Establishment of a controlled insulin delivery system using a glucose-responsive double-layered nanogel

Article abstract of DOI:10.1039/c4ra16656f

Glucose-responsive insulin delivery systems have been proposed as a promising alternative to conventional intramuscular administration methods, which causes low patient compliance due to the requirement of multiple administration. In addition, protein-based glucose-responsive systems using glucose oxidase and lectin have not achieved success in clinical trials because of their low biostability and potential cytotoxicity. In order to overcome these issues, the phenylboronic acid (PBA)-derivatives converted to hydrophilic moieties with an elevated glucose level play a key role in controlled insulin delivery systems due to their better biostability and high biocompatibility. In order to endow glucose-responsiveness to insulin delivery carriers using PBA derivatives, a glycol chitosan (GC)/sodium alginate (SA)-poly(L-glutmate-co-N-3-L-glutamylphenylboronic acid) (PGGA) graft polymer double-layered nanogel is synthesized by N-carboxyanhydride (NCA) polymerization and carbodiimide coupling reactions. The GC/SA-PGGA double-layered nanogel controllably releases insulin at diabetic glucose levels in vitro, and shows high biocompatibility, determined by cell viability and a hemolysis assay. Moreover, controlled insulin release at high glucose levels can be accomplished using the GC/SA-PGGA double-layered nanogel in mouse studies. Therefore, the GC/SA-PGGA double-layered nanogel characterized by glucose-sensitivity and superior biocompatibility may act as a potential platform for advanced insulin delivery systems.

Full text of DOI:10.1039/c4ra16656f

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Establishment of a controlled insulin delivery
system using a glucose-responsive double-layered
Cite this: RSC Adv., 2015, 5, 14482
nanogel†
a
b
a
a
c
DaeYong Lee, Kibaek Choe, YongJun Jeong, Jisang Yoo, Sung Mun Lee,
d
b
a
Ji-Ho Park, Pilhan Kim and Yeu-Chun Kim*
Glucose-responsive insulin delivery systems have been proposed as a promising alternative to conventional
intramuscular administration methods, which causes low patient compliance due to the requirement of
multiple administration. In addition, protein-based glucose-responsive systems using glucose oxidase
and lectin have not achieved success in clinical trials because of their low biostability and potential
cytotoxicity. In order to overcome these issues, the phenylboronic acid (PBA)-derivatives converted to
hydrophilic moieties with an elevated glucose level play a key role in controlled insulin delivery systems
due to their better biostability and high biocompatibility. In order to endow glucose-responsiveness to
insulin delivery carriers using PBA derivatives,
a
glycol chitosan (GC)/sodium alginate
(SA)-poly(L-glutmate-co-N-3-L-glutamylphenylboronic acid) (PGGA) graft polymer double-layered
nanogel is synthesized by N-carboxyanhydride (NCA) polymerization and carbodiimide coupling
reactions. The GC/SA-PGGA double-layered nanogel controllably releases insulin at diabetic glucose
levels in vitro, and shows high biocompatibility, determined by cell viability and a hemolysis assay.
Moreover, controlled insulin release at high glucose levels can be accomplished using the GC/SA-PGGA
double-layered nanogel in mouse studies. Therefore, the GC/SA-PGGA double-layered nanogel
characterized by glucose-sensitivity and superior biocompatibility may act as a potential platform for
advanced insulin delivery systems.
Received 18th December 2014
Accepted 23rd January 2015
DOI: 10.1039/c4ra16656f
www.rsc.org/advances
patients, there has been an ongoing impetus to create alterna-
tive routes of therapeutic administration that are non-invasive
1
. Introduction
Insulin, a peptide hormone produced by Langerhans beta cells and convenient, such as intradermal, oral, nasal and drug
4
–6
in the pancreas, is essential to control a normal glucose level in delivery system. However, as of yet, most of these recently
the blood by activating the conversion of glucose to larger devised alternative drug delivery routes have been incapable of
1
molecules such as glycogen and fatty acids. Diabetic patients being implemented into clinical practice due to several tech-
3
lose the capability of producing insulin or disorderly secrete nological limitations. The concept of stimuli-responsive drug
insulin from the pancreas, caused by innate genetic defects or delivery systems have recently gained much interest because of
2
7
excessive intake of carbohydrates or fat. For the regulation of its distinct controlled drug release characteristic. Such systems
an abnormal glucose level in the blood in diabetic patients, can respond to external stimuli, such as pH, temperature,
insulin is generally administrated by intravascular injection on redox-potential and glucose, by converting their chemical or
a frequent basis, causing great inconvenience and low patient physical properties, thereby establishing a framework for self-
1,3,4
7
compliance.
In order to resolve these issues for diabetic regulatory drug release. Glucose-responsive systems have
been widely used in various biomedical applications such as
controlled insulin delivery and glucose concentration diag-
a
Department of Chemical and Biomolecular Engineering, Korea Advanced Institute of
Science and Technology (KAIST), Daejeon 305-701, Republic of Korea. E-mail:
7,8
nosis.
