Glucose- and H2O2-Responsive Polymeric Vesicles Integrated with

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Glucose- and H2O2-Responsive Polymeric Vesicles Integrated with Microneedle Patches for Glucose-Sensitive Transcutaneous Delivery of Insulin in Diabetic Rats Zaizai Tong, Junyi Zhou, Jiaxing Zhong, Qiuju Tang, Zhentao Lei, Haipeng Luo, Pianpian Ma, and XiangDong Liu ACS Appl. Mater. Interfaces, Just Accepted Manuscript • Publication Date (Web): 22 May 2018 Downloaded from http://pubs.acs.org on May 22, 2018

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ACS Applied Materials & Interfaces

Glucose- and H2O2-Responsive Polymeric Vesicles Integrated with Microneedle Patches for Glucose-Sensitive Transcutaneous Delivery of Insulin in Diabetic Rats Zaizai Tong*,1,2,3,4 Junyi Zhou,1 Jiaxing Zhong,1 Qiuju Tang,1 Zhentao Lei,1 Haipeng Luo,1 Pianpian Ma,2,3,4 Xiangdong Liu1,2,3,4 1. Department of Polymer Materials, Zhejiang Sci-Tech University, Hangzhou 310018, China 2. Key Laboratory of Advanced Textile Materials and Manufacturing Technology (ATMT), Ministry of Education, Hangzhou 310018, China 3. National Engineering Laboratory for Textile Fiber Materials and Processing Technology (Zhejiang), Hangzhou 310018, China 4. Institute of Smart Fiber Materials, Zhejiang Sci-Tech University, Hangzhou 310018, China

Keyword:

Polymeric

vesicles;

Stimuli-responsive

Transcutaneous microneedles; Biomaterials

*Corresponding author. Tel: +86 571 86843527; E-mail address: [email protected] (Z.Z. Tong)

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release;

Insulin;

Diabetes;

ACS Applied Materials & Interfaces 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60

ABSTRACT Herein, a dual-responsive insulin delivery device by integrating glucose- and H2O2responsive polymeric vesicles (PVs) with transcutaneous microneedles (MNs) has been designed. This novel microneedle delivery device achieves a goal of a fast response, excellent biocompatibility and painless administration. The PVs are self-assembled from a triblock copolymer involving with polyethylene glycol (PEG), poly(phenylboronic acid) (PPBA, glucose-sensitive block) and poly(phenylboronic acid pinacol ester) (PPBEM, H2O2-sensitive block). After loading with insulin and glucose oxidase (GOx), the drugloaded PVs display a basal insulin release as well as a promoted insulin release in response to hyperglycemic states. The insulin release rate responds quickly to elevated glucose and can be further promoted by the incorporated GOx, which will generate the H2O2 at high glucose levels and further break the chemical links of PBEM group. Finally, the transdermal delivery of insulin to the diabetic rats ((insulin+GOx)-loaded MNs) presents an effective hypoglycemic effect compared with that of subcutaneous injection or only insulin-loaded MNs, which indicates the as-prepared MNs insulin delivery system could be of great importance for the applications in the therapy of diabetes.


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1. INTRODUCTION Diabetes mellitus is a serious chronic metabolic disorder disease characterized by the elevated blood glucose levels, which has become one of the most challenging global health issues.1-4 Insulin therapy is essential in the treatment of patients with insulindependent diabetes (type 1) and for many patients with non-insulin-dependent diabetes (type 2).5 The subcutaneous injection of insulin is one of the most frequently adopted methods of insulin administration for the traditional treatment of diabetic patients. However, this method always accompanies with many problems such as pain, tenderness, local tissue necrosis, microbial contamination, and nerve damage.6-8 Therefore, the administration method of directly injected insulin is inconvenient, uneconomic and uncontrolled, leading to poor patient compliance.9 To overcome the aforementioned limitations and drawbacks of conventional delivery, a number of insulin delivery systems involving micro- or nano-technologies have been developed in recent decades.10-15 Particularly, stimuli-responsive drug delivery systems based on amphiphilic copolymers have received great attention due to the smart, controllable and long-term insulin release behavior.16-24 Among various stimuli-responsive systems, glucoseresponsive copolymers have gained great attention due to the high blood glucose levels in diabetes mellitus patients, which could be utilized on controlled release of insulin.25-26 Phenylboronic acid (PBA) is a typically functional group to design glucose-responsive copolymers, which is due to the linkages with glucose. The nano-vehicles self-assembled from the glucose-responsive amphiphilic copolymers could disassemble under glucose stimulus and further release the incorporated insulin.27-28 On the other hand, glucose oxidase (GOx) is an enzyme that can convert glucose to gluconic acid and further



