Fabrication and evaluation of dissolving bird-bill microneedle arrays

doi:10.21203/rs.3.rs-4966848/v1

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Fabrication and evaluation of dissolving bird-bill microneedle arrays

https://doi.org/10.21203/rs.3.rs-4966848/v1

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Coated microneedles (MNs) have some disadvantages, such as low mechanical strength, the risk of clogging and infection due to repeated application, and denaturation at high temperatures. We aimed to fabricate a dissolving bird-bill MN (dBB MN) with a vertical groove between two thin plate-shaped needles and evaluated its ability of transdermally deliver a large-molecular-weight insulin drug into systemic circulation. Hydrogels with various concentrations of polyvinylpyrrolidone (PVP) or sodium hyaluronate (HA) were prepared, and dBB MN arrays were fabricated by micromolding under negative pressure for potential mass production. The needle height of the dBB MN was maximum when the hydrogel was 25 w/w% PVP, with a viscosity of 8–9 Pa∙s. Furthermore, the buckling force of dBB MNs made from 25 w/w% PVP was 130.6 ± 51.0 mN, which increased to 195.6 ± 65.3 mN when insulin was added at 1 w/w%. The blood glucose concentration in diabetic rats decreased slowly and significantly after a 3-h application of the insulin-loaded dBB MN array. Therefore, the dBB MN array demonstrated sufficient ability to puncture rat skin and transdermally deliver a large-molecular-weight drug into the systemic circulation. These findings suggest that the dBB MN array holds promise as a minimal invasive drug delivery platform, with potential applications in improving patient adherence and expanding access to essential therapies, particularly in resource-limited settings.

bird-bill microneedle

dissolving microneedle array

transdermal drug delivery

biocompatible polymer

polyvinylpyrrolidone

Microneedle (MN) arrays are a physical penetration enhancement method that allows macromolecular weight and highly hydrophilic compounds to be delivered through the stratum corneum (SC) into the body. The MN array comprises hundreds of micrometer-high MNs and has the advantage of being a painless device that can be delf-administrated by patients. MNs can be classified into five types: hollow, solid, coated, dissolving, and hydrogel-forming MNs. They can be also used for diagnosis and therapeutic drug monitoring [1–3].

Dissolving MNs (dMNs) are fabricated from various biodegradable and biocompatible macro- and high-molecular-weight polymers such as sodium chondroitin sulfate [4], sodium hyaluronate (HA) [5], poly-γ-glutamic acid [6], polyvinyl alcohol (PVA), and polyvinyl pyrrolidone (PVP) alone or a mixture of PVA and PVP [7–9]. dMNs are expected to quickly dissolve in the skin after skin insertion and enable efficient and complete penetration of drugs encapsulated in the MNs across the SC. In addition, they are not associated with blood contamination because they are single-use and incinerable devices. dMNs can contribute to the spread of vaccination in areas with insufficient physicians [10]. However, the release profile of drugs from dMNs and needle hardness depends on the dissolving and degradable properties of the polymers in the skin. The needle hardness of dMNs can be improved by adding additives such as nonionic surfactants [11]; however, it generally tends to be lower than that of other types of MNs. In the case of hollow and solid MN arrays, drug dosage can be increased without changing the shape of the needle. Moreover, it is necessary to change the number of needles, needle density, and the other structural factors to increase drug dosage in the dMN array.

dMNs generally have conical and pyramidical shapes, which ensure the ability of skin insertion and facilitate the manufacturing process of casting/molding to pull out the dMN from the female mold. Researchers have developed various manufacturing processes to fabricate dMNs with pointed tips without a mold. Lithography was performed to prepare maltose dMNs [12]. The prepared MNs had a high aspect ratio and different shapes but were limited to polymers and drugs that do not denature at high temperatures. Kim et al. reported a method called droplet-born air blowing, which controlled the amount of drug by the pressure and time of the droplet dispenser and formed an MN shape by blowing air [13]. Inkjet printing [14], 3D printing [15], and spray-filled molding [16] have also been reported; however, most researchers have used microcasting/micromolding. For manufacturing dMNs by molding, the polymer hydrogel should be of low viscosity enough to be poured into a female mold that replicats the MNs shape, and the tips of the dMNs should be fine and hard enough to puncture the skin.

Bird-bill MNs reportedly overcome the disadvantages of coated MNs [17]. Bird-bill MNs have a vertical groove between two thin plate-shaped needles; the resulting skin insertion ability is improved, and the amount of drug that is not only coated on the needle but also filled in the vertical groove is increased. However, it is difficult to fabricate dissolving bird-bill (dBB) MNs. Therefore, we aimed to fabricate a dBB MN array by microcasting/micromolding with mass-production capability, improved skin puncture properties, and sufficient hardness for insertion into the skin.