The schematic architecture of currently existing
glucose-responsive systems can be narrowed down to three
dohnanyi@kaist.ac.kr
b
Graduate school of Nanoscience and Technology, Korea Advanced Institute of Science main categories: glucose oxidase, lectin, and phenylboronic
8
and Technology (KAIST), Daejeon 305-701, Republic of Korea
acid (PBA)-modied systems. Glucose oxidase is an enzyme
c
Department of Biomedical Engineering, Khalifa University of Science, Technology and
which can catalyze the oxidation process from glucose to glu-
Research (KUSTAR), Abu Dhabi 127-788, UAE
8,9
conic acid. Because of this characteristic, pH-sensitive poly-
d
Department of Bio and Brain Engineering, Korea Advanced Institute of Science and
mers have been combined with glucose oxidase by conjugating
Technology (KAIST), Daejeon 305-701, Republic of Korea
8
–10
glucose oxidase to the side chain or backbone of the polymer.
†
Electronic supplementary information (ESI) available. See DOI:
0.1039/c4ra16656f
Conversion to gluconic acid lowers the pH of the solution, and
1
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Fig. 1 Schematic illustration of glucose-responsive GC/SA-PGGA double-layered nanogel formation and controlled insulin release by
complexation between PBA derivatives and glucose.
1
6
triggers swelling or deswelling of pH-sensitive groups, physiological pH value. Hence, EWG-conjugated PBA facil-
8
depending on the specic types of the groups used. Lectins are itates the conversion of the complex from a hydrophobic to
5
,15,17
carbohydrate-affinitive proteins, which have highly strong hydrophilic property at physiological pH.
Herein, we
3
affinity with sugar moieties. Concanavalin A (Con A) is a lectin propose that glycol chitosan (GC)/sodium alginate (SA)-
protein that has four saccharide binding sites, and is strongly poly(L-glutmate-co-N-3-L-glutamylphenylboronic acid) (PGGA)
1,11
responsive to glucose concentration.
In glucose-responsive gra polymer double-layered nanogel can exhibit excellent
systems, Con A is generally complexed with a saccharide- bio-compatibility with potential glucose-responsive charac-
modied polymer chain. The complex between Con A and teristic. We hypothesize that GC/SA-PGGA double-layered
saccharide can be dissociated by the replacement of free nanogel being glucose-responsive in nature, would enhance
1,11–13
glucose, allowing release of encapsulated agents.
However, control release of insulin at elevated glucose levels (Fig. 1).
the glucose oxidase and lectin protein-based glucose responsive The glucose molecules may bind to the boron atoms of the
systems have shown many limitations in practical applications PBA derivative, and the PGGA chains can be converted to
8
due to low biostability and relatively high toxicity.
hydrophilic structures, resulting in swelling and the conse-
PBA-modied glucose-responsive systems have recently gained quent release of insulin (Fig. 1). The endowment of this
much interest as a promising approach in that they exhibit glucose-responsive characteristic was conrmed by degree of
better biostability and low toxicity compared to the other two swelling and cumulative insulin release test with various
3,6,14–21
types of glucose-responsive systems.
PBA and its deriva- glucose concentrations. The biocompatibility of GC/SA-PGGA
tives possess strong affinity with 1,2-diol compounds, resulting double-layered nanogel was conrmed by cell viability assay
in the formation of ve- and six-membered rings, which is with Henrietta Lacks (HeLa) cells. Furthermore, mouse study
15
thermodynamically a favorable binding event. In an aqueous was conducted to demonstrate the controlled insulin
solution, PBA is equilibrated between two forms: dehydrated release and reduced frequency of insulin administration for
and hydrated forms. With increasing glucose level, glucose enhanced patient compliance. Thus, we conclude that
molecules specically bind to the boron of PBA, and the PBA GC/SA-PGGA double-layered nanogel may contribute to the
becomes negatively ionized. Hence, the equilibrium between future development of novel controlled release of insulin
the two originally existing forms is broken, and the newly delivery systems with improved patient compliance.
5,15
formed glucose–PBA complex becomes the dominant form.
However, PBA cannot solely endow a glucose-responsive char-
acteristic at physiological pH, since the electron density of the
boron is relatively high by the resonance effect. Electron
withdrawing group (EWG)-conjugated PBA derivatives lower the
electron density of the boron, thereby enabling the PBA deriv-
atives to have strong affinity with glucose molecules at a
2
. Materials and methods
15
Materials
NMR solvents, anhydrous tetrahydrofuran (THF), anhydrous
dimethylformamide (DMF), hexamethyldisilazane (HMDS),
sodium alginate (SA), glycol chitosan (GC), bovine pancreas
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insulin, methyl thiazolyl tetrazolium (MTT), Dulbecco's Modied basic conditions, 35% hydrochloride solution (1 mL) was slowly
Eagle's Medium (DMEM), heat-inactivated fetal bovine serum, added into the cloudy suspension. The suspension was stirred
and antibiotic antimycotic solution were purchased from Sigma at RT for 30 min. It was dialyzed against deionized water for
Aldrich. 5-Benzyl-L-glutamate and triphosgene were purchased 1 day. The dialyzed suspension was lyophilized to obtain the
25
from Alfa Aesar. 3-Aminophenylboronic acid (APBA), morpholine, poly-L-glutamate (PLG) (1.8 g) (yield: 86%).
uorenylmethyloxycarbonyl (Fmoc) chloride, N-hydrosuccinimide
NHS), and 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide
EDC)$HCl were purchased from Tokyo Chemical Industry Co.,

(
(
Fmoc-protection of PLG
Ltd. Precipitating and high performance liquid chromatography PLG (2.4 g) was dissolved into anhydrous DMF. Fmoc chloride
HPLC) grade solvents were purchased from Daejung Chemical (0.06 g) was added into the mixture. The reaction mixture was
and Metals Co., Ltd. Molecular weight cut off (MWCO) stirred at RT for 3 h. The mixture was precipitated with diethyl
(
ꢀ1
10 000 g mol
dialysis bag was purchased from Spectrum ether. The precipitate was washed with diethyl ether three times
Laboratories, Inc. Reduced serum medium Opti-MEM® was and ltered by vacuum ltration to obtain the Fmoc-protected
purchased from Gibco® by Life Technologies™.
