Smarter glucose-sensitivity of polymeric micelles formed from phenylborate ester-co-pyrenylboronic ester for insulin delivery at physiological pH

Abstract

Exploring an intelligent system capable of releasing insulin in response to glucose level changes would improve the therapeutic potential in diabetes. Herein, we developed a dispersing glucose-responsive group strategy to effectively enhance glucose-sensitivity of the system for the self-regulated delivery of insulin in response to physiological need. One kind of amphiphilic block polymer, which was named poly(ethylene glycol)-block-poly[(2-phenyl boronic esters-1,3-dioxane-5-ethyl) methylacrylate-co-(2-pyrenyl boronic esters-1,3-dioxane-5-ethyl) methylacrylate] MPEG-b-P(PBDEMA-co-PyBDEMA), was fabricated by one-step atom transfer radical polymerization (ATRP) of two monomers PBDEMA and PyBDEMA with MPEG5000-Br as a macroinitiator. These amphiphilic polymers MPEG-b-P(PBDEMA-co-PyBDEMA) were self-assembled to form polymeric micelles with a hydrophobic core composed of both glucose-responsive PPBDEMA and strongly hydrophobic PPyBDEMA. As a result, these polymeric micelles exhibited very slow insulin release at glucose concentration of 1.0 mg mL⁻¹ (normoglycemia) and relatively rapid release behavior at 3.0 mg mL⁻¹ (hyperglycemia) at pH 7.4. Moreover, the glucose-triggered on-off release of insulin was further investigated at pH 7.4 with alternate 1.0 and 3.0 mg mL⁻¹ glucose incubation to exhibit an effective self-regulated insulin delivery in response to physiological glucose level fluctuation. In addition, the encapsulation efficiency (EE) and loading capacity (LC) of insulin were distinctly enhanced to nearly 65% and 20%, respectively. This type of nanocarrier may be a promising candidate for in vivo insulin delivery.

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1. Introduction

Insulin-dependent diabetes is a serious disease affecting many people in the world.^1,2 Although different strategies for insulin delivery such as oral, nasal, buccal and transdermal drug-delivery systems have been widely explored,^3–6 these non-invasive routes known as open-loop insulin delivery,^7,8 like insulin injection, still require frequent blood sampling to measure the blood glucose level and adjust the dose and dosing time of insulin, which essentially cannot shake off the major disadvantages of pain, inconvenience, cost, infection, inability to handle insulin and the local hypertrophy.⁹ Great efforts are devoted to explore glucose-responsive materials for smart insulin delivery systems which can detect elevated glucose levels and automatically dispense insulin as needed.^10–12 Amongst, purely synthetic glucose-responsive components have special superiority including their versatility for different designs and better stability, and attract increasing attention for self-regulated insulin release systems.

Because phenylboronic acid (PBA) can complex with glucose and form stable phenylborates,¹³ a series of studies on PBA-based glucose-responsive materials have been explored for the construction of self-regulated insulin delivery systems.^14,15 Two critical issues, which is effective glucose-responsiveness at pH 7.4 and at reasonable concentration of glucose (1–3 mg mL⁻¹), are demanded for one ideal boronic acid-based glucose responsive insulin delivery system (GRIDS). By far, great progress has been achieved on decreasing the apparent pK_a of PBA-based glucose-responsive materials to realize insulin release at neutral pH. For example, Kataoka pioneered to explore phenylborate derivatives with electron-withdrawing groups at para- or ortho-position of boronic acid group to possess an appreciably glucose-responsive behavior at pH 7.4.^16–19 The construction of boronate ester between PBA and diol-containing molecules to reduce pK_a of aryl boronic acid achieved significant glucose-responsiveness under physiological pH, as well.^20–26

Besides, the reasonable glucose sensitivity is another hot challenge. One ideal GRIDS is demanded to finely tune the system so it shows a gated response to the change in glucose concentration critically at the level of normoglycemia, ca. 1 mg mL⁻¹. Based on this object, Miyahara reported one kind of smart gels capable of undergoing intelligent glucose-responsiveness under conditions closely related to human glucose homeostasis.^19,27 Very recently, Zhang's group introduced novel glucose-responsive system with the diblock glycopolymers self-assembling into nanoparticles that exhibited different glucose-responsiveness under the physiological and pathologic condition.^24,28 Shi et al. also developed one class of complex micelles showing notable glucose-responsiveness under physiological conditions.^29,30 Despite such exciting progress satisfying the substantial need for GRIDS, the principles to solve this critical issue are relatively insufficient comparing to decreasing the apparent pK_a of PBA-based materials. Therefore, it is still an extremely important and urgent job to explore more alternative strategies to realize insulin release only after the glucose level exceeding normal level and no release under normoglycemic condition.

