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Author's personal copy
Journal of Controlled Release 132 (2008) 2–11
Contents lists available at ScienceDirect
Journal of Controlled Release
j o u r n a l h o m e p a g e : w w w. e l s e v i e r. c o m / l o c a t e / j c o n r e l
Review
Chemically controlled closed-loop insulin delivery
Valérie Ravaine a,b,⁎, Christophe Ancla a,b, Bogdan Catargi c,⁎
a
b
c
Université Bordeaux, Institut des Sciences Moléculaires, ENSCPB, 16 Av. Pey Berland, 33607, Pessac Cedex F33607, France
CNRS, Institut des Sciences Moléculaires, Talence Cedex, F33405, France
Université Bordeaux, Hôpital Haut-Lévêque, CHU de Bordeaux, Pessac, France
a r t i c l e
i n f o
Article history:
Received 15 May 2008
Accepted 6 August 2008
Available online 23 August 2008
Keywords:
Insulin
Diabetes
Hydrogel
Glucose-responsive
Phenylboronic acid
a b s t r a c t
Alternative treatments for diabetes are currently being investigated to improve both patient comfort and
avoid complications due to hyperglycaemia episodes. In the absence of a cure like pancreas or beta-islets
transplants, the ideal method would be an artificial “closed-loop” system able to mimic pancreas activity.
This would operate continuously and automatically, causing appropriate response to losses and gains in
glucose levels. Chemically controlled closed-loop insulin delivery has been explored by two main strategies.
The first one consists in delivering insulin with a glucose-responsive matrix. Polymeric hydrogels that swell
or shrink according to the glucose concentration allow delivering insulin doses adapted to the glucose
concentration. The second strategy consists in modifying insulin itself with glucose-sensitive functional
groups that trigger its activity. Recent developments made in these areas represent significant progress in
terms of biocompatibility, selectivity, pharmacokinetics, and easiness of administration, as required for in
vivo applications. Although some issues still have to be overcome, this field of research is promising as a
possible alternative to other approaches for diabetes treatment.
© 2008 Elsevier B.V. All rights reserved.
Contents
1.
2.
Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
Glucose-responsive hydrogels . . . . . . . . . . . . . . . . . . . . . .
2.1.
Glucose oxidase modified hydrogels . . . . . . . . . . . . . . . .
2.1.1.
Insulin release by diffusion through a hydrogel membrane .
2.1.2.
Insulin release by other mechanisms . . . . . . . . . . .
2.1.3.
Towards biocompatible hydrogels . . . . . . . . . . . . .
2.2.
Lectin-modified hydrogels . . . . . . . . . . . . . . . . . . . .
2.3.
Hydrogels and microgels modified with a phenylboronic acid moiety
2.3.1.
Glucose-swelling hydrogels . . . . . . . . . . . . . . .
2.3.2.
Glucose-shrinking hydrogels . . . . . . . . . . . . . . .
2.3.3.
Glucose-responsive microgels . . . . . . . . . . . . . .
3.
Modified insulin . . . . . . . . . . . . . . . . . . . . . . . . . . . .
4.
Conclusion and perspectives. . . . . . . . . . . . . . . . . . . . . . .
4.1.
Treating diabetes with self-delivery materials . . . . . . . . . . .
4.2.
Challenges for the future . . . . . . . . . . . . . . . . . . . . .
References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .
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1. Introduction
⁎ Corresponding authors. Ravaine is to be contacted at Université Bordeaux, CNRS,
Institut des Sciences Moléculaires-UMR 5255, site ENSCPB, 16 Av. Pey Berland, 33607
Pessac, France. Tel.: +33 5 40 00 27 30; fax: +33 5 40 00 27 17. Catargi, Université
Bordeaux, Hôpital Haut-Lévêque, CHU de Bordeaux, Pessac, France. Tel.: +33 5 57 65 68
40; fax: +33 5 57 65 68 06.
E-mail addresses: [email protected] (V. Ravaine), [email protected]
(B. Catargi).
0168-3659/$ – see front matter © 2008 Elsevier B.V. All rights reserved.
doi:10.1016/j.jconrel.2008.08.009
Diabetes mellitus is a disorder of glucose regulation, characterized
by an accumulating glucose concentration in the blood. It is a major
public health problem affecting 8% of the United States population.
The global prevalence of type 2 diabetes continues to rise [1]. The
breakdown of glucose regulation can be attributed to the inability of
the endocrine pancreas to secrete insulin or to the body's inability to
properly use insulin. In the case of type 1 diabetes, the insulin secreting pancreatic beta-cells undergo an autoimmune-mediated destruction.
Author's personal copy
V. Ravaine et al. / Journal of Controlled Release 132 (2008) 2–11
Type 2 diabetes arises from disorders of both insulin resistance and
secretion [2,3]. The treatment of type 2 diabetes concerns mainly oral
antidiabetic drugs, but in some cases, insulino-therapy is needed when
insulinopenia arises [4]. The usual treatment for type 1 diabetes is
achieved by multiple subcutaneous insulin injections, administered
daily. However, this method does not maintain normoglycaemia. In 1993,
the results of a 10-year landmark study Diabetes Control and Complications Trial (DCCT) were published [5]. There were two important
conclusions: first, that tight control of glucose closer to the normal level
resulted in significant reductions in complications such as limb
amputation, blindness, and kidney failure. Second, the tight control led
necessarily to increased incidence of low blood sugar (severe hypoglycemia). To avoid such abnormal episodes, curative therapy can be
considered. It consists in the replacement of functional insulinproducing pancreatic beta-cells, with pancreas or islet-cell transplant.
