A Hierarchically Structured, Stretchable, Anti‐Biofouling Encapsulation for Biodegradable Electronics

Abstract

Biodegradable polymers have been employed as encapsulants for transient, resorbable implantable devices due to moderate water permeability, mechanical flexibility, and biocompatibility, however most of them relatively lack inherent anti‐biofouling properties. This limitation can lead to undesired protein adsorption, cell adhesion, and fibrotic encapsulation, compromising device function and biocompatibility, particularly for long‐term implantation scenarios. Here, this study introduces a soft, stretchable, and anti‐biofouling encapsulant engineered by integrating self‐assembled organosilicon nanowire networks onto micropatterned biodegradable elastomers. The resulting hierarchical surface architecture imparts superhydrophobicity while preserving mechanical integrity, improving water barrier performance by up to 420% compared to unmodified films and retaining stability under cyclic strains. Integration into a transient, stretchable optoelectronic device enables prolonged operation in aqueous environments, and in vitro and in vivo evaluations demonstrate suppressed cell adhesion, reduced fibrotic tissue formation, and excellent biocompatibility, highlighting the potential for long‐lasting, bioresorbable electronic implants.

A soft, stretchable, anti‐biofouling encapsulant is developed by integrating organosilicon nanowire networks with microstructured biodegradable elastomers. The hierarchical surface exhibits superhydrophobicity and enhances water barrier properties by 420% over pristine polymers, while preserving mechanical integrity. In vitro and in vivo studies confirm inhibited cell adhesion, reduced fibrosis, and excellent biocompatibility, validating its potential for clinically scalable, long‐lasting bioresorbable electronic implants.

1. Introduction

Soft, bioresorbable implantable devices have attracted significant attention as next‐generation therapeutic platforms capable of performing a wide range of functions—including biosignals and chemical sensing,^{[ 1} , ² , ³ , ⁴ , ^{5 ]} neuroregeneration,^{[ 6} , ⁷ , ^{8 ]} pacemaking,^{[ 9} , ¹⁰ , ^{11 ]} drug delivery,^{[ 12} , ¹³ , ^{14 ]} and tissue/bone repair^{[ 15} , ¹⁶ , ^{17 ]}–within a predefined time window, followed by complete degradation in the body. These transient devices eliminate the need for secondary removal surgeries, reduce the risk of infection, and alleviate patient discomfort, offering evident advantages in clinical contexts. However, their successful operation in physiological environments critically depends on robust encapsulation layers that protect transient electronic components from biofluid‐induced degradation while preserving mechanical compliance and biocompatibility. In this context, various natural and synthetic polymers, such as silk fibroin, wax, poly(lactic‐co‐glycolic acid) (PLGA), poly(glycolide‐co‐ε‐caprolactone)(PGCL), poly(buthanedithiol 1,3,5‐triallyl‐1,3,5‐triazine‐2,4,6(1H,3H,5H)‐trione pentenoic anhydride) (PBTPA), and bis(3‐mercaptopropyl) succinate copolymers, have been explored as encapsulant materials.^{[ 18} , ¹⁹ , ²⁰ , ²¹ , ²² , ²³ , ^{24 ]} While these materials provided flexibility/stretchability, biodegradability, and ease of processing, they often exhibited limited barrier performance against water and ions, leading to premature device failures. As a promising alternative, polymer composites incorporating water‐impermeable particles have been developed to extend the water diffusion path, thereby enhancing encapsulation performance.^{[ 25} , ^{26 ]} Nevertheless, most existing encapsulants lack inherent anti‐biofouling functionality, leading to undesirable protein adsorption or cell adhesion. This biofouling can not only trigger immune responses and fibrosis but also impair device function and biocompatibility.^{[ 27} , ²⁸ , ^{29 ]} Given the complexity of in vivo environment—especially for long‐term implantation—effective biofouling prevention is as critical as ensuring water resistance.

