Bio-inspired MN array

To create a stimulus responsive MN platform with low penetration force and strong adhesion, we pursued a double-layered MN design with selective localization of swellable material in the tip region. Following preferential distal swelling, each MN had both soft (outer) and stiff (inner) regions. In addition to having a MN platform comprising both swellable and non-swellable components, we sought to control the mechanical and water permeation properties of the swellable material, while promoting significant interaction between the soft outer layer and the stiff inner layer to prevent delamination. For this purpose, we considered an amphiphilic block copolymer (BCP) design that exhibits selective responsiveness to stimuli such as the presence of aqueous or organic solvents20,21. We envisioned that dual hydrophilic–hydrophobic properties of the outer layer would enable rapid absorption of water and promote intimate interaction with the non-swellable inner hydrophobic region. Block copolymers contain two or more chemically different polymers connected by covalent bonds that offer a means of combining the desirable characteristics of different polymers into a hybrid material22. The mechanical properties and swellability of the outer region of the MNs can be controlled by manipulating the overall average molecular weight of the polymer and the weight fraction of each block20,23. In this study, we chose a polystyrene-block-poly(acrylic acid) (PS-b-PAA) block copolymer as the swellable material and PS homopolymer as a non-swellable material. PAA is a well-known super-absorbent polymer used in several biomaterials-based strategies and in consumer products such as diapers that possesses COOH groups that quickly become ionized in the presence of water. In contrast, PS exhibits mechanical strength and structural integrity without swelling. As predicted, high volume expansion occurred when PAA was the major block component (over 70% in weight fraction). The mechanical robustness of the block copolymer could be strengthened by increasing the weight fraction of the PS block (Supplementary Fig. S1). Interestingly, the stiffness of the block copolymer in its swollen state, with PS of ~25% weight fraction, approximates that of skin or intestine tissue (Supplementary Fig. S1), which reduces the risk of underlying tissue damage without significantly compromising the interlocking mechanism. To permit rapid water penetration into the PS-b-PAA layer when the PS block preferentially assembles at the air-copolymer surface, we selected a PS block with a relatively low-molecular weight (<100,000 g mol−1) that minimizes interfacial surface area coverage due to short polymer chains. To achieve this, PS-b-PAA was prepared from the hydrolysis of polystyrene-block-poly(tert-butyl acrylate) (PS-b-PtBA) with number-average-molecular weights (M n ) of 26,000 g mol−1 for PS and 128,000 g mol−1 for PtBA. Complete conversion of PtBA to PAA was confirmed by nuclear magnetic resonance (NMR) (Supplementary Fig. S2). For functional testing in tissue, we selected swellable tips fabricated with PS-b-PAA with M n of 26,000 g mol−1 and 76,000 g mol−1 for PS and PAA, respectively, and a PS weight fraction of ~25%.

As shown in Fig. 1a, b a double-layered MN array having water-swellable tips was fabricated by solvent-casting a PS-b-PAA solution in N, N-Dimethylformamide (DMF) on a PDMS (polydimethylsiloxane) mould with a 10 × 10 array of conical cavities within a 1 cm2 area. Drying of the solvent prompted the formation of a thick film at the tip region of the conical cavity via capillary forces and a thin film on the remainder of the mould. To form the MN inner core and backing layer of the film, the PS homopolymer was melted on top of the continuous PS-b-PAA layer at 180 °C under vacuum. After cooling to room temperature, the double-layered conical MN array was carefully peeled from the PDMS mould.