Recently, we reported a kind of a glucose-responsive system self-assembling from poly(ethylene glycol)-block-poly[(2-phenylboronic esters-1,3-dioxane-5-ethyl) methyl acrylate] (MPEG-b-PPBDEMA) based on the competition mechanism between sugar molecule and acyclic diol and the characteristics of boronate ester with low pK_a, where significant glucose-responsiveness was obtained under pH 7.4.^25,26 However, these nanocarriers from MPEG-b-PPBDEMA had too sensitive glucose-responsiveness, and could completely release insulin even below 1.0 mg mL⁻¹ glucose, which stayed away from the needed gate response. Moreover, the polymeric micelles formed by the hydrophobic driving force of only PPBDEMA block were not stable enough and resulted in some insulin releasing in the absence of glucose over the span of measurement.

In this study, a strategy of dispersing glucose-responsive groups was exerted by strongly hydrophobic monomer randomly copolymerizing with glucose-responsive monomer PBDEMA to strengthen the ability to distinguish glucose concentration and enhance the stability of the overall nanocarriers. A boronate ester-containing new glucose-responsive amphiphilic polymers MPEG-b-P(PBDEMA-co-PyBDEMA), comprised of PBDEMA and (2-pyrenylboronic esters-1,3-dioxane-5-ethyl) methylacrylate (PyBDEMA) were synthesized. PPBDEMA endowed the resulting polymers with glucose-responsiveness. While, the introduction of strongly hydrophobic PPyBDEMA into the polymer architecture would strengthen the hydrophobic driving force of the whole polymers, benefiting to the formation of solid and stable micelles in aqueous environment and avoiding the overall collapse of the polymeric micelles in response to glucose. More importantly, the copolymerization of PyBDEMA with PBDEMA in random would disperse glucose-responsive PBDEMA in the hydrophobic core, which was dedicated to enhancing their glucose selectivity and realized controllable insulin release. In addition, the driving force enhancement of self-assembling would also provide one promising to increase the encapsulation efficiency (EE) and loading capacity (LC) of insulin in these nanocarriers.

2. Materials and methods

2.1 Materials

Pyrene (98%), acryloyl chloride (98%) and n-butyllithium (C = 2.5 M) were purchased from J&K chemical regent company. 2-Bromopropionyl bromide (98%) from Aldrich and pentamethyl diethylenetriamine (PMDETA, 99%) from Acros Organics were used without further purification. CuBr purchased from Aldrich was purified by stirring in acetic acid overnight, followed by washing with ethanol and diethyl ether, and dried in vacuum. Tetrahydrofuran (THF) was dried with sodium using benzophenone as an indicator. Triethylamine (TEA) and methylene dichloride (CH₂Cl₂) were dehydrated with KOH and CaCl₂ overnight and distilled, respectively. Insulin (27 UI mg⁻¹) purchased from Genview was labeled by fluorescein isothiocyanate (FITC) based on the previous report.³¹ Macroinitiator MPEG5000-Br and monomer (2-phenylboronic esters-1,3-dioxane-5-ethyl) methylacrylate (PBDEMA) were prepared according to the reported methods, respectively.²⁵ Other analytical reagents such as 1,1,1-tris(hydroxymethyl) propane, N-bromosuccinimide (NBS, 98%) were used as received. Synthesis of monomer (2-pyrenylboronic esters-1,3-dioxane-5-ethyl) methylacrylate (PyBDEMA) and amphiphilic diblock polymer MPEG-b-PPyBDEMA were described in ESI.†

2.2 Characterization

The chemical structures of the copolymers were characterized by ¹H and ¹³C NMR carrying out on a 400 MHz NMR instrument (Bruker Corporation, Germany) at room temperature using CDCl₃ as solvent. The chemical shifts were measured against the solvent signal of CDCl₃ as internal standard. The molecular weights and polydispersity index of the polymers were determined with Waters 515-2410 gel permeation chromatograph (GPC) instrument equipped with Styragel HT6E-HT5-HT3 chromatographic column following a guard column and a differential refractive-index detector. The measurements were performed using THF as eluent (flow rate of 1.0 mL min⁻¹ at 30 °C) and a series of narrow polystyrene standards for the calibration. The fluorescence spectra were recorded by a Hitachi F-4600 Fluorescence instrument (Hitachi High-Technologies Corporation, Tokyo Japan) at 37 °C. The particle size and zeta potential of FITC-insulin loading and unloading nanoparticles were both measured with Nano-ZS (ZEN3600, Malvern) equipped with Zetasizer software and with 35 mW solid state laser operated at a laser light wavelength of 660 nm. The size measurements were carried out at 25 °C at a scattering angle of 90°. The polymeric micelles were imaged on a Hitachi H800 transmission electron microscopy (TEM) (Hitachi High-Technologies Corporation, Tokyo, Japan) operated with 100 kV.

