A synthetic approach toward a self-regulated insulin delivery system

Article abstract of DOI:10.1002/anie.201106252

Protein-free: A hydrogel containing phenylboronate was optimized so as to undergo rapid glucose-dependent changes in the state of hydration under physiological aqueous conditions. A localized dehydration of the gel surface to form a "skin layer" enabled control of the release of insulin from the gel. This dehydration is induced by fluctuations in the glucose concentration in the range between normo- and hyperglycemia. Copyright

Full text of DOI:10.1002/anie.201106252

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Angewandte
Communications
DOI: 10.1002/anie.201106252
Insulin Delivery Gel
A Synthetic Approach Toward a Self-Regulated Insulin Delivery
System**
Akira Matsumoto, Takehiko Ishii, Junko Nishida, Hiroko Matsumoto, Kazunori Kataoka, and
Yuji Miyahara*
Stimuli-responsive “smart gels” that undergo physicochem-
ical property changes in response to external stimuli can
provide both a functional and structural basis for self-
regulated materials and systems. Of particular interest are
those materials sensitive to chemical stimuli, that is, the
fluctuation in the concentration of specific biomolecules, that
can mimic biofeedback systems, thus providing many attrac-
tive platforms for biomedical applications. Glucose has been
a molecule of interest for decades owing to its relevance to the
treatment of diabetes. Diabetes is not an infectious disease,
but its rapidly increasing worldwide prevalence has been
recognized as a pandemic, and thus diabetes poses a serious
global health threat much like epidemics of truly infectious
diseases, such as HIV/AIDS.[1] Despite the necessity for the
continuous and accurate glycemic control in the management
of insulin-dependent diabetes mellitus (IDDM), the current
palliative treatment relies almost solely on the patients self
injection of insulin; this self injection not only impinges on the
quality of life of the patient but also fails to precisely control
the dose of insulin, where an overdose must be strictly
avoided as it causes serious hypoglycemia.
These protein-based materials eventually denature and lose
activity, and thus are not practical for long-term use and
storage. Another problem is the strong cytotoxicity of Con A;
this cytotoxicity has, thus far, prevented any clinical applica-
tions of these materials.
Herein we describe a totally synthetic alternative to these
materials. Phenylboronic acid (PBA), a synthetic molecule
capable of reversibly binding with 1,2- or 1,3-cis-diols includ-
ing glucose, is utilized as the molecular basis.[3] Our previous
studies have shown that the glucose-dependent shift in the
equilibrium of PBA between the uncharged and anionically
charged forms, when coupled with a properly amphiphilic
three-dimensional backbone (or gel), could induce a reversi-
ble change in the volume of the gel.[4] The resultant abrupt
and rapid change in the hydration, under certain conditions,
could cause a localized dehydration of the gel surface, that is,
a so-called skin layer, thus offering a method to instantly
control the permeation of gel-loaded insulin.[4a,c]
This goal is a great challenge and relies on achieving
sufficient glucose sensitivity under physiological pH and
temperature, that is, pH 7.4 and 378C, while also fine tuning
the system so it shows a gated response to the change in
glucose concentration critically at the level of normoglyce-
mia, i.e., ca. 1 glÀ1. To achieve this goal, we have systemically
explored the structure–property correlation by focusing on
(meth)acrylamide-based hydrogels. The major efforts have
been directed to controlling the apparent pKa of PBA; the
pKa depends not only on chemical structure of PBA, but also
on the state of hydration.[4e] To make the situation more
complicated, in a hydrogel environment the degree of
hydration is a function of (as well as temperature, ionic
strength, pH, and sugar concentration) the hydrophilicity,
rigidity, and density of the main chain components and the
amount of PBA. Herein we describe a gel that meets all the
above criteria; the chemical structure of this gel turned out to
be a remarkably simple copolymer system. A smart gel has
been shown to act as an artificial pancreas under conditions
closely related to human glucose homeostasis.
For the preparation of self-regulated insulin delivery
systems, the following two approaches have historically
prevailed.[2] One is based on enzymatic reactions between
glucose oxidase (GOD) and glucose, a similar rationale to
those exploited in commercial glucose sensors. The other type
utilizes carbohydrate-binding lectin proteins, such as conca-
navalin A (Con A), as a complementary binder to glucose.