Here, we present a soft, stretchable, and anti‐biofouling encapsulant that can suppress cell adhesion, minimize inflammation, and extend the functional lifespan of biodegradable medical implants. Self‐assembly of organosilicon nanowire (OSN) networks onto microstructured biodegradable elastomers created a hierarchical surface that achieves superhydrophobicity without compromising mechanical integrity. Such approach for dissolvable electronic components significantly enhanced water barrier performance—by up to 420 % compared to pristine polymers—and demonstrated broad applicability across various biodegradable polymer systems. Comprehensive in vitro and in vivo evaluations validated complete inhibition of cell attachment and reduced fibrosis, along with excellent biocompatibility, which demonstrates the potential for a scalable and clinically relevant solution for the long‐term reliability of bioresorbable electronics.

2. Results and Discussion

2.1. Organosilicon Nanonetworks‐Based Hierarchical Structure for Hydrophobic, Anti‐Biofouling Encapsulants

Figure 1a presents the concept of an anti‐biofouling encapsulant for stretchable and biodegradable electronic devices. The encapsulant consisted of a hybrid structure that integrates self‐assembled organosilicon nanowire (OSN) networks with a biodegradable elastomer, poly(L‐lactide‐co‐ε‐caprolactone) (PLCL), featuring precisely engineered microscale surface topographies including micro‐pillars or ‐holes with a diameter of ≈5 µm, an aspect ratio of 2, and a center‐to‐center pitch of ≈7.5 µm. Here, the dimensions were chosen to achieve the Cassie‐Baxter (CB) wetting state, in accordance with previous findings.^{[ 30} , ^{31 ]} The OSN networks were synthesized through the controlled self‐assembly of methyltrichlorosilane (MTCS) in humid conditions, forming interpenetrating nanomeshes across the microstructured polymer surface. Details of the fabrication process are provided in Figure S1 (Supporting Information). The incorporation of OSN networks onto the microstructured PLCL generated a hierarchical dual‐scale roughness, which significantly enhanced water repellency by promoting the CB wetting state. Figure 1b demonstrates a superhydrophobic behavior of the resulting PLCL‐OSN composite with pillar patterns (referred to as PLCL‐p‐OSN). When a water droplet was gently applied to the surface at a tilt angle of only 1°, it immediately rolled off without leaving any footprint, indicating a high water contact angle and an ultralow sliding angle. This performance was attributed to the minimal solid‐liquid contact area and air entrapment within the surface textures.^{[ 32 ]} To evaluate the anti‐biofouling performance, mouse embryonic fibroblast cells (NIH‐3T3) were cultured on both pristine PLCL and PLCL‐p‐OSN films at 37 °C for 72 h (Figure 1c). The pristine PLCL surface supported strong cell adhesion and proliferation, forming a dense cell monolayer. In contrast, the PLCL‐p‐OSN surface completely inhibited cell attachment, indicating excellent resistance to biofouling, possibly attributed to the combined effects of reduced effective contact area and physical surface topography that impedes stable cell adhesion and spreading.^{[ 33} , ^{34 ]} It is noteworthy that the degradation process primarily involves hydrolytic chain cleavage of the PLCL substrate, which also contributes to the gradual disintegration of the OSN network into smaller fragments, although the clearance pathways remain to be fully elucidated.

Figure 1.

Soft, stretchable, biodegradable, and anti‐biofouling encapsulant. a) Schematic illustration of a poly(l‐lactide‐co‐ε‐caprolactone) (PLCL) micropillar array decorated with organosilicon nanowire network (PLCL‐p‐OSN), designed to inhibit biofouling and water permeation for stable operation of stretchable, biodegradable electronics, along with photograph of a representative sample in the inset. b) Superimposed time‐lapse images capturing a water droplet rapidly rolling off a PLCL‐p‐OSN surface tilted at 1°, without adsorption, along with the fabricated PLCL‐p‐OSN film in the inset. c) Fluorescence microscopy images of pristine PLCL and PLCL‐p‐OSN films after 24 h culture with NIH‐3T3 cells.

