Journal of Controlled Release 331 (2021) 1–6
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Journal of Controlled Release
journal homepage: www.elsevier.com/locate/jconrel
Facile preparation of multi-stimuli-responsive degradable hydrogels for
protein loading and release
Syuuhei Komatsu a, Moeno Tago a, Yu Ando a, Taka-Aki Asoh b, Akihiko Kikuchi a, *
a
b
Department of Materials Science and Technology, Tokyo University of Science, 6-3-1 Niijuku, Katsushika-ku, Tokyo 125-8585, Japan
Department of Applied Chemistry, Osaka University, 2-1 Yamadaoka, Suita, Osaka 565-0871, Japan
A R T I C L E I N F O
A B S T R A C T
Keywords:
Hydrogels
Multi-stimuli-responsive
Thermoresponsive
Redox-responsive
Protein release
Functional materials that can recognize the tumor microenvironment, characterized by acidic or reducing
conditions, are needed for the designing of drug delivery carriers for cancer treatment. Hydrogels are potential
protein drug carriers because they contain a large amount of water and stimuli-responsive functions can easily be
introduced in them. However, it is difficult to introduce multi-stimuli-responsive functions and degradability at
the same time. Here, we synthesized thermo- and pH-responsive hydrogels via a coupling reaction between poly
(ethylene glycol) diglycidyl ether (PEGDE) and cystamine (CA). The prepared hydrogels showed lower critical
solution temperature-type thermoresponsive behavior and pH-responsive swelling changes due to the proton­
ation of secondary and/or tertiary amino groups arising from the crosslinking agent CA. Under reducing con­
ditions, the hydrogels were degraded via the thiol exchange reaction in the presence of dithiothreitol or
glutathione. The loading and release properties of FITC-labeled model proteins from the hydrogels were inves­
tigated. The loaded amount of the protein increased with decreasing molecular weight or hydrodynamic radius,
which is based on the size of the network structure of the hydrogels. Notably, loaded proteins in the hydrogels
were released only under reducing conditions, which mimic the tumor microenvironment. Thus, the prepared
multi-responsive degradable hydrogels are expected to be used as functional drug delivery carriers for cancer
treatment.
1. Introduction
Biopharmaceuticals, including proteins, peptides, and oligonucleo­
tides, which are derived from living organisms, such as cells, viruses,
and bacteria, are used for the treatment of potentially life-threatening
diseases, such as cancer [1–4]. However, these drugs have potential
risks of denaturation or loss of activity in the living body before reaching
the target site. Therefore, drug delivery system (DDS) carriers, such as
hydrogels, are required to suppress denaturation and release drug
molecules at the target site.
Proteins, peptides, and oligonucleotides loaded into hydrogels do not
show denaturation, as hydrogels have a 3D network structure and a
relatively high water content, similar to that of soft biological tissues
[5–8]. Moreover, properties such as response to various stimuli, such as
pH [9] and temperature [10], as well as biodegradability [11], can be
incorporated into hydrogels in their molecular design. To date, various
physical and/or chemical stimuli-responsive hydrogels for controlled
drug release have been reported [9,10,12–14]. Wang et al. reported
thermo-, light-, redox- and guest molecule-responsive hydrogels pre­
pared via a multi-step reaction and complex formation with guest
molecules [15]. Qu et al. reported redox-, thermo-, and pH-responsive
hydrogels based on poly{(N-isopropylacrylamide)-co-vinyl(γ-benzyl-Lglutamate)} crosslinked by N,N′ -bis(acryloyl)cystamine [16]. However,
the preparation of these hydrogels required multi-step reactions and was
rather complicated. In addition, although various stimulus-responsive
hydrogels have been reported, it is important to design hydrogels that
are responsive to a stimulus that can be acquired in a biological envi­
ronment. Thermo- and pH-responsive hydrogels are widely studied as
DDS carriers, as these hydrogels are responsive to the body environ­
ments; thus, such stimuli are utilized to modify the drug loading and
release behaviors of hydrogels through mesh size control and hydro­
phobic/electrostatic interactions of drugs with the network polymers of
the hydrogels [17,18].
Cancer microenvironments are known to be slightly acidic with a
high glutathione (GSH) concentration of approximately 10 mmol L− 1,
compared with that in normal tissues (~3 mmol L− 1). Thus, hydrogels
* Corresponding author.
