Introduction

Double emulsions (DEs) are complex multiphase systems in which one liquid phase is dispersed within another immiscible liquid phase1,2. The most commonly studied configurations include water-in-oil-in-water (W/O/W) and oil-in-water-in-oil (O/W/O) emulsions, both of which enable the simultaneous encapsulation of hydrophilic and hydrophobic compounds3,4. This structural versatility makes DEs highly attractive for pharmaceutical5,6, cosmetic7,8, and food applications6,9,10. A variety of methods have been developed for DE generation, including spontaneous formation11, two-step emulsification12, and microfluidic emulsification6,13,14. Among these, microfluidic techniques provide superior control over emulsion size and composition, enabling the formation of monodisperse DEs13,14,15,16. The ability to precisely tune droplet properties has positioned DEs as a promising platform for drug delivery. However, their inherent instability presents significant challenges for practical applications17,18. Also, the structural fragility of DEs leads to rapid coalescence, Ostwald ripening, and phase separation, which severely limit their long-term storage stability19. As a result, conventional DEs are unsuitable for applications requiring extended shelf life or controlled drug release. Moreover, encapsulation efficiency remains a critical concern, particularly for hydrophilic drugs, as they tend to diffuse into the external aqueous phase over time, leading to unintended drug loss and reduced therapeutic effectiveness20.

Double emulsion capsules (DECs) are advanced drug delivery carriers derived from DEs, designed to enhance the stability and controlled release of encapsulated substances12,21,22,23. These capsules incorporate a protective shell, typically composed of polymers, formed through various solidification techniques, including solvent evaporation12,21,24, solvent diffusion22, UV polymerization23,25,26, and ionic cross-linking27. This structural reinforcement effectively mitigates the inherent instability of conventional DEs, preventing phase separation and droplet coalescence16. By encapsulating the emulsion within a semi-permeable or cross-linked polymeric shell, DECs provide improved control over drug release kinetics, enabling prolonged therapeutic effects and targeted drug delivery. However, significant passive leakage through the polymer shell remains a major challenge for long-term drug delivery with DECs. This issue primarily arises from the intrinsic porosity of the polymeric capsule shell, allowing unintended diffusion of encapsulated drugs into the surrounding medium28. Passive leakage is further exacerbated by osmotic pressure differentials, variations in shell thickness, and the molecular size-dependent permeability of the polymer network29,30. To address this challenge, polymer blending strategies have been explored, incorporating hydrophobic and hydrophilic polymers into the shell material to fine-tune permeability and reduce leakage31. Multilayered capsule structures have also been investigated to reduce leakage by introducing sequential polymer coatings with distinct diffusion barriers32,33. However, these methods often require intricate microfluidic device fabrication or multiple processing steps, limiting their scalability and practical implementation in drug delivery applications.

Polydopamine (PDA), which mimics the adhesive proteins found in mussels, exhibits strong adhesion properties due to its catechol and amine functional groups34,35. This robust adhesion arises from multiple interaction mechanisms, such as hydrogen bonding, π–π interactions, and covalent bonding35. As a result, PDA can be easily deposited onto various substrates, including metals34,36,37, polymers34,38,39,40, and ceramics34,41,42, making it a versatile coating material. Beyond its adhesion properties, PDA is also well known for its exceptional biocompatibility, chemical stability, and functional reactivity. These characteristics have led to its widespread exploration in surface modification and biomedical applications43. However, despite its advantages, the application of PDA in drug carrier systems remains limited. The Fu group demonstrated PDA coating on drug-loaded mesoporous silica nanocarriers, but the pore size of the carrier dictated the type and quantity of encapsulated drugs, imposing inherent constraints on loading capacity44. A recent study reported PDA coating on microcapsules, but it primarily focused on improving interfacial adhesion between the capsule and surrounding matrices, rather than investigating its effects on capsule leakage and controlled drug release45. Although a few other PDA-coated drug delivery systems have been introduced, many of them still face significant limitations, such as high passive leakage, low drug loading capacity (typically less than 10−7 mg), and dependence on gradual drug release through polymer dissolution in aqueous environments, rather than enabling on-demand release (Supplementary Table S1).