2.3 Synthesis of poly (ethylene glycol)-block-poly[(2-phenylboronic esters-1,3-dioxane-5-ethyl methylacrylate-co-(2-pyrenylboronic esters-1,3-dioxane-5-ethyl) methylacrylate] MPEG-b-P(PBDEMA-co-PyBDEMA)

The synthesis of MPEG-b-P(PBDEMA₂₅-co-PyBDEMA₈) (P-2) was exampled to describe ATRP procedure in this work. A Schlenk flask charged with CuBr (12.9 mg, 0.09 mmol) and macroinitiator MPEG5000-Br (462 mg, 0.09 mmol) was degassed using three vacuum–nitrogen cycles. The degassed materials including ligand PMDETA (15.5 mg, 0.09 mmol), monomers PBDEMA (3.75 g, 13.67 mmol), PyBDEMA (2.16 g, 5.4 mmol) and anisole (1.5 mL) were introduced into the reaction flask using syringes under nitrogen atmosphere. The reaction system was further degassed using three freeze–pump–thaw cycles, and then immersed in an oil bath at 95 °C under thermostat control. After a predetermined polymerization time, the cooled reaction solution was diluted with CHCl₃ (3 mL) and passed through a neutral alumina column to remove the catalyst. The concentrated reaction solution was dropwise added into the mixed solvent of hexane/diethyl ether (v/v = 4/1). ¹H NMR (400 MHz, CDCl3) δ (ppm): 8.98 (br, 1H, o-pyrenyl), 8.43 (br, 1H, from pyrenyl), 8.08 (br, 7H, other hydrogens from pyrenyl ring), 7.75 (br, 2H, o-C₆H₅), 7.75 (br, 2H, o-C₆H₅), 7.32 (m, 1H, p-C₆H₅), 7.29 (m, 2H, m-C₆H₅); for the overlapped proton signals of PBDEMA and PyBDEMA: 3.76–3.99 (m, 6H, CH₂OOC, CH₂OBOCH₂ and CH₂OBOCH₂), 2.29 (br, 1H, CHCH₂), 1.60–1.91 (br, 2H, CHCH₂), 1.35 (m, 2H, CH₂CH₃), 0.79 (m, 3H, CH₂CH₃); for MPEG block: 3.64 (s, 4H, CH₂CH₂O), 3.38 (s, 3H, CH₃O). The preparation procedures of P-1and P-3 were similar to P-2 except the molar ratio of PBDEMA, PyBDEMA and MPEGBr (as listed in Table 1).

Table 1 Molecular characteristics of amphiphilic copolymers^a

Sample no.	Amphiphilic copolymer^b	PBDEMA/PyBD EMA/MPEGBr^c	Mn^d (NMR)	GPC^e
				Mn	Mw/Mn
a The amphiphilic copolymers were fabricated via ATRP with MPEG5000-Br as macroinitiator in anisole at 95 °C for 24 h.b The numbers at the footnote showed the indicative repeating unit number of each segment, the repeating unit numbers of PPyBDEMA and PPBDEMA were determined by ^1H NMR.c Indicating the molar ratio of PBDEMA, PyBDEMA and MPEGBr.d The number-average molecular weights of the copolymers were calculated by ^1H NMR results.e Determined by GPC.
P-1	MPEG-b-P(PBDEMA11-co-PyBDEMA8)	100 : 60 : 1	11 200	8800	1.41
P-2	MPEG-b-P(PBDEMA25-co-PyBDEMA8)	150 : 60 : 1	15 000	10 500	1.37
P-3	MPEG-b-P(PBDEMA33-co-PyBDEMA8)	180 : 60 : 1	17 200	11 600	1.39

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2.4 Preparation of polymeric micelles from MPEG-b-P(PBDEMA-co-PyBDEMA)

Briefly, 5.0 mg MPEG-b-P(PBDEMA-co-PyBDEMA) dissolved in 0.5 mL THF was added into 10 mL PBS (0.02 M, pH = 7.4) under vigorous stirring. The mixed solution was dialyzed against PBS using dialysis membranes (MWCO 3500 Da), and PBS was refreshed every 6 h for 2 days. The morphology of the nanoparticles was measured by TEM and DLS, respectively.

2.5 Evaluation of insulin-loading capacity of the polymeric micelles

As example, a stock of FITC-insulin (2.0 mg) dissolved in 10.0 mL HCl (0.01 M) was adjusted to pH 6.0 using NaOH (0.1 M), and dropwise added into the polymer (5.0 mg) in 0.5 mL THF with strong stirring in ice bath overnight. The purification of the polymeric micelles encapsulating FITC-insulin was performed by dialysis method (MWCO, 14 000 Da), and monitored by fluorescence technique. The similar operation happened to the other applied insulin concentrations listed in Table 2.