generate hydrogen peroxide (H2O2) in the presence of oxygen,29 which could be employed to construct a glucose-mediated and H2O2-responsive system. To our best knowledge, phenylboronic acid pinacol ester (PBEM) is the most sensitive groups which could be used to design a H2O2 stimulus-responsive drug delivery system,30-31 especially for insulin delivery.32-33 The ester linkages of PBEM will be hydrolyzed under H2O2 stimulus,34 which endows the PBEM-containing block copolymers with valuable functional groups for stimuli-responsive drug delivery.35-37 Therefore, a glucose- and H2O2-responsive drug delivery system via the single glucose stimulus could be established by coating glucose oxidase and insulin into the self-assembled nano-objects. On the other hand, microneedles (MNs) are sharp, needle-type structure assembled as an array with the micron size and up to 1000 µm lengths of each needle, which could be used as a promising transdermal drug delivery tool onto the skin.38-41 MNs can be transiently inserted into the stratum corneum of the skin, and subsequently deliver their cargo across the skin barrier with no bleeding and limited pain.42 The penetration depth of MNs is very limited (typically < 500 µm) and thus penetrating the epidermis will not injure subcutaneous tissues.43 Up to now, solid MNs,44 dissolving MNs,36 coated MNs,45 and hollow MNs,46 have been developed for the transdermal drug treatment. Compared with other MNs, dissolving MNs are paid to more attention since they could be dissolved in the skin tissue completely and safely, and release the incorporated cargo. Watersoluble polymers, involving synthesized polymers (polyvinylpyrrolidone (PVP) and polyvinyl alcohol (PVA))47 and natural polymers (gelatin, hyaluronate),48-49 are usually used to prepare dissolving MNs. Herein, consideration of the high glucose levels at physiological environment of the


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diabetes mellitus patients, a dual-responsive insulin delivery system by integrating glucose- and H2O2-responsive polymeric vesicles (PVs) with a painless transcutaneous MN array patch has been designed (Scheme 1). This novel microneedle delivery device achieves a goal of a fast response, excellent biocompatibility and painless administration. Polymeric vesicles usually held great promise for controlled drug release due to their robust hollow structure and large loading capacity of hydrophilic drugs.50-52 In present work, the polymeric vesicles are self-assembled from an amphiphilic triblock copolymer, poly(ethylene

glycol)-b-poly(3-acrylamidophenylboronic

acid)-b-poly(4-(4,4,5,5-

tetramethyl-1,3,2-dioxaborolan-2-yl)benzyl acrylate) (PEG-b-PPBA-b-PPBEM). The vesicles present glucose- and H2O2-responsiveness for the selected sensitive functional groups PBA (glucose sensitive) and PBEM (H2O2 sensitive). After loading insulin and GOx, the PVs have glucose- and H2O2- dual responsive ability via the single glucose stimulus, which are then further integrating into the PVP/PVA MNs for transdermal drug delivery of insulin (Scheme 1). The transdermal delivery of insulin to the diabetic rats ((ins+GOx)-loaded MNs) presents an effective hypoglycemic effect compared with that of subcutaneous injection or only insulin-loaded MNs. This work suggests that the MNs integrated with insulin-loaded polymeric vesicles may have a promising application in diabetes treatment.



Scheme 1. Cartoon representations of fabricated MN patches, and transdermal drug delivery in diabetic rat.