However, in addition to the difficulty to find adequate donor cells,
pancreatic graft rejection has to be avoided. Therefore, transplantation
has to be coupled to lifelong immunosuppression, which also presents
severe side effects. This could be improved by encapsulating islets in a
semi-permeable membrane, which could protect the grafts from the host
immune systems [6]. A field of research has recently been developed in
this area [7,8].
In the absence of a cure for diabetes, the next best alternative
would be a “closed-loop” system, able to mimic the pancreatic activity
of healthy people. An artificial glucose-responsive insulin delivery
system that works by an automated feedback would therefore have
obvious advantages and has been explored. In this concept, the insulin
dosage and delivery would be governed by the blood glucose levels on
an automatic and continuous basis. It is tempting to develop a closedloop system based on the association of an insulin pump, capable of
delivering insulin continuously, and a monitoring device, capable of
sensing glucose continuously. However, another issue arises here,
which is the delays in glucose sensing and insulin delivery. An insulin
delivery algorithm is therefore required to link an existing glucose
sensor and an external insulin pump. Recent technological progress in
continuous glucose sensors [9] has led to prototypes of closed-loop
systems based on the combination of a continuous glucose monitor, a
control algorithm and an insulin pump [10,11]. However, the feasibility
of this solution remains to be confirmed.
Among various strategies aimed at mimicking pancreas activity, a
great deal of research was devoted to glucose-responsive hydrogels,
which undergo a modification of their physical properties in response
to a variation of glucose concentration. In particular, the variation of
the mechanical and permeation properties, associated to a modification of their swelling state and/or viscosity, make them suitable for
delivering amounts of insulin according to the glucose concentration.
Therefore, the first part of this review article will provide an overview
of current research in the field of glucose-responsive materials and
their potential use as insulin self-delivering systems. In a second part,
another strategy will be presented where insulin itself is modified and
its availability depends on glucose concentration.
3
magnetic field or light, have been studied both experimentally and
theoretically. More recently, the concept of biomolecule-sensitive
hydrogel [16] has been developed as biomaterials aiming at mimicking natural feedback systems. Just like natural ones, they can
detect specific ions or biomolecules, which induce conformational
changes and modify their biological functions. Hydrogels can sense a
biomolecule (glucose) and respond to it by a release of the hormone
(insulin).
Insulin release from hydrogels can occur mainly through two
different pathways (Fig. 1). The hydrogel can be used as a membrane
with a controlled permeability, which is triggered by glucose. The
membrane separates a reservoir full of insulin from the outside [17–
20]. It can have different geometries, for example planar in the case of
membranes or spherical in the case of capsules [21]. The hydrogel can
also be the reservoir itself thanks to its porous structure [22–24]. In
both pathways, insulin is released when the hydrogel swells. More
sophisticated release mechanisms can be imagined when the hydrogel
is used in association with a device, for example when it is grafted in
the pores of a membrane [25] or when it is inserted as a microvalve in
a microfluidic delivery device [26].
The equilibrium volume of a neutral hydrogel is a balance between
the osmotic pressure of the polymer network, which is governed by
polymer–solvent interactions and the elasticity of the polymer
network. If the polymer is soluble within the solvent, the corresponding hydrogel tends to swell in order to expand the contact between the
polymer and the solvent. The elastic restoring force is controlled by
the amount of cross-linkage: the higher the cross-linker density, the
lower the swelling ratio is. When the hydrogels are constituted with
charged polymers, they can present a very high swelling ratio. Their
large swelling volume is mainly governed by a third contribution
arising from the electrostatic repulsion between the charged monomer units and the osmotic pressure exerted by the mobile counter-ion
concentration inside the gel [27,28]. The swelling ratio depends on the
charge density of the gel but also on the ionic strength of the
surrounding medium. Therefore, any factor affecting one of these
contributions will induce a modification of the gel swelling volume.
At least, three issues have to be addressed by “intelligent” insulin
delivery.
- Possession of glucose sensitivity, hydrogels have to contain a molecule that specifically interacts with glucose and senses its levels in
the physiological range;
2. Glucose-responsive hydrogels
Hydrogels are cross-linked polymeric networks that absorb large
amounts of water and swell. To maintain the three-dimensional
structure, they can be either chemically or physically cross-linked.
Some hydrogels belong to the class of the stimuli-responsive materials
and present the unique property of undergoing abrupt volume
changes from their collapsed to swollen states in response to environmental changes [12–15]. Sometimes called “intelligent” materials, stimuli-responsive materials display both sensor and effector
functions. They can sense a stimulus as a signal and transduce it via
the structural changes, but the structural change itself can play the
role of effector. Various stimuli might be responsible for such modifications. The effect of stimuli such as pH, temperature, electric or
Fig. 1. Illustration of the two possible ways for self-regulated insulin release from
glucose-responsive hydrogels: (a) across a membrane; (b) the hydrogel plays the role of
a reservoir. Insulin diffusion occurs when the hydrogel is swollen.
Author's personal copy
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V. Ravaine et al. / Journal of Controlled Release 132 (2008) 2–11
- Modulation of insulin according to blood glucose levels;
- Biocompatibility since diabetes is a lifelong, chronic disease.
2.1. Glucose oxidase modified hydrogels
2.1.1. Insulin release by diffusion through a hydrogel membrane
Historically, the first glucose-responsive materials were made from
combining glucose oxidase (GOD) as a sensing element with a pHsensitive hydrogel. When GOD reacts with glucose and converts it into
gluconic acid, the pH within the hydrogel is lowered. The gel responds
to pH because its constituting element will capture protons and the
charge density of the polymer is modified. This gives rise to the
glucose-triggered response in terms of volume change.