Table 2 Morphological characteristics of the nanoparticles without and with FITC-insulin

Polymersample	CMC (mg mL^−1)	Without FITC-insulin			With FITC-insulin^c
		Diameter^a (nm)	PDI	Zeta potential^b (mV)	Diameter^a (nm)	PDI	Zeta potential^b (mV)
a The particle average dimension determined by DLS at 25 °C.b Measured by zeta potential analyzer at 25 °C and pH 7.4.c The applied insulin concentration was 0.15 mg mL^−1.
P-1	0.017	35	0.124	−4.74	15	0.409	−18.4
P-2	0.011	34	0.126	−7.26	14	0.391	−20.3
P-3	0.009	32	0.153	−8.97	14	0.366	−21.0

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The entrapment efficiency of insulin was determined using a fluorescence instrument after the isolation of free insulin from the polymeric micelle solution. The encapsulated insulin mass was calibrated according to the measured standard curve of fluorescence intensity against insulin concentration. The insulin entrapment efficiency (EE) and loading capacity (LC) were calculated using the following equations.

2.6 Responsive release of FITC-Insulin

The in vitro release test of FITC-insulin from the polymeric micelles was evaluated by dialysis method. The insulin-loaded solution diluted to 1.0 mL was sealed by a dialysis membrane (MWCO 14 000 Da) and immersed into 25 mL PBS (0.02 M, pH = 7.4). 1.0 mL of the external medium was sampled every determined time. The emission intensity was measured by fluorescence spectroscopy at an emission wavelength of 525 nm upon excitation at 494 nm. The cumulative release percentage at determined time was calculated by the ratio of the fluorescence intensity at that time to the sum of the released intensity.

2.7 Circular dichroism spectroscopy

The stability of the released insulin was determined by analysis of the conformation of released insulin using circular dichroism (CD) and the resulting spectrum was compared to standard insulin. The standard insulin was dissolved in PBS (0.02 M, pH = 7.4) at the final concentration of 40 μg mL⁻¹. CD measurements were carried out on a Jasco J-810 CD spectropolarimeter at 25 °C with a cell length of 1.0 cm. For the far-UV CD spectra, samples were scanned from 190 to 250 nm and accumulated 5 times, at a resolution of 0.2 nm and scanning speed of 700 nm min⁻¹.

2.8 In vitro cytotoxicity of polymeric micelles

NIH3T3 (mouse embryo fibroblasts) were selected and seeded onto 96-well plates at a density of 10 000 cells per well with 100 μL of growth medium. The pre-prepared micelle solutions were diluted to give a range of final concentration from 12.5 to 200 mg L⁻¹ using culture medium. The plates were maintained in the incubator for 24, 48 and 72 h, respectively. Cell viability (%) = (Ni/Nc) × 100, where Ni and Nc are the absorbance of surviving cells treated with or without polymeric micelles, respectively.

3. Results and Discussion

3.1 Synthesis and self-assembly of MPEG-b-P(PBDEMA-co-PyBDEMA)

PyBDEMA had strongly hydrophobic structure and π–π stacking interaction between the adjacent units on the polymer chain, which would significantly enhance the stability of the whole nanoparticles in aqueous environment. Moreover, similarity of chemical structure between PBDEMA and PyBDEMA would benefit to the copolymerization of these two monomers and generating the self-assembly behavior. Based on these thoughts, PyBDEMA was selected as comonomer in this study. MPEG-b-P(PBDEMA-co-PyBDEMA) was fabricated by one-step ATRP with MPEG5000-Br as macroinitiator (Scheme 1). By changing the ratio of PBDEMA to PyBDEMA, three distinct copolymers having different PPBDEMA content were obtained, as listed in Table 1. The chemical structures of the block copolymers were confirmed by ¹H NMR, and the representative spectrum of MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈) (P-3) is displayed in Fig. 1. The chemical component of the hydrophobic blocks was estimated by comparing to the peak integral ratio between phenyl ring of PBDEMA moiety at 7.76 ppm and pyrenyl ring signal at 8.94 ppm, and the corresponding molecular weight of the copolymer calculated based on ¹H NMR spectrum was shown in Table 1. Furthermore, GPC measurements exhibited unimodal and relatively narrow distribution without MPEG5000-Br and free monomer (Fig. S5†), revealing the successful synthesis of structurally well-defined block copolymers MPEG-b-P(PBDEMA-co-PyBDEMA).

Scheme 1 Synthesis route of the block copolymers MPEG-b-P(PBDEMA-co-PyBDEMA).

Fig. 1 ¹H NMR spectrum of MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈).

The amphiphilic block copolymers MPEG-b-P(PBDEMA-co-PyBDEMA) consists of MPEG segment which is soluble in water and P(PBDEMA-co-PyBDEMA) block which is insoluble in neutral solution. Therefore, MPEG-b-P(PBDEMA-co-PyBDEMA) would self-assemble to form polymeric micelles with hydrophobic core composed of both PPBDEMA and PyBDEMA and a hydrophilic shell composed of MPEG in neutral solution. Critical micellar concentration (CMC) was firstly investigated to measure the driving force of the copolymers in aqueous environment. Although the reported fluorescence technique³² was utilized to determine CMC of these copolymers in this study, differently, CMC was obtained by directly measuring the fluorescence ratio of these copolymers themselves at 338 and 333 nm, due to PyBDEMA itself having fluorescence property instead of additional fluorescence probe. As listed in Table 2, CMCs of the copolymers MPEG-b-P(PBDEMA-co-PyBDEMA) was below 0.020 mg mL⁻¹, which is lower 3 orders of magnitude than our previously reported MPEG-b-PPBDEMA series (Their CMCs were in the range of 50–120 mg mL⁻¹). This indicates that the introduction of hydrophobic PPyBDEMA moiety provided strong driving force for the self-assembly of the whole copolymers in aqueous solution, benefiting to the formation of stable polymeric micelles at relatively low concentration of copolymers. Moreover, an increase in PPBDEMA content of the overall polymers resulted in slight decrease of their CMC value, further suggesting the dominant contribution of PPyBDEMA moiety to the self-assembly.