2.2. Formation and Characterization of OSN Networks on Polymer Films

It is important to conduct nanoscale inspections of chemically modified products and characterize the resulting features using various types of tools. Figure 2a illustrates the formation process of OSN networks on a PLCL film. Initially, the non‐polar methylene (─CH₂─) and ester (─COO─) groups in PLCL were converted into hydroxyl (─OH) and carboxyl (─COOH) groups via radical‐mediated oxidation using a strong oxidizing agent, ammonium persulfate (see detailed reaction scheme in Figure S2, Supporting Information).^{[ 35} , ^{36 ]} In contrast, oxygen plasma treatment—commonly used for surface modification—proved less effective in introducing hydroxyl groups, likely due to the low reactivity of functional groups in PLCL toward plasma‐generated radicals such as •O, •OH, and O₂•⁻. The hydroxylated film was then exposed to MTCS under humid conditions to form OSN structures. Notably, humidity played a critical role in this process: water vapor reacted with Si─Cl bonds in MTCS to generate silanol (Si─OH) groups, which then spontaneously condensed to form siloxane (Si─O─Si) crosslinked networks, as shown in the following reactions:

$$\mathrm{C}{\mathrm{H}}_3\mathrm{SiC}{\mathrm{l}}_3+3{\mathrm{H}}_2\mathrm{O}\to \mathrm{C}{\mathrm{H}}_3\mathrm{Si}{\left(\mathrm{OH}\right)}_3+3\mathrm{HCl}$$ (1)

$$-\mathrm{Si}-\mathrm{OH}+\mathrm{HO}-\mathrm{Si}-\to -\mathrm{Si}-\mathrm{O}-\mathrm{Si}-+{\mathrm{H}}_2\mathrm{O}$$ (2)

Figure 2.

Morphological, physical, and mechanical characterization of OSN network formed on PLCL with and without pillar and hole structures. a) Schematic illustration of the formation mechanism of OSN networks via hydroxylation of PLCL surface followed by reaction with methyltrichlorosilane (MTCS). b) Fourier‐transform infrared (FT‐IR) spectra of pristine PLCL, hydroxylated PLCL (PLCL‐OH), and silanized PLCL (PLCL‐OH‐OSN). c) Scanning electron microscopy (SEM) images of PLCL with pillar and hole (inset) structures before (left) and after (middle) OSN formation, and magnified image (right) of the red box, showing detailed nanowire morphology. d,e) Optical images (d) of water droplets on PLCL, PLCL‐OH, and PLCL‐OH‐OSN with and without pillar or hole structures, and the corresponding water contact angles (WCAs) (e). n = 3. f) Stress‐strain curves of PLCL and PLCL‐OH‐OSN films (200 µm‐thick).

Under controlled humidity, this process enabled nucleation and anisotropic growth, leading to 1D or fibrous nanostructures. Figure 2b confirms successful OSN formation. Compared to pristine PLCL, hydroxylated PLCL (PLCL‐OH) exhibited a slightly enhanced broad absorption band between 3200 and 3700 cm⁻¹, attributed to O‐H stretching vibrations.^{[ 37 ]} In MTCS‐treated films (PLCL‐OH‐OSN), distinct new peaks appeared at 1271, 1014, and 779 cm⁻¹, corresponding to Si‐CH₃ bending, Si─O─Si stretching, and Si─C stretching vibrations, respectively. Figure 2c displays the surface morphology of micropatterned PLCL films (with pillars or holes) before and after MTCS treatment. Appropriate treatment durations produced rough, densely packed OSN structures (≈15 nm in diameter) conformally covering the pillar and hole surfaces. However, deviations from the optimal duration resulted in either sparse or excessively dense OSN formation (Figure S3, Supporting Information). Figure 2d,e shows measured water contact angles of different PLCL films (flat, hole‐patterned, and pillar‐patterned) before and after hydroxylation and silanization. The contact angles at the intrinsic state were 84 °, 120 °, and 135 ° for each geometry, while hydroxylation slightly reduced these angles due to the introduction of hydrophilic groups. Subsequent OSN formation significantly increased the contact angles to over 150 °, indicating that the surface was modified to be superhydrophobic. It should be noted, however, that excessive OSN coverage resulted in a slight reduction in the water contact angle. Figure 2f and Figure S4 (Supporting Information) demonstrate that the surface modification process had negligible impact on the mechanical integrity of the PLCL film, suggesting the strong potential for robust and functional surface engineering.