PLG (1.8 g) (yield: 75%).
Characterization
Poly(L-glutmate-co-N-3-L-glutamylphenylboronic acid) (PGGA)
synthesis
GPC analysis was performed using HPLC system (YL9100,
Younglin, Korea) consisting of quaternary pump, refractive
3
Fmoc-protected PLG (3.0 g) was dissolved into DMSO. EDC$HCl
(11.0 g) and NHS (4.5 g) were added into the mixture. Aer 20
index detector, adiabatic oven, vacuum degasser and 10 to
5
ꢀ1
1
0 g mol DMF GPC column (KD-803, Shodex, Japan). 0.01 M
3
ꢁ
min, APBA (2.7 g) was added into the mixture. The mixture was
LiBr DMF solution was used as a mobile phase at 35 C with a
ꢀ1
stirred at RT for 1 day. Then, the mixture was precipitated with
methanol and ltered by vacuum ltration. The white brown
powder was dried at reduced pressure. Fmoc-protected PGGA
was dissolved into DMF. Morpholine 20% (v/v) in DMF was
added into the solution. The reaction mixture was stirred at
room temperature for 3 h. The reaction mixture was precipi-
tated with diethyl ether, washed several times, and dried at
reduced pressure. PGGA (3.2 g) was obtained (yield: 71%).

ow rate of 1 mL min . A 400 MHz NMR spectrometer
1
(
400 MHz 54 mm NMR DD2, Agilent, USA) was used for
spectra measurements. Deuterated dimethyl sulfoxide
DMSO-d ) and D O were used as a NMR solvent. Fourier
H
(
6
2
transform infrared (FT-IR) spectrometry analysis was conducted
using FT-IR spectrometer (Alpha-p, Bruker, Germany). The
particle size and size distribution were measured with a
dynamic light scattering (DLS) method (Zetasizer nano zs,
ꢁ
Malvern, UK) conducted with a He–Ne (633 nm) and 90 col-
lecting optics at room temperature (RT). The samples were
SA-PGGA synthesis
ꢀ
1
prepared with 1 mg mL concentration. An element analyzer
Flash 2000, Thermo scientic, USA) was used for element
SA (10 mg) was dissolved into distilled water. EDC$HCl (50 mg)
and NHS (10 mg) were dissolved into the SA solution and stirred
at RT for overnight. PGGA (4.4 g) was dissolved in DMSO and the
PGGA DMSO solution was slowly added into the solution. The
(
analysis of carbon, nitrogen and oxygen weight percent. The
samples were prepared with 1 mg mass. TEM was conducted
using 120 kV FEI Technai G2 Bio-TEM. To prepare for the TEM
ꢁ
ꢀ1
reaction mixture was stirred at 40 C for 3 days. The reaction
sample, 0.01 mg mL of the double-layered nanogel solution
was dropped onto a 230 mesh copper grid with carbon, and then
negatively stained with 2% uranyl acetate solution. The copper
grid was dried by a heater within a minute.
mixture was dialyzed (MWCO 10 000) against deionized water
for 3 days. The large aggregate was ltered and the suspension
was lyophilized. SA-PGGA (1 g) was obtained (yield: 3.8%).
Poly(5-benzyl-L-glutamate) (PBLG) synthesis
Insulin-loaded GC/SA-PGGA nanogel preparation
5
-Benzyl-L-glutamate NCA was synthesized by methods
GC/SA-PGGA double-layered nanogel was prepared by an
isotropic gelation method and electrostatic interactions.
22,23
described in the previous literature.
l-glutamte NCA (4 g) was dissolved into anhydrous DMF. HMDS
In a glove box, 5-benzyl-
26
SA-PGGA (10 mg) was dissolved into DI water (10 mL) and
(63 mL) was added into the mixture. The feeding molar ratio of
ꢀ1
insulin solution (1 mL in 1 mg mL 0.01 N HCl solution) was
24
5-benzyl-L-glutamate NCA to HMDS was 50. The mixture was
slowly added into SA-PGGA solution while pipetting for 5 min.
stirred at RT for 2 days. Aer polymerization, the mixture was
poured into methanol. The precipitate was ltered, washed with
methanol and dried at RT for overnight. PBLG (3.2 g) was
obtained (yield: 96%).