[*] Dr. A. Matsumoto,[+] J. Nishida, Dr. H. Matsumoto, Dr. Y. Miyahara
Institute of Biomaterials and Bioengineering
Tokyo Medical and Dental University
Kanda-surugadai 2-3-10, Chiyoda-ku, Tokyo 101-0062 (Japan)
E-mail: miyahara.bsr@tmd.co.jp
Dr. T. Ishii[+]
Department of Bioengineering, The University of Tokyo
Hongo 7-3-1, Bunkyo-ku, Tokyo 113-8656 (Japan)
Dr. K. Kataoka
Department of Materials Science and Engineering
The University of Tokyo
Hongo 7-3-1, Bunkyo-ku, Tokyo 113-8656 (Japan)
As illustrated in Figure 1a, PBA derivatives in water exist
in equilibrium between uncharged (i) and anionically charged
(ii) forms. Upon addition of glucose, only the charged PBA (ii)
can form a stable complex with glucose (iii), and the
formation of this complex results in an apparent decrease in
the amount of anionically charged PBA (ii). On consumption
of the phenylboronate (ii), the equilibrium between (i) and (ii)
shifts toward the latter (ii). Since the complexed PBA (iii) is
also anionically charged, further addition of glucose leads to
a shift in the equilibrium (i + ii + iii) toward the phenyl-
boronate anions (ii + iii), and a decrease in the amount of
[+] These authors contributed equally to this work.
[**] This research is supported by the Japan Science and Technology
Agency (JST), Core Research of Evolutional Science & Technology
(CREST), and, in part, by the Japan Society for the Promotion of
Science (JSPS) through the “Funding Program for World-Leading
Innovative R&D on Science and Technology (FIRST Program),”
initiated by the Council for Science and Technology Policy (CSTP).
Supporting information for this article is available on the WWW
under http://dx.doi.org/10.1002/anie.201106252.
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both the structure and position of
the substituents (see the Support-
ing Information, Scheme S2).[6]
Accordingly, a new derivative 4-
(2-acrylamidoethylcarbamoyl)-3-
fluorophenylboronic
acid
(AmECFPBA), possessing para-
carbamoyl and meta-fluoro sub-
stituents, was synthesized (the
Supporting
Information;
Scheme S1). Although it is often
the case that strongly electron-
withdrawing substituents (repre-
sentatively, fluorocarbons) are
very hydrophobic, thus limiting
the use of the resultant com-
pounds in an aqueous environ-
ment, the combination of the two
substituents in the new derivative
could minimize such a side effect.
The pKa of the new derivative was
determined to be 7.2 (see the
Supporting Information, Fig-
ure S1),
a
value appreciably
lower than physiological pH
(7.4), thus suggesting glucose sen-
sitivity under such a condition.
Figure 1b and c show structure
and phase diagrams (obtained at
pH 7.4), respectively, of a copoly-
mer gel with a poly(N-isopropyl-
methacrylamide) (PNIPMAAm)
main chain and bearing an opti-
mized (in terms of obtaining the
glucose sensitivity under physio-
logical conditions) molar content
Figure 1. a) Glucose-dependent equilibria of phenylboronic acid. b) Structures of the monomers
contained in the gel and their optimized molar amounts in the feed solution so as to obtain glucose
sensitivity under physiological pH and temperature (pH 7.4 and 378C). c) Phase diagram of the gel
showing the equilibrium volume changes as a function of temperature for various glucose concen-
trations investigated at pH 7.4. The green arrow indicates physiological temperature (378C). d) Volume
change of the gel as a function of the glucose concentration at 378C. The green arrow indicates the
normoglycemic concentration of glucose (1 glÀ1).
glucose shifts the equilibrium toward the uncharged phenyl-
boronate (Figure 1a). Such a glucose-dependent change in
the phenylboronate equilibrium, when coupled with an
amphiphilic three-dimensional polymeric backbone (or gel),
gives rise to a reversible volume change (change in hydration)
of the gel in response to a change in glucose concentration.[4]
This volume change, termed volume phase transition, is
primarily driven by the change in the counterionsꢀ osmotic
pressure in the gel synchronized with the equilibrium shift of
PBA.