E-mail address: kikuchia@rs.tus.ac.jp (A. Kikuchi).
https://doi.org/10.1016/j.jconrel.2021.01.011
Received 31 August 2020; Received in revised form 6 January 2021; Accepted 7 January 2021
Available online 9 January 2021
0168-3659/© 2021 Elsevier B.V. All rights reserved.
S. Komatsu et al.
Journal of Controlled Release 331 (2021) 1–6
having responsive properties to local pH and reduction conditions may
have potential applications as carriers for drug delivery to the cancer
environment [19–22].
We have previously prepared redox-responsive trisoligo(ethylene
glycol) (trisOEG) hydrogels [23]. TrisOEG hydrogels are made of threearmed, low-molecular-weight oligo(ethylene glycol) crosslinked with
disulfide bonding, which allows regulation of the mesh size to protect
proteins from degradation. Once the hydrogels are degraded with the
reducing agent dithiothreitol (DTT), rapid protein release is achieved. In
addition, we have reported a thermo- and redox-responsive polymer
made of trisOEG interconnected with disulfide bonding [24]. The ther­
moresponsive property depends on the hydrophobic/hydrophilic bal­
ance and/or chemical structure of the polymers [25–27]. The solubility
of the prepared polymers was tuned through alterations in the chemical
composition of trisOEG and cystamine (CA). However, the preparation
of multiple stimuli-responsive hydrogels that can maintain the proper­
ties of degradation is difficult and complicated with multi-step reactions
[28,29]. To the best of our knowledge, there is currently no report on a
one-pot preparation of multi-stimuli-responsive hydrogels capable of
controlled release and rapid excretion after degradation. Therefore,
facile preparation of multi-stimuli-responsive hydrogels that can
degrade into hydrophilic oligomers in a redox-responsive manner for use
as DDS carriers would be highly desirable.
In this study, we aimed to prepare multi-stimuli-responsive (thermo-,
pH-, and redox-responsive) hydrogels in one pot through an epoxy ringopening reaction between poly(ethylene glycol) diglycidyl ether
(PEGDE) and CA to introduce redox-responsive disulfide bonds and pHresponsive amino groups in the hydrogels. The prepared hydrogels were
characterized and their physicochemical properties were measured. In
addition, we investigated the potential of using the prepared hydrogels
as a drug delivery carrier. Specifically, we loaded and immobilized
proteins of various sizes into the hydrogels and determined the drug
release profiles. The prepared hydrogels are expected to serve as po­
tential DDS carriers for use in cancer treatment.
Fig. 1. Schematic illustration
responsive hydrogels.
of
the
preparation
of
multi-stimuli-
degradable hydrogels were also prepared by the same method using
PEGDE with 1,6-hexamethylenediamine (Supplementary Information).
2.3. Swelling behavior of the hydrogels
The disk gels were thoroughly soaked in distilled water at 50 ◦ C for
24 h. Thereafter, the disk gels were immersed in 500 mL of PBS (pH 7.4, I
= 0.15 mol L− 1) or phosphate buffer (PB) (pH 5.0–8.4, I = 0.15 mol L− 1)
for 24 h, and the gel was swollen with each solution at 10–45 ◦ C for 24 h.
The gels were removed, excess water on the surface of the gels was
sucked with Bemcot® (Asahi KASEI, Tokyo, Japan), and the weight of
the swollen disk gels was measured. Subsequently, the swollen gels were
lyophilized overnight to obtain dry gels. The swelling ratio (SR) of the
disk gels was calculated from the weight of the swollen gel (Ws) and dry
gel (Wd) using eq. (1):
SR = (Ws − Wd )/Wd
(1)
2. Materials and methods
2.4. Degradation of the hydrogels under reducing conditions
2.1. Materials
Degradation of the hydrogels was investigated in the presence of a
reductant, DTT in PBS solution, in an accelerated test. The disk gel with
equilibrium swelling at 25 ◦ C in PBS (pH 7.4, I = 0.15 mol L− 1) was
immersed in 50 mL PBS containing 3 mmol L− 1 DTT at 37 ◦ C. At pre­
determined time intervals (20 min), the hydrogel disk was removed, and
the residual weight was measured after excess medium was removed
with Bemcot®.