2.3. Applications of OSN Network‐Based Hierarchical Structures in Stretchable, Biodegradable Electronics

The resulting materials and properties can be applied to electrical and optical devices to determine the practical utility. Figure  3a illustrates an experimental setup to evaluate the encapsulation performance of OSN network‐decorated polymers. A 300 nm‐thick magnesium (Mg) trace was deposited on a glass substrate, and a 300 µm‐thick encapsulant film was laminated onto the trace, followed by attachment of a polydimethylsiloxane (PDMS)‐based chamber. The chamber was then filled with phosphate‐buffered saline (PBS, pH 7), and changes in electrical resistance of the dissolvable metal were monitored. Here, the time at which the resistance began to increase was defined as functional lifetime. Figure 3b shows the resistance profiles of Mg traces encapsulated with various materials. Pristine PLCL exhibited a functional lifetime of ≈18 h, whereas the introduction of surface textures, particularly OSN networks, significantly extended this duration. A hierarchical structure combining micropillars and OSN networks (PLCL‐p‐OSN) achieved enhanced service lifetime of ≈75 h (≈420 % improvement over pristine PLCL), which was further increased to ≈100 h by incorporating a 100 nm‐thick SiO₂ layer. Encapsulation performance of hole‐patterned samples appears in Figure S5 (Supporting Information). Functional lifespan linearly increased with thickness of the encapsulant (Figure 3c), consistent with the extension of the expected water diffusion path.^{[ 23} , ^{26 ]} Notably, a 500 µm‐thick PLCL‐p‐OSN film achieved a functional lifetime exceeding 130 h. To assess the mechanical durability, PLCL‐p‐OSN films were subjected to cyclic tensile strains of 30 %, 50 %, and 100 %, followed by functional lifetime measurements. As depicted in Figure 3d, increased strain levels and cycle counts led to a slight reduction in performance. However, the degradation remained well below 10 %, likely due to the elasticity of the polymer matrix and the porous, compliant nature of the OSN layer. The versatility of this surface modification approach was further demonstrated on various biodegradable polymers, including polycaprolactone (PCL), poly(lactic acid) (PLA), polyurethane (PU), and a PLCL composite containing silicon dioxide (SiO₂) nanoparticles (Figure 3e). While the inherent water barrier properties varied depending on polymer hydrophobicity and water‐impermeable nanoparticle content, OSN treatment consistently enhanced the functional lifetimes. For example, pristine PLCL+nanoparticle(NP) exhibited a lifetime of ≈5.6 days, which extended to ≈16.7 days after OSN treatment. We note that excessive thickness not only increases the mechanical rigidity of devices but also reduces soft, stretchable properties, which may lead to functional impairment and mechanical mismatch with surrounding tissues or organs. In this context, the surface‐engineering strategy improved protection without changing thickness or the form factor of devices, thereby retaining softness, flexibility, and mechanical compatibility with tissues. As a proof of concept, PLCL‐p‐OSN and PLCL‐h‐OSN films were employed as encapsulation and substrate layers for a stretchable, transient optoelectronic device composed of an infrared (IR) light–emitting‐diode (LED) connected via 300 nm‐thick Mg interconnects (Figure 3f). Thermographic imaging of the devices submerged in PBS (pH 7, at 37 °C) revealed stable device operation for 4 days, with complete functional failure on day 5, consistent with the encapsulation performance of PLCL‐p‐OSN and PLCL‐h‐OSN. Data plots of the emission intensity at 850 nm (Figure 3g) further confirmed this behavior, showing steady outputs through day 4, followed by a rapid decline to negligible levels. Although the proposed approach was validated for the device‐level performance, long‐term experiments using practical, system‐level devices could be a valuable way to explore the clinical applicability.