In order to use an isotropic gelation, calcium chloride solution
ꢀ1
(
470 mL in 1.1 mg mL deionized water) was added in the
solution and then SA-PGGA solution was mixed by pipetting for
0 min. To form GC/SA-PGGA double-layered nanogel, GC
1
ꢀ
1
ꢀ1
solution (1 mg mL in 1 mg mL deionized water) was slowly
added into the SA-PGGA solution while pipetting for 10 min. To
Deprotection of PBLG
PBLG (3.0 g) was suspended into deionized water. Sodium remove unloaded-insulin, the solution was dialyzed against
hydroxide (1.16 g) was added into the suspension. The reaction deionized water for 1 day and then, lyophilized. Insulin-loaded
mixture was stirred at RT for 1 day. For the neutralization of GC/SA-PGGA double-layered nanogel (15 mg) was obtained.
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ꢀ
1
Conrmation of biocompatibility and blood compatibility
poured into ve dialysis bags (MWCO 10 000 g mol ). Each
dialysis bag containing the nanogel solution was soaked in
The biocompatibility of synthetic nanogels was evaluated by
MTT assay using HeLa cells. The HeLa cells were seeded in
ꢀ
1
6
various glucose solutions (0, 1, 3, 5 and 10 mg mL ). Each
glucose solution was stirred at 150 rpm. Each insulin-released
solution (100 mL) was extracted from the corresponding
glucose levels at a predetermined time point. The amount of
released insulin was quantied by uorescence spectropho-
tometer (RF5301-PC, Shimadzu, Japan). The degree of insulin-
release was expressed as the ratio of the amount of release
9
1
6-well plate at 10 000 cells per one well in DMEM (100 mL) with
0% heat-inactivated fetal bovine serum and 1% antibiotic
ꢁ
antimycotic solution, and incubated at 37 C in 5% CO
2
atmo-
sphere. Aer 1 day-incubation, the medium was removed, and
each well containing HeLa cells was treated with different
nanogel solutions (100 mL): SA, SA-PGGA, GC, and GC/SA-PGGA
for 24 h (in reduced serum medium (90 mL) and each nanogels
5
insulin to that of total encapsulated insulin. All the experi-
ꢀ1
ments were triplicated.
in phosphate buffered saline (PBS) solution (10 mL, 10 mg mL
each sample dissolved in PBS)). Following 24 h of incubation
with nanogel solutions, the medium was removed, and MTT
solution (20 mL, 5 mg mL in PBS solution) was added into
Cumulative insulin release curve with different glucose levels
in vitro
ꢀ
1
each well. MTT solution-treated cells were incubated for 3 h. FITC-insulin-loaded GC/SA-PGGA (10 mg) was suspended in
ꢀ
1
Aer that, solubilization buffer (10% triton-X in isopropanol deionized water. The nanogel suspension (1 mL, 10 mg mL
containing 35% hydrochloride solution (2.5 mL) was added into FITC-insulin-loaded GC/SA-PGGA suspension) was poured into
ꢀ1
each well. The absorbance of each solution in each well was ve dialysis bags (MWCO 10 000 g mol ). Each dialysis bag
quantied by UV-Visible (UV-Vis) spectrometer (MultiSkan™ containing the nanogel solution was soaked in various glucose
ꢀ
1
Go, Thermo Scientic, USA) at 590 nm. The cell viability was solutions of 0, 1, 3, 5 and 10 mg mL . Each glucose solution
expressed as the ratio of absorbance with non-treated cells to was stirred at 150 rpm. Each insulin-released solution (100 mL)
absorbance with sample-treated cells. All the experiments were was extracted from the corresponding glucose level at a pre-
triplicated. The blood compatibility demonstration of nanogels determined time point. The amount of released insulin was
3
was performed using hemolysis assay. Blood was drawn from a quantied by uorescence spectrophotometry. The degree of
mouse and stored in an ethylenediaminetetraacetate-treated insulin release was expressed as the ratio of the amount of
3
tube. In the rst step, the blood sample (0.5 mL) was centri- release insulin to that of total encapsulated insulin. All the
fuged at 1200g for 10 min, and then the supernatant was dis- experiments were triplicated.
carded. In the following step, saline solution (1 mL) was added
into the blood sample, and the red blood cells (RBCs) were
In vivo demonstration of controlled insulin release
isolated by centrifugation at 1200g for 10 min, followed by
All the animal experimental procedures were performed under
washing with saline solution twice. The isolated RBCs were
approval from the Animal Care Committee of KAIST. C57BL/6
diluted to saline solution (5 mL). Then, diluted RBCs suspen-
mice aged eight weeks were used. Overnight-fasted mice were
sion (0.4 mL) was added into PBS, SA, SA-PGGA, GC,
anesthetized by intramuscular injection of a mixture of Zoletil®
GC/SA-PGGA and Triton X-100 saline solutions (0.4 mL). All the
ꢀ
1
ꢀ1
ꢀ
1
(10 mg kg ) and xylazine (11 mg kg ). The level of anesthesia
was continuously monitored during the experiments by using a
toe pinch and maintained by intramuscular injection of half the
dose of initially injected Zoletil–xylazine mixture whenever a
polymer sample concentrations were 4 mg mL , and each
sample (0.1 mL) was added into saline solution (0.3 mL). PBS
(0.1 mL) and Triton X-100 (0.1 mL) were independently added
into each saline solution (0.3 mL). Saline and Triton X-100 were
used as negative and positive controls, respectively. All the
ꢀ
1
response was observed. Glucose (0.3 g kg ) dissolved in PBS
ꢁ
solution (100 mL) was administrated by retro-orbital injection at
treated RBCs samples were incubated at 37 C for 1 h, and then
ꢀ
1
0, 70 and 260 min. Blank GC/SA-PGGA, free insulin (0.5 IU kg ),
centrifuged at 3000g for 10 min. Each supernatant (100 mL) was
transferred into a 96-well plate. The UV-Vis absorbance of each
supernatant was quantied by UV-Vis spectrometer (Multi-
Skan™ Go, Thermo Scientic, USA) at 540 nm. Hemolysis
activity of RBCs was expressed as the percentage of UV-Vis
absorbance difference between sample and negative control
and insulin-loaded GC/SA, prepared by the procedure described
2
6,28
ꢀ1
in literature
(
(0.5 IU kg ), and insulin-loaded GC/SA-PGGA
ꢀ
1
0.5 IU kg ) were dissolved in PBS solution (100 mL) each and
administrated by retro-orbital injection immediately aer the
measurement of glucose level at 0 min. Blood samples were
collected from the tail veins of mice aer initial injection of
glucose (0 min), immediately followed by each sample injection.