Here, we present the first PDA-coated DECs, fabricated through three sequential processes: (1) DE generation using microfluidics, (2) encapsulation by solvent evaporation, and (3) one-pot PDA coating (Fig. 1a). A custom-built sequential co-flow capillary microfluidic system is employed to generate monodisperse DEs with precise control over droplet size and composition (Fig. 1a and Supplementary Fig. S1). This system consists of a series of coaxially aligned capillaries, allowing for the controlled formation of W/O/W DEs through hydrodynamic flow focusing. The inner aqueous phase, containing fluorescein isothiocyanate–dextran (FITC-dextran) as a drug model, is sheared into droplets by an immiscible polymer-containing middle phase, which is then emulsified within the outer aqueous phase (Fig. 1b). Polycaprolactone (PCL) was selected as the polymer shell material due to its excellent biocompatibility, slow degradation rate, and controlled permeability, making it an ideal candidate for long-term drug release applications46. The resulting DEs are stored in a collecting solution identical to the outer phase. After DE generation, the collecting solution containing DEs is exposed to air overnight at room temperature to allow for solvent (dichloromethane, DCM) evaporation from the middle phase. During this process, the polymer shell solidifies, ensuring the formation of stable DECs (Fig. 1c,d). Cross-sectional imaging of DEC reveals a well-defined core–shell structure, confirming the successful encapsulation of the aqueous drug (Fig. 1e). A conformal PDA coating is then applied to the DEC surface through the one-pot coating method, significantly enhancing capsule stability and reducing passive drug leakage (Fig. 1f). In addition to improving stability, PDA also offers an additional function: programmable light-triggered drug release (Fig. 1g and Supplementary Video S1). As a well-known photothermal agent, PDA efficiently absorbs near-infrared (NIR) light and converts it into localized heat through thermal vibrations47,48. Upon NIR laser irradiation, the PDA-coated DEC experiences a localized temperature elevation, which disrupts the polymer shell integrity and facilitates controlled drug release. To optimize PDA-coated DECs, we systematically investigated the effects of microfluidic parameter, specifically the inner phase flow rate, and polymer concentration on DE size and stability. Furthermore, the influences of PDA coating condition (coating time) and dopamine hydrochloride (DA) concentration on passive leakage were thoroughly examined to reduce passive leakage. Based on these studies, we successfully demonstrated an on-demand drug release system with prolonged stability and reduced passive leakage, triggered by NIR laser irradiation.

Fig. 1
figure 1

Fabrication of PDA-coated DECs for on-demand drug release. (a) Schematic illustration of the fabrication and rupture processes for PDA-coated DECs, including microfluidic DE generation, solvent evaporation, one-pot PDA coating, and DEC rupture through laser irradiation. (b) DEs generated using a microfluidic system. (c) A DEC formed through the solvent evaporation process. (d) SEM image of an intact DEC, showing its surface morphology. (e) Cross-sectional SEM image of a DEC, revealing the core–shell structure. (f) PDA-coated DEC created by the one-pot PDA coating method. (g) On-demand drug release through NIR laser irradiation.

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Results and discussion

Double emulsion (DE) generation using microfluidics

The passive leakage of DECs is closely related to the diffusion rate of the drug within the polymer shell. Increasing the polymer shell density can effectively lower the diffusion rate, thereby reducing passive leakage49,50. To achieve a denser polymer shell, the polymer concentration in the middle phase was increased from 0.5 to 6.0%, and the effect of polymer concentration on DE formation was analyzed using an optical microscope (Fig. 2a and Supplementary Fig. S2). At polymer concentrations exceeding 7.0%, the viscosity of the middle phase became excessively high, leading to the breakage of the microfluidic chip. To generate DEs, the flow rates of the inner, middle, and outer phases were set to 0.15, 0.2, and 1.0 mL/min, respectively. As shown in Fig. 2a, the inner diameter, outer diameter, and middle phase thickness of the DEs remained consistent across different polymer concentrations (or middle phase viscosities), indicating that polymer shell density can be effectively modulated while maintaining a uniform inner volume and overall DE size. This morphological consistency in DEs was also observed across varying inner phase flow rates (Supplementary Figs. S2 and S3). These findings suggest that polymer concentration can be adjusted to control shell density without significantly altering DE morphology. It is noteworthy that an increase in middle phase viscosity could influence both the inner and outer diameters of the DEs due to the enhanced drag force exerted on the inner phase51. However, under the material and concentration conditions used in this study, the change in drag force was not significant, resulting in minimal alterations to the DE morphology.