Figure 3.

Encapsulation capability of OSN‐formed PLCL films for stretchable, biodegradable electronics. a) Experimental setup for evaluating the encapsulation performance of OSN‐coated PLCL films. The encapsulant delays water penetration to the underlying dissolvable magnesium (Mg) electrode, and the time at which a substantial increase in resistance occurs is defined as the functional lifetime. b) Resistance changes of Mg electrodes encapsulated with various film types: PLCL, PLCL‐p, PLCL‐OSN, PLCL‐p‐OSN, and PLCL‐p‐OSN with a 100 nm of SiO₂ layer. c) Functional lifetimes of PLCL‐p‐OSN films as a function of film thickness. n = 3. d) Relative encapsulation performance of PLCL‐p‐OSN films after repeated deformation at 30 %, 50 %, and 100 % strains. n = 3. e) Comparison of functional lifetimes for various polymer matrices before and after Si nanowire formation. f) Schematic (left) and time‐dependent infrared (IR) thermographic images (right) of a stretchable, transient optoelectronic device consisting of an IR light‐emitting‐diode (LED) device supported on a PLCL‐h‐OSN and encapsulated with PLCL‐p‐OSN, immersed in phosphate‐buffered saline (PBS, pH 7) at 37 °C. g) Measured IR intensities of the LED over time.

2.4. Anti‐Biofouling Effect and Biocompatibility of OSN Network‐Based Hierarchical Structures

To confirm whether PLCL‐p‐OSN affects cell adhesion and exhibits an anti‐biofouling effect, we inoculated NIH‐3T3 mouse fibroblast cell lines onto the material and observed cell adhesion. As shown in Figure  4a, cells in the Control, PLCL, and PLCL‐p groups adhered to the surface on day 1, however cell adhesion was not observed in the PLCL‐p‐OSN group, similar to the PDMS group, which is known to have inherently low surface energy. On Day 3, prominent cell proliferation was observed in the Control group, while no additional cell attachment was observed in the PDMS and PLCL‐p‐OSN groups. In the quantitative analysis of cell attachment (Figure 4b), the PLCL‐p‐OSN group showed 97.5%, 95.4%, and 96.1% less cell attachment than the Control, PLCL, and PLCL‐p groups, respectively, on Day 1. Similar tendency was observed on Day 3, and the PLCL‐p‐OSN group showed 99.5%, 98.9%, and 98.2% less cell attachment compared to the Control, PLCL, and PLCL‐p groups.

Figure 4.

Evaluation of anti‐biofouling effect and biocompatibility. a) Fluorescence microscopy images of NIH‐3T3 cells cultured on a standard culture dish (control) and on PDMS, PLCL, PLCL‐p, and PLCL‐p‐OSN films, where live and dead cells were stained with green and red, respectively. Scale bars: 100 µm. b) Quantification of NIH‐3T3 cell numbers after 1, 2, and 3 days of culture on each surface. n = 4. c) Hematoxylin and eosin (H&E) and Masson's trichrome (MT)‐stained histological images of subcutaneous tissues in rats implanted with polydimethylsiloxane (PDMS), PLCL, PLCL‐p, and PLCL‐p‐OSN films for 3 weeks. Scale bars: whole mount (400 µm), 200x (50 µm), Masson's Trichrome (400 µm). d) Quantitative comparison of fibrous tissue thickness for each sample at 1 and 3 weeks post‐implantation. n = 3. e) H&E‐stained histological sections of major organs, including heart, lung, liver, kidney, and spleen, collected at 1 and 3 weeks after implantation of PLCL‐p‐OSN films. Scale bars: 100 µm.