Several groups have developed glucose-responsive hydrogels based
on GOD entrapment or immobilization within the gel. In order to achieve
insulin release, hydrogel swelling is required when glucose concentration increases, i.e. when the pH decreases due to the production of
gluconic acid (Fig. 2). Ishihara et al. [17] were the first to report the
regulation of insulin permeation across a glucose-sensitive membrane,
which was itself made of two associated membranes: the pH-responsive
one, a copolymer made of N,N-diethylaminoethyl methacrylate
(DEAEM) and 2-hydroxypropyl methacrylate (HPMA), and a polyacrylamide one containing glucose oxidase. They further built up polymer
capsules containing insulin using the interfacial precipitation method.
The capsules, with a diameter of about 1.5 mm, were packed in a
thermostated glass column studied. Insulin was shown to be released at
a higher rate when glucose was present in the eluent compared to the
buffer solution in the absence of glucose [21].
To achieve high insulin permeability, Albin et al. [18] designed a
macroporous membrane made of a similar copolymer hydrogel
entrapping glucose oxidase. The porosity of the gel was achieved
through phase separation occurring during polymerization. However,
the membrane did not show any experimental response for rises
above normal blood glucose levels. A theoretical model taking into
account the kinetics of glucose reaction was developed [29]. It
described the steady state behavior of glucose-sensitive membrane.
It was shown that the poor availability of the oxygen required for the
enzymatic reaction was responsible for the limitation of the response
to a particular range of glucose concentrations. In the case of thin
membranes, the optimal range was found to be below the pathophysiological range. However, other device configurations were designed to create pathways for oxygen distribution inside the gel. Four
configurations using silicone rubber in contact with the glucosesensitive membrane were found to enlarge the range of glucose concentration sensitivity [30].
Oxygen depletion might also be overcome by the addition of
catalase. Catalase reacts with hydrogen peroxide produced during the
conversion of glucose into gluconic acid. This reaction liberates oxygen
which can be used for glucose oxidation. The presence of catalase
seemed to increase the swelling kinetic as observed by Traitel et al.
with poly(HEMA-co-DMAEM) [31] and by Zhang et al. with p(NIPAMco-MAA) nanoparticles dispersed within a hydrophobic polymer [32].
Another glucose-sensitive hydrogel, based on sulfonamide chemistry
with covalently conjugated glucose oxidase and catalase showed a
reversible glucose dependent swelling [33]. Traitel et al. also reported
the first in vivo studies [31]. Matrices were implanted in the peritoneum of rats. No tissue encapsulation was observed and the released
insulin seemed to keep at least part of its bioactivity. This result was
confirmed very recently by Satish et al. [34].
2.1.2. Insulin release by other mechanisms
In the release mechanisms mentioned so far, insulin is released by
diffusion through a hydrogel when the hydrogel expands. However, a
more rapid response is expected from a squeezing effect. Therefore,
pH-sensitive hydrogels with a shrinking behavior at low pH were
combined with GOD. At low pH, poly(methacrylic acid-g-ethylene
glycol), abbreviated as p(MAA-g-EG) exhibit intrapolymer complexation due to hydrogen bonds of the carboxylic acid groups and the ether
group of the PEG chain [35,36]. Although oscillatory pH-swelling
studies showed a promising rapid collapse supposed to squeeze insulin out of the gel, insulin release was studied via a model but not
experimentally [37].
Instead of a hydrogel, a gating membrane can be used to deliver
insulin in a self-regulated fashion. Here, the polymer expansion is
used to close the pores of a membrane whereas the polymer collapse
allows insulin diffusion through the pores of the membrane. In this
case, polymers with carboxylic functions were used because they
collapse when pH decreases, i.e. when gluconic acid is produced. This
was achieved by Ito et al. using cellulose membranes grafted with poly
(acrylic acid) (PAA) and immobilized GOD [25]. A poly(vinylidene)
membrane was also modified by an acrylic acid copolymer derivative
with covalently bound GOD and associated to a pressurized chamber
[38]. Unfortunately, it did not satisfy the medical requirements due to
a too high basal flow rate. The importance of the grafting yield was
reported by Chu et al. [39]. It appeared necessary to adjust the grafting
yield to design the ideal gating response.
All the examples listed above use the gel permeability variation to
modulate insulin release. Another method consists in degrading the
gel to release the peptide. Uchiyama et al. [40] synthesized a hydrogel
which was degraded by hydroxyl and/or hydroperoxy radicals. Thus,
the liberation of H2O2 produced by the enzymatic reaction between
GOD and glucose, induces a degradation which is synchronized with
glucose concentration.
2.1.3. Towards biocompatible hydrogels
Efforts were made to improve the formulation of glucose-swelling
hydrogels in order to get more biocompatible materials. Incorporation of
poly(ethylene glycol) grafts on the main chains of glucose-responsive
Fig. 2. Schematic representation of pH-responsive hydrogel with entrapped GOD. After glucose diffusion inside the gel, glucose is enzymatically converted into gluconic acid. The pH
decreases, and the polymer becomes charged due to the protonation of pendant amino groups. The hydrogel swells, facilitating the release of insulin by the diffusion mediated process.
Author's personal copy
V. Ravaine et al. / Journal of Controlled Release 132 (2008) 2–11
hydrogels was developed by Podual et al. [41–44]. These grafts were
expected to retard the degradation of enzymes and proteins within the
gel and to minimize the immunoreaction and subsequent rejection by
the body [45]. The size of these hydrogels was scaled down to
microparticles having a diameter between 30 µm and 300 µm [44].
They showed rapid swelling/deswelling dynamics in response to
changes in pH, with a faster response for the smallest particles. A
reversible swelling was achieved in less than 5 min in response to
pulsatile variations of glucose [44].