Fig. 2
figure 2

Optimization of PCL concentration and DE generation process for high DEC yield. (a) Effects of PCL concentration on the inner and outer diameters and middle phase thickness at a fixed inner phase flow rate of 0.15 mL/min. (b) Influence of inner phase flow rate on the inner and outer diameters and middle phase thickness at a PCL concentration of 6%. (c) Yield map of DEC production under different PCL concentrations and inner phase flow rates. The purple, green, and orange regions in the map represent unstable conditions, low-yield conditions (< 40%), and high-yield conditions (> 40%), respectively. (d) DEC yields under high-yield conditions. The hatched bar indicates the optimized conditions yielding the highest DEC production.

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Controlling the drug load within a single DE enables the fabrication of DECs with varying drug dosages, thereby expanding their potential for more diverse and efficient drug delivery applications. In a microfluidic system, increasing the inner phase flow rate allows for a greater volume of inner phase material to be encapsulated within each DE, enhancing the drug-loading capacity of the resulting DECs. To investigate the effect of inner phase flow rate on DE morphology, DEs were generated by adjusting the inner phase flow rate from 0.05 to 0.30 mL/min, while maintaining constant flow rates of 0.2 mL/min for the middle phase (6% PCL) and 1.0 mL/min for the outer phase (Fig. 2b). As expected, DEs produced at higher inner phase flow rates exhibited larger inner diameters, indicating an increase in drug-loading capacity. This trend was consistently observed across different PCL concentrations (Supplementary Figs. S2 and S4). Furthermore, in agreement with a previously reported study51, the outer diameter of the DEs remained nearly unchanged despite variations in the inner phase flow rate. These findings suggest that the inner phase flow rate can be adjusted to precisely regulate drug loading without changing the overall size of DEs. This capability provides fine control over both DE morphology and drug dosage, which is essential for reproducible and efficient drug delivery applications.

Encapsulation by solvent evaporation

The stability of DEs plays a crucial role in achieving a high DEC yield, as DEs have to remain intact throughout the evaporation process to ensure successful capsule formation. Any instability during evaporation leads to droplet rupture, resulting in incomplete or defective DECs. Since a high DEC yield is essential for large-scale production and cost-effective manufacturing, we identified the optimal conditions that produced the highest DEC yield based on DEs generated in previous studies. All DEs were converted into DECs through solvent evaporation, and their yield was quantified using an image analysis method (see details in “Materials and methods” section). As shown in Fig. 2c, a significant portion of DEs ruptured during the evaporation process, resulting in a low DEC yield of less than 40%. Only DEs, prepared using a middle phase containing 5% and 6% PCL and an inner phase flow rate of 0.15 and 0.30 mL/min, exhibited relatively high DEC yields, exceeding 40%. Among these, DEs generated with a 6% PCL concentration and an inner phase flow rate of 0.15 mL/min demonstrated the highest DEC yield, reaching approximately 85.2% (Fig. 2d). As these conditions allow for the formation of a dense polymer shell due to the high PCL concentration while also achieving the highest DEC yield, they are particularly well-suited for the on-demand drug release system with low passive leakage. Therefore, these optimized conditions were used for DEC fabrication in all subsequent studies.

One-pot polydopamine (PDA) coating

DECs were coated with PDA using a simple one-pot coating method, in which the DECs were immersed in a PDA coating solution, consisting of DA and Tris–HCl buffer, and stirred at room temperature (Fig. 3a). Upon PDA coating, an additional PDA shell (or layer) was successfully deposited onto the polymeric DEC shell, effectively increasing the diffusion path length for encapsulated aqueous drugs. This structural modification played a significant role in reducing passive leakage, as the increased diffusion barrier hindered drug permeation through the polymer shell52. Moreover, the PDA coating covered the surface of pores in the polymer shell, thereby reducing their effective size and further limiting passive drug leakage. However, while PDA introduces a physical barrier against passive leakage, its hydrophilic functional groups, including amine and catechol moieties, modify the pore surfaces, rendering them hydrophilic53,54. This increased hydrophilicity enhances the diffusion rate of aqueous inner phase materials, potentially counteracting the reduction in leakage. Therefore, careful optimization of the PDA coating process is necessary to maximize passive leakage reduction.