To investigate the behaviors in vivo, we implanted the experimental groups (PDMS, PLCL, PLCL‐p, PLCL‐p‐OSN) subcutaneously in mice and analyzed the thickness of fibrous tissue and the presence or absence of systemic toxicity at 1 week and 3 weeks post‐implantation. Histopathological evaluation results showed that none of the implants exhibited a prominent local inflammatory response (Figure 4c; Figure S6, Supporting Information). However, in all groups, tissue stained eosinophilic by hematoxylin and eosin (H&E) staining was observed surrounding the implants, and these tissues were confirmed to be fibrous tissue (blue) showing a positive reaction to Masson's trichrome staining. Interestingly, the PLCL‐p‐OSN group showed a significant reduction in the thickness of the fibrous capsule surrounding the implant at 1 week compared to the other groups. This trend continued until 3 weeks, with the PLCL‐p‐OSN group showing the thinnest fibrous capsule thickness at 3 weeks compared to the other groups. In the quantitative analysis of fibrous tissue, the PLCL‐p‐OSN group showed 55.3% (PDMS), 67.3% (PLCL), and 30% (PLCL‐p) thinner fibrous tissue thickness at week 1, and 53.2%, 53.8%, and 27.4% at week 3, compared to the other groups. (Figure 4d). In the histopathological analysis of the heart, lung, liver, kidney, and spleen to assess the toxic effects of the implant, all experimental groups (PDMS, PLCL, PLCL‐p) including the PLCL‐p‐OSN group showed no significant toxic reactions (Figure 4e; Figure S7, Supporting Information). These results showed that PLCL‐p‐OSN exhibits effective anti‐biofouling control at both the in vitro and in vivo levels and does not cause noticeable toxic reactions in vital organs.

3. Conclusion

The concepts, materials, and fabrication approach reported here propose a robust surface engineering strategy to develop soft, superhydrophobic, and anti‐biofouling encapsulants for stretchable, bioresorbable electronics. The hierarchical structure, constructed by integrating micropatterned biodegradable elastomers with self‐assembled OSN networks, significantly enhanced water repellency while preserving mechanical properties. This dual‐scale architecture reduced the effective contact area and fluid permeability, thereby enabling substantial improvements in encapsulation performance across various biodegradable polymers. Integration of the OSN‐treated films into a stretchable, transient optoelectronic device successfully maintained device functionality over extended periods in aqueous environments. Systematic evaluations under physiological conditions, including cell adhesion, tissue integration, and fibrotic response, validated the anti‐biofouling efficacy and biosafety of the proposed encapsulant. These findings suggest the promise of this hierarchical surface modification as a universal platform for enhancing the functional lifespan and immune tolerance of transient biomedical implants.