Macroscopically, the double-layered BCP MN exhibits a ‘needle-in-needle’ structure consisting of swellable outer PS-b-PAA layer and supporting PS inner core (Fig. 1c). Microscopically, the PS-b-PAA layer self-assembles into multiple phases whose location and properties are dependent on the interaction with the mould and air surface. To examine the composition of the swellable polymer surface at the air and PS core interfaces, we performed X-ray photoelectron spectroscopy (XPS) on PS-b-PAA films that were casted onto PS and PDMS substrates (Supplementary Fig. S3). Given that the theoretical C/O ratio for PAA is 1.5, the C/O ratio at the surface of the PS-b-PAA layer on PDMS (14.9) and PS (22.3) suggests that PS was the dominant interfacial block. Thus, we speculate that a loosely packed PS film with a thickness of ~10 nm forms at the surface of the PDMS mould (air interface following removal from the mould) and modulates the rate of diffusion, preventing a rapid modulus drop caused by immediate water absorption, which would result in insertion failure of the BCP MN array. Furthermore, during filling of inner core by melted PS, the PS block in the PS-b-PAA layer can entangle with a PS homopolymer core and the COOH groups in the PAA chains thermally crosslink via intermolecular anhydride formation24,25. By varying the concentration of the casting solution, the height of the PS-b-PAA layer was controlled to be either 20, 40 or 70% of BCP MN height (Fig. 1d–f). Importantly, the PS base material completely filled the inner cores showing good contact with the PS-b-PAA layer (Fig. 1g). Figure 1h shows macroscale images of a double-layered BCP MN array with high-pattern fidelity having a density of 10 × 10 cm−2, and having a swellable tip height of 20%, a MN base diameter of 280 μm and a MN height of 700 μm. An overall needle height of 700 μm was selected to ensure that the needle tip swelling would occur beneath the surface of the tissue (for example, thickness of epidermis is less than 100 μm at typical needle insertion sites)26 and to minimize the penetration depth.

Reversible shape change of a water-responsive swellable MN

To visualize the real-time shape change of swellable BCP MN, we monitored the swelling behaviour immediately following insertion into a transparent agarose hydrogel (see Supplementary Fig. S4 for experimental setup). Once BCP MNs were inserted into the hydrogel (tip height ~40% of total MN height), the volume of the PS-b-PAA tip layer rapidly increased and reached ~60% of the maximum swollen state within 1 min. Subsequently, swelling equilibrated to a maximum swollen state within 10 min (Fig. 2a and Supplementary Movie 1). The kinetics of water absorption of micron-sized PS-b-PAA were much faster than those of bulk material with mm-scale thickness (Supplementary Fig. S1), which was likely the result of the high surface area per volume ratio for the MN tips and short water diffusion length. At equilibrium, the volume of PS-b-PAA MN layer had expanded to ~9 times its initial volume (Supplementary Fig. S5). Volume expansion of BCP MN by swelling predominantly occurred in the tip region and swollen BCP MN interlocked with the hydrogel. The interlocked BCP MN could be removed from agarose gel without breakage or delamination. The swollen tips immediately began to deswell following retrieval from the hydrogel substrate and fully recovered to their original conical structure within 15 min (Supplementary Fig. S6 and Supplementary Movie 2).

Figure 2: Swelling of BCP MN following insertion into a hydrogel and muscle tissue. (a) Time-dependent swelling of the BCP MNs (40% swellable tip height fraction) following insertion into a 1.4 wt% agarose hydrogel (0, 60 and 600 s). (b) OFDI images showing swelling of the same BCP MNs following insertion into muscle tissue (0, 120, 360 and 600 s). Scale bar, 500 μm. Full size image

When PS or PAA homopolymers were used as the MN tip material, they produced opposing results. The non-swellable PS MN was easily inserted into the hydrogel and freely removed without significant applied force, whereas the PAA MN did not penetrate the hydrogel because of low mechanical strength (Supplementary Fig. S7). Although it is possible to make stiff PAA MN that can be inserted into hydrogel (or tissue) by crosslinking following harsh thermal treatment, pure PAA MN is also prone to delamination during removal of the MN due to poor interfacial adhesion with the backing layer.

To verify the stimulus-responsive mechanical interlocking of the BCP MN in animal tissue, we utilized optical frequency domain imaging (OFDI) that is capable of providing non-invasive cross-sectional views of internal tissue structures to a depth of 1–2 mm with high resolution27. While the BCP MN inserted into muscle tissue showed less initial volume expansion rate compared with the agarose hydrogel, it reached the same final swollen state within 10 min (Fig. 2b and Supplementary Movie 3). Interestingly, insertion into muscle tissue induced radial expansion of the BCP MN compared with axial swelling within the homogeneous agarose hydrogel. As observed within muscle tissue via OFDI, the swollen BCP MN formed a mushroom-shaped structure that provides a more favourable geometry to support mechanical interlocking with tissue.