Very recently, Kashyap et al. [22] reported glucose-responsive
hydrogels made of a nontoxic, nonimmunogenic, biocompatible and
biodegradable polymer. Chitosan, a structural amino polysaccharide
found in invertebrate animals, can form an in situ gelling system when
combined to a polyol counter-ionic dibase salt such as β-glycerol
phosphate disodium. A sol–gel transition occurs when the solution is
heated to body temperature. A gel containing GOD and peroxidase –
which converts H2O2 into oxygen and increases GOD activity – can be
obtained. This gel is glucose-responsive and swells to a maximum when
glucose concentration is around 3 mg/mL. The swelling response is not
linear with the glucose concentration. However, these authors showed
that loaded insulin was released from the gels in a pulsatile manner over
2-hour periods. The formulation was evaluated in vivo in diabetic rats. A
dose of 3 IU/kg insulin was administered subcutaneously to the rats, as a
loaded gel form to a group of animals and as plain insulin to a control
group. A third group of animals was treated with a blank gel. Plasma
insulin levels and glucose insulin levels were measured. The insulin-geltreated group displayed a sustained and better hypoglycemic activity
which illustrated the pharmacodynamic effect of the formulation. The
pharmacokinetic profile was also better, since the residence time of
insulin was longer for the insulin-gel-treated group (up to 16 h). A
second oral glucose challenge was given after 24 h. Although the plasma
glucose level of the insulin-gel-treated group was lower than the blankgel-treated group, no difference could be evidenced in the plasma
insulin levels of the two groups. The pulsatile delivery of insulin could
not be proved in vivo.
2.2. Lectin-modified hydrogels
Lectins are a family of carbohydrate binders, which can be used as
natural receptors in the development of glucose-responsive hydrogels. These proteins are known to interact with glycoproteins and
glycolipids on the cell surface, where they play a role in cellular
adhesion or hormone regulation. The most studied of these lectins is
concanavalin A (Con A), which presents four binding sites [46].
Glucose-regulated insulin release was first achieved using a
glycosylated insulin derivative (and still biologically active) complexed with Con A. Insulin release occured in the presence of glucose
thanks to the binding competition between glucose and glycosylated
insulin with Con A which led to the breakdown of the complexes [47–
49]. This system was further enclosed within a polymer microcapsule
[50]. Succinyl-amidophenyl-glucopyranoside insulin (SAPG-insulin)
was shown to be released according to glucose variations. Both
membranes and microcapsules were optimized to operate in in vivo
studies.
Different polymers bearing saccharide residues were complexed
with Con A to form a gel and encapsulate insulin. Natural saccharide
polymers [51] as well as novel polymers containing saccharide
residues were synthesized [52] and their complexation with Con A
was studied. For example, the complexation of poly(2-glucosyloxyethyl methacrylate) (pGEMA) with the lectin receptors was shown to
be broken down by the addition of monosaccharides [52]. pGEMA was
further cross-linked in order to entrap Con A within a hydrogel [53].
Their swelling ratio increased when free glucose concentration
increased, because Con A played the role of an additional cross-linker.
As it was progressively dissociated from the gel upon adding glucose,
the cross-linking density decreased and the gel swelled (Fig. 3).
5
Fig. 3. Schematic representation of hydrogel swelling upon glucose addition for a
polymer bearing glucose groups complexed with Con A.
Based on the concept of sol–gel transition, another type of polymer
with glucose moieties, vinylpyrrolidinone-allylglucose (VP/AG) copolymer was synthesized. The complex with Con A formed a hydrogel, but it
became a sol upon glucose addition. This sol–gel transition was shown to
be reversible upon glucose removal [54,55]. This hydrogel was then
sandwiched between two porous membranes and used as a gate
between two diffusion chamber cells [19]. The release of model proteins
(insulin and lysozyme) through the hydrogel membrane was studied as
a function of the free glucose concentration in the environment. The
release rate of model proteins through the glucose-sensitive hydrogel
membrane was dependent on the concentration of free glucose. This
study demonstrated that the glucose-sensitive phase-reversible hydrogels can be used to regulate the insulin release according to the free
glucose concentration in the environment.
A severe limitation to these systems was the progressive loss of
activity due to Con A leaking through the mesh, which also caused a
problem of toxicity. To circumvent this problem and obtain a reversible
glucose-responsive hydrogel, the lectin was covalently bound. Several
strategies were developed. Miyata et al. obtained a covalent link with
Con A by copolymerizing GEMA with Con A having vinyl groups [56].
Concanavalin A was covalently coupled to glycogen using derivatives of
Schiff's bases [57,58] or using carbodiimide chemistry with carboxylic
acid on Carbopol [59,60] or carboxylic acid modified dextran [20,61,62].
The carbodiimide coupling chemistry has the advantage that it
introduces no potentially cytotoxic groups into the gels formed, rendering them more suitable for potential in vivo applications. Such
materials showed a differential delivery of insulin in response to glucose
with in vitro diffusion experiments [20,61].
2.3. Hydrogels and microgels modified with a phenylboronic acid moiety
As discussed above, glucose-responsive materials can be developed from the combination of a hydrogel matrix and a natural recognition element of glucose such as lectin or glucose oxidase. A very
different method was opened by the pioneering work of Kataoka et al.
[23,63], who first used phenylboronic acid (PBA) as a non natural
receptor of glucose. PBA and its derivatives are known for their ability
to form complexes with polyol molecules such as glucose in aqueous
solution [64,65].
2.3.1. Glucose-swelling hydrogels
The pioneering studies reported a glucose-responsive system
based on the binding competition between glucose and another
polyol [63,66–68]. They synthesized copolymers bearing PBA moieties,
Author's personal copy
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V. Ravaine et al. / Journal of Controlled Release 132 (2008) 2–11
Fig. 4. Representation of the complexation between the (alkylamido)phenylboronic acid and glucose in aqueous solution.