The average hydrodynamic diameters of the polymeric micelles were measured by DLS. DLS diagram of MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈) was exampled in Fig. 2A, and the other results were summarized in Table 2 and Fig. S7.† It is obvious that the average micelle sizes of these amphiphilic polymers were maintained below 50 nm. Interestingly, these micellar diameters were only slightly different with an increase in the PPBDEMA content of the whole hydrophobic moiety, which could be attributed to that the dominant PPyBDEMA weakened the effect of PPBDEMA content changing on the whole size of the nanoparticles. The pristine insulin-unloading nanoparticles further characterized by zeta potential at pH 7.4 showed some negative charge, which was probably resulted from the anionic boronate ester structure in aqueous solution.³³ The particle morphology of the amphiphilic copolymers in the dry state was detected by TEM, and one representative TEM image of polymeric micelles from polymer P-3 was exhibited in Fig. 2B, in which the nanoparticles exhibited the closely spherical polymeric micelles in the diameter range of 30 and 50 nm.

Fig. 2 (A) DLS diagram of MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈) (P-3) micelles and (B) TEM image of P-3.

3.2 Insulin-loading of the polymeric micelles from MPEG-b-P(PBDEMA-co-PyBDEMA)

Fluorescence-labeled insulin such as FITC-insulin was selected to elucidate the glucose-responsiveness of the amphiphilic copolymers MPEG-b-P(PBDEMA-co-PyBDEMA) in this study. By adjusting pH value of the self-assembly aqueous solution to 6.0 close to its isoelectronic point (pI, 5.35–5.45), FITC-insulin was encapsulated in the hydrophobic core of the polymeric micelles MPEG-b-P(PBDEMA-co-PyBDEMA) due to the low solubility of insulin in the proximity of pI and the hydrophobic interactions between insulin and hydrophobic block of the polymers. According to the previous purification method,²⁶ the free insulin in the pristine self-assemblied solution was thoroughly removed by multiple ultrafiltration centrifuge (MWCO 30 kDa). Subsequently, the insulin-loaded polymeric micelles were further characterized by DLS and zeta potential, respectively. As shown in Table 2, comparing to the polymeric micelles without FITC-insulin, the sizes of insulin-loaded nanoparticles became smaller, and their zeta potential dramatically decreased to about −20 mV, as well. Insulin as a protein containing kinds of amino acids could have interactions of hydrogen bonding, van der Waals force and hydrophobic interaction with the hydrophobic core of the nanoparticles. For MPEG-b-P(PBDEMA-co-PyBDEMA) series, the interaction between insulin and hydrophobic core of polymeric micelles probably led to size shrinkage of the polymeric micelles and the comparatively lower zeta potential. In addition, no free insulin was detected at −7.2 mV, further substantiating successful encapsulation of insulin in the polymeric micelles and the complete removal of free insulin from the solution.

Furthermore, the entrapment efficiency (EE) and loading capacity (LC) of FITC-insulin in the polymeric micelles were examined depending on the hydrophobic PyBDEMA/PBDEMA content ratio and the given mass of FITC-insulin, as well. When all of the polymer concentration were determined at 0.5 mg mL⁻¹ and the given insulin concentration was 0.2 mg mL⁻¹, as shown in Fig. 3a, EE of the polymeric micelles formed from P-1 with lowest hydrophobic content could reach above 40%, which is greatly higher than the reported MPEG-b-PPBDEMA system having similar hydrophobic length. Moreover, with an increase in PPBDEMA content of the polymer structure, EE of FITC-insulin in the polymeric micelles was distinctly enhanced to nearly 65%, and their LC was also from 11% up to 20%. To our knowledge, few works was reported to obtain such high EE and LC of insulin at the same time, in particular, LC was mostly reported about 10%.²⁴ The cause to result in the high loading efficiency could be attributed to the hydrophobic PPyBDEMA moiety. Due to its strong hydrophobicity and π–π stacking interaction, PPyBDEMA associating with PPBDEMA in the hydrophobic core would provide more effective interaction with the encapsulated insulin, contributing to insulin entrapment. What's more, the introduction of PyBDEMA into the polymer architectures significantly strengthened the hydrophobic driving force of the whole polymers, which benefits to the self-assembly at the low polymer concentration such as 0.5 mg mL⁻¹, less than 1.0 mg mL⁻¹ used previously. The given polymer mass impacted the overall LC of insulin, and lowering polymeric micelle concentration would be helpful to increasing LC. Additionally, the effect of given insulin concentration on EE and LC was detected with fluorescence technique. When MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈) as an example was controlled at 0.5 mg mL⁻¹ (Fig. 3B), EE of insulin was dramatically enhanced from 54% to 73% with insulin concentration increasing, and LC was also increased to 19%. Notably, such high entrapment and loading efficiency are less reported, and the polymeric micelles from MPEG-b-P(PBDEMA-co-PyBDEMA) as nanocarriers are conducive to effective delivery of insulin.