Subsequently, blood samples were collected at various time
points to measure the blood glucose level. The blood glucose
levels were measured by a glucose meter (ACCU-CHEK Active,
(PBS) over the difference between positive control (Triton X-100)
and negative control. All the experiments were conducted ve
times.
FITC-insulin-loaded GC/SA-PGGA double-layered nanogel
preparation
Roche Diagnostics GmbH, Germany). Additionally, free insulin
ꢀ1
(
0.25 IU kg ) in PBS solution (100 mL) was administrated by
FITC-insulin was prepared by the previous literature proce- retro-orbital injection twice, at 0 and 70 min. Free insulin in PBS
2
7
ꢀ1
dure. FITC-insulin-loaded GC/SA-PGGA (10 mg) was sus- solution (100 mL) was injected at 0 min (0.5 IU kg ) and 70 min
ꢀ
1
pended in deionized water. The nanogel suspension (1 mL, (0.25 IU kg ). Each datum was expressed as normalized
0 mg mL FITC-insulin-loaded GC/SA-PGGA suspension) was glucose level, which is the ratio of blood glucose level at a
ꢀ
1
1
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predetermined time point to that at 0 min. All the experiments 8.1 ppm correspond to the protons of APBA. The degree of APBA
were quadruplicated.
conjugation was obtained by the ratio of the protons of APBA
group a to that of the methylene groups b in the side chains
based on the peak integration (Fig. 3a). The modication degree
was 37.5% (Fig. 3a). SA-PGGA gra polymer was synthesized by
carbodiimide coupling reaction between the carboxyl groups of
3. Results and discussion
Synthesis and characterization of SA-PGGA gra polymer
PBA derivatives, which have highly strong affinity with glucose SA and the end amine groups of PGGA (Fig. 2). The degree of
and are converted to the negatively charged forms aer a PGGA conjugation was quantied by element analysis (Fig. 3c
binding event, are widely used for insulin delivery systems and and Table S1†). The mole fraction of PGGA chains was deter-
14–16
as glucose-measuring sensors.
For this reason, SA-PGGA was mined based on the amine weight percent, which indicates the
synthesized by NCA polymerization and carbodiimide coupling weight percent of PGGA chains, the number of PGGA repeating
reactions. In the rst step, 5-benzyl-L-glutamate NCA units and the degree of APBA modications. The mole fraction
was controlled-polymerized by using hexamethyldisilazane of SA segments was obtained by the ratio of the total carbon
2
4
(
HMDS) as a living initiator (Fig. 2). The chemical structure of moles, excluding PGGA chains' carbon moles, to the number of
1
poly(5-benzyl-L-glutamate) (PBLG) was characterized by H NMR carbons in one SA segment. Therefore, the degree of PGGA
spectrometry (Fig. 3a). The molecular weight and polydispersity conjugation was expressed as the ratio of the mole fraction of
index of PBLG, determined by gel permeation chromatography PGGA chains to that of SA chains, and was approximately 30%
GPC), were 6800 g mol and 1.1, respectively (Fig. 3b). The (Fig. 3b and c). Moreover, SA-PGGA gra polymer was qualita-
benzyl groups of PBLG were deprotected by transesterication tively determined by Fourier transform (FT)-IR spectrometry
ꢀ
1
(
ꢀ
1
25
method (Fig. 2), and the chemical structure of poly-L-gluta- (Fig. 3d). The IR absorption at 3284 cm (amine groups) and
1
ꢀ1
mate (PLG) was conrmed by H NMR spectrometry (Fig. 3a). the three absorption peaks at 1647, 1447 and 1232 cm , cor-
Compared to the PBLG spectrum, the phenyl peak at 7.2 ppm responding to the phenyl ring in APBA, implied that PGGA
and the benzyl position at 4.8 ppm disappeared aer depro- chains were successfully bonded to SA (Fig. 3d).
tection (Fig. 3a). To synthesize glucose-responsive PGGA chains,
3
-aminophenylboronic acid (APBA) was conjugated to the
carboxyl groups of PLG, using carbodiimide coupling reaction
Fig. 2). The chemical structure of PGGA and degree of APBA
Optimization and characterization of SA-PGGA nanogel and
GC/SA-PGGA double layered nanogel
(
1
conjugation were conrmed by H NMR spectrometry (Fig. 3). Blank SA-PGGA nanogel was prepared by gel isotropic
As shown in the PGGA spectrum (Fig. 3a), the peaks at 7.4 to method.