(poly(N-vinyl-pyrolidone-co-PBA), and formed a complex with a long
polyol, poly(vinyl alcohol) (PVA) [66]. The complex between the PBA
entity and the polyol was dissociated in the presence of glucose, which
acted as a competing polyol. Each glucose molecule bound to one PBA
receptor, whereas the poly(vinyl alcohol) was linked to several sites
and played the role of cross-linker. The gel swelled and/or erodes, and
became a sol at high glucose concentrations. The concept was applied
to the design of a glucose-sensitive electrode [69]. The electrode was
coated with the poly(NVP-co-PBA)/PVA complex. The complex
dissociation upon glucose addition induced current changes proportional to glucose concentration. However, this complex was not stable
at physiological pH. The same team further showed that it could be
stabilized by the presence of amino groups in the vicinity of the PBA
moiety [63]. Therefore, a hydrogel with both functions was prepared
and formed a complex with gluconated insulin (insulin chemically
modified with gluconic acid) which dissociated upon glucose competition [68]. Insulin release in response to the glucose concentration
was achieved at physiological pH.
As well as the binding competition, PBA-derived hydrogels might
swell upon glucose addition following another mechanism, which
relies on an increase of their charge density upon glucose complexation. It is well-known that charged hydrogels exhibit higher swelling
ratios than neutral ones, when the swelling medium has a low ionic
strength. This is due to the presence of mobile counterions within the
gel which exert an osmotic pressure [27]. Therefore, the volume of the
gel increases when the charged density of a gel increases. This concept
was applied to PBA-derived hydrogels to give a glucose-responsive
material. Indeed, both the uncharged and the charged form of PBA
exist in equilibrium (Fig. 4). Upon complexation, this equilibrium is
shifted towards the charged form because the complex in the charged
form is more stable than that in the neutral form, which is highly
susceptible to hydrolysis [70]. Therefore, polymers bearing PBA
moieties undergo an increase of their charge density when the pH is
close to their pKA and PBA-modified hydrogels can swell upon glucose
complexation.
Although this electrostatic mechanism should be limited to media
with low salt concentrations, the PBA-derivatives could keep their
glucose-responsive properties at physiological salinity, when combined
with thermoresponsive polymers. Thermoresponsive polymers, such as
the family of poly-alkylacrylamides, are soluble at low temperature in
water and precipitate when heated above a critical temperature, called
critical solution temperature (LCST). When cross-linked, they are
swollen at low temperature and shrink upon heating. The transition
occurs at the volume phase transition temperature (VPTT). Kataoka et al.
synthesized copolymers with N,N-dimethylacrylamide and PBA [63].
They showed that the presence of the PBA moiety decreased the LCST of
the polymer, but the LCST increased upon glucose addition. This
indicates clearly that the PBA derivative plays the role of a hydrophobic
monomer but complexation with glucose imposes hydrophilicity. In the
case of a cross-linked hydrogel, it tends to shrink without glucose but
swells in the presence of glucose to expand the contact with water.
Likewise, hydrogels obtained from N-isopropylacrylamide (NIPAM) and
PBA swelled and shrank according to the glucose concentration [23].
This system was used to demonstrate that insulin could be released
repeatedly with an on–off regulation in response to stepwise changes in
the glucose concentration (Fig. 5).
This important proof of principle was demonstrated at pH 9.
The system was more recently improved to operate at physiological
pH conditions by modifying the chemical structure of the receptor
with an electron-drawing group in the phenyl ring [71,72]. Beads of
this hydrogel could be obtained using an inverse emulsion as a
template for polymerization. Their diameter was about hundreds
of micrometers [93]. In an attempt to design a glucose-responsive
chemical-valve, cylindrical gels were also prepared and shown to
effectively control the flow rate of the applied solution in response to a
change in the glucose concentration [73].
The pioneering work of Kataoka et al. was followed by many
developments in the area of PBA-derived hydrogels, in particular,
autonomous colorimetric sensors that could be engineered from such
hydrogels. Several research teams developed the concept of colorimetric sensors using a periodic array within the gel. This could be
obtained via three strategies: the inclusion of a colloidal crystal within
the gel [74,75], or a periodic array of voids within the gel [76–78], or
the gel itself forms an array in the case of a hologram [79]. The array
diffracts light according to Bragg's law. When the distance between
particles has the same order than the wavelength of light, one can
observe a colored material. The color depends on the interparticle
distance and therefore on the gel swelling. Any factor affecting gel
swelling thus modifies the crystal color. In fact, colorimetric sensors
can be obtained when combining the existence of a periodic array
Fig. 5. Repeated on–off release of fluorescently-labelled insulin from the pNIPAM-coAAPBA gel at 28 °C, pH 9.0, in response to external glucose concentration (from ref [23]
Fig. 3).
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7
within a molecule-sensitive hydrogel. This was achieved for glucosesensing materials using the inclusion of latex particles [80,81], using
an inverse opal [82,83] and making a holographic grating with the
hydrogel [84–86]. In all cases, the glucose-sensitive hydrogel
constituted the matrix. The use of an inverse opal was chosen to
accelerate the response time resulting from the quicker diffusion
through large voids. Holographic sensors are attractive for their mass
fabrication. An optical sensor could also be obtained using an inverse
opal without glucose sensitivity, which was filled with glucoseresponsive nanoparticles [87].