Fig. 3 (A) EE and LC of insulin in the polymeric micelles dependence of PyBDEMA and PBDEMA content ratio (the given insulin concentration was 0.2 mg mL⁻¹); (B) the effect of the insulin mass on EE and LC with MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈) as a representative, the polymeric micelles concentration was maintained at 0.5 mg mL⁻¹.

3.3 Glucose-responsive release of FITC-insulin

In the human body, the blood glucose level is tightly regulated in the range of 0.6–1.2 mg mL⁻¹. When the glucose level exceeds 2.0 mg mL⁻¹, it is generally considered as hyperglycemia, which has to be treated. During exploring intelligent glucose-responsive delivery systems, scientific attention is improved to focus on not only glucose-sensitivity under hyperglycemia conditions but also the real-time response to glucose under normoglycemia (1 mg mL⁻¹ glucose) to avoid a hypoglycemia effect induced by excess insulin released from the drug carriers. In this study, PPyBDEMA moiety showed inert response to glucose even at pH 8.6 (Fig. S11†), due to π electron cloud delocalization of pyrene ring weakening Lewis-acidity of boron atom on PyBDEMA structure. Based on the consideration of reasonably adjusting glucose-sensitivity, the copolymerization of PBDEMA and strongly hydrophobic PyBDEMA was used to prepare glucose-responsive polymeric micelles. In order to evaluate the glucose-responsiveness of the polymeric micelles at pH 7.4, FITC-insulin loading micelles were prepared at pH 6.0, and the cumulative release of insulin from the polymeric micelles was calculated by monitoring the changes in the fluorescence intensity of the external fluid at pH 7.4. As shown in Fig. 4, the resultant polymeric micelles with different compositions were generally stable in the absence of glucose, and less 15% insulin was released from MPEG-b-P(PBDEMA-co-PyBDEMA)s for 65 h. When FITC-insulin loaded micelles were incubated at the glucose concentration of 1.0 mg mL⁻¹ (normoglycemic condition), only about 20% insulin release occurred in similarly measured time. As the glucose concentration was increased to 3.0 mg mL⁻¹ (hyperglycemic condition), the remarkable and sustained insulin release was observed at a relatively stable release rate, and the cumulative release from these polymeric micelles was above 90%. These polymeric micelles formed from MPEG-b-P(PBDEMA-co-PyBDEMA)s exhibited comparatively stable state under normoglycemia while apparent glucose-responsiveness under hyperglycemic conditions. It is noteworthy that such great disparity of glucose-responsiveness between normoglycemia and hyperglycemia is a significant progress compared with our recent work (as shown in Fig. S12†), which indicates these nanocarriers formed from MPEG-b-P(PBDEMA-co-PyBDEMA)s could be useful for delivery of insulin in response to the fluctuation of blood glucose concentration.

Fig. 4 The insulin release dependence of glucose concentration at pH 7.4 and 37 °C for MPEG-b-P(PBDEMA11-co-PyBDEMA8) (P-1) (A); MPEG-b-P(PBDEMA25-co-PyBDEMA8) (P-2) (B); MPEG-b-P(PBDEMA33-co-PyBDEMA8) (P-3) (C); release profiles of insulin from P-1, P-2 and P-3 without glucose and under 3.0 mg mL⁻¹ glucose (D).

Furthermore, it is interestingly found that variation of composition generated different glucose-responsiveness of the polymeric micelles as indicated in Fig. 4D. At the glucose concentration of 3.0 mg mL⁻¹, the polymeric micelles from MPEG-b-P(PBDEMA₁₁-co-PyBDEMA₈) (P-1) showed long release retard period. Insulin release from P-1 was not observed until 25 h incubation, and continuous release was detected in the time range of 30–90 h and then leveled off. As PBDEMA content was increased in the hydrophobic blocks, the glucose-responsiveness for MPEG-b-P(PBDEMA₂₅-co-PyBDEMA₈) (P-2) and MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈) (P-3) was distinctly accelerated. The polymeric micelles from P-3 with the highest PBDEMA content displayed relatively rapid release behavior, and 90% release was completed in 50 h. Additionally, no burst release was detected at the initial time when the polymeric micelles from these three polymers were treated by various concentrations of glucose, further supporting that no free insulin was absorbed onto the surface of the complex micelles after thorough purification.

