Stimuli‐responsive delivery of therapeutics for diabetes treatment

Abstract Diabetic therapeutics, including insulin and glucagon‐like peptide 1 (GLP‐1), are essential for diabetic patients to regulate blood glucose levels. However, conventional treatments that are based on subcutaneous injections are often associated with poor glucose control and a lack of patient compliance. In this review, we focus on the different stimuli‐responsive systems to deliver therapeutics for diabetes treatment to improve patient comfort and prevent complications. Specifically, the pH‐responsive systems for oral drug delivery are introduced first. Then, the closed‐loop glucose‐responsive systems are summarized based on different glucose‐responsive moieties, including glucose oxidase, glucose binding protein, and phenylboronic acid. Finally, the on‐demand delivery systems activated by external remote triggers are also discussed. We conclude by discussing advantages and limitations of current strategies, as well as future opportunities and challenges in this area.

regulation of meals can effectively delay disease progression. 7 But for advanced type 2 diabetic patients, injection of insulin or other diabetic therapeutics, like glucagon-like peptide 1 (GLP-1), are also required to control the BGLs. 10,11 However, the traditional subcutaneous injection is painful and inconvenient. Moreover, this open-loop administration is often associated with inadequate glucose control. 10 In recent decades, a number of technologies have been developed to overcome the limitations and drawbacks of such injection therapy. Among them, strategies utilizing stimuli-responsive delivery systems are highly desirable to regulate glycemia with minimal patient effort and improve the quality of life for diabetic patients. 12,13 In this review, we will focus on the different stimuli-responsive systems for diabetic therapeutic delivery, including pH-responsive, glucose-responsive, and external physical triggeractivated systems ( Figure 1). The future opportunities and challenges will also be discussed.