CA dihydrochloride, sodium chloride (NaOH), and DTT were pur­
chased from FUJIFILM Wako Pure Chemical Corporation (Osaka,
Japan). PEGDE (Mn 500), bovine serum albumin (BSA) fluorescein iso­
thiocyanate conjugate (FITC-BSA) (M = 66,000, pI 4.8), FITC-labeled
polyclonal rabbit anti-human lysozyme (FITC-lysozyme) (M = 14,300,
pI 11.1), and FITC-labeled insulin from bovine pancreas (FITC-Insulin)
(M = 5700, pI 5.4) were purchased from Sigma-Aldrich (MO, USA).
2.5. Preparation of protein-loaded hydrogels and determination of drug
release profiles
2.2. Preparation of multi-stimuli-responsive hydrogels
Protein-loaded hydrogels were prepared based on the pH- and
temperature-responsive behaviors of the prepared hydrogels. First, the
gels were swollen in PBS solution at 5 ◦ C. Thereafter, each gel disk was
immersed in 5 mL PBS solution containing 0.025 mmol L− 1 FITC-labeled
protein (FITC-Lysozyme, FITC-BSA, or FITC-Insulin) and allowed to
stand for 2 days at 5 ◦ C with mild shaking in a thermostated shaking bath
(Neslab RTE7; Thermo Fisher, Tokyo, Japan). Subsequently, the disk gel
was heated to 37 ◦ C, and the gel was allowed to stand for 3 days to retain
proteins inside the hydrogels. Next, the protein-loaded disk gel was
immersed in PBS and washed by replacing PBS every day for 3 days.
Finally, the disk gel was immersed in PBS containing 3 mmol L− 1 DTT at
37 ◦ C and completely decomposed to measure the loading amount of
protein per disk gel. The release of FITC-labeled proteins from hydrogels
was evaluated at 37 ◦ C in 500 mL PBS containing 3 mmol L− 1 DTT at
37 ◦ C (pH 7.4, I = 0.15) as the external solution. In this assay, three disks
of protein-loaded hydrogels were placed in a dissolution test apparatus
(Dissolution Tester PJ-12 N; Miyamoto Riken Ind. Co., Ltd., Osaka,
Japan), according to the rotating basket method described in the 16th
Multi-stimuli-responsive hydrogels were prepared via coupling re­
action between PEGDE and CA dihydrochloride (Fig. 1). PEGDE was
dissolved in ultrapure water, and CA dihydrochloride was dissolved in
an aqueous NaOH solution (CA:NaOH = 1:2, mol ratio). NaOH aqueous
solution was added to neutralize the dihydrochloride and isolate CA, and
ultrapure water was added to PEGDE and CA to achieve PEGDE con­
centration of 20 wt% to 50 wt%. The PEGDE solution was added drop­
wise to the CA solution with stirring until complete mixing was
achieved. The prepared pre-gel solution was injected between two glass
plates backed with polypropylene films and gasketed with a 0.5 mmthick polydimethylsiloxane spacer at 25 ◦ C for 24 h. The obtained bulk
hydrogels were immersed in 10 mmol L− 1 NaOH solution to reduce
unreacted epoxy groups by a ring-opening reaction at 25 ◦ C for 24 h.
Subsequently, the gel sheet was purified by immersing it in ultrapure
water for 5 days and in phosphate-buffered saline (PBS) solution (Ca2+and Mg2+-free, pH 7.4, I = 0.15 mol L− 1) for 2 days. It was then cut to
disk-type gels with a diameter of 13 mm using a cork borer. Non2
S. Komatsu et al.
Journal of Controlled Release 331 (2021) 1–6
Japanese Pharmacopeia. At predetermined time intervals (15 min), 5
mL of external sample solution was retrieved, and 5 mL of PBS was
added. The amount of released FITC-labeled proteins was determined
using a fluorophotometer (FP-6500; JASCO, Tokyo, Japan) with exci­
tation and emission wavelengths of 495 nm and 520 nm, respectively,
and a slit width of 2 nm.
3. Results and discussion
3.1. Preparation of multi-stimuli-responsive hydrogels
Multi-stimuli-responsive hydrogels were prepared by the ringopening reaction of the epoxide groups in PEGDE with the amino
groups of CA as the crosslinking agent (Fig. 1). Tertiary amino groups
and disulfide bonds were introduced into the three-dimensional network
structure of the hydrogels.