26,28
To nd the lowest diameter of SA-PGGA nanogel,
Fig. 2 Synthetic scheme of SA-PGGA graft polymer.
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1
Fig. 3 (a) H NMR spectra of PBLG (DMSO-d
6 2 6
), PLG (D O) and PGGA (DMSO-d ). (b) Gel permeation chromatograph of PBLG. (c) Elemental
weight ratio of SA and SA-PGGA by element analysis. (d) FT-IR spectra of SA and SA-PGGA.
SA-PGGA was treated with four different weight ratios of CaCl2 with the swelling nature of GC/SA-PGGA in aqueous
2
9,30
(
Fig. S1a†). It was found that the mean diameter of SA-PGGA solution.
nanogel was the lowest at 20 : 1 weight ratio of SA-PGGA to
CaCl . With decreasing weight ratio SA-PGGA to CaCl , the
2
2
Demonstration of biocompatibility of SA-PGGA and GC/SA-
mean diameter of SA-PGGA became increasingly higher. In
PGGA double layered nanogel
addition, the zeta potential of SA-PGGA nanogel with different
weight ratios of SA-PGGA to CaCl
2
ranged from ꢀ25 to ꢀ10 mV
To demonstrate the biocompatibility of SA-PGGA nanogel and
GC/SA-PGGA double-layered nanogel, methyl thiazolyl tetrazo-
lium (MTT) assay was performed using HeLa cells. Firstly, the
(
Fig. S1b†). Blank GC/SA-PGGA double-layered nanogel was
prepared by electrostatic interactions between the two poly-
mers. To optimize the size of GC/SA-PGGA double-layered
nanogel, SA-PGGA was treated with four different weight
ratios of GC. The mean diameter of the double-layered
nanogel was the lowest at 0.8 : 1 weight ratio of GC to SA-
PGGA (Fig. S1c†). However, with increasing or decreasing
weight ratio of GC away from this optimal point, the mean
diameter of GC/SA-PGGA double-layered nanogel became
progressively higher. Furthermore, the zeta-potential of GC/
SA-PGGA double-layered nanogel with different ratios of GC
to SA-PGGA ranged from 5 to 12 mV (Fig. S1d†). As shown in
Fig. 4a, the hydrodynamic radius of insulin-loaded GC/SA-
PGGA double-layered nanogel was 767.9 nm, which was
much larger than that of blank GC/SA-PGGA double-layered
nanogel (Fig. S1c†). The zeta-potential of insulin-loaded GC/
SA-PGGA double-layered nanogel was 15 mV (Table S2†),
6
cells were treated with SA, GC, blank SA-PGGA nanogel and
blank GC/SA-PGGA double-layered nanogel (100 mg) for 1 day.
Here, phosphate buffered saline (PBS) was used as a positive
control (Fig. 5a). From MTT analysis, it was found that both
SA-PGGA nanogel and GC/SA-PGGA double-layered nanogel
were considerably biocompatible with the HeLa cells, exhibiting
more than 80% cell viability (Fig. 5a). In the case of SA-PGGA
double-layered nanogel, the biocompatibility was somewhat
enhanced compared to SA (Fig. 5a). Furthermore, the fact that
GC/SA-PGGA double-layered nanogel shows the highest cell
viability in vitro conrmed that GC may allow to reduce cyto-
toxicity (Fig. 5b). It also observed that the GC/SA-PGGA double-
layered nanogel does not affect adversely the viability of cells
7
due to the moderate zeta-potential, which is approximately
10 mV (Table S2†). Finally, to demonstrate the blood compati-
7
which may be suitable for biomedical uses. Insulin loading
bility of GC/SA-PGGA double-layered nanogel, a blood compat-
efficacy quantied by the insulin standard curve was 71 ꢂ
3
ibility test was performed using hemolysis activity assay. The
3
.5% (Table S2†). Transmission electron microscopy (TEM)
quantied value of hemolysis activity indicates the degree of
lysed RBCs by a specic substance in contact with the blood.
Therefore, a smaller percentage of hemolysis activity indicates
better blood compatibility of particular substances (Fig. 5b). In
this case, the hemolysis percentage of all the nanogel samples
measurement was employed to demonstrate the morphology
and size of insulin-loaded GC/SA-PGGA double-layered
nanogel. The TEM images demonstrated that the
morphology was spherical (Fig. 4b). The huge difference in the
nanogel diameter between DLS and TEM may be associated
ꢀ1
were less than 2%, even with high concentrations at 1 mg mL
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Fig. 4 (a) Hydrodynamic radius distribution of insulin-loaded GC/SA-PGGA double-layered nanogel. (b) TEM images of insulin-loaded
GC/SA-PGGA double-layered nanogel.
3,5,6,15
(Fig. 5b). Therefore, based on MTT and hemolysis assay, it can swelling of the nanogel.