2. EXPERIMENTAL SECTION 2.1 Synthesis of the Samples The detail synthetic routine of 3-acrylamidophenylboronic acid (PBA), 4-(4,4,5,5Tetramethyl-1,3,2-dioxaborolan-2-yl)benzyl acrylate (PBEM), PEG-DDMAT macroRAFT agent and the block polymer PEG-b-PPBA, PEG-b-PPBA-b-PPBEM were described in Supporting Information in detail. 2.2 Preparation of the Polymeric Vesicles and Drug Loaded Vesicles The polymeric vesicles were prepared from a conventional self-assembly method by using an amphiphilic block copolymer referred to our previous reports.53-55 In detail, the as prepared product, PEG45-b-PPBA38-b-PBEM17, was first dissolved in DMF to obtain a homogeneous solution with a polymer concentration of 1 mg mL−1. And then 9.0 mL of the solution was added very slowly (2 mL/h) into a 36.0 mL of deionized water without or with 1.8 mg porcine insulin and 0.18 mg glucose oxidase. The mixture solution was 6 ACS Paragon Plus Environment

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further stirred for another 3 h. Subsequently the mixture was dialyzed against deionized water for 72 h to remove DMF by using MWCO 3500 dialysis bag. Then the vesicles or drug loaded-vesicles ((ins+GOx)-PVs) were obtained by centrifugation (12000 rpm, 30 min, 37 oC) and washed 3 times with deionized water. The supernatant liquid (unloaded insulin and unloaded GOx) was collected to measure the un-loaded insulin and GOx. For comparison, insulin loaded-vesicles without GOx (ins-PVs) were also prepared using the same method. Finally, the drug-loaded vesicular solution was freeze-dried for 2 days to afford the drug-loaded vesicles. The drug loading properties were measured using a UV-Vis spectrophotometer. Hitachi U-3010 was used to measure the absorbance value of unloaded insulin and unloaded GOx at a reference wavelength of 276 nm and 450 nm, respectively. Drug loading content (DLC) and encapsulation efficiency (EE) were calculated according to the following formulas: DLC ( wt %) =

EE ( wt %) =

Weight of loaded drug × 100% Weight of drug loaded micelles

Weight of loaded drug × 100% Weight of drug initially added

(1)

(2)

2.3 The Controlled Release of the Drug-Loaded Vesicles in Vitro In vitro insulin release, 15.0 mL of each insulin-loaded with ((ins+GOx)-PVs) or without GOx solution (ins-PVs) was placed into a MWCO 7000 dialysis bag and immersed into 50 mL of PBS (pH=7.4) at 37 °C under gentle shaking. The insulin-loaded vesicular solution was exposed to 200 or 400 mg/dL of glucoses or different concentrations of H2O2 (0.1 mM and 0.5 mM) to reveal the drug release behavior of polymeric assemblies upon glucose-stimuli. At certain time intervals, a 4.0 mL dialysate 7 ACS Paragon Plus Environment


was removed and added by 4.0 mL of fresh PBS solution. The insulin release properties were measured using a UV-Vis spectrophotometer.

2.4 The Skin insertion ability of MNs FITC-insulin loaded PVs MN patches were applied on separated SD rat skins to evaluate the skin insertion ability of the MNs in vitro. The FITC-insulin was prepared according to the previous research.28 The rat skins were shaved and cleaned with 75% alcohol and air-dried for 10 min after wetting in 37 oC water. A confocal laser scanning microscope (CLSM, C2, Nikon Corporation, Japan) was used to observe the insertion and the dissolution of the MNs into the skin after the FITC-ins-PVs MN patch pierced into the rat skin. To further study the insertion and the drug release of the MNs, histological specimens for the insertion of MNs were excised from rat skins using a scalpel. The isolated skin was embedded in an opti-mum cutting temperature compound (OCT) for histological sectioning using a cryotome (CryoStar NX50, Thermo Fisher Scientific, USA). The frozen OCT-skin specimens were sliced into 7 mm-thick sections to be observed using a CLSM. The insertion depth of MNs, the dissolution of MNs and the release of drug could be observed directly from the bright-field and fluorescence images.

2.5 In vivo Insulin Delivery on SD Rats All animal procedures were adopted and guided by Animal Ethics Committee of Zhejiang Sci-Tech University and Experimental Animal Center of Zhejiang Academy of Medical Sciences (China). Streptozotocin (STZ) was used to establish a type 2 diabetic animal model according to previous literature.56 The blood glucose meter (Sinocare Inc., Changsha, China) was employed to measure the blood glucose concentrations by taking blood samples from the tail vein of the rats. The experiment began when the blood


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glucose levels of rats reached hyperglycemia levels and remain after streptozotocin injection for 3 days. The diabetic SD rats were divided into 4 groups (n=3 for each group) to analyze the drug administrations: (a) control group, without other treatment on diabetic rats; (b) injection group, with a single dose of insulin injection on diabetic rats (20 IU/kg) ; (c) ins-PVs MNs group, with ins-PVs MNs transdermal drug delivery (40 IU/kg) on diabetic rats; (d) (ins+GOX)-PVs MNs group, with ) (ins+GOX)-PVs MNs transdermal drug delivery (40 IU/kg) on diabetic rats. The blood glucose levels (BGLs) were monitored immediately by a blood glucose meter (Sinocare Inc., Changsha, China).