Table 1 shows the phase diagram of gelation with respect to the feed
ratio of PEGDE and CA as well as PEGDE concentration. Both the
PEGDE/CA ratio and PEGDE concentration strongly affected the gela­
tion behavior. No gelation occurred at the low PEGDE/CA ratio of
1.0:1.0 and 1.2:1.0 (mol ratio), regardless of the PEGDE concentration.
As CA that has reacted once with the epoxy groups of PEGDE is a sec­
ondary amine, it can react with another epoxy group to form the
network structure of the hydrogels. A relative increase in the amount of
CA induced insufficient network formation, resulting in weak or no
gelation after the coupling reaction. Using a PEGDE concentration of
30–50 wt% and crosslinker PEGDE:CA ratio of 1.5:1.0 (mol ratio) or
above resulted in the successful formation of the hydrogels. In subse­
quent experiments, we mainly characterized the hydrogels prepared
with 30 wt% PEGDE concentration and crosslinker ratio of PEGDE:CA =
2.0:1.0 (mol ratio), as these hydrogels showed sufficient mechanical
properties for handling and for measurement of physicochemical
properties.
Fig. 2. pH-responsive swelling changes in the prepared hydrogels (PEGDE:CA
= 2.0:1.0, mol ratio; 30 wt% PEGDE). a) Macroscopic observations of the pHresponsive swelling (pH 5.4) and shrinking (pH 7.4) behaviors of the hydro­
gels prepared at 37 ◦ C. Scale bars, 10 mm. b) pH-dependent change in the
swelling ratio of the hydrogels prepared at 37 ◦ C (PEGDE:CA = 2.0:1.0, mol
ratio; 30 wt% PEGDE). Data are expressed as mean with standard deviation (n
= 3).
3.2. pH- and temperature-dependent swelling behavior of the hydrogels
The as-prepared hydrogels have tertiary or secondary amino groups
at the crosslinking point with PEGDE. The hydration property of the
hydrogels is expected to change in response to pH changes. Fig. 2 shows
the pH-dependent swelling behavior of the prepared hydrogels at 37 ◦ C.
As shown in Fig. 3a, the hydrogels were deswollen and shrunken at pH
7.4, compared with those at pH 5.4. As the amino groups in the
hydrogels are protonated at low pH and protonated amino groups repel
each other, more water molecules would have penetrated the hydrogels.
Next, the pH-dependent change in swelling was evaluated (Fig. 2b). The
SR of the hydrogels decreased with increasing pH. The pKa value of the
hydrogels was 6.21, as determined by acid-base titration of the hydro­
gels (Fig. S1). In acidic conditions (pH 5.0–6.2), the tertiary amino
groups derived from CA were protonated, and thus the SR increased. In
contrast, under neutral to alkaline conditions (pH 6.8–8.4), deprotona­
tion of amino groups in the hydrogels caused water molecules to be
expelled from hydrogels, eventually causing the gels to shrink. Inter­
estingly, at pH 6.5, the hydrogels showed a sharp deswelling behavior. A
large and distinct change in SR occurred just above the pKa value,
Table 1
Preparation condition of hydrogels.
PEGDE:CA
(mol:mol)
Concentration of PEGDE (wt%)
20
30
40
50
2.0: 1.0
1.7: 1.0
1.5: 1.0
1.2: 1.0
1.0: 1.0
Y
Y
N
N
N
Y
Y
Y
N
N
Y
Y
Y
N
N
Y
Y
Y
N
N
Fig. 3. Thermoresponsive changes in the swelling ratio of the prepared
hydrogels (PEGDE:CA = 2.0:1.0, mol ratio; 30 wt% PEGDE). a) Macroscopic
findings of the thermoresponsive swelling (10 ◦ C) and shrinking (45 ◦ C) be­
haviors of the prepared hydrogel at pH 7.4. Scale bars, 10 mm. b) Temperaturedependent change in the swelling ratio of the hydrogels at pH 7.4. Data are
expressed as mean with standard deviation (n=3).
Y: gelation occurred, N: no gelation occurred.
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S. Komatsu et al.
Journal of Controlled Release 331 (2021) 1–6
probably owing to the charge balance of the protonated and deproto­
nated amino groups within the hydrogels. This result suggests that a
proton sponge effect may occur in the hydrogels owing to the differences
in pH inside (pH 7.4) and outside (pH 5.4) the cells or near cancer cells
[28].