GC/SA-PGGA double-layered
be conrmed that GC/SA-PGGA double-layered nanogel is suit- nanogel was treated with ve different glucose concentrations:
ꢀ
1
able for many potential biomedical applications due to its 0, 1, 3, 5 and 10 mg mL . The degree of swelling was expressed
strong biocompatibility.
as the ratio of the mean diameter with glucose to the mean
diameter without glucose. As shown in Fig. 6a, GC/SA-PGGA
double-layered nanogel was increasingly swollen with elevated
glucose-concentration. At 1 mg mL glucose concentration, the
mean diameter of GC/SA-PGGA double-layered nanogel was
Conrmation of GC/SA-PGGA double-layered nanogel's
glucose-responsive characteristic in vitro
ꢀ1
To conrm the endowment of glucose-responsive characteristic, increased by 125% compared to the mean diameter without
the degree of swelling and cumulative insulin release test was glucose (Fig. 6a), suggesting that a small portion of glucose
carried out with different glucose levels (Fig. 6a and b). We molecules have bound to the boron atom of PBA deriva-
3,5,6,15
ꢀ1
hypothesized that the hydrophobic PGGA segments conjugated tives.
At 5 mg mL
glucose concentration, the mean
to SA would be converted from hydrophobic to hydrophilic by diameter of GC/SA-PGGA double-layered nanogel was almost
the negative ionization of the PBA boron atom, resulting in doubled over the non-treated nanogel, indicating that more
Fig. 5 (a) Determination of biocompatibility of SA, SA-PGGA, GC and GC/SA-PGGA nanogel using HeLa cell with PBS as a positive control. All the
experiments were triplicated (n ¼ 3). The error bars were expressed as mean standard deviation. (b) Blood compatibility of SA, SA-PGGA, GC and
GC/SA-PGGA using hemolysis assay with saline and Triton X-100 as negative and positive controls, respectively. All the experiments were
conducted five times (n ¼ 5). The error bars were expressed as mean standard deviation.
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ꢀ1
Fig. 6 (a) Determination of degree of swelling with different glucose levels (1, 3, 5 and 10 mg mL ). (b) Cumulative insulin release curve from
ꢀ1
FITC-insulin-loaded GC/SA-PGGA double-layered nanogel in deionized water with five different glucose levels (0, 1, 3, 5 and 10 mg mL ). All the
experiments were triplicated. The error bars were expressed as mean standard deviation (n ¼ 3).
complexes between PBA derivatives and glucose molecules were meaning that PGGA chains are strongly bound together by
3
,5,6,15
ꢀ1
formed.
Accordingly, at 10 mg mL glucose concentra- hydrophobic interaction due to incomplete complexation
15
tion, GC/SA-PGGA double-layered nanogel was much more between PBA-derivatives and glucose molecules (Fig. 6b).
swollen than the previous groups. These analysis were indica- Therefore, based on the cumulative insulin release curves, it is
tive of the excellent swelling capacity of GC/SA-PGGA double- reasonable to conclude that GC/SA-PGGA double-layered
layered nanogel with elevated glucose concentrations. Thus nanogel clearly exhibits a glucose-responsive characteristic
conrming the glucose responsive property of the double and controlled release of therapeutic doses of insulin at diabetic
ꢀ
1
layered nanogel system. Precisely, it can be concluded that, the glucose levels (3, 5 and 10 mg mL ). Hence, we have demon-
encapsulated insulin can be released out of core in GC/SA-PGGA strated the potential use of GC/SA-PGGA double-layered nano-
double layered nanogels by swelling in response to the glucose gel for in vivo study by investigating the in vitro insulin release
concentration, followed by destabilization of hydrophobic prole.
15
moieties within the GC/SA-PGGA nanogels. To visually
demonstrate the controlled insulin release characteristic at
different glucose levels, Fluorescein isothiocyanate (FITC)- Demonstration of GC/SA-PGGA double-layered nanogel's
insulin-loaded GC/SA-PGGA was poured into dialysis bags, controlled insulin release in vivo
and treated with ve different glucose concentrations: 0, 1, 3, 5
To conrm the controlled insulin release capability of
ꢀ
1
and 10 mg mL
in double deionized water (Fig. 6b). The
GC/SA-PGGA nanogel in vivo, mice were treated with four
different samples: blank GC/SA-PGGA, free insulin, insulin-
loaded GC/SA, and insulin-loaded GC/SA-PGGA. Blank
GC/SA-PGGA and free insulin were used as negative and positive
controls, respectively. Insulin-loaded GC/SA was employed as a
negative comparison group to conrm the endowment of
resulting cumulative insulin release curve is shown in Fig. 6b.
As observed higher the glucose level, the more would be the
rapid release of insulin. In the case of the two high glucose
ꢀ1
levels, 5 and 10 mg mlL , more than 50% of encapsulated
FITC-insulin was rapidly released from GC/SA-PGGA double-
layered nanogel until 12 h (Fig. 6b). As shown in the cumula-
glucose-responsive characteristic on insulin-loaded GC/SA-
ꢀ1
tive insulin release curves (0 and 1 mg mL glucose concen-
trations), the amount of released insulin was moderately
increased until 12 h (Fig. 6b). Aer 1 day, most of the encap-
ꢀ1
PGGA. The dosage of insulin in each sample was 0.5 IU kg
.