3. RESULTS AND DISCUSSION 3.1 Characterization of the PEG45-b-PPBA38-b-PBEM17 The block copolymer, poly(ethylene glycol)-b-poly(3-acrylamidophenylboronic acid)b-poly(4-(4,4,5,5-tetramethyl-1,3,2-dioxaborolan-2-yl)benzyl acrylate) (PEG-b-PPBA-bPPBEM), was synthesized using a two-step polymerization process by reversible addition–fragmentation transfer polymerization (RAFT), as shown in Scheme 2. The diblock copolymer, poly(ethylene glycol)-b-poly(3-acrylamidophenylboronic acid) (PEG-b-PPBA), was synthesized by polymerization of PBA in the presence of the PEG macro-chain transfer agent. The polymerization degree of the PPBA block was determined to be 38 from the 1H-NMR spectrum (Figure S4). Then, the triblock copolymer, PEG-b-PPBA-b-PPBEM, was synthesized by polymerization of PBEM in the presence of the PEG45-b-PPBA38 macro-chain transfer agent. The polymerization degree of the PPBEM block was estimated to be 17 (Figure S5). From the result of 1H-NMR spectra, the triblock copolymer, PEG45-b-PPBA38-b-PPBEM17 was successfully prepared.



Scheme 2. The synthetic route of PEG45-b-PPBA38-b-PPBEM17. 3.2 Characterization of Stimuli-Responsive Polymer Vesicles (PVs) Then the self-assembly behavior of the obtained triblock copolymer, PEG-b-PPBA-bPPBEM, was first investigated. The transmission electron microscopy (TEM) and dynamic light scatting (DLS) were employed to characterize the morphology and size distribution of the self-assembled aggregates. Since the PEG block is hydrophilic, while the other blocks, PPBA and PPBEM, are hydrophobic, the triblock copolymer forms polymeric vesicles (PVs) with a typical hollow structure as shown in Figure 1a. The average diameter of the obtained PVs is characterized as about 200 nm from the TEM image, which is consistent with the result determined from DLS (217 nm, Figure 1d). Based on the molecular weight of PEG45-b-PPBA38-b-PPBEM17, the weight fraction of hydrophilic segment (PEG) is calculated to be about 13.8 wt %. According to the theory proposed by Hammer based on block copolymers,57 this fraction is facilitated to forming a vesicular morphology. Therefore, the formation of vesicles is reasonable for the present triblock copolymers. On the other hand, since PBA group is known to be in an equilibrium between an uncharged form and a charged form with glucose, the hydrophobic PBA group is changed 10 ACS Paragon Plus Environment

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to hydrophilic group after reaction with glucose (Scheme S6A). Therefore, the changes between the uncharged hydrophobic form and the charged hydrophilic form could be used to form the glucose-responsive vesicles. As shown in Figure 1b, The PVs are totally disassembled and transformed to irregular aggregates after incubation in 200 mg mL-1 glucose solution. DLS result shows the average diameter of the PVs after glucose stimulus for 24 h increases to 2669 nm, which confirms the morphological change of the PVs. Because of the alteration of the hydrophilicity/hydrophobicity ratio of the PVs with glucose stimuli, the size of the aggregates is gradually increased with time (Figure S6), indicating the disassociation of the PVs. On the other hand, PBEM unit is very sensitive to H2O2, whose chemical structure can be damaged by H2O2 since the reaction between PBEM group and H2O2 is irreversible. (Scheme S6B). As a result, the PVs are totally disassociated and irregular aggregates are formed when the PVs are exposed to 0.5 mM H2O2 solution (Figure 1c). Meanwhile, DLS result shows the average diameter of the aggregates is about 3090 nm (Figure 1d). Based on the above two factors, a glucose and H2O2 dual-responsive PV is successfully construct from a self-assembly process.