Fig. 3
figure 3

PDA coating on DECs using the one-pot coating method. (a) Schematic illustration of the PDA coating process on DECs. The inset schematic shows the deposition of PDA on the DEC shell, increasing the diffusion path length and modifying the pore surfaces. (b) Passive leakage study of DECs coated for different coating times at a fixed DA concentration of 2 mg/mL. (c) Passive leakage study of DECs coated with varying DA concentrations at a constant coating time of 2 h. (d) Overall shell thickness of PDA-coated DECs, comprising both the polymer shell and the PDA layer, increases with higher DA concentrations.

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In the PDA coating process, the thickness of the PDA layer on the DEC surface can be controlled by modifying the chemical composition of the coating solution or adjusting the coating duration (or coating time). Generally, higher DA concentrations in the coating solution result in thicker PDA layers, while extending the coating time also increases the coating thickness55,56. To investigate the relationship between coating duration and passive leakage, DECs were coated using a PDA coating solution containing 2 mg/mL of DA, with coating times set at 1, 2, or 3 h. A total of 20 coated DECs from each condition were placed in deionized (DI) water for passive leakage measurement. Since FITC-dextran, the drug model used in this study, is a fluorescent dye, passive leakage was quantitatively analyzed by measuring the fluorescence intensity of the surrounding solution using a microplate reader over time. The FITC-dextran concentration was determined using a calibration curve (Supplementary Fig. S5), which correlates the FITC-dextran concentration with fluorescence intensity. Passive leakage was then calculated as the ratio of the FITC-dextran concentration released over time to the total concentration measured after rupturing all DECs at the end of the test. Passive leakage of coated DECs was monitored over an 8-day period, with uncoated DECs used as a control. As shown in Fig. 3b, DECs coated for 1 h exhibited higher passive leakage compared to uncoated DECs, suggesting that a shorter coating time was insufficient to form a thick outer PDA layer, resulting in minimal changes to the diffusion length. Moreover, due to the hydrophilic nature of PDA, the coated pores facilitated increased diffusion of the aqueous inner phase, leading to enhanced passive leakage. When the coating duration was extended to 2 h, a sufficient PDA layer was deposited on the DEC surface, effectively increasing the diffusion path length. At this stage, the effect of increased coating thickness outweighed the hydrophilic influence of PDA, leading to an overall reduction in passive leakage. Notably, DECs coated for 3 h showed a passive leakage trend similar to that of the 2-h coated DECs, suggesting that all available DA in the solution had been consumed within 2 h, and further coating did not contribute to additional layer formation. Therefore, 2 h was determined to be the optimal coating time for a coating solution containing 2 mg/mL of DA as it resulted in sufficient PDA deposition to effectively reduce passive leakage while preventing unnecessary processing time.