Materials

Biocompatible and biodegradable polymers, polyvinylpyrrolidone K90 (PVP, average molecular weight; 360,000, special grade), and sodium hyaluronate FCH-80 (HA, molecular weight; 600,000–1,000,000, cosmetic grade) to fabricate dBB MN arrays were purchased from Fujifilm Wako Pure Chemical Co. (Osaka, Japan) and Kikkoman Biochemifa Co. (Tokyo, Japan), respectively. Bovine pancreatic insulin and sodium fluorescein (FL) as model drugs were purchased from Sigma-Aldrich Co. (Tokyo, Japan). Streptozotocin (STZ) and other reagents were purchased from Fujifilm Wako Pure Chemical Co.

Manufacturing of a dBB MN array

A female mold from polydimethylsiloxane (PDMS) was prepared for the dBB MN array designed to have 52 needles / the base of 0.44 cm² with 1000 µm needle height (700 µm groove and 300 µm needle-pedestal) (original male mold, Fig. 1a). After submerging the original mold into a container filled with PDMS resin and hardener (Sylgard® 184, Dow Chemical Co., MI, USA), the container was put under reduced pressure for 5 min using a vacuum pump Da-20D (Ulvac Kiko Inc., Miyazaki, Japan) to let the air out from PDMS and subsequently kept at 80 ˚C for 1 h.

HA and PVP hydrogels were prepared at concentrations of 1–4 w/w% and 15–30 w/w%, respectively. The viscosities of the hydrogels were measured using a tuning-fork vibration viscometer SV-10 (viscosity measurement range; 0.3–10,000 mPa∙s, A&D Co., Tokyo, Japan). FL and insulin, as model drugs, were added to the hydrogels at concentrations of 0.5 w/w% and 1.0 w/w%, respectively. The insulin-loaded hydrogels were prepared under ice-cold conditions because heat was generated by mixing PVP and distilled water (DW). The hydrogel (0.8–1.0 g) was placed into the PDMS mold and vacuumed at 9.5 torr for 20 min using Da-20D. The air remaining at the tip of the needle after filling the hydrogel was not removed from inside the female mold, even after centrifugation [18]. Therefore, in this study, a vaccum process was used to fill the hydrogel in the mold. The dBB MN array was dried in a desiccator OH-3S (As One Co., Osaka, Japan) at a temperature of 25˚C and humidity of 20% for 24–48 h. The dBB MN arrays were observed under a Leica MC170 HD digital microscope (Leica Microsystems GmbH, Wetzlar, Germany). After measuring the needle height and weight, the array was stored in a desiccator. The insulin-loaded dBB MN arrays were prepared immediately before the experiments to prevent insulin denaturation.

Dissolution experiments of dBB MN

FL-loaded dBB MN (4% HA and 25% PVP) was placed in 50 mL phosphate-buffered saline (PBS, Sigm-Aldrich Co.) preheated at 37 ˚C. PBS solutions of 500 µL were collected at 1 min intervals until the MN array was completely dissolved. FL concentrations were assayed using a fluorescein spectrometer FP-8550 (Jasco, Tokyo, Japan).

Evaluation of skin insertion capability of dBB MN

The buckling forces of three types of dBB MNs, 4% HA, 25% PVP, and 25% PVP + insulin, were measured using a compression testing machine MCT-510 (Shimadzu Co., Kyoto, Japan) with a loading speed of 20.7 mN/sec and a maximum compression force of 1000 mN. The compression test was repeated thrice using a single needle and different needles.

Swine skin (landrace strain) with a 2–3 mm thickness was purchased from DARD Co. (Tokyo, Japan). The skin was placed on a Kimwipe® (Nippon Paper Crecia, Co., Tokyo, Japan) and moistened with DW. FL-loaded dBB MN (25% PVP) array was punctured into the skin for 10 s. After 10, 30, 60, and 180 min, the skin surface was photographed, and the height of the needles was measured using a digital microscope.

In vivo experiments with diabetic rats

All animal experiments were conducted in accordance with the guidelines of the Kyushu Institute of Technology and approved by the Animal Care and Use Committee of our institution.