Figure 1. Characterization of stimuli-responsive polymer vesicles (PVs): (a) TEM 11 ACS Paragon Plus Environment


images of PVs construct from a self-assembly process; (b) TEM image of PVs after incubation in 200 mg mL-1 glucose solution for 24 h; (c) TEM image of PVs in the presence of 0.5 mM H2O2 for 24 h; (d) corresponding DLS result of PVs at different stimuli.

3.3 In Vitro Insulin Release and Cytotoxicity of the PVs The in vitro insulin release behavior from the PVs in response to glucose was assessed by incubating the PVs with PBS containing a high concentration of glucose. First, the insulin loading content (DLC) and encapsulation efficiency (EE) are calculated to be 16.8 wt% and 84.4 % based on Equation 1 and 2, respectively. On the same time, the DLC and EE of GOx are also measured to be 14.8 wt% and 87.0%, respectively, which indicated both the insulin and GOx are encalpulated into the PVs. As shown in Figure 2a, the release behaviors of the drug-loaded vesicles are quite different under the influence of glucose stimulus. The insulin-loaded vesicles display a significantly rapid insulin release rate at the hyperglycemic level, while limited insulin release is observed at the control level. The released percentage of insulin from ins-PVs is about 20% within 5 h without external stimuli. For a comparison, when the drug-loaded PVs are incubated in a 200 or 400 mg/dL (a significant high blood sugar levels) glucose solution, it causes a faster release behavior. Here, the threshold value for diabetes patients is 400 mg dL-1. And 200 mg dL-1, which is slight higher than normal blood level, is selected as a comparison. The percentages of the released insulin from ins-PVs can reach to 52% and 70% in the presence of 200 and 400 mg/dL glucose concentration, respectively (Figure 2a). This implies a higher glucose concentration is beneficial to not only a faster release of insulin but also a higher level of accumulative amount, which is contributed to the glucosesensitive group of PBA in the triblock copolymer.


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The behaviors of the insulin release are also investigated in the presence of different concentrations of H2O2. As shown in Figure 2b, about 40% of insulin has been released in a 0.1 mM solution of H2O2 within 5 hours, which is faster than that without H2O2stimulus. When the concentration of H2O2 reaches to 0.5 mM, a significantly rapid release is detected. The percentages of insulin released from ins-PVs can achieve to be about 89% within 5 h, indicating the drug delivery system possesses a H2O2 concentrationdependent performance. This is mainly due to the presence of the H2O2-sensitive group of PBEM in PVs, which results in the disassembly of the vesicles. On the other hand, it is well known that GOx can interact with glucose and then further generate H2O2.29, 32 Thus, a glucose-promoted insulin release system could be established when GOx is encapsulated into PVs. As presented in Figure 2c, the (ins+GOx)-PVs shows a slow drug release behavior without glucose-stimulus, and only 30% of the insulin is released from the PVs. By contrast, when the (ins+GOx)-PVs are exposed to the 400 mg/dL glucose solution, a significantly fast release behavior of insulin is determined, and about 91% of the insulin is released from the vesicles. Compared with a controlled release profile (ins-PVs exposed to 400 mg/dL glucose solution), the drug release rate of (ins+GOx)-PVs presents a faster behavior (Figure 2c), implying the presence of GOx can further promote the disassembly rate of the vesicles and accordingly a faster insulin release rate. This manner can be attributed to the presence of both glucose- and H2O2stimuli in the (ins+GOx)-PVs system, while it contains only glucose-responsiveness in ins-PVs. Therefore, a rate-controlled insulin release can be established when the drugloaded PVs are exposed to different external stimuli. Finally, in order to evaluate the potential toxicity of the drug-loaded PVs for drug



delivery applications, the in vitro cell viability tests of drug-loaded PVs is measured. Human breast adenocarcinoma (MCF-7) cells are selected to evaluate the potential toxicity of the vesicles by the CCK-8 cell survival assay. The relative cell viability of MCF-7 cells cultured with PVs, ins-PVs or (ins+GOx)-PVs indicates that there is negligible influence on cell viability at all concentrations ranging from 50-250 µg mL-1 as shown in Figure 2d. The viability is over 93% after incubated with drug-loaded PVs for 48 h at all tested concentrations, which shows the ins-PVs and (ins+GOx)-PVs have very low cytotoxicity towards MCF-7 cells and can be used as a potential drug delivery vehicle.