As discussed above, these polymeric micelles with enhanced glucose-sensitivity were obtained due to the introduction of PyBDEMA. For MPEG-b-P(PBDEMA-co-PyBDEMA) series, the hydrophobic core of polymeric micelles consisted of both PPBDEMA and PPyBDEMA segments. Amongst, PPyBDEMA in the core worked for strengthening the hydrophobic interaction and dispersing PPBDEMA moiety. On one hand, the presence of strongly hydrophobic PPyBDEMA moiety significantly enhanced the stability of the overall polymeric micelles without glucose for the measured time. On the other hand, PPyBDEMA and PPBDEMA randomly distributed in the hydrophobic block, and hydrophobic PPyBDEMA with inert glucose-responsiveness at pH 7.4 interrupted the continuous aggregation of glucose-responsive PPBDEMA in the core, which is favorable to maintain the polymeric micelles with good stability and avoid the concentrated response of PPBDEMA in the presence of glucose. Such dispersed PBDEMA form is helpful to distinguish glucose concentration. At low glucose concentration such as 1.0 mg mL⁻¹, the response of some dispersed PBDEMA units is not capable of generating enough force to change the strongly hydrophobic aggregation to help insulin escaping from the core. In contrast, when the glucose concentration was increased to 3.0 mg mL⁻¹, more dispersed PBDEMA realized the polarity transfer from hydrophobic to hydrophilic, effectively resulting in the disintegration of the whole polymeric micelles. Moreover, the more content of PBDEMA in the polymer, the higher aggregation density in the core, thus exhibiting the faster glucose-responsiveness at 3.0 mg mL⁻¹ glucose.

To further explore their accurate glucose-sensitivity of these polymeric micelles under normoglycemia and hyperglycemia, the pulsed release of insulin in response to glucose was studied with MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈) (P-3) as an example at 1.0 mg mL⁻¹ (normoglycemia) and 3.0 mg mL⁻¹ (hyperglycemia) glucose as a trigger (Fig. 5). When FITC-loaded polymeric micelles were incubated in pH 7.4 and at 1.0 mg mL⁻¹ glucose for the first 10 h, the insulin release was very slow. Subsequently, the polymeric micelles were moved and incubated in the environment of 3.0 mg mL⁻¹ glucose for the second 10 h, and the rapid release of insulin from the polymeric micelles was detected. After the polymeric micelles were placed back to 1.0 mg mL⁻¹ glucose solution, inert insulin release appeared again. Five apparent cycles of on-off insulin release were reproducibly obtained via such alternate incubation in 1.0 and 3.0 mg mL⁻¹ glucose as indicated in Fig. 5, and the cumulative release of insulin was close to 100% after 5-time pulsed stimuli of 3.0 mg mL⁻¹ glucose, consistent with the effect of the polymeric micelles from P-3 continuously incubating in 3.0 mg mL⁻¹ glucose as above discussed. From these results, these polymeric micelles from MPEG-b-P(PBDEMA-co-PyBDEMA) as drug carriers not only displayed excellent glucose-responsiveness under hyperglycemia conditions, importantly, they exhibited intelligent insulin release at basal release rates under normoglycemic conditions at physiological pH. These self-regulated insulin delivery system in vitro studies may have potential for the treatment of diabetes.

Fig. 5 Glucose-triggered on-off release of insulin from P-3 at the glucose concentration of 1.0 and 3.0 mg mL⁻¹ and 37 °C.

In addition, to elucidate possible glucose-responsiveness mechanism of MPEG-b-P(PBDEMA-co-PyBDEMA), the size change of these polymeric micelles formed from MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈) was monitored by DLS measurement in the process of glucose response. As shown in Fig. 6A, the size of the nanoparticles in the absence of glucose had little change over 65 h, indicating that the polymeric micelles were stable in aqueous solution. After treatment with 1.0 mg mL⁻¹ glucose, the size of these micelles was only slightly increased from 31 to 37 nm, suggesting that these nanoparticles still remained relatively stable during the determined time span, in accordance with very low insulin release as above mentioned. Differently, the nanoparticle bulk expansion occurred at 3.0 mg mL⁻¹ glucose. After treatment with 3.0 mg mL⁻¹ glucose for identical time span, the size was gradually swollen to about 68 nm. The morphology of the nanoparticles after 65 h glucose exposure at 3.0 mg mL⁻¹ was observed by TEM, and displayed larger polymeric micelles with core–shell structure (as shown in Fig. S13†). The chemical structure of these polymeric micelles after glucose response was further characterized by ¹H NMR technique. Comparing to the chemical structure of the block polymer without glucose, it was obviously vanished that the signals between 7.25 and 7.67 ppm were associated with phenyl ring signals (i, j, k) from the PBDEMA moiety (Fig. 6B), which demonstrates the exhaustive removal of phenylborate ester moiety from the polymer architecture. Notably, the characteristic signals attributed to pyrene structure still appeared at 8.83 and 8.31 ppm, indicating that PPyBDEMA segment was remained in the resulting structure. Combined with the above all experimental results, one possible reason resulting in insulin release from the hydrophobic core of the polymeric micelles is supposed that the phenylborate ester endowed the polymeric micelles with glucose responsiveness and detached from the polymer structure in the stimuli of 3.0 mg mL⁻¹ glucose, which led to the polarity change of the original hydrophobic core and the nanoparticles' swelling. While, PPyBDEMA with inert glucose-responsive property at pH 7.4 and strong hydrophobicity devoted to maintaining the morphological completeness of the polymeric micelles despite of the bulk swelling during the process of glucose response and controlling the glucose-sensitivity and responsive rate of the whole nanoparticles.