Firm adhesion of the swellable MN adhesive

The adhesive properties of the BCP MN were evaluated by testing the normal adhesion strength of a 10 × 10 MN array (1 cm2 cross-sectional area) on fresh cadaveric porcine skin. We investigated the effect of swelling-induced MN shape change on adhesion, compared with the adhesion of non-swellable PS MN (Fig. 3a). Adhesion of flat PS-b-PAA (BCP) and PS films were also measured. Both flat-surfaced BCP and PS films showed low adhesion strength to skin (<0.1 N cm−2). While, BCP MN and PS MN showed similar force versus displacement profiles during skin insertion (Supplementary Fig. S8) and exhibited similar adhesion strength to skin immediately after insertion, after 10 min the maximum swollen state for BCP MN was achieved and the adhesive strength of the BCP MN dramatically increased. BCP MN with a swellable tip height of 20% (BCP MN (20%)) showed ~7 times higher adhesion strength (0.69±0.17 N cm−2) than PS MN adhesive (0.098±0.015 N cm−2). Furthermore, for BCP MN with a swellable tip height of 40% (BCP MN (40%)), the adhesive strength increased ~12-fold (1.23±0.26 N cm−2). However, MN with a swellable tip height of 70% (BCP MN (70%)) exhibited reduced adhesive performance (Supplementary Fig. S9), likely due to a less favourable configuration for interlocking with tissue. No significant increase in adhesive strength was observed for the BCP MN for longer than 30 min swelling (Fig. 3b). For most intended uses, a short application time is desirable; therefore, the swelling time was standardized at 10 min for all additional experiments.

Figure 3: MN adhesive firmly adheres to skin. (a) Normal adhesion strength for PS MN and BCP MN adhesives with 20 and 40% swellable tip height fractions following insertion into skin (2 and 10 min). Flat PS and PS-b-PAA films were used as controls. (b) Effect of swelling time within skin on adhesion for a BCP MN adhesive with a swellable tip height fraction of 40%. (c) Representative force-displacement curve during BCP MN insertion into, and removal from pig skin. (d) Photograph of flexible BCP MN adhesive (2 × 2 cm) prepared by using a thermoplastic elastomer as the base material (in place of PS). (e) Adhesion on a dynamic surface. Flexible BCP MN adhesives applied to shaved skin on top of the pig wrist joint showed firm attachment during ~100 cycles of bending motion. All error bars represent standard deviation. Full size image

The shape change adhesive mechanism of the BCP MN within porcine skin is shown in Fig. 3c. Following insertion into skin (Fig. 3c(i)), the tissue recoils and applies a mechanically compressive force against the BCP MN shaft. The position of the BCP MN is held fixed as the inserted BCP MN immediately swells in the presence of interstitial fluid and rapid volume expansion nears completion within 10 min, where the shape of the swollen tip stabilizes. The concentrated volume expansion at the tip of the BCP MN (Fig. 3c(ii)) leads to an arrow-head-like structure facilitating mechanical interlocking with the tissue. Removal of the BCP MN from the skin can be safely performed without fracture or delamination due to the deformability of the swollen tip and strong interfacial interaction between the swollen layer and base material (Fig. 3c(iii)). Significant adhesive strength was maintained even at high levels of strain between BCP MNs and skin until BCP MNs are fully removed from the skin, indicating that the BCP MN exhibits high adhesion energy. Owing to the reversible responsiveness in water (Fig. 3c(iv)), the swollen BCP MNs quickly return to their original conical structure after removal. Following recovery of shape and stiffness (Supplementary Fig. S10), the BCP MN exhibited reversible adhesive properties during multiple swelling/deswelling cycles; therefore it is possible that the BCP MN could be reused following sterilization by ethanol or autoclave treatment.