Figure 2. In vitro stimuli-responsive insulin release at 37 °C and cell viability tests: in vitro accumulative release of insulin from ins-PVs with (a) glucose-stimulus and (b) H2O2-stimulus; (c) effect of GOx on the in vitro accumulative release of insulin with glucose-stimulus; (d) cell viability tests of PVs, ins-PVs and (ins+GOx)-PVs against MCF-7 cells. 14 ACS Paragon Plus Environment

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3.4 Fabrication and Characterization of PVs-Loaded Microneedle (MN) Patch The drug-loaded PVs are then further deposited in the tips of PVP/PVA-based MN array patch using a two-step molding approach to achieve convenient and painless administration as described in Scheme 1. The prepared pyramid-shaped MNs are arranged in a 10 × 10 array in with 500 µm tip-to-tip spacing in a 100 mm2 patch. The morphology characterization of (ins+GOx)-PVs MNs is shown in Figure 3 and Figure S7, indicating the successful replication of the PDMS mold. The height of each microneedle is about 500 µm and the bottom length of each needle is about 250 µm. In order to evaluate the mechanical behavior of the MNs, the universal static testing system is hired to measure the resistances of MNs to tensile compression. As shown in Figure 3d, the drug-loaded PVs/PVP/PVA MNs exhibit good mechanical performance. A failure force for drug-loaded PVs/PVP/PVA MNs is determined as 0.168 N per needle with a displacement of 0.75 mm. According to the previous research, a microneedle with a tip diameter of 20 µm requires a minimum force of about 0.15 N to penetrate into the skin.5859

Therefore, the mechanical strength of the drug-loaded PVs/PVP/PVA MNs prepared in

this study is high enough to penetrate into the skin for the transdermal drug delivery. The dissolved ability of prepared (ins+GOx)-PVs/PVP/PVA MNs was further characterized by time-dependent digital camera. Figure S8 presents the digital images of dissolved process of MNs in rat skin after incubation for different times. It can be seen that the soluble MNs can be totally dissolved within 10 min after the MNs are inserted into the skin tissue. On the other hand, the inflammatory reaction is a key problem when the MNs penetrate to the skins, which should carefully be taken into consideration. We also conducted the skin



recovery experiment after the MNs are applied on the skins (Figure S9). It can be observed that the skin will recover its original state within about 10 min. And no other side reactions are observed in the skin tissue in later time. Based on this factor, we think the MNs are safe in current stage as a transcutaneous drug delivery system.

Figure 3. Characterizations of the prepared MNs: (a) the SEM image of the MNs; (b) the corresponding partial enlarged SEM image of (a); (c) the digital image of MNs and (d) the force-displacement curve of (ins+GOx)-PVs/PVP/PVA MNs.

3.5 The Skin Insertion Ability and Transdermal Delivery Property of MNs in Vitro The dissolved ability of the MNs in skin tissue is confirmed by the confocal laserscanning microscopy (CLSM) images of the separated skins treated with FITC-labelled (ins+GOx)-PVs/PVP/PVA MNs for different time. The 3D reconstruction image of the 16 ACS Paragon Plus Environment

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prepared FITC-labelled (ins+GOx)-PVs/PVP/PVA MNs shows the height of the each needle is about 550 µm (Figure S10), which is in accordance with the SEM characterization of the MNs (Figure 3). The basically consistent fluorescence intensity of each needle indicates the drug-loaded PVs are uniformly distributed in the needles. After insertion into skins for 5 min, the pyramid micro-cavities with a depth at 268 µm can be observed (Figure 4a). Elongating the insertion time to 30 min, the depth of the pyramidshaped needles is further decreased to 192 µm, indicating more and more fraction of the needles is dissolved by the interstitial fluid. Meanwhile, the bottom of the needles is also swollen to some extent by the tissue fluid, which can be confirmed from the irregular morphology of the needles (Z=0). This result in turn reveals the encapsulated drug can be released from the MNs and further diffused to deeper tissues. Therefore, the insulinloaded PVs in the MNs can be applied on transcutaneous delivery of insulin and would not cause significant damage to the skin.



Figure 4. Confocal micrographs of the FITC-(ins+GOx)-PVs MNs after insertion into the skin for 5 min (a) and 30 min (b), and the corresponding 3D reconstruction images (scale bar: 200 µm).