2.3.2. Glucose-shrinking hydrogels
The totally synthetic gels discussed above, swell according to the
target concentration and also according to the complexation constant
with the PBA group. Competition might arise between different
saccharides. For example, fructose has a stronger affinity than glucose
with the PBA group. Phenylboronates are usually not selective and
bind all molecules with a cis-diol group as mono-bidendate (1:1)
complexes. However, bis-bidendate (2:1) complexes can form involving one sugar with two cis-diols and two boronates. Among the
relevant physiological saccharides, glucose is the only molecule that
presents two diols which can be involved in a bis-bidentate complex
(Fig. 6). Therefore, glucose can be recognized selectively. The bisbidentate complex forms additional cross-links within a hydrogel.
Glucose-responsive hydrogel can thus shrink upon glucose addition.
Asher et al. were the first authors who reported a selective glucoseresponsive hydrogel which shrinks when the glucose concentration
increases [80,88]. The complex was stabilized by the presence of
sodium ions in the vicinity of the boronate group. Using the concept of
colorimetric sensors with a colloidal crystal inclusion, these authors
developed a colorimetric sensor which blue shifted at glucose concentrations in the patho-physiological range, at physiological pH and
salinity. Glucose determination could even be performed in tear fluid,
which made it suitable for the development of an integrated colorimetric sensor in soft contact lenses [89].
Samoei et al. synthesized another hydrogel sensor which shrank
selectively when glucose concentration increased and operated in human
blood plasma [90]. Their synthesis consisted in the chemical modification
of a commercial polymer, PMMA, via a very simple and modular batch
procedure. Hydrogel shrinkage occurred upon the formation of bisbidendate complexes between the PBA receptors and glucose. This
complex was thought to be favored by the presence of emerging cationic
charges from water and glucose bound to boron, which may reduce the
electrostatic repulsion and solvation by water that would lead to swelling.
They would therefore stabilize the cross-linking through bis-bidendate
binding to boron. The sensitivity and response time could be easily
optimized to selectively sense glucose within the range of 0 to 10 mM,
within 20 minutes.
The two previous examples were devoted to sensor applications.
Glucose-shrinking hydrogels have been also investigated for their
Fig. 6. Bis-bidendate (2:1) complex involving a glucose molecule and two boronates
(adapted from [80]).
Fig. 7. Response of microvalve to changes in glucose concentration between 0 mM and
20 mM in PBS solutions at pH 7.4, ionic strength 0.155 mM and at room temperature.
Fluid pressure at inlet is 6 kPa relative to outlet. From ref [91], Fig. 11.
ability to deliver insulin. Siegel et al. have also shown that acrylamide
hydrogels derived with PBA were able to either shrink or swell when
the glucose concentration increased, depending on the pH conditions
[26]. The hydrogels swelled upon glucose addition at pH 7.4 and
shrank at pH 9.0 where they recognized glucose selectively compared
to fructose. With this hydrogel, these authors designed a simple
microvalve prototype inserted into a microfluidic device [91]. Insulin
was delivered with an on–off regulation according to the glucose
concentration. However, insulin flow was stopped when the gel was
swollen, i.e. when glucose concentration increased at physiological pH
and the response time was too long (20 min) (Fig. 7). They further
improved the system and designed another microvalve, the basic
structure of which is a silicon membrane having an array of orifices in
which a hydrogel is anchored [92]. The hydrogel allows gating the
flow across the membrane. In the swollen state, the hydrogel
completely occupies the void space of the orifice, blocking pressuredriven fluid flow. In the shrunken state, the hydrogel collapses and
leaves a void, allowing fluid to flow through an opened annular gap.
The response to glucose variations occurs within 10 min.
2.3.3. Glucose-responsive microgels
Most of the hydrogels cited above present the defect of exhibiting
very long response times. Even beads (with a diameter of several
hundred of micrometers) displayed response times of several hours
[93]. As mentioned previously, the kinetics could be improved to
20 min by incorporating voids or macropores within the gel [83]. The
same order of magnitude was obtained by Samoei et al. [90]. It is wellknown that the shrinking rate of the gel is inversely proportional to
the square of the smallest dimension of the gel [27,28]. Therefore, the
kinetics properties of macroscopic hydrogels do not seem suitable for
fast insulin controlled release but these systems can be scaled down to
colloidally stable particles made from hydrogels, which are referred to
as microgels or nanogels. Their size is generally comprised between a
few tens of nanometers to a few micrometers.
The first report concerning PBA-modified microgels was published
by Hazot et al. [94]. They structure achieved the copolymerization of a
PBA monomer with a thermoresponsive monomer. They proved the
presence of the phenylboronate group at the microgel surface but did
not show any glucose-triggered swelling response. Later on, several
groups reported the synthesis of glucose- and thermoresponsive
microgels bearing PBA functions [95–97]. Two synthesis methods are
available: the PBA receptor can be grafted to a thermoresponsive
microgel bearing carboxylic groups [95,97] via a carbodiimide coupling, or a PBA monomer can be synthesized and copolymerized with
other monomers to be statistically incorporated within the microgel
[96]. The post-modification method has the drawback that not all the
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carboxylic groups are substituted because the PBA cannot diffuse deep
inside the microgel and the chemical modification is effective mainly
at the surface. However, the particles obtained by the two methods
had a similar size, in the range of 200–400 nm. They swelled upon
glucose addition at pH above 8.5. In addition, our group showed that
they were still responsive at physiological salinity [96]. The synthesis
of the microgels has been extended to hollow nanocapsules [98].
These objects were obtained by synthesizing core-shell microgels
with a degradable core and a glucose-responsive shell. The core was
made of cross-linked hydrogel with a degradable cross-linker. The
debris of the degraded polymer was released through the shell,
leaving a hollow capsule which could be further filled with insulin.