4. Experimental Section

Synthesis of Soft, Stretchable, Biodegradable, Anti‐Biofouling Encapsulants

The synthesis of poly(L‐lactide‐co‐ε‐caprolactone) (PLCL, Mn ≈160k) followed a previously established protocol.^{[ 9 ]} PLCL was dissolved in dimethylformamide (DMF) to a concentration of 15% (w/v) and stirred overnight to ensure complete dissolution. The resulting solution was then cast into polydimethylsiloxane (PDMS, 10:1) molds that were prefabricated on silicon (Si) master molds featuring micro‐patterns of holes or pillars (diameter, 5 µm; aspect ratio, 2; pitch, 2.5 µm), which were produced using deep reactive ion etching (DRIE). After casting, the polymer solution was dried on a hotplate at 80 °C for 24 h to remove solvent, yielding flexible films with ≈300 µm in thickness. These films were gently peeled from the molds for further processing. To create surface nanostructures, the films were initially immersed in a 30 w/v% aqueous solution of ammonium persulfate (APS) at 80 °C for 10 h, followed by thoroughly rinsing with deionized (DI) water and drying at 80 °C for 1 h. Then, the hydroxylated films were soaked in 0.05 m methyltrichlorosilane (MTCS) dissolved in petroleum ether under ambient conditions (25 °C, ≈60% humidity) for 24 h. Afterward, the samples were sequentially washed with petroleum ether, ethanol, and deionized water, and finally dried at 120 °C for 10 min. Fourier‐transform infrared (FT‐IR; Cary 630, Agilent, USA) spectroscopy was used to analyze changes in chemical structure at each stage. The same process was applied to other polymer matrices, including poly(caprolactone) (PCL, Sigma‐Aldrich, USA), poly(lactide) (PLA, Sigma‐Aldrich, USA), polyurethane (PU),^{[ 9 ]} and PLCL composite containing 20 v/v% silicon dioxide (SiO₂) nanoparticles to evaluate the versatility of the fabrication strategy.

Characterization of Morphological, Physical, and Mechanical Properties

The surface structures of the prepared films were analyzed using scanning electron microscopy (SEM; S‐4700, Hitachi Hi‐Tech, Japan). To assess surface property, water contact angle (WCA) measurements were performed by dispensing ≈6 µL of DI water onto the film surface under ambient conditions, using a contact angle analyzer (Phoenix‐MT(M), SEO, South Korea). For mechanical testing, the prepared films were cut into a dumbbell‐shaped geometry following ASTM D638 standards. These specimens were then subjected to tensile testing using a universal testing machine (Instron 5900 series, Instron, USA) operated at a constant elongation rate of 6 mm min⁻¹.

Assessment of Encapsulation Performance

To evaluate the protective capability of encapsulant layers, a 300 nm‐thick magnesium (Mg) layer was deposited on a glass substrate using an electron beam evaporator (VER5004, South Korea) and patterned via photolithography to form a resistor configuration. Various types of films were applied over the Mg resistor, and a PDMS chamber was adhered to the substrate. This chamber was then filled with 0.1 m phosphate‐buffered saline (PBS, pH 7, Sigma‐Aldrich, USA). The operational stability of the Mg resistor was monitored by measuring its resistance over time using a source meter (Keithley 2636b, Tektronix, USA) at a data acquisition rate of 1 kHz. The functional lifetime of each encapsulant was defined as the duration until the resistance began to increase.

Fabrication of a Transient Infrared (IR) Micro‐Light‐Emitting‐Diode (µ‐LED) Device

A temporary substrate was prepared by sequential spin‐coating of poly(methyl methacrylate) (PMMA, MicroChem, USA; ≈100 nm thick) and polyimide (PI, Sigma‐Aldrich, USA; ≈1.2 µm thick) onto a p‐type silicon wafer (Silicon Technology Co., Japan). A 300 nm‐thick Mg layer was deposited onto this substrate using electron beam evaporation and subsequently patterned via photolithography to form the electrode. A diluted polyimide (D‐PI) layer (≈400 nm thick) was spin‐coated as a temporary insulating top layer. Electrode patterning was completed using reactive ion etching (RIE; JVAC, South Korea). The PMMA layer was dissolved in acetone to release the entire device from the wafer. Removal of the bottom PI layer via RIE allowed the device to be transferred onto a PLCL‐h‐OSN film (thickness, ≈500 µm) as substrate. Then, the D‐PI insulating layer was also etched away using RIE. A µ‐LED (0402 surface‐mount package) was then bonded to the Mg electrode using silver conductive epoxy, and a wire was affixed to the contact pads. The device was then encapsulated with a PLCL‐p‐OSN film (thickness, ≈500 µm), which was chemically bonded via UV/ozone surface activation (PSDP‐UVT, Novascan, USA).