In our previous work, trisOEG crosslinked with CA showed ther­
moresponsive deswelling behavior around normal body temperature
[22]. The reason might be that the trisOEG is crosslinked with CA threedimensionally, which limits the hydration behavior of the trisOEG seg­
ments. Linear poly(ethylene glycol) (PEG) is well known to show
dehydration and lower critical solution temperature-type thermores­
ponsive behaviors owing to modification of both termini by small hy­
drophobic moieties [29]. In addition, the branched chemical structure
affects the thermoresponsive behavior; the branched polymer shows
decreased hydration compared with the linear polymers, owing to lower
molecular mobilities [24]. Thus, the branched polymer shows both
thermoresponsive behavior and dehydration. Accordingly, the prepared
hydrogels are expected to show thermoresponsive dehydration. Fig. 3
shows the temperature-dependent swelling-deswelling behavior of the
prepared hydrogels. Fig. 3a shows the macroscopic images of the
hydrogels immersed in a water bath at 10 ◦ C and 45 ◦ C until equilibrium
swelling was reached. The diameter of the gel disk was 12.5 mm at
10 ◦ C, but decreased to 11 mm at 45 ◦ C. Thus, the prepared hydrogels
showed temperature-dependent swelling behavior. We next evaluated
temperature-dependent changes in the SR in the range of 5–45 ◦ C, and
the results are shown in Fig. 3b. The SR of the hydrogels gradually
decreased with increasing temperature, especially around 30 ◦ C at pH
7.4. The hydrogels were composed of a hydrophilic, short-chain-length
PEG, in which both termini were connected with CA, limiting the mo­
lecular mobility of the PEG segments. Such mobility limitation of the
PEG apparently reduced the hydration of the PEG chains, resulting in
temperature-dependent deswelling of the hydrogels. As the PEG seg­
ments within the prepared hydrogels maintained hydrogen bonding
with water molecules even at 45 ◦ C, deswelling behavior was gradual
and did not show volume-phase transition, as is often observed in poly
(N-isopropylacrylamide) (PNIPAAm)-based hydrogels [10]. The above­
mentioned results indicate that the prepared hydrogels composed of
PEGDE and CA showed swelling changes in response to alterations in
external pH and temperature [30,31].
Fig. 4. Degradation of hydrogels (PEGDE:CA = 2.0:1.0, mol ratio; 30 wt%
PEGDE) via reduction of disulfide bonds in the hydrogel network structure. a)
Macroscopic findings of the hydrogels during degradation from 0 to 200 min.
Scale bars, 10 mm. b) Percentage of remaining weight and swelling ratio of the
hydrogels during degradation from 0 to 200 min at 37 ◦ C. Data are expressed as
mean with standard deviation (n=3).
surface. After 80 min, the SR increased with increasing degradation time
owing to the expansion of the hydrogel network through degradation.
After complete degradation, the molecular weight of the decomposed
products was examined using gel permeation chromatography
(Fig. S3a). The average molecular weight was below 1000, which is the
molecular weight of PEGDE with CA after degradation at both chain
ends. Moreover, the decomposition products showed no thermores­
ponsive property in PBS owing to the loss of balance of hydrophobicity/
hydrophilicity in the polymer chain (Fig. S3b). It would thus be excreted
by the renal system, as the resultant materials are water-soluble, lowmolecular-weight materials [32]. Cancer cells produced a high amount
of the reducing substance GSH of approximately 10 mmol L− 1 both in­
side and outside the cells (GSH concentration in normal tissues: 1–2
mmol L− 1) [20]. In preliminary experiments, we observed increased SR
of the hydrogel sheet incubated with cultured HeLa cells for several
days. This is probably due to the reduction of disulfide bonds caused by
the GSH present in the environment (data not shown). Therefore,
hydrogels that are responsive to reducing agents can be used as anti­
cancer drug carriers for the prevention of tumor recurrence after sur­
gical treatment.