In our effort to find a system that is sufficiently glucose
sensitive under physiological pH and temperature (pH 7.4
and 378C) a primary challenge stemed from the relatively
high intrinsic pKa of PBA, which is typically around 9 (see the
Supporting Information, Scheme S2). Under the physiolog-
ical pH (7.4) the weak acidity of PBA results in a limited
fraction of charged phenylboronate, which is responsible for
the interaction with glucose. A strategy to decrease the pKa of
PBA has been reported previously.[4b,e] Briefly, it involves the
introduction of strongly electron-withdrawing substituents to
the phenyl ring.[5] Such a structural change results in an
electron-poor boron atom of PBA, thus making it more acidic
(i.e. a lower pKa); the magnitude of this effect is dependent on
of AmECFPBA (0.075 molar fraction of the total amount of
monomer in the feed solution; see also the Supporting
Information, Figure S2). It is shown in Figure 1c that the gel
shrinks with an increase of the temperature as a result of the
thermo-sensitivity of the PNIPMAAm main chain. When
observed at constant temperatures the gel becomes more
swollen with an increase in the amount of glucose owing to
the increased fraction of pheneylboronate anions, as men-
tioned above. Separately shown in Figure 1d (extracted from
Figure 1c) is the volume change of the gel as a function of the
glucose concentration under physiological pH and temper-
ature (pH 7.4 and 378C). Notably, at physiological temper-
ature the gel is completely shrunken (dehydrated) for
a glucose concentration of up to 1 glÀ1 (a value correspond-
ing to normoglycemia), whereas with any further increase in
glucose concentration it becomes markedly swollen
(hydrated). These observations thus demonstrate the ability
of the gel to undergo a volume phase transition in response to
a change in glucose concentration, the threshold value of
which is in agreement with that of normoglycemia (1 glÀ1
)
under physiological conditions (pH 7.4, 150 mm NaCl, 378C).
Importantly, a 0.075 molar fraction of AmECFPBA was
critical to achieve the desirable glucose sensitivity as obtained
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in Figure 1d. That is to say,
any further increase in the
fraction of AmECFPBA,
although leading to a more
significant
glucose-depen-
dent volume change, caused
an excessive decrease in the
hydrophilicity of the overall
polymer network owing to
the strong hydrophobicity of
AmECFPBA (see the Sup-
porting Information Fig-
ure S2), thus making the gel
insensitive under the physio-
logical temperature. Such
precision in the structure con-
trol would not have been
realistic without the use of
(ready-to-copolymerise)
PBA-containing monomers,
as opposed to strategies
involving polymeric state
reactions.
In addition to the above-
disclosed states of equilibri-
um, (only) when the magni-
tude of the change in hydra-
tion is substantial, the gel can
Figure 2. a) Time-course (left) transmittance and (right) 8-anilino-1-naphthalene sulfonic acid (ANS)
fluorescence top-view images of a cylinder-shaped piece of gel when changing the glucose concentration
under pH 7.4 and 378C. Indicated in each transmittance image are the elapsed times after the initial
lowering of the concentration from 2 glÀ1 to 1 glÀ1. In the fluorescence images, the appearance of the skin
layer is seen as an increase in the intensity of the blue color of ANS, thus correlating with the dehydration
on the gel surface. Conversely, a reduction in the intensity upon an increase of the glucose concentration
(t=90 min.) is indicative of the disappearance of the skin layer. b) Changes in the ANS intensity profiles of
a 150 mm depth cross-section (from the gel surface) during the growth of the skin layer. c) Schematic
representation of the cross-section density distribution of the polymer network in the gel when reversibly
forming the skin layer.