The FITC-insulin loaded PVs/PVP/PVA MNs are then further applied on separated full-thickness skins of SD rat to evaluate the transdermal delivery property of the MNs in vitro. The successful skin insertion and the release of the incorporated (ins+GOx)-PVs for MNs are confirmed from the observations on the histological sections of the diabetic and healthy SD rats skin (Figure 5). As shown in the bright field image, sharp cavities with a depth of 250 µm can be observed after applying the MNs on the skin for 5 min, which proves the MNs successfully puncture into the skin. On the other hand, strong green


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fluorescence signal of FITC-insulin can be observed around the punctured sites from the fluorescence image and merge image, which implies the successful penetration of the (ins+GOx)-PVs into skin. However, the strong green fluorescence can only be observed around the puncture sites for the health rats after insertion for both 5 and 30 min. No green fluorescence can be found in deeper tissues (Figure 5A), indicating almost no insulin has been diffused into the tissues. By contrast, in the case of diabetic rats, strong green fluorescence can be observed in the deeper tissues after application of (ins+GOx)PVs MNs on diabetic rats for 5 and 30 min (Figure 5B). Especially, after administration for 30 min, a large amount of insulin is observed under the puncture sites with a deeper depth, indicating more and more insulin are released from the nano-objects, followed by diffusing into deeper tissues. It is believed that the high glucose level in the tissue of diabetic rats could trigger the rapid release of insulin from the PVs, which results in the diffusion of insulin to a deeper tissue. This is consistent with the previous observations based on the morphological transition, drug release in vitro and 3D reconstruction images.



Figure 5. The confocal laser-scanning microscopy images of the histological sections of the healthy SD rat skin (A) and the diabetic SD rat skin (B) after application of FITCinsulin loaded PVs/PVP/PVA MNs for 5 min and 30 min (scale bar: 200 µm).

3.6 Transdermal Delivery of Insulin by MNs in Vivo The MN-array patch was then further administered on streptozotocin-induced diabetic


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SD rats to evaluate its in vivo transdermal delivery of insulin. The diabetic SD rats were randomly divided into four groups and treated with following different MN samples: (a) control group, blank PVs MNs transdermal on diabetic rats; (b) injection group, with a single dose of insulin injection on diabetic rats (20 IU/kg); (c) ins-PVs MNs group, with ins-PVs MNs transdermal drug delivery (40 IU/kg) on diabetic rats; (d) (ins+GOx)-PVs MNs group, with (ins+GOx)-PVs MNs transdermal drug delivery (40 IU/kg) on each diabetic rats. It should be noted that despite the effective drug release from the PVs under the external stimuli, the drug-loaded PVs/PVP/PVA MNs cannot release its full amount of loaded insulin, which is due to a viscous solution in the MN proximity.60 Accordingly, it eventually slows down the release rates of insulin from the MNs. In order to increase the release of insulin and produce a hypoglycemic effect, a higher amount of insulin has been reasonably adopted in the MNs. On the other hand, when the insulin is injected directly into the diabetic rates with a high value (40 IU/kg), it definitely causes a fast decline of the blood glucose concentration with an extremely low level, which causes the diabetic rats a quick death. Taken this phenomenon into consideration, we finally use 20 IU/kg instead of 40 IU/kg in injection group, which is also frequently adopted in other reported work.36 Figure 6 presents the plasma glucose levels of treated rats (in each group) over time. One can see that there is negligible change in the blood glucose levels in the blank group. However, the blood glucose level in injection group significantly decreases from 510 to 80 mg/dL within 2 h, then rebounds to the initial level of 500 mg/dL after 7 h. This result shows injection of insulin presents a short and fast hypoglycemic effect. However, it is quite different for the application of the ins-PVs MNs or (ins+GOx)-PVs MNs on the diabetic rats. The blood glucose levels of ins-PVs MNs



group decrease to 140 mg/dL within 4 h and increase to 500 mg/dL after 8 h. On the other hand, the blood glucose levels of (ins+GOx)-PVs MNs group rapidly decrease to 110 mg/dL within 4 h and then a subsequent gradual increase is detected. The blood glucose levels of this group can be retained for about 4 h in the normoglycemic range (

Glucose- and H2O2-Responsive Polymeric Vesicles Integrated with

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