Glucose was administrated to mice at 0 (all groups), 70
excluding blank GC/SA-PGGA group) and 260 min (excluding
(
ꢀ
1
sulated insulin was released at 5 mg mL , a typical diabetic
glucose concentration indicating that the formation of
complexes between PBA derivatives and glucose molecules
weakens hydrophobic interactions between PGGA segments
blank GC/SA-PGGA group). Once glucose was administered to
mice, their blood glucose levels immediately boosted from
normal ranges to diabetic glucose ranges. Before the second
glucose injection, all the blood glucose levels that of excluding
blank GC/SA-PGGA followed a similar behavior in that their
glucose levels were progressively lowered (Fig. 7a). Compared to
1
5
(
(
Fig. 6b). On the other hand, at low glucose concentrations
ꢀ
1
0 and 1 mg mL ), the encapsulated insulin was not released,
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Fig. 7 (a) Normalized blood glucose levels in mice by retro-orbital administration of blank GC/SA-PGGA, free insulin, insulin-loaded GC/SA and
ꢀ1
insulin-loaded GC/SA-PGGA. (b) Normalized blood glucose levels in mice by the administration of free insulin injected twice (0.25 IU kg ), free
ꢀ1
ꢀ1
insulin injected (0.5 and 0.25 IU kg ), and insulin-loaded GC/SA-PGGA (0.5 IU kg ). All the error bars were expressed as mean standard error (n
4, *p < 0.1, **p < 0.05 compared with Insulin-loaded GC/SA alone).
¼
insulin-loaded GC/SA, insulin-loaded GC/SA-PGGA-treated mice because of the difference in insulin dosage (Fig. 7b). Aer re-
ꢀ1
showed that the blood glucose levels were more dramatically administration of 0.25 IU kg free insulin, it was shown that
decreased until 60 min, meaning that GC/SA-PGGA nanogel can the blood glucose levels of insulin-retreated mice groups of
ꢀ1
controllably release the encapsulated insulin by its glucose- 0.25 IU kg free insulin resemble the plot of insulin-loaded
responsiveness (Fig. 7a). Aer the second glucose injection, GC/SA-PGGA from 70 min to 130 min (Fig. 7b). The similar
ꢀ
1
the remarkable differences in normalized glucose levels tendency of the graphs of free insulin (0.5 and 0.25 IU kg ),
between the insulin-loaded GC/SA-PGGA group and the other and insulin-loaded GC/SA-PGGA illustrated that GC/SA-PGGA
group are displayed in Fig. 7a. Notably, the blood glucose levels double-layered nanogel could stabilize the encapsulated
of two particular groups, free insulin and insulin-loaded GC/SA, insulin, and successfully achieve glucose-triggered insulin
were no longer dramatically decreased. Moreover, the blood release (Fig. 7b). Consequently, a reduction in the number of
glucose levels were maintained, indicating that the originally insulin injection can be attained using glucose-responsive
administrated insulin have been nearly physiologically GC/SA-PGGA double-layered nanogel and contribute to the
consumed and the insulin entrapped inside the nanogel may enhancement of patient compliance. Based on these results, it
have stayed inside the particle core. On the other hand, the was found that insulin-loaded GC/SA-PGGA double-layered
blood glucose levels were still kept relatively low, meaning that nanogel can controllably release the encapsulated insulin at
the encapsulated insulin were released more controllably high glucose levels and maintain low blood glucose levels for up
(
Fig. 7a). Aer the third glucose injection, the encapsulated to 3 h (Fig. 7a and b and Table S3†). Furthermore, the decreased
insulin in the GC/SA-PGGA nanogel was almost consumed aer frequency of insulin administration was demonstrated by the
50 min (Table S3†). To demonstrate the reduced frequency of experimental observations as shown in Fig. 7b, followed the
insulin injection, one group of mice was given two adminis- same behavior. Therefore, it can be concluded that GC/SA-PGGA
2
ꢀ
1
trations of 0.25 IU kg insulin at 0 and 70 min, and another double-layered nanogel may act as a promising platform for
ꢀ
1
group was treated with 0.5 and 0.25 IU kg insulin at 0 and next generation controlled insulin release systems.
0 min, respectively (Fig. 7b). Prior to the second glucose
administration from 0 to 70 min, the blood glucose levels of free
7
ꢀ
1
insulin (0.5 and 0.25 IU kg ), and insulin-loaded GC/SA-PGGA 4. Conclusions
exhibited a similar decreasing tendency (Fig. 7b). Whereas, the
ꢀ
1
To conclude, we have succeeded in demonstrating the feasi-
bility of GC/SA-PGGA double-layered nanogel as a potential
controlled insulin drug delivery system. For developing this
blood glucose levels of mice injected twice with 0.25 IU kg of
insulin were slightly lowered compared to those of other groups
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system, initially SA-PGGA gra polymer possessing a glucose-
responsive moiety was synthesized by NCA polymerization
and carbodiimide coupling reactions. GC/SA-PGGA double-
layered nanogel was formed by gelation method and electro-
static interactions between two polymers. The developed
nanogel had size of 700 nm as per DLS, and TEM images
showed much reduced size of 200 nm with a sponge-like
morphology. In addition, it was found that GC/SA-PGGA
double-layered nanogel showed good biocompatibility and
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Acknowledgements
This study was supported by the KUSTAR-KAIST Institute at
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