3.3. Degradation of hydrogels under reducing conditions
We then investigated the degradation behavior of the prepared
hydrogels in 3 mmol L− 1 DTT containing PBS in an accelerated test
(Fig. 4). Under reducing conditions, the hydrogels showed gradual size
change at 40 min, followed by complete degradation after 200 min of
incubation in 3 mmol L− 1 DTT containing PBS through disruption of the
disulfide bonds in the CA units of the hydrogels (Fig. 4a). After 40 min,
the diameter of the prepared hydrogels decreased from 12.0 mm to 10.5
mm via degradation. In the absence of DTT, the hydrogels did not show
any degradation, and the disk shape remained constant. Moreover, the
hydrogels made of PEGDE crosslinked with hexamethylenediamine
(which had no disulfide bond) instead of CA showed no degradation
(Fig. S2). These results suggest that the degradation of the hydrogels
depended on the cleavage of disulfide bonds derived from CA in the
network structure.
Fig. 4b shows the time-dependent changes in swelling and the
remaining weight of the hydrogels that underwent reduction reaction in
PBS solution containing 3 mmol L− 1 DTT. The gels gradually collapsed
and became water-soluble at approximately 3 h, as the remaining weight
decreased linearly along with gel degradation during this period.
Moreover, hydrogel SR and degradation decreased for up to 40 min, and
the hydrogels reswelled after 80 min (remaining weight 60%). Between
40 and 80 min, the SR showed a constant value, whereas the remaining
weight changed from 80% to 60% at the same time intervals. This result
suggested that the hydrogels degraded starting from near the hydrogel
3.4. Preparation of protein-loaded hydrogels
As discussed in the previous section, the prepared hydrogels may be
utilized as drug delivery carriers. Therefore, we investigated the loading
of proteins into the hydrogels. Fig. 5 shows the appearance and loaded
amount of proteins of various molecular sizes within the prepared
hydrogels. Figs. 5a-c show the appearance of FITC-labeled proteinloaded disk gels, which were generated by utilizing the thermores­
ponsive swelling and shrinking behaviors of the hydrogels. The hydro­
gels were swollen in PBS solution at 5 ◦ C, at which protein molecules
were incorporated within the hydrogels. The hydrogels turned from
clear transparent to yellow after 2 days of incubation at 5 ◦ C, indicating
the introduction of FITC-labeled proteins within the hydrogels (Fig. 5a).
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Journal of Controlled Release 331 (2021) 1–6
Fig. 5. Protein loading based on the temperature- and pH-responsive properties of the hydrogels. a) Macroscopic image of each protein-loaded hydrogel at 5 ◦ C for 2
days. b) Macroscopic image of each protein-loaded hydrogel at 37 ◦ C for 5 days. c) Macroscopic image of the complete degradation of each protein-loaded hydrogel.
All scale bars in a)-c) are 10 mm. d) Loaded amount of each protein in the hydrogels. Data are expressed as mean with standard deviation (n = 3).
Each protein molecule was diffused into the swollen hydrogels. Upon
increasing the temperature to 37 ◦ C, the protein-loaded hydrogels
shrunk and maintained a yellow color, indicating that the drug mole­
cules were entrapped inside the gels (Fig. 5b). Finally, these hydrogels
were immersed in PBS containing 3 mmol L− 1 DTT at 37 ◦ C and
decomposed to measure the amount of protein encapsulated per disk gel
(Fig. 5c, d).
The above results (Fig. 5d) indicated that the loaded amount of each
protein increased with decreasing molecular weight and hydrodynamic
diameter of the proteins: insulin [33] (MW 5700, hydrodynamic diam­
eter (Dh): 3.0 nm), lysozyme [34] (MW 14,300, Dh: 3.8 nm), and BSA
[35] (MW 66,000, Dh:7.2 nm). According to the molecular sizes of the
hydrogel components, the calculated mesh size of the hydrogel was 5.0
nm × 1.2 nm, with an ideal reaction that allowed smaller-molecularweight proteins, such as insulin and lysozyme, to be entrapped within
the hydrogels, as these proteins have similar or smaller hydrodynamic
diameter as the mesh size. In contrast, BSA is larger than the mesh size of
the hydrogels; thus, only small amounts of BSA were loaded into the
hydrogels.
above the pI values of the amino groups. In contrast, under reducing
conditions, 100% BSA release was observed after 1 h because of the
decomposition of the hydrogels. The BSA structures before loading and
after release from the hydrogels were determined by circular dichroism
(Fig. S4). These CD spectra showed the same curves between 200 and
250 nm, suggesting that the protein structure was not affected by the
reducing environment of DTT, pH, or temperature in the loading and
releasing experiments. The release profiles of lysozyme and insulin in
reducing conditions showed the same tendency as that of BSA (Fig. 6b,
c). In PB, the release percentages of lysozyme and insulin after 24 h were
approximately 20%, even though the hydrogels were in the swollen and
expanded state in acidic conditions through protonation of amino
groups. Assuming uniform gelation, the mesh size of the gel was
calculated to be 5.0 nm × 1.2 nm. Therefore, the diffusion of lysozyme
and insulin from the hydrogels might have occurred slowly, which ex­
plains why only 20% release was observed after 24 h. Together, these
results indicated that complete release of the proteins inside the gels
occurred only under reducing conditions.