Performance Evaluation of IR µ‐LED Device in Aqueous Environment

The device was immersed in PBS (pH 7) at 37 °C and operated under a constant 4 V supply. To monitor functional degradation over time, infrared thermographic imaging was performed at an interval of 8 h, measuring peak emission intensity at 850 nm.

In Vitro Anti‐Biofouling Assay

To verify the in vitro anti‐biofouling effect, NIH‐3T3, a mouse fibroblast cell line was used. Experimental materials, PLCL, PLCL‐p, and PLCL‐p‐OSN were attached to a 60 mm cell culture dish in 2 cm x 2 cm sizes (n = 4 per each material). Subsequently, 2.5 x 10⁵ NIH‐3T3 cells were seeded on each material. Cells were stained using a live/dead staining kit for visualization and biocompatibility assessment. Images were taken on days 1, 2, and 3 using an inverted fluorescence microscope. Live cells were stained with calcein‐AM and observed as green, while dead cells were stained with ethidium homodimer‐1 and observed as red. The acquired images were counted and quantitatively analyzed using Image J software.

In Vivo Assay

A subcutaneous mouse model was used to analyze the antibacterial adhesion effect of the experimental material in vivo. All animal experiments were conducted in accordance with the IACUC guidelines of Korea University (Approval Number: KUIACUC‐2024‐0087). Twenty‐four 7‐week‐old C57Bl/6 male mice were classified as follows: three mice per group, implanted with PDMS, PLCL, PLCL‐p, and PLCL‐p‐OSN at 1 week and 3 weeks, respectively. On the day of surgery, each animal was administered with antibiotics and analgesics to prevent infection and control pain. For subcutaneous implantation, the animals were anesthetized with 2% isoflurane inhalation anesthesia (2 L mi⁻¹n2 oxygen). Prior to implantation, the dorsal skin hair was removed, and the surgical area was disinfected with 70% ethanol and povidone to prevent infection. Following this, a 1–1.5 cm skin incision was made, and a subcutaneous pocket was created using blunt separation with Metzenbaum scissors. Experimental materials measuring 1 cm × 1 cm were implanted into the subcutaneous pocket and sutured using surgical clips and non‐absorbable sutures. One week and three weeks later, the animals were euthanized using CO₂, and the skin and vital organs (heart, lung, liver, kidney, spleen) were collected and fixed in 10% neutralized buffered formalin.

Histopathological Analysis

The organs (skin, vital organs) fixed with fixative were processed using the general paraffin embedding technique. Briefly, the tissues underwent dehydration, clearing, and paraffin infiltration, followed by paraffin block preparation. Each paraffin block was cut into 4–5 µm thick sections, which were then subjected to hematoxylin & eosin staining (HE) and Masson Trichrome staining (MT). The stained slides were examined using an inverted optical microscope, and quantitative analysis (fibrous thickness) was performed using Image J software.

Statistical Analysis

Statistical analyses were performed using at least three independent samples unless otherwise noted. Data were presented as mean ± standard deviation (SD), with N indicating the number of experimental replicates. One‐way ANOVA followed by Tukey's Honestly Significant Difference (HSD) post‐hoc test was applied to determine statistical significance. Error bars represent SD, and results were derived from a minimum of three independent experiments. Differences were considered statistically significant at p <0.05 (p <0.05, #p <0.01, **p <0.005, ***p <0.0005).

Conflict of Interest

The authors declare no conflict of interest.

Supporting information

Supporting Information

Han W. B., Han S., Ko G.‐J., et al. “A Hierarchically Structured, Stretchable, Anti‐Biofouling Encapsulation for Biodegradable Electronics.” Adv. Healthcare Mater. 15, no. 7 (2026): e03622. 10.1002/adhm.202503622

Data Availability Statement

The data that support the findings of this study are available from the corresponding author upon reasonable request.