also exhibit
a metastable
phase known as a skin layer
by a process of shrinkage
(dehydration).[7] Note that
none of the other materials
(including those in our pre-
vious reports) have shown
such an abrupt response that would be sufficient to induce
the skin layer under physiological conditions. A skin layer is
a thin surface layer of a collapsed (dehydrated) polymer
network, which, once formed, upon a decrease in the glucose
concentration, inhibits both further penetration of glucose
and further loss of water.[4a,c] As a result, two coexistent
phases, that is, an inner hydrated and a surrounding dehy-
drated phase, can stably prevail for an appreciably long period
of time, during which permeation of any preloaded molecules
such as insulin can also be suppressed. Our previous study
showed that a skin layer formed on a micro gel bead
composed of a poly(N-isopropylacrylamide) main chain and
3-acrylamidophenylboronic acid could be maintained for at
least 10 hours (before further shinking occurred), despite its
relatively small size (several hundred mm in diameter) and
a large change in glucose concentration (from 5 glÀ1 to
0 glÀ1), although it was observable only under nonphysio-
logical conditions, such as pH 9 at 288C.[4c] As already
mentioned, our strategy to achieve the controlled release of
insulin capitalizes on the skin layer. Figure 2 summarizes
kinetic observations of the skin layer that elaborates upon
distinctive for the skin layer, the fluorescence reporter
agent 8-anilino-1-naphthalenesulfonic acid (ANS) was uti-
lized for its ability to increase fluorescence intensity when
exposed to an environment of decreased local polarity
(implicated in dehydration). Figure 2a shows both trans-
mittance (left) and fluorescence (right) top-view images of
the gel that had been initially equilibrated in hyperglycemic
conditions (2 glÀ1) and then transferred, back and forth, to
normoglycemic conditions (1 glÀ1); the prompt and rever-
sible kinetics of the skin layer can be seen as changes in the
intensity of the blue color of ANS (see also the Supporting
Information, Figure S3). When the gel was kept under
normoglycemic conditions (1 glÀ1) for 24 hours, a slight
decrease in volume was observed but not complete shrinkage;
this result supports the preservation of the skin layer, and thus
the partitioning between the inner-hydrated and the outer-
dehydrated phases, even after 24 hours. Figure 2b shows
cross-sectional profiles of the ANS fluorescence intensity
during the growth of the skin layer. These profiles reveal
a progressive increase in the density of the collapsed
(dehydrated) polymer chains; this increase in density
becomes more prominent over time and at the vicinity of
the edge (surface) of the gel, a feature consistent with the
switching the glucose concentration between normo- (1 glÀ1
)
and hyperglycemic (2 glÀ1) levels. In order to visualize the
dynamic and reversible process of dehydration that is
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formation of a skin layer. These observations are schemati-
cally presented in Figure 2c.
to the previous experiments
(1.5m). As expected, the leakage
of insulin became barely visible
owing to the increased network
density or suppressed diffusion
of insulin, thus reducing the
induction times. This observa-
tion is of critical importance to
prevent overdosing of insulin
once a normoglycemic condition
is reached. Remarkably enough,
this gel did not allow leakage of
insulin for a few days while
normoglycemia was maintained,
and upon re-increase of the
glucose concentration, the gel
did allow a burst release of
insulin with no visible lag time
(see the Supporting Informa-
tion, Figure S5). Neither deteri-
oration in the shape of the gel
nor cracks were observed after
repeated formation of the skin
layer over two weeks. Also
worthy of mention is that the
FITC-labeled (bovine) insulin was loaded onto the gel by
first allowing the gel to be hydrated (swollen) in a solution
containing insulin and then introducing the skin layer onto its
surface. Figure 3a–c show time-course profiles of the released
insulin under different addition patterns of glucose under
physiological conditions. It is clearly demonstrated that the
release of insulin from the gel can be continuously controlled
Figure 4. Improved insulin
release using a gel with
a dense polymer network
with twice the monomer
concentration compared to
Figure 3. (Top) Time-course
change in fluorescence inten-
sity of FITC-labeled bovine
insulin released from the gel
under physiological aqueous
conditions. (Bottom) Tempo-
ral pattern of the glucose
concentration investigated in
the experiment.
Figure 3. (Top) Time-course changes in the fluorescence intensity of
FITC-labeled bovine insulin released from the gel under physiological
conditions (pH 7.4, I=0.15, 378C). (Bottom) Temporal patterns of the
fluctuation in glucose concentration, investigated in each experiment.
normal pattern of physiological response to a stepwise
increase in blood glucose concentration is initially a rapid
and pulsatile release of insulin followed by a ramp in
secretion.[9] Indeed, this type of pattern could also be imitated
by lowering the amount of insulin remaining in the gel (see
the Supporting Information Figure S4).
by the skin layer (see also the Supporting Information,
Figure S4) with close correspondence to each addition pattern
of glucose. In Figure 3a, the glucose concentration was varied
between 1 glÀ1 and 2 glÀ1, which correspond to normogly-
cemia and the typical cut-off value for the diagnosis of
diabetes, respectively. Note that the profile in Figure 3b, for
which the glucose concentration is increased to up to 3 glÀ1, is
displayed at 10 times the scale as for Figure 3a and Figure 3c.