In order to investigate the effect of DA concentration in the coating solution on passive leakage, DECs were coated for 2 h using PDA solutions containing 0.2, 2, and 20 mg/mL of DA, and their passive leakage was measured over time (Fig. 3c). The results show that as the DA concentration increased up to 2 mg/mL, passive leakage decreased. This trend can be attributed to the formation of a thicker PDA layer on the DEC surface, which increases the diffusion path length, thereby reducing drug loss. The thickness of the DEC shell was analyzed using SEM images, revealing that uncoated DECs had a shell thickness of 3.74 µm, while DECs coated with 0.2, 2, and 20 mg/mL of DA exhibited shell thicknesses of 4.00, 4.61, and 5.14 µm, respectively (Fig. 3d and Supplementary Fig. S6). Despite the increased shell thickness in DECs coated with 20 mg/mL DA, these capsules exhibited higher passive leakage compared to those coated with 2 mg/mL DA. This paradoxical result suggests that while an excessively thick PDA layer increases diffusion length, it also enhances the hydrophilicity of the DEC shell and pore surfaces, thereby facilitating diffusion of the aqueous inner phase material and leading to greater passive leakage. The effect of DA concentration on shell porosity was also analyzed using BET measurements (Supplementary Fig. S7). All PDA-coated PCL films exhibited very low pore volumes, with small variation across different DA concentrations. These results indicate that DA concentration has little influence on shell porosity, further supporting the conclusion that variations in passive leakage from PDA-coated DECs are more likely governed by the complex interplay between PDA layer thickness and surface hydrophilicity rather than by porosity. To determine the optimal coating conditions with minimal passive leakage, various combinations of DA concentration and coating duration were systematically tested (Supplementary Fig. S8). The results indicated that DECs coated with a 2 mg/mL of DA solution for 2 h showed the lowest passive leakage (56% at Day 8), approximately 20% lower than that of uncoated DECs (74.0% at Day 8). While the absolute leakage may still be considerable, these DECs demonstrated a meaningful improvement in leakage suppression relative to previously reported drug delivery systems (Supplementary Table S1). Accordingly, these optimized coating conditions were applied for PDA coating in all subsequent studies.

Characterization of PDA-coated polymer surfaces

The addition of functional groups on PCL surface after PDA coating was studied using FTIR spectroscopy. Rectangular PCL film samples (25 mm × 23 mm × 0.2 mm) were prepared using a glass mold, followed by solvent evaporation and cutting (see details in “Materials and methods” section). PDA-coated film samples were then prepared using the one-pot coating method under optimized conditions. As shown in Fig. 4a, two new peaks appeared in the FTIR spectrum of the PDA-coated sample (red line in Fig. 4a), compared to the uncoated PCL sample (black line in Fig. 4a). These peaks at 3450 and 1650 cm−1 correspond to the stretching vibration of hydroxyl (–OH) and N–H functional groups, respectively. The appearance of –OH and N–H peaks came from hydrophilic groups (phenolic hydroxyl and amino groups) in PDA, indicating the successful coating of PDA onto the PCL surface.

Fig. 4
figure 4

Characterization of PCL surfaces coated with PDA under optimized conditions. (a) FT-IR spectra of uncoated (black line) and PDA-coated PCL (red line) films. The PDA-coated film exhibits two additional peaks corresponding to hydroxyl (–OH) and N–H functional groups, representing the presence of PDA. (b) Contact angle measurements over time for an uncoated PCL film (black square) and PCL films coated with 0.2 (red circle), 2 (blue triangle), and 20 mg/mL (green diamond) of DA. Films coated with higher DA concentrations exhibited lower contact angles, indicating that PDA coating enhances the surface hydrophilicity of the films. (c) UV–Vis absorption spectra of uncoated (black line) and PDA-coated PCL (2 mg/mL, red line) films. The PDA-coated film shows higher absorbance across the entire spectral range, confirming the deposition of light-absorbing PDA.

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The presence of hydrophilic functional groups in PDA increases the hydrophilicity of the PCL surface following PDA coating. To evaluate the hydrophilicity of both coated and uncoated PCL surfaces, we measured the water contact angles on each surface over time (Fig. 4b and Supplementary Fig. S9). Rectangular PCL film samples were prepared using the same method as in the FTIR analysis, and a water droplet was deposited on the film surfaces. For all film surfaces, the contact angle exhibited a decreasing trend over time. However, the PCL sample coated with higher concentrations of DA consistently showed lower contact angles, confirming an increase in surface hydrophilicity due to PDA deposition. These results further validate the successful deposition of PDA onto the PCL surface.

PDA is one of the well-known photothermal agents and exhibits high photothermal conversion (> 40%) through strong NIR absorption57. Accordingly, PDA coatings have been extensively employed in photothermal therapy (PTT) by modifying the surfaces of various particles and fibers, enabling their use in cancer treatment58,59,60 and targeted hyperthermia61. To evaluate the photothermal effect of PDA coating, both PDA-coated and uncoated PCL disk samples were prepared in a well plate (see details in “Materials and methods” section), and their absorption spectra were obtained using a microplate reader. As shown in Fig. 4c, the PDA-coated sample exhibited higher absorption across the entire spectral range (300–900 nm), including the NIR region. Given that 808 nm NIR laser radiation can penetrate tissue relatively deeply62,63, it has been extensively utilized for photothermal therapy58. Therefore, an 808 nm NIR laser was employed in all subsequent photothermal and rupture tests to investigate the thermal response and controlled release behavior of PDA-coated DECs.