Diabetes was induced in healthy rats (8 weeks old; Jcl:SD strain, male; CLEA Japan Inc., Tokyo, Japan), and blood glucose concentration (BGC) was measured as previously described [17]. Hair on the dorsal skin of diabetic rats induced by an intraperitoneal injection of STZ (65 mg/kg in ice-cold 20 mM sodium citrate buffer, pH: 4.6) was removed 48 h prior to the experiment. Insulin-loaded dBB MN (25% PVP) was applied to the shaved area using a spring-driven applicator (5.3 m/s, 17–19 N) and maintained for 10 s. Subsequently, the MN was fixed with surgical tape (Multipore™ Dry, 3M, MN, USA) on the skin for 6 h. The BGC in a drop of blood collected from the tail vein was measured using a Medisafe® Mini GR-102 (Terumo Co., Tokyo, Japan). To ascertain the efficacy of insulin-loaded dBB MN, BGCs of rats treated with 200 IU insulin solution (soaked in cotton) and treated with the MN array without insulin were used as controls.

Fabrication of dBB MN arrays

PVP and HA were selected as the dMN materials in this study. PVP is a water-soluble, inert, non-toxic, pH-stabile, biocompatible, and biodegradable polymer that has been widely used as an excipient in various pharmaceutical applications [19]. HA has also been established as a material for dMN fabrication [5, 20]. The viscosity of a hydrogel varies based on its chemical structure, molecular weight, and polymer material concentration [21]. A mixture of polymers was used for the microcasting of dMN to adjust the hydrogel’s characteristics [8, 21]. We also attempted to fabricate a dMN array using a mixture of polymers, HA, PVP, and PVA, and concluded that it was difficult to fabricate a dBB MN array using this mixture because of its geometric complexity [18].

Table 1 summarizes the needle height of the dBB MN and their viscosity at various polymer concentrations. Needle height was maximal at 4% HA (954.2 ± 8.6 µm, Fig. 1b) and 25% PVP (982.7 ± 2.4 µm, Fig. 1c). The shape of original male mold was well reproduced using 25% PVP; however, plate-shaped needles of 4% HA were noticeably thinner than those of the original mold (Fig. 1a). The viscosities of 4% HA and 25% PVP were 9.4 Pa∙s and 8.5 Pa∙s, respectively. For the polymer concentrations with viscosities ≥ 10.0 Pa∙s (6% HA and 30% PVP) and ≤ 5.0 Pa∙s (2% HA and 15% PVP), the needle height was ≤ 950 µm. The needle tips did not have sharp edges when the hydrogel had a viscosity of > 10.0 Pa∙s, The viscosity of the hydrogel influenced the needle height and sharpness of the dBB MN. Therefore, the property of the hydrogel needed to fabricate the dBB MN array was found to be in the range of 8–9 Pa∙s in viscosity.

Table 1

Needle height and viscosity in various concentrations of polymers
	Concentration [%]	Needle height [µm]	Viscosity [Pa∙s]
HA	2	932.2 ± 4.2	1.9
	4	954.2 ± 8.6	9.4
	6	932.8 ± 4.0	> 10.0
PVP	15	953.6 ± 2.8	3.4
	20	956.1 ± 9.6	5.4
	25	982.7 ± 2.4	8.5
	30	941.1 ± 3.9	> 10.0

Evaluation of dBB MN arrays

The influence of water content on the performance of the dMN was measured to evaluate the dBB MN array. 4% HA and 25% PVP were heated at 85˚C for 6 h after storage in a desiccator; the resulting water contents were 6.25 ± 1.46% and 2.16 ± 1.42%, respectively. The MN height after drying decreased to 883.0 ± 8.2 µm for 4% HA and 936.3 ± 16.9 µm for 25% PVP, suggesting that 4% HA dBB MNs would be affected by the drying that occurred during long-term storage. Thereafter, we measured the needle hardness of 4% HA and 25% PVP dBB MN (Fig. 2a). The buckling force for 4% HA and 25% PVP was 19.8 ± 11.5 mN and 130.6 ± 51.0 mN, respectively (Fig. 2b and 2c). The buckling force of a single dMN mixed with two polymers was reported to range from 100 mN to 480 mN, and the hardness of the dMN was sufficient to withstand the force required for skin puncture [11]. The depth of micro-holes is at least 200 µm using an optical coherence tomography device [11]. Thus, 4% HA has no skin insertion capability and may not be suitable as a base polymer for dBB MNs.

Figure 3 illustrates the release profiles of FL from 4% HA and 25% PVP dBB MN arrays in PBS. The dBB MN array with 4% HA completely dissolved within 7 min, whereas the 25% PVP dBB MN array exhibited a sustained release behavior and fully dissolved within 18 min. We assumed that the high polymer concentration in the hydrogel favored the maintenance of the shape and increased the hardness of the dBB MN. Based on the results of fabrications and dissolution experiments, we decided to use 25% PVP as the material of dBB MN in the following experiments.