Fig. 6 (A) Diameter of the polymeric micelles from MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈) (P-3) at different glucose concentration at pH 7.4 and 37 °C; (B) ¹H NMR spectra of block polymer MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈) before and after incubation at 3.0 mg mL⁻¹ glucose and at pH 7.4 and 37 °C.

3.4 Stability of the released insulin and cell toxicity assays

As reported,³⁴ CD spectroscopy is an efficient technique to evaluate the conformational changes in insulin. Based on the principle that the ratio of the band at 208 nm arising from α-helix structure to that at 223 nm from β-structure ([Φ]₂₀₈/[Φ]₂₂₃) can qualitatively measure the overall conformational structure of insulin, the standard insulin and the released insulin suspended in the aqueous solution were analyzed using far UV-CD spectropolarimeter, respectively. As indicated in Fig. 7A, no significant conformational change was detected for the insulin released from the polymeric micelles at pH 7.4 in comparison with the standard insulin, and [Φ]₂₀₈/[Φ]₂₂₃ for standard insulin and released insulin was 1.11 and 1.08, respectively. Furthermore, the spectral characteristics indicate that tertiary structure of released insulin was well preserved.

Fig. 7 (A) UV-CD spectra of standard and released insulin; (B) cell viability assay in NIH3T3 mouse fibroblast cell lines. The cells were treated with polymeric micelles formed from MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈) at various concentrations at 37 °C for 24, 48 and 72 h, respectively.

To evaluate the potential toxicity of these copolymers, in vitro cytotoxicity assays of the polymeric micelles from the copolymers were performed using familiar NIH3T3 mouse fibroblast cell line and analyzed by MTT method. As shown in Fig. 7B, dose–response study was conducted by exposing NIH3T3 cells to various concentrations of polymeric micelles from MPEG-b-P(PBDEMA₃₃-co-PyBDEMA₈). The good cytocompatibility (>80 ± 8.9%) was detected as increasing the concentration of polymeric micelles from 12.5 mg L⁻¹ to 200 mg L⁻¹ within 24 h. Moreover, the NIH3T3 cell viability was still maintained at least 50% after 48 and 72 h cultivation in the concentration range of polymeric micelles below 50 mg L⁻¹, suggesting that the introduction of PPyBDEMA into the polymeric micelles didn't bring great cytotoxicity to this system.

4. Conclusions

In summary, glucose-responsive polymeric micelles were prepared via the self-assembly of the amphiphilic polymers MPEG-b-P(PBDEMA-co-PyBDEMA) having different PBDEMA and PyBDEMA contents. The random distribution of PyBDEMA and PBDEMA in the hydrophobic core provided three aspects of advantages to this delivery system of insulin: (1) because the strongly hydrophobic PyBDEMA with inert glucose-responsiveness at pH 7.4 effectively dispersed glucose-responsive PPBDEMA segment, the resultant polymeric micelles displayed notable glucose-responsiveness under physiological pH 7.4 with 3.0 mg mL⁻¹ glucose (hyperglycemia), while was comparatively stable with 1.0 mg mL⁻¹ (normoglycemia); the on-off release of insulin was successfully realized under pulsed trigger of 1.0 mg mL⁻¹ and 3.0 mg mL⁻¹ glucose. (2) The introduction of PyBDEMA in the hydrophobic core significantly enhanced the loading capacity of insulin, and their EE and LC could reach 73% and 20%, respectively; (3) due to the presence of strongly hydrophobic PyBDEMA, the morphology of polymeric micelles was maintained in response to glucose but some swelling. This type of polymeric micelles constructed by such strategy of dispersing glucose-responsive groups has both lower apparent pK_a and better sensitivity to glucose. This strategy may be a promising method to develop self-regulated insulin delivery in the treatment of diabetes. Further study will focus on in vivo insulin delivery and will be reported in the future.

Acknowledgements

This work was supported by National Natural Science Foundation of China (NSFC, Grant no. 21374005 and 81273631) and Natural Science Foundation of Beijing (Grant no. 7122090), Scientific Research Foundation for the Returned Overseas Chinese Scholars, State Education Ministry and Chinese Universities Scientific Fund.

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Footnote

† Electronic supplementary information (ESI) available. See DOI: 10.1039/c4ra08593k

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