We compared the adhesive strength of the BCP MN with currently used commercial pressure-sensitive adhesive bandages on semi-dry and wet porcine skin surfaces (Supplementary Fig. S11). While commercial bandages show higher adhesive strength on dry skin surfaces compared with the BCP MN adhesive, the adhesive strength of commercial bandages on wet skin decreased by more than 50%. However, BCP MNs showed similar adhesive strength regardless of the presence of water on the tissue surface. To demonstrate the versatility of the BCP MN adhesive, the stiff PS base material was substituted for a flexible thermoplastic PS-based elastomer (Fig. 3d). The flexible BCP MN adhesive (Flex BCP MN) with 10 × 10 array in a 1 cm2 was applied to a wet skin surface on top of a pig wrist joint that was cycled through an angle of ~60° to stretch and compress the tissue. Flex BCP MNs maintained strong attachment without migration during 100 cycles of bending motion (Fig. 3e and Supplementary Movie 4).

MN adhesive for use in skin graft fixation

With regard to potential immediate applications of BCP MN adhesives, we envision fixation of skin grafts as a strong candidate. Skin grafts are often employed for closure of open wounds as a result of burns, trauma or surgical resections28,29. For successful engraftment, continuous contact between the skin graft and underlying tissue is essential to assure graft survival by directed diffusion of wound bed nutrients as well as to prevent haematoma or seroma formation30. Sutures or staples applied to the perimeter of the skin graft represent the current standard of care, and frequently this results in separation between the tissue layers, as direct fixation only occurs along the periphery of the wound. Furthermore, fixation by staples is not rigid enough to avoid motion and shear resulting frequently in partial or complete loss of the graft. The likelihood of failure is diminished by secondary dressings, and grafted body part immobilization.

From dye diffusion tests, we found that staples provided a surface contact area of ~50% between the skin graft and underlying hydrogel, and that the staples produced significant damage to the hydrogel during insertion and removal (Fig. 4a). In contrast, the BCP MN adhesive showed continuous contact of ~100% with minimal damage to the hydrogel (Fig. 4b). As marked by an arrow in Fig. 4a, the depth of penetration (~3 mm) of staples causes significant tissue damage and results in a higher risk of bacterial infiltration. In adhesion testing, although stapled skin grafts did not fully detach from underlying tissue during pull-off, central regions of the skin grafts were separated due to low adhesion strength (Fig. 4c). Skin grafts fixed using the BCP MN adhesive on muscle tissue showed significantly higher adhesive strength (0.93±0.23 N cm−2) than stapled skin grafts (0.28±0.11 N cm−2) and non-fixed skin grafts (0.22±0.09 N cm−2) (Table 1). The attachment was maintained following an elongation of ~4 mm, which is more than five times the height of the BCP MN (Fig. 4d). To prevent the accumulation of fluid at the interface between BCP MN adhesive and skin graft, drainage holes can easily be placed within the backing material of the adhesive array (Supplementary Fig. S12). An ideal tissue adhesive should also provide a biological barrier to reduce the risk of surgical site infection. In the case of staple fixation, stress concentrations localized around the staple legs can cause excessive tearing of a skin graft and produce a larger hole than a diameter of the staple (Fig. 4e). Staple holes serve as a pathway for bacterial infiltration12 and as expected, Escherichia coli (E. coli) transformed to express the green fluorescent protein (GFP), infiltrated through the staple holes and formed colonies (Fig. 4f). For the BCP MN adhesive, the interlocked and swollen MNs tightly seal the punctured holes, thus preventing the physical passage of microorganisms (Fig. 4g). The BCP MN adhesive was removed without a significant damage to the agar layer (left image in Fig. 4h) and GFP signal was not detected in the agar plate placed beneath the incised skin grafts covered by the BCP MN (right image in Fig. 4h). As a result of the non-permeable backing layer and tight sealing of the punctured holes by the swollen MNs, the BCP MN adhesive effectively prevents the infiltration of bacteria, which is a main cause of infection.