Skin insertion and transdermal delivery capability of dBB MN arrays

The skin insertion capability of dBB MN was assessed in vitro using swain skin. Figure 4a shows the relationship between application time and needle height. The MN height decreased to 471.6 ± 54.1 µm within 10 min after insertion; subsequently, it decreased at a constant rate and was 200.4 ± 48.2 µm at 180 min. Two thin plate-shaped needles of dBB MN inserted in the SC were dissolved within approximately 10 min (Fig. 4b), and then, not only the needle pedestal but also the base of the MN array was dissolved slowly (Fig. 4c).

The buckling force of 25% PVP with insulin (195.6 ± 65.3 mN) was 1.50 times greater than that without insulin (Fig. 2a), which may result in insulin filling the space among PVP molecules. The time course of BGC in diabetic rats is shown in Fig. 5. No significant difference was observed between the BGCs of rats treated with the insulin solution and those treated with MN without insulin. In contrast, BGC decreased slowly and significantly after 3 h of application of the insulin-loaded dBB MN array (p < 0.05, two-tailed Student’s t-test with dMN without insulin). The number of dMNs administered is proportional to the net weight of the MNs. By measuring the weight of the base removed from the MNs, the net weight of the dBB MNs was measured at 1.01 ± 0.14 mg. Thus, the amount of insulin in the MNs was estimated to be 10.1 µg/MNs. Although this was not sufficient to immediately decrease BGC in diabetic rats, the insulin included in the base was also delivered into the body, ultimately resulting in a reduction in BGC. Thus, the dBB MN array demonstrated sufficient ability to puncture rat skin and transdermally deliver a large-molecular-weight drug into the systemic circulation. However, it had the disadvantage of taking 3 h to reduce BGC. This issue could be addressed by optimizing the composition and concentration of biodegradable polymers.

In this study, dMNs with pointed needle tips and sufficient hardness for skin insertion were fabricated using 25% PVP, with viscosity in the range of 8–9 Pa∙s, and the BGC in diabetic rats was significantly decreased by delivering a large-molecular-weight drug, insulin, into the systemic circulation. The two plate-shaped needles dissolved quickly in the SC after producing microholes, and the drug included in the needles and the base of the dBB MN array penetrated the skin. However, the BGC did not show an immediately reduction after applying the dBB MN. This could be solved by investigating the quickly dissolution of biodegradable polymers in the skin and optimizing the needle geometry for deeper insertion into the skin. The dBB MN array offer a promising advancement in minimal invasive drug delivery systems, potentially improving patient compliance and access to therapies, particularly in underserved areas.

Competing interests

Ms. Amano, Mr. Takaki, Mr. Takei, Mr. Matuo, Mr. Hara, Mr, Tashiro, and Mr. Oniki declare they have no relevant financial interests. Dr. Hikima and Dr. Ito has a patent pending regarding the method of fabricating the dissolving bird-bill MN.

Ethics approval

All animal experiments were conducted in accordance with the guidelines of the Kyushu Institute of Technology and approved by the Animal Care and Use Committee of our institution.

Funding

This work was supported by the Japan Society for the Promotion of Science (KAKENHI, Grant Number 22K12846). Dr. Hikima has received a partial research support from Mishima Kosan Co.

Authors contribution

All authors contributed to the study conceptions and design. The experiments were completed by Ms. Amano, Mr. Takaki, and Mr. Takei. The in vivo studies were carried out by Ms. Amano. The manuscript was designed, written, and revised by Dr. Hikima and Ms. Amano. All authors read and approved the final manuscript.

Acknowledgement

This work was supported by the Japan Society for the Promotion of Science (JSPS), KAKENHI, Grant Number 22K12846.

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Editorial decision: Major Revisions Needed
12 Oct, 2024
Reviewers agreed at journal
03 Sep, 2024
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You are reading this latest preprint version

Fabrication and evaluation of dissolving bird-bill microneedle arrays

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Version 1

Abstract

Figures

Introduction

Materials and Methods

Materials

Manufacturing of a dBB MN array

Dissolution experiments of dBB MN

Evaluation of skin insertion capability of dBB MN

Results and discussion

Fabrication of dBB MN arrays

Evaluation of dBB MN arrays

Skin insertion and transdermal delivery capability of dBB MN arrays

Conclusions

Declarations

Competing interests

Ethics approval

Funding

Authors contribution

Acknowledgement

References

Status:

Version 1