| P H -R E S P O N S IV E SY S T E M S FOR ORAL D RUG DE LIVE RY
Oral delivery is considered to be the most patient-friendly method for insulin administration. [14][15][16][17] However, the low pH of gastric medium in the stomach and various digestive enzymes in the gastrointestinal tract may degrade insulin, leading to low efficacy of therapy. 18,19 To enhance the bioavailability and treatment efficacy, numerous pH-responsive based systems have been exploited, such as enteric capsules and particles, 20 which are able to protect protein drugs from the harsh gastric environment, as well as allow specific release in the intestine. [21][22][23] These drug carriers are stable under acidic conditions in the stomach, but can rapidly release cargoes under neutral pH in the intestine. In the following section, we will introduce the recent pH-responsive oral delivery systems based on polymeric and inorganic materials.
Peppas and coworkers pioneered the utilization of pH-responsive complexation gel for oral delivery of insulin. 22,[24][25][26] They loaded insulin into poly(methacrylic-g-ethylene glycol) (P(MAA-g-EG)) hydrogels, and orally administered these polymeric microspheres to streptozotocininduced type 1 diabetic rats. 22 The hydrogels were prevented from swelling under the acidic condition in the stomach because of the formation of intermolecular polymer complexes by hydrogen bonding between the carboxylic acid protons and the etheric groups on the grafted chains, which efficiently protected insulin form degradation. In the neutral environment of the intestine, the intermolecular polymer complexes disassociated and the gel swelled, resulting in a rapid insulin release (Figure 2a, b). They demonstrated that the insulin was sufficiently absorbed in the upper small intestine with an obvious hypoglycemic action; the bioavailability was 4.6-7.2% in a diabetic rat model. 24 They further conjugated insulin with transferrin to enhance the bioavailability of oral insulin. 25 Transferrin can be uptaken by the epithelial cells to increase the permeability of insulin across the epithelial barrier, and it can also prevent insulin degradation by intestinal enzymes. Integrating the transferrin with the pH-responsive hydrogels resulted in a 22-fold net increase in insulin permeability. In another work, they modified P(MAA-g-EG) with wheat germ agglutinin (WGA) to improve mucoadhesive ability. 26 WGA is a class of lectins, which can bind to the luminal surface of the small intestine. 27 In addition, WGA showed minimal binding to mucin at a low pH in the stomach. They revealed this system could significantly improve the overall adhesion to a cellular monolayer in an in vitro study.
In addition to synthetic polymers, natural polymers such as chitosan (CS) also show a potential role in pH-responsive oral insulin delivery. 23,[28][29][30] CS is a natural cationic polysaccharide, which exhibits good mucoadhesiveness and the capability to open tight junctions. 28 More importantly, its physicochemical properties also depend on the surrounding pH value. Sung and coworkers developed a pH-responsive NP consisting of CS and poly(c-glutamic acid) (c-PGA) for oral insulin delivery. 31 Since both CS and c-PGA were ionized at pH 2.5-6.6, the prepared NPs were stable by the electrostatic interaction between CS and c-PGA. In contrast, at pH 7.4, CS was deprotonated, which led to the collapse of NPs and subsequent insulin release. Furthermore, the positive charged CS shell could increase the paracellular permeability by transiently opening the tight junctions. After oral administration in diabetic rats, these NPs could effectively decrease the BGLs. To improve the stability of the NPs in a broader pH range, magnesium sulfate (MgSO 4 ) and tripolyphosphate sodium (TPP) were further introduced to construct a multi-ion-crosslinked NPs [32][33][34] (Figure 2c). The introduction of MgSO 4 and TPP also significantly increased their loading efficiency and content of insulin. The in vivo results demonstrated a significant hypoglycemic effect in diabetic rats by oral administration, and the corresponding relative bioavailability of insulin was about 15%.
Through conjugating thiol groups on the CS chain, its mucoadhesive capability can be enhanced due to the disulfide formation between thiolated CS and the cysteine residues of the mucin. 35 36 These polyelectrolyte NPs exhibited a significant enhanced mucoadhesion and permeability in rat intestine, which allowed a better hypoglycemic effect.
Enteric capsules or enteric coating are also utilized to improve the drug efficiency by protecting insulin from the digestive enzymes in the stomach. Mitragotri and coworkers loaded mucoadhesive intestinal patches in an enteric capsule for oral insulin delivery. 37 The capsule could protect insulin-loaded patches in the acidic environment of the stomach, while release them in the intestine. The released patches adhered to the intestinal mucosal layer to promote insulin absorption. Recently, they further loaded the intestinal devices with a permeation enhancer into a capsule coated with a pH-responsive enteric coating to improve oral absorption of insulin. 38 Sung and coworkers also filled the freeze-dried insulin-loaded NPs into an enteric-coated capsule. 39 The Eudragit ® S100 or Eudragit ® L100-55-coated capsules remained intact in the acidic environment of the stomach, while rapidly dissolved and released NPs in the neutral environment of the intestine. In vivo studies in diabetic rats indicated that relative bioavailability of insulin was approximately 20%.
Apart from polymeric materials, inorganic NPs have also been explored as insulin carriers due to their high loading capability and good compatibility with insulin. Sun et al. utilized mesoporous silica NPs (MSN) to increase the loading capacity of insulin. 40 These inorganic NPs were further coated with pH-sensitive dextran-maleic acid (Dex-MA) and then grafted with glucose-sensitive 3-amidophenylboronic acid (APBA). The Dex-MA shell could shrink and block the pores to inhibit insulin release in the acidic pH of the stomach, but swell to allow insulin diffuse at neutral pH in the intestine. Verma et al.
designed a vitamin B12 functionalized layer by layer calcium phosphate NP to improve oral delivery of insulin. 41 Vitamin B12 was chosen due to its high paracellular and receptor mediated uptake efficiency in the intestine, as well as its low pKa (1.8) that changed the surface charge of NPs as a function of pH.

| G L U C O S E -R E S P O N S IV E D R U G D E LI V E R Y
Closed-loop based smart insulin delivery that can mimic the pancreas' b-cells to secrete insulin in response to hyperglycemia has gained an increasing amount of attention. 12,13,45,46 Usually, such closed-loop delivery systems are composed of a glucose monitoring module and a glucose-triggered insulin releasing module. 12,45 One notable example is the electronic/mechanical insulin pump, which consists of a continuous glucose sensor and an external insulin infusion pump. 47,48 The insulin infusion rate can be adjusted in response to the BGL signal from the glucose sensor in these wireless, portable, and wearable systems. 49,50 Conversely, synthetic smart insulin delivery systems have been heavily explored to achieve closed-loop delivery from the use of material design and formulation engineering. 12,13,45,46 Typically, these chemical closed-loop systems, which based on glucose-responsive materials, can sense an increase in BGLs and respond accordingly to release a certain amount of insulin for glycemia regulation. 12,45 Glucose oxidase (GOx), glucose binding proteins (GBPs), and phenylboronic acid (PBA) are the three most used glucose-sensing moieties. 45,46 We will introduce the mechanism of these diverse systems, respectively, and discuss the most recent advances in this section.