Recently, Hoare and Pelton [24] produced amphoteric thermoresponsive microgels with PBA functional groups that were glucoseresponsive at physiological temperature, salinity and pH. The composition of the microgels could be adapted to produce microgels either
swelling or shrinking in response to glucose. These amphoteric microgels also showed interesting properties concerning insulin incorporation based on electrostatic interactions. Insulin uptake was high
when the net charge of the microgels and insulin was opposite. A
larger amount of insulin was released when the glucose concentration
was raised to 2 g/L compared to no glucose.
3. Modified insulin
Self-regulated insulin delivery was addressed by another approach
developed by Hoeg-Jensen et al. [99,100]. These authors used the glucose
sensor as part of the insulin molecule. Insulin was modified with a
phenyl boronate group. Their idea was to modify the diffusion rate of
insulin according to the glucose concentration. It is known that the
diffusion rate of insulin is strongly dependent on the form of the injected
depot, because it is proportional to its half-time. For example, fast acting
insulin is monomeric (t1/2 ~ 1 h) [101], hexameric insulin is intermediate
acting (t1/2 ~ 2–4 h), whereas long acting insulin is obtained from binding
to high molecular weight proteins [102] or from soluble high molecular
complexes (t1/2 N10 h) [103]. Therefore, Hoeg-Jensen et al. developed the
idea that the molecular weight of boronated insulin self-assemblies
could be controlled by the glucose concentration. Insulin was modified
with a pair of boronate and carbohydrate to provide sufficient driving
force for the formation of hexamer–hexamer self-assembly, as illustrated
on Fig. 8 [100]. The self- assembly of hexamers was achieved only when
the N-terminal B-chain of insulin was alpha-helical, i.e. in the presence of
phenol. The addition of D-sorbitol disassembled the complex in a dosedependent manner in less than 5 min. The effect of D-glucose was much
less pronounced.
In a related approach, the same authors synthesized a series of
novel insulins using sulfonamide derivatives of phenylboronates,
compatible with pH physiological conditions. They were immobilized
Fig. 8. Illustration of insulin self-assembly and disassembly under carbohydrate control
(from ref [100]).
Fig. 9. Insulin release upon batch-wise washings with buffer or buffered D-glucose
solutions (5, 25, 50 mM) at pH 7.4, from monomer (squares) and Zn(II) insulin hexamer
(triangles) formulation immobilized on D-glucamine polymer, (from [99], Fig. 3).
on a glucamine-derived polyethylene glycol polyacrylamide as a
model. Their dose-responsive glucose-triggered release was demonstrated (Fig. 9) [99]. By formulating insulin as a traditional Zn(II)
hexamer, a steeper glucose sensitivity and release profile was gained
compared to monomeric insulin formulation.
4. Conclusion and perspectives
4.1. Treating diabetes with self-delivery materials
The goal of type 1 diabetes treatment is to achieve tight glucose
control, to avoid chronic complications, while limiting the frequency
of hypoglycemic episodes in day-to-day life. Although considerable
efforts have been made to improve pharmacokinetics of insulin and to
develop user-friendly monitors, this goal remains difficult to achieve
and an automated artificial pancreas is still required. Considerable
work remains to be done to develop algorithms for the automated
regulation of glucose by insulin. Furthermore, even if insulin pumps
and glucose sensors have been miniaturized, such devices remain
cumbersome for the patient.
Other strategies for the cure of diabetes are also on their way. The
transplantation of islets of Langerhans was made possible by developing methods for the isolation of islets. Nevertheless, their rate of
success is still low and the number of donors (two to three pancreases
per patient) remains insufficient. With the progress of molecular
biology, insulin secreting cells can be created but these cells are still
far from having the fine-tuning of the genuine ß-cell in response to
glucose. The use of human embryonic stem cells still represents a
scientific challenge and an ethical issue.
Alternative strategies have therefore to be considered. Progress in
the field of biomaterials, drug delivery and insulin provide tools for
constructing an “intelligent” insulin vector. The ideal vector would
enable insulin delivery in a self-regulated and long acting but modulated manner. Self-regulated since in such a system the necessary
amount of insulin could be administered in response to the blood
glucose concentration and therefore replace the need of blood glucose
monitoring for the patient. Long acting and modulated, to avoid the
need of multiple day insulin injections. In other words, such biosystems aim at mimicking natural feedback by combining sensor
function (to perceive glucose modification) and effector function
(insulin delivery) not to cure the patient, but at least to deliver him
from the heavy burden of the disease in terms of day-to-day treatment
but also of the possibility of chronic complication.
However, these glucose-responsive materials still require further
research before any clinical application could be envisaged.
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4.2. Challenges for the future
The challenges remain to find a device which would combine:
- sensor specificity, which means that the sensor has to respond only
to glucose variations;
- pharmacokinetics similar to normal pancreatic activity ;
- easiness of administration, possibly by simple long lasting
injections (ideally a weekly or monthly subcutaneous injection);
- biocompatibility without any in vivo toxicity, and possibly no long
term side effect.
A classification of the existing methods is proposed in Table 1. It
relates their advantages and disadvantages with regard to the above
criteria. Comments about the future are given below.
It has been shown that three types of receptors could be used. It is
generally admitted that natural receptors like proteins recognize
specifically target molecules, whereas artificial receptors are non
specific. However, this statement has to be refined in the area of
glucose sensing. Among natural glucose receptors, only glucose oxidase specifically recognizes D-glucose, whereas lectins are non specific
receptors. Artificial receptors are represented by the family of
phenylboronate derivatives, able to form mono-bidendate (1:1) complexes as well as bis-bidendate (1:2) complexes with 1,2-diols such as
carbohydrates. The 1:2 complexation promotes glucose selectivity
over other common blood sugars such as fructose and galactose. In the
case of 1:1 complexes, fructose is known to bind to boronate with a
higher affinity constant than glucose. However, since the physiological
concentration of glucose is approximately 100 times higher than
the fructose one [104], this competition is likely not to be a major
issue.