By comparing these spectra, it can be determined that under
such more severe hyperglycemic conditions the rate of release
is 10 times greater (Figure 3b; 3 glÀ1) as compared to the
other glucose concentrations (Figure 3a; 2 glÀ1). On the
other hand, under normoglycemic conditions (shown in
Figure 3c) the release is effectively halted over the timescale
of a day (at least 1000 min). Overall, sharp onsets of the
release in response to increases in glucose concentration are
readily achievable, whereas during the offset processes, that is
upon decreased glucose concentration, somewhat substantial
induction times are seen depending on the addition pattern of
glucose. This delayed shut-off response can be attributed to
a leakage of insulin through the skin layer at the premature
stage of the polymer-network contraction. Based on the fact
that the diffusion coefficient of a solute molecule in a gel is
dependent on the density of the polymer network, that is,
effective mesh size relative to the size of solute molecule, in
the present case insulin,[8] we anticipated that the leakage
could be reduced by increasing the density of the cross-linker
(N,N’-methylene-bis(acrylamide)) or the total concentration
of monomers at preparation (with consideration to the
molecular weight of insulin). Figure 4 verifies the validity of
such a remedy. For this particular experiment, the gel was
prepared with twice as high a polymer-network density
(monomer concentration at preparation: 3m) as compared
Our results demonstrate that a nonprotein-based, totally
synthetic smart gel can achieve “lasting” control over the
provision of insulin under conditions closely associated with
human glucose homeostasis. Owing to the significant rate and
magnitude of the glucose-induced phase transition, the gel
can promptly form a skin layer, thus providing a rationale for
the surface-controlled provision of insulin. This non-equilib-
rium mechanism leads to a remarkably shortened response
time (owing to the short diffusion distance across the skin
layer: typically on the order of 100 mm); this shortened
response time is an important criterion for the tight control of
the insulin dosage. Such a surface-limited mechanism (as
opposed to more bulk-dependent systems) in theory allows
the prediction of the release profile, independent of the size
and topological structure of the gel. This feature may
facilitate ease and accuracy in the management of the
administration of insulin. Furthermore, the size independency
may lead to compatibility with other well-developed tech-
nologies such as dwelling needles and semi-embedded
devices. Aside from the skin-layer-based strategy described
herein, the material can also be exploited as other mechan-
ically functional elements such as chemical valves or pumps
for related applications.[10] In addition, with its ability to cause
dramatic changes to other parameters including size, charge
density, conductivity, and permittivity (owing to the change in
hydration), the material provides a unique platform for
electrochemical sensors and actuators. There is also room for
size control (e.g., preparation in micro- or nanoporous
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structures which are already well established),[11] through
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Further efforts are directed towards these possibilities along
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Experimental Section
Gels were prepared by radical copolymerization in a 1 mm diameter
glass capillary using 2,2’-azobisisobutyronitrile (AIBN) as an initiator
in the presence of N,N’-methylene-bis(acrylamide) (MBAAm) as
a cross-linking agent in DMSO. The concentrations of the total
monomers, AIBN, and MBAAm in feed were 1.5m, 7.5 mm, and
15 mm, respectively. The reaction was carried out at 608C for 24 h.
The obtained gels were taken out of the capillaries and washed with
an excess amount of DMSO and distilled water to remove any
unreacted reagents and to thoroughly exchange the solvent from
DMSO to water. Diameter changes of the capillary gels were
observed under an optical microscope (SZX16, OLYMPAS, Japan).
To control the temperature, a thermostated water-flow chamber was
utilized on the microscope stage, to which the gel capillaries (pieces of
5 mm length in the shrunken state) were placed with each exper-
imental buffer solution. The equilibrium volume changes when
decreasing the temperature were recorded under various glucose
concentrations to obtain the phase diagrams. Based on the cylindrical
symmetry of the gel, the degree of swelling (relative volume) of the
gel was defined as (d/d0)3, where d0 is the diameter obtained at 458C
(with no glucose) whereas d is those observed under various
conditions.
More detailed methods for material preparation, optimization,
analysis of the skin layer, release experiment along with pKa data of
various PBAs as determined by titration, additional data of skin layer
observation, and the release experiments are provided in Supporting
Information.
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Matsukawa, I. Ando, Macromolecules 1996, 29, 7136 – 7140.
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Received: September 4, 2011
Published online: December 9, 2011
Keywords: gels · insulin · phenylboronates · polymers · skin layer
.
[1] United Nations Resolution 61/225: World Diabetes Day (from
83rd plenary meeting, 20 December 2006).
[12] T. Tanaka, D. J. Fillmore, J. Chem. Phys. 1979, 70, 1214 – 1218.
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