| GOx-based systems
GOx is the most prevalent among the various glucose-responsive moieties described in the literature. 13 GOx presents a high level of specificity for glucose, and it converts glucose into gluconic acid in a biological environment: Accompanying the oxidation of glucose, a local acidic, hypoxic, and high H 2 O 2 concentration environment will be generated. 51 Based on these changes, several stimuli-sensitive materials have been developed to achieve glucose-responsive action. [52][53][54][55] With the generation of gluconic acid, local pH decreases rapidly, which can serve as the trigger of insulin release. [56][57][58][59] For example, Ishihara and coworkers developed a pH-sensitive poly(amine) membrane for glucose-responsive delivery. 60 In the presence of glucose, the protonation of tertiary amino groups under acidic conditions increased the water content of this poly(amine) membrane, further enhancing the permeability of insulin. Langer and coworkers designed a GOximmobilized polymeric system to achieve the controlled release of insulin. 61 The reduced pH due to the enzymatic reaction of glucose increased the solubility of trilysyl insulin dramatically, leading to a higher drug diffusion rate in response to glucose concentration. Peppas and coworkers utilized a pH-sensitive hydrogel for glucose-responsive insulin delivery. 62 to stable GOx. The mechanical properties of crosslinked BSA were further improved due to the introduction of MnO 2 NPs. Under a high glucose concentration, a decrease in pH generated by glucose oxidation caused the pH-sensitive hydrogel NPs to shrink, leading to a fourfold increase in insulin release. In addition, the same group utilized this bioinorganic nanocomposite membrane as a glucose-responsive plug to regulate insulin release from a reservoir made of medical grade silicone. 65 In vitro experiments showed that the insulin release profile exhibited a typical pulsatile pattern when changing glucose concentration between normal and hyperglycemic levels for several cycles.
Through intraperitoneal implantation, the BGLs of diabetic rats could be controlled in the normal range for 4 days. Furthermore, the glucoseresponsiveness of the device was also demonstrated in diabetic rats by a glucose challenging. The BGLs return from a hyperglycemic state to the normal level within 60 min.
Anderson's group also reported a pH-sensitive sponge-like matrix as an insulin reservoir. 52 The glucose-responsive microgels were prepared by crosslinking chitosan using TPP, a medical biocompatible polymer, insulin, and glucose-specific enzymes (GOx and CAT), which were entrapped in this non-covalent crosslinked polymeric matrix. To improve stability and immunogenicity of enzymes, both GOx and CAT were encapsulated into the nanocapsules. When subjected to hyperglycemic conditions, hydrogen ions were generated during glucose oxidation by GOx protonated chitosan and the microgel system swelled more than fivefold in volume, subsequently enhancing insulin release from the swelling sponge.
Besides pH-induced volume change, acid hydrolysis is also applied in GOx-mediated glucose-responsive systems. Langer, Anderson, and coworkers developed an injectable and acid-degradable polymeric network for self-regulated insulin delivery. 53 Insulin, GOx, and CAT were entrapped in an acid-sensitive nanoparticle made of acetal modified dextran (m-dextran) through a double emulsion-based solvent evaporation/ extraction method. In this system, m-dextran was hydrolyzed into water-soluble dextran in the presence of produced gluconic acid, and caused the dissociation of the NPs, leading to subsequent insulin release.
To make the formulation injectable and overcome burst release, the NPs were coated with oppositely charged polymers, chitosan, and alginate, respectively, and mixed to form a nanocomposite-based porous network. The gel-like nano-network effectively dissociated to release insulin at a hyperglycemic level (400 mg/dL), but showed insignificant release at a normal level (100 mg/dL). After a single subcutaneous injection into type 1 diabetic mice, the nano-network was demonstrated to provide improved glycemia regulation for up to 10 days.
Using a similar response mechanism, Gu and coworkers synthesized a pH-sensitive diblock copolymer, PEG-Poly(Ser-Ketal), and pre-pared the polymersome-based nanovesicles to encapsulate insulin and glucose-specific enzymes. 66  was bioreduced to a water-soluble product, resulting in the dissociation of GRVs and subsequent insulin release. In vitro results revealed that the GRVs with a novel trigger mechanism presented a fast reversible control of insulin release between a normal and hyperglycemic state.
To achieve ease of administration, GRVs were loaded in to a microneedle (MN)-array patch for painless and continuous insulin delivery 69,70 ( Figure 3b, c). Under a normal BGL, the GRVs were stable in the MN; however, they quickly disassembled to release insulin when exposed to high interstitial fluid glucose in vascular and lymph capillary networks.
In vivo experiments displayed that one patch could decrease BGL to a normoglycemic state within 30 min and could be maintained for up to 4 hr without any risk of hypoglycemia in a type 1 diabetic mouse model