Figure 4: MN adhesive achieves effective fixation of skin grafts and resists bacterial infiltration. (a,b) Comparison of the contact area between a skin graft and tissue-like hydrogel (4 wt% agarose gel) after (a) applying staples and (b) a BCP MN adhesive. Staples showed less contact area between the skin graft and underlying hydrogel and on removal induced significant damage to the underlying hydrogel (region marked by arrow in Fig. 4a), while the BCP MN adhesive showed continuous contact ~100% with minimal damage to the hydrogel. (c,d) Force-displacement profiles and photographs acquired during pull-off tests of skin graft on muscle tissue fixed by (c) staples and (d) BCP MN adhesive. While the stapled skin graft was easily separated from the underlying muscle with a low pull-off strength, the BCP MN adhesive provided continuous contact between the skin graft and muscle tissue via mechanical interlocking with underlying muscle tissue. (e–h) Comparison of the bacterial barrier property of incised skin grafts following application of (e,f) staples and (g,h) a BCP MN adhesive. (e(i)) Cartoon illustrating the primary site of bacterial infiltration through the gaps between staple legs and skin grafts. (e(ii)) A photograph showing stapled skin grafts after bacteria infiltration; cyanoacrylate glue was used to tightly seal the incised region (dark area outlined by dotted black line). (f) GFP-expressing E. coli colonies formed near deep staple holes (marked by red dots) where skin grafts did not appose the underlying agar layer (left, bright field image). The infiltration of E. coli through the staple holes was confirmed by green fluorescence (right, fluorescent image). (g(i)) Cartoon showing bacterial barrier resistance of the BCP MN adhesive resulting from tight sealing of holes by swollen MNs. (g(ii)) Photograph of the BCP MN adhesive applied on the incised skin grafts. (h) BCP adhesive prevented bacterial infiltration (left, bright field image) with minimal damage and (right, fluorescent image) no green fluorescence was detected on the agar plate. Scale bar, 1 mm. Full size image

Table 1 Tissue adhesion for skin grafts fixed by BCP MN adhesive and staples. Full size table

While the micron-sized puncture marks created after removal of the MNs from the wound could serve as a source of infection, it has been reported that the skin recovers its barrier function within a few hours following MN removal31. The disappearance of puncture marks was confirmed within 1 h of removal of the swellable MN adhesive (Supplementary Fig. S13). In addition, it has been reported that microbial penetration through the holes created by MNs is minimal32. Considering its rapid and long-term adhesion, this superior bacterial barrier provided by the BCP MN adhesive could reduce the risk of infection to open wounds including burn tissue.

Firm adhesion of the MN adhesive to intestinal tissue

To evaluate the capacity for BCP MN adhesives to be used to seal intestine tissue, which may be useful to prevent leaks following gut anastomosis procedures, we examined adhesion to the outer serosal surface (Fig. 5a) that is a relatively smooth surface (roughness of several micrometres) and inner wrinkled mucosal surface (roughness on the sub-mm scale) that is covered by a sticky mucin layer (Fig. 5b). The non-swellable PS MN adhesive applied to the outer surface of the intestine tissue (Fig. 5c) showed an adhesive strength of 0.48±0.18 N cm−2, whereas a swellable MN adhesive with a PS base (BCP MN) showed an adhesive strength of 1.62±0.17 N cm−2. Interestingly, the BCP MN adhesive exhibited an adhesive strength of 3.83±1.35 N cm−2 to the inner mucosal surface, possibly resulting from non-covalent interactions with mucin. Moreover, the Flex BCP MN showed further improvement in adhesive strength to the mucosal surface, up to 8 N cm−2 (mean value: 4.53 N cm−2). The enhanced adhesion of the Flex BCP MN on the mucosal surface may result from increased intimate contact during removal that may be facilitated by increased absorption of energy by the flexible backing material. The BCP MN adhesive appears to achieve adhesion to tissues that cannot be achieved effectively using conventional tape-based approaches33. As shown in Fig. 5d, inserted PS MNs into the outer surface of intestine tissue showed an average maximum torque of 0.85±0.28 Ncm during torsion tests, through a 60° angle of rotation. After removing PS MNs from the tissue, it was observed that ~30% MNs remained broken. When the PS MNs were exposed to 100° of rotation before removal from tissue, they showed a higher maximum torque (1.22±0.41 Ncm) and nearly all tips or whole bodies were broken (Fig. 5e). Comparatively, less than 5% of the BCP MN and none of the Flex BCP MN were broken following 60° of rotation, and less than 5% of Flex BCP MN were broken following 100° of rotation (Fig. 5d). Interestingly, puncture marks resulting from the BCP MN insertion into tissue disappeared within 2 h of removal (Supplementary Fig. S14).