Based on their thermal- and pH-responsive behaviors, the prepared
multi-stimuli-responsive hydrogels are expected to be DDS carriers onto
which proteins can be loaded and then released in a controlled manner
via biodegradation due to their redox-responsive behavior in the living
body. The pH in the tumor tissue microenvironment is slightly lower
than the physiological pH, while GSH concentration may be increased
than that in the normal tissue environment. Thus, the drug-containing
bulk hydrogel sheet can be placed at the tumor site just after excision
of tumor tissues to eliminate the remaining cancer cells.
3.5. Stimuli-responsive protein release
The release profiles of BSA, lysozyme, and insulin from the hydrogels
in PBS containing 3 mmol L− 1 DTT (pH 7.4) or PB (pH 5.4) are sum­
marized in Fig. 6. Fig. 6a shows the BSA release profiles. In PB (pH 5.4),
10% BSA release was observed after 24 h, despite the swelling and
network expansion due to electrostatic interactions between the pro­
tonated hydrogels and BSA (pI 4.8). Moreover, BSA was not released in
PBS (pH 7.4) because of the smaller mesh size of the shrunken hydrogels
Fig. 6. Release profiles of proteins from hydrogels at 37 ◦ C. a) BSA release: circle plot, in PBS containing 3 mmol L− 1 DTT (pH 7.4); triangle plot, in PB (pH 5.4);
diamond plot, in PBS (pH 7.4). b) Lysozyme release: circle plot, in PBS containing 3 mmol L− 1 DTT (pH 7.4); triangle plot, in PB (pH 5.4). c) Insulin release: circle
plot, in PBS containing 3 mmol L− 1 DTT (pH 7.4); triangle plot, in PB (pH 5.4). All data are expressed as mean with standard deviation (n=3).
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Journal of Controlled Release 331 (2021) 1–6
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In this study, one-pot synthesis of multi-responsive degradable
hydrogels was achieved via the coupling reaction between PEGDE and
CA. The prepared hydrogels showed thermoresponsive behavior in
aqueous media, with changes in the SR. In addition, the SR increased
with decreasing pH owing to the protonation of the amino groups in the
hydrogel network. The hydrogels showed degradability under reducing
conditions, which was due to thiol exchange reaction. With increased
incubation time under reducing conditions, these hydrogels swelled and
collapsed. We further investigated the FITC-BSA-, FITC-Lysozyme-, and
FITC-Insulin-loading and release profiles of the hydrogels. The proteinloaded hydrogels were obtained based on the swelling and shrinking
behaviors of the hydrogels, which were attributed to their thermores­
ponsive property. The amount of loaded protein increased with
decreasing molecular weight or hydrodynamic radius, as the diffusion of
these molecules depended on the mesh size of the hydrogels. Notably,
the FITC-labeled proteins were released only under reducing conditions
owing to hydrogel degradation. Moreover, in acidic (pH 5.4) and neutral
(pH 7.4) conditions, these proteins remained almost intact in the
hydrogels, despite the swelling and network expansion, owing to the
combination of electrostatic interaction and size of the hydrogel
network structure. Accordingly, the synthesized multi-responsive
degradable hydrogels are expected to serve as DDS carriers that can be
used for controlled drug release in cancer treatment. In particular, the
drug-containing bulk hydrogel sheet can be placed at the tumor site just
after excision of tumor tissues to eliminate the remaining cancer cells.
Credit author statement
• The corresponding author is responsible for ensuring that the de­
scriptions are accurate and agreed by all authors.
• T-AA and AK designed researches, SK, MT, and YA conducted re­
searches, SK, MT, YA, T-AA, and AK analyzed data, and SK, T-AA and
AK wrote the manuscript.
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