| Glucose-binding protein-based systems
Another non-enzymatic strategy to achieving glucose-triggered insulin delivery is based on the reversible interaction between glucose and glucose binding moieties. GBPs, such as lectins, are a group of natural carbohydrate-binding proteins. 74,75 One of the most commonly used lectin for glucose-responsive insulin delivery is concanavalin A (Con A), which is derived from the jack bean. 76 Con A is formed with two dimers and has four binding sites for D-glucose, D-mannose, and polysaccharides. [77][78][79] Kim and coworkers utilized gluconic acidmodified insulin (G-Ins) to complementarily bind to Con A to achieve a glucose-induced release of insulin. 79-81 When exposed to glucose solution, insulin can be released through competitive binding.
An alternative glucose regulated insulin delivery matrix is based on the affinity between Con A and natural polysaccharide polymers. [82][83][84] Ying and coworkers utilized Con A as a crosslinker to form a dextranbased NP. 85,86 Once exposed to an elevated glucose concentration, these NPs quickly dissolved due to the competition binding of glucose to Con A, leading to the release of encapsulated insulin. Nakamae and coworkers developed a glucose-responsive hydrogel consisting of Con A and poly(2-glucosyloxyethyl methacrylate) (poly(GEMA)). 87 The introduction of Con A increased crosslinking density due to the complexation between Con A and poly(GEMA). In the presence of glucose, the

| PBA-based systems
Besides natural glucose-binding molecules, synthetic molecules have also been used in diabetes diagnosis, glucose sensors, and glucoseresponsive insulin delivery. 92,93 PBA is a synthetic molecule that is able to reversibly bind to 1,2-or 1,3-cis-diols, including many kinds of sugar, to form cyclic esters. 92,94,95 Since the first discovery of the reversible interaction between PBA and sugar in 1959, 96  encapsulate cAMP in the mesopores through a PBA-diol interaction.
Once exposed to glucose, the capped G-Ins were replaced by free glucose molecules, and the loaded cAMP was simultaneously released to achieve the regulation of glycaemia.
Glucose-responsive systems can also be achieved by complexation between PBA-containing polymer and polyol polymers like polyvinyl alcohol (PVA) and glycopolymers. 104  Besides competitive binding with PBA by free glucose, an alternative approach to achieve glucose responsiveness is to utilize the charge change of PBA when interacting with glucose. 45

| Ultrasound-triggered drug delivery
Painless transdermal drug delivery is a potential means of drug administration. However, macromolecular drugs, like proteins, are difficult to transport by this method due to the extremely low permeability of skin. 70 Mitragotri and coworkers employed low-frequency ultrasound to enhance the permeability to these drugs. [115][116][117]  was readily diffused into body (Figure 5a and b). Moreover, the application of ultrasound did not decrease the barrier properties of the skin after 24 hr. In vivo studies showed after transdermal delivery of insulin with ultrasound, the BGLs could significantly reduce to a normal range in rats (Figure 5c and d), and even in a large animal model-pig. 118,119 Kwok et al. designed a polymeric monolith coated with an ultrasound-responsive shell for insulin delivery. 120 Relatively impermeable and ordered methylene chains formed this ultrasound-controlled "on-off" switch by self-assembling. In the absence of ultrasound, the orderly structure was able to keep the drug inside the polymer carrier.
While with the assistance of ultrasound, the encapsulated insulin

| Other physical signals-triggered drug delivery
Electrical potentials have also been exploited as a trigger to activate insulin release by Shi and coworkers. 128  Friedman and coworkers applied a modified thermo-sensitive channel, transient receptor potential cation channel subfamily V member 1 (TRPV1) for radio wave-mediated secretion of insulin. 113 The temperature-sensitive TRPV1 was decorated with antibody-coated iron oxide NPs. When exposed to radio waves, the iron oxide NPs increased the local temperature, further stimulating TRPV1 to gate calcium and subsequent release of bioengineered insulin driven by a Ca 21 -sensitive promoter. More recently, the same group utilized radio waves or magnetic fields to control glucose homeostasis without the exogenous ferritin NPs. 130 The activation of a glucose-sensing neuron depot-based systems may result in hypoglycemia, which may cause a variety of symptoms including behavioral and cognitive disturbance, seizure, loss of consciousness, brain damage, and even death. 136 A combination delivery system 137 that can simultaneously release glucagon in response to low BGLs is envisioned to avoid risk of hypoglycemia. Third, it is important to achieve real-time glucose monitoring for external remote triggered system. A non-invasive glucose sensor may act as an ideal activator to control the external signals to promote drug release. Last but not least, biocompatibility and non-toxicity of stimuliresponsive formulations should be carefully assessed. Since diabetes treatment is a long-term, even lifetime, administration any possible side effects may lead to serious problems. 138