instead of
normal proliferation part of the seeding cells in this
medium should be dead. The profiles of cell viability
of E5 and control are very close, although at some culture intervals the cytoviability of E5 is still lower than
that of the control. Figure 7 compares the morphology
919
Figure 6. Cytoviability of 3T3 fibroblasts as a function of
culture time. Cell seeding number was 1 105/well.
[Color figure can be viewed in the online issue, which is
available at www.interscience.wiley.com]
of 3T3 fibroblasts after cultured in mediums of control
[Fig. 7(a,d)], E5 [Fig. 7(b,e)] and E10 [Fig. 7(c,f)] for 1d
[Fig. 7(a–c)] and 4d [Fig. 7(d–f)]. Confluent cell layers
have been formed for the control and the E5 since the
culture time of 1d. No apparent difference of cell morphology for the control and the E5 can be identified.
By contrast, a fewer cells were measured for the E10,
particularly after cultured for 4d. Since clinically the
cell number (>2 million/mL) is far beyond the highest
value used in this study, the hydrogel formed at an
initiator’s concentration of 5 mM can be roughly
regarded as nontoxic to cells, or at least that the toxicity can be neglected.
In vivo inflammatory reaction
Figure 5. Cytoviability of 3T3 fibroblasts as a function of
cell seeding number. Cells were cultured in (a) control
DMEM medium and (b) E5 medium. Inset is the preserved
ratio of cytoviability, which is calculated by [(ab)/a] 100%. [Color figure can be viewed in the online issue,
which is available at www.interscience.wiley.com]
The in vivo inflammatory reaction was assessed by
subcutaneous injection of a mixture of 1% CS–MA–
LA/PBS solution and 5 mM APS/TEMED solution
into white mice. All the mice survived throughout
the implantation period with normal performance.
No malignant infection, tissue necrosis, and abscess
were found in the implanted sites. Histological sections of the skin–hydrogel interfaces and the chitosan
hydrogels are shown in Figure 8. After implantation
for 1d, a large number of neutrophils (inflammation
cells) infiltrated through the skin–hydrogel interface
[Fig. 8(a)] and the hydrogel [Fig. 8(d)], implying that
acute inflammatory reaction occurred at this stage. At
day 3, the inflammation was aggravated. The skin–
hydrogel interface became loose, and the exudation
and the edema could be observed as well [Fig. 8(b)].
Meanwhile, histocytes infiltrated into the hydrogel
[Fig. 8(e)]. After implantation for 10d, the large numJournal of Biomedical Materials Research Part A DOI 10.1002/jbm.a
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HONG ET AL.
Figure 7. CLSM images to show morphology of 3T3 fibroblasts after cultured in mediums of control (a,d), E5 (b,e) and
E10 (c,f) for 1d (a–c) and 4d (d–f). Cell seeding number was 1 105/well. Viable cells were stained by FDA, thus exhibit
bright color. [Color figure can be viewed in the online issue, which is available at www.interscience.wiley.com]
ber of neutrophils had significantly decreased in the
interface [Fig. 8(c)] and the hydrogel [Fig. 8(f)]. Instead, a large number of histocytes appeared to form a
foreign body granuloma in the hydrogel, indicating
that the acute inflammation has shifted to the foreign
body reaction. As a general rule in evolution of inflammatory reaction, the granuloma will be finally assimilated to eliminate the inflammation. Thus the emergence of the granuloma is a positive sign indicating
that the inflammation caused by the hydrogel is only
temporary. In conclusion, although the hydrogel could
cause acute inflammation and foreign body reaction,
no tissue necrosis and malignant infection are evidenced in vivo, demonstrating that the material has
better histocompatibility.
CONCLUSIONS
Water-soluble and thermocrosslinkable chitosan is
successfully synthesized via sequentially grafting of
MA and LA under the catalysis of water-soluble carbodiimide. FTIR, 1H NMR, and elemental analysis
Journal of Biomedical Materials Research Part A DOI 10.1002/jbm.a
confirm the molecular structure and the substitution
degree of the as-synthesized CS–MA–LA. Gelation of
the CS–MA–LA is performed at very mild conditions
by using a redox initiation system, APS/TEMED. At a
concentration of 5/5 (mM) (APS/TEMED) the gelation
times are *20 and *6 min at 25 and 378C, respectively. No apparent temperature elevation is recorded
during the gelation process. The property of the chitosan solution, i.e. rather stable at room temperature
while can be gelated at body temperature, is very
promising for clinical application. The swelling ratio
of the chitosan hydrogels decreases along with the
increase of the APS/TEMED concentration initially,
then reaches a constant value. This matches inversely
to and thus can be explained by the crosslinking yield
of the hydrogels. 3T3 fibroblast culture demonstrates
that with sufficient number of cells, the cytotoxicity introduced by the initiators is very minimal and neglectable. Since clinically the cell number (>2 million/mL)
is far beyond the highest value used in this study
(1 105/200 mL), the hydrogel formed at an initiator’s concentration of 5 mM can be roughly regarded
as nontoxic to cells, or at least the toxicity can be
CHITOSAN HYDROGEL FORMED AT NEUTRAL pH AND BODY TEMPERATURE
921
Figure 8. Histological evaluation of the skin–hydrogel interfaces (a–c) and the chitosan hydrogels (d–f) after subcutaneous injection of a mixture of 1% CS–MA–LA/PBS solution and 5 mM APS/TEMED solution into white mice for 1d (a,d),
3d (b,e), and 10d (c,f). N and H represent the neutrophils and histocytes, respectively. Magnification of left side and right
side in each image is 100 and 400, respectively. [Color figure can be viewed in the online issue, which is available at
www.interscience.wiley.com]
neglected. In vivo test shows that the hydrogel system can induce foreign body reaction with sufficient
histocompatibility. The study has demonstrated that
the as-synthesized water-soluble chitosan has the potential capability to be used as injectable and in situ gelable scaffold.
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