Several physical factors are known to affect the insulin release
kinetics through hydrogels. Some of them have already been tested: it
is well-known that insulin diffusion depends on the gel structure such
9
as porosity or dimensions. Therefore macroporous hydrogels have
been investigated and showed an improved response. Due to their
small dimensions, microgels represent another solution to improve
insulin delivery kinetics. Up to now, very few data are available
concerning insulin delivery due to the very recent publication of the
material synthesis. Their size should influence the pharmacokinetics
but also their internal structure and their chemical nature, as it could
influence the strength of insulin interaction with the gel matrix. It is
expected that this field will become more active within the next few
years.
Insulin self-regulated delivery systems can also present several
forms which will have to be related to the route of administration. A
cross-linked gel is a highly elastic material, which cannot be administered via a simple injection but probably as an implant. The
easiest route of administration is certainly the subcutaneous injection
of a liquid, such as a suspension of microgels, a pre-gel solution or a
glucose-responsive insulin depot. Ideally, this injection should enable
the insulin vehicle to reach blood circulation, which would prevent
from any time lag between glycemia variations and insulin release.
Microgels are the best candidates to achieve this goal. Their size below
0.5 µm makes them suitable to go through blood capillaries without
any obstruction. In return, microgels should have a prolonged circulation time to deliver insulin during several hyperglycemia
episodes. A fast clearance from blood circulation has to be avoided
and a major issue will be to render them stealthy. Another interesting
form of insulin delivery is in situ gelling systems, which can be
injected in the liquid state and form a gel at body temperature. This
method has recently shown promising results. Nevertheless, if administered subcutaneously, the issue of time lag remains unsolved.
Insulins with built-in boronated glucose sensors might also be injected subcutaneously. They are designed to be administered similarly
to long acting commercially available insulin depots, except that their
action time should depend on the glycemia. This field of research is
Table 1
Summary and classification of the different methods for chemically closed-loop insulin release
Sensor
Mechanism for self-regulated insulin delivery
GOD
Swelling of a pH-sensitive hydrogel bearing amino groups
(with or without catalase [31–34])
With PEG grafts [41–44]
Gating membrane : cellulose [25] or poly(vinylidene) [38,39]
membrane having pores grafted with PAA
Degrading matrix : cross-linked poly(methacryloyloxyethyl
phosphorylcholine) [40]
In situ gelling system: chitosan + β-GPS [22]
Advantages
Blood compatibility and biocompatibility
Biocompatible Suitable for subcutaneous injection
Con A
Disruption of the complex with SAGP-insulin [47–50]
Dissolution of the complex with polysaccharides or polymers
with saccharides residues such as p(GEMA) [51] or
p(VP-co-AG) [54,55]
Swelling of a hydrogel by removal of Con A (cross-linker) [53] Insulin release obtained in the patho-physiological range
with Con A leaking or Con A covalently attached [56–62]
for glucose [61]
PBA
Disadvantages
Selectivity over other carbohydrates
Exists as a membrane [17], a capsule [21], or macroporous
membrane [18]
Biocompatible membrane [31,34]
Minimizes immunoreactions and subsequent rejections by
the body
Mechanical resistance
Allows a wide variety of chemistries
Binding competition with PVA complexed with p(NVP-co-PBA)
[66]
Binding competition with gluconated insulin complexed with
amine containing PBA gel [68]
Change in the LCST of thermoresponsive polymers
Exist as bulk hydrogels, beads [93] microgels [24,94–97] or
[23,63,71,72]
nanocapsules [98] with possible faster response time and
fluid preparation suitable for injection
Microgel size is well-adapted to join blood circulation
Selectivity compared to other physiological sugars
Used as a microvalve in an insulin delivery device [26]
Obtained also for microgels [24]
Dissociation of the hexamer–hexamer self-assembly of insulin No need of polymer matrix
modified with a PBA-carbohydrate pair [100]
Competition with other saccharides
Possible toxicity
Requires a specific insulin analog
Not reversible
Not suitable for injection
Hydrogel polymerized by UV irradiation
Not suitable for injection
Competition with other polyols
Not reversible
Requires a specific insulin analog
Possible clearance from blood circulation
Hydrogel shrinkage by bis-bidendate formation
[24,26,80,88,90]
Uncertain bioactivity of the insulin analog
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still in its infancy and will require significant further work to become
applicable.
Biocompatibility and non toxicity are finally the last but not the
least requirements. Diabetes treatment is a long term treatment,
which does not tolerate any possible side effects. Many hydrogels are
known to be highly biocompatible. This can be explained by their high
water content and their similarity of structure with the extracellular
matrix, in particular for polysaccharide hydrogels. More generally, all
natural biopolymers possessing a high degree of functional groups can
be cross-linked to obtain three-dimensional structures, and further
scaled down to the design of microgels. This is the case of chitosan-,
hyaluronan-, or polyaminoacid-based hydrogels [105]. Successful in
vivo applications have already been obtained with hydrogels [106].
Biodegradability is also desirable, at a larger time scale than that used
for insulin delivery. Biodegradation of hydrogels can be achieved by
several mechanisms including enzymatic, hydrolytic or environmental (like pH, temperature, etc). A number of degradable cross-linkers
have already been reported, such as acetals, which can be degraded
under acidic environment (pH b 6.5) [107], or disulfildes, which can be
degraded in the presence of a reducing agent including a tripeptide
glutathione [108,109]. They might be incorporated into glucose-responsive hydrogels to build up the ideal insulin delivery system.
Therefore, many tools are now available to design a chemically controlled closed-loop device for in vivo use. Although some issues still
have to be overcome, this field of research is promising as a possible
alternative to other approaches for diabetes treatment.
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