On-demand drug release by NIR laser irradiation

Since our objective is to achieve on-demand drug release by rupturing PDA-coated DECs using NIR laser irradiation, it is important to understand their photothermal conversion capability. To evaluate this, a custom-built laser system was employed (Fig. 5a), and the maximum temperatures of the samples were monitored under NIR irradiation using a thermal imaging camera (Fig. 5b and Supplementary Fig. S10). Nine rectangular PDA-coated PCL film samples were prepared, with varying in coating times and DA concentrations. Upon exposure to an NIR laser (1 W) for 5 s, the maximum temperatures of all PDA-coated samples increased, reaching higher temperatures than that of the uncoated sample (29.1 °C). More specifically, as DA concentration increased and coating duration was extended, the maximum temperature also increased proportionally, confirming a direct correlation between PDA coating parameters and photothermal conversion efficiency. To rupture PCL-based DECs, the temperature of the DEC shell has to exceed the melting point of PCL (62.7 °C)64. Under optimized coating conditions (2 mg/mL of DA, 2-h coating time), the maximum temperature of the coated PCL sample reached 88.7 °C, well above the threshold for PCL shell rupture. This result confirms that the optimized PDA coating effectively enables DEC rupture. To further validate this, a rectangular PCL film sample was coated under optimized conditions, and its rupture upon 808 nm NIR laser irradiation was successfully demonstrated (Fig. 5c).

Fig. 5
figure 5

On-demand drug release from PDA-coated DECs by NIR laser irradiation. (a) A custom-built laser system consists of a linear stage, a focusing lens, an NIR laser (808 nm), a thermal imaging camera, and a switchable sample holder. (b) Maximum temperatures of PDA-coated PCL films reached under 5 s of laser irradiation. Red and black dashed lines indicate the melting temperature (Tm) of PCL (62.7 °C) and the maximum temperature of the uncoated film (28.2 °C), respectively. (c) Surface morphology of films after 5 s of laser irradiation. The laser generated large holes on the PDA-coated film, while no visible change was observed on the uncoated film. (d) On-demand drug release using a PDA-coated DEC embedded within a soft tissue model (agarose gel). Three stages of DEC rupture: (i) intact DEC, (ii) DEC rupture by laser irradiation, and (iii) diffusion-based drug release from the ruptured DEC.

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Based on our findings, we successfully demonstrated on-demand drug (FITC-dextran) release from PDA-coated DECs using NIR laser irradiation. First, a PDA-coated DEC, immersed in DI water, was successfully ruptured after 3 min of NIR laser exposure (Fig. 1g and Supplementary Video S1). Although the DEC was surrounded by water, which has a high specific heat capacity, the photothermal conversion was sufficient to raise the local temperature above the melting point of PCL, enabling DEC rupture. To further investigate the feasibility of drug release in a tissue-mimicking environment, an agarose gel (0.5%) was prepared as a soft tissue model, and a DEC was embedded within the gel matrix (Fig. 5d(i)). Upon NIR laser irradiation, drug release was monitored using a digital camera. As FITC-dextran is a fluorescent dye, drug release was quantitatively analyzed by measuring the fluorescence intensity of the DEC over time (Supplementary Fig. S11). Blue light was used to enhance visualization of the released drug. As shown in Fig. 5d(ii), fluorescence began to be released from the ruptured portion of the DEC shell after 2 min of irradiation, confirming the initiation of drug release. The release process then occurred solely through diffusion, with more than 80% of the encapsulated drug diffusing out of the capsule within the first hour, and nearly complete release achieved within 6 h, as indicated by the disappearance of fluorescence within the DEC (Fig. 5d(iii)). These results clearly demonstrate the versatility of our PDA-coated DEC as an on-demand drug delivery platform, capable of functioning effectively in various surrounding environments, including liquid phases and hydrogels.

Conclusion

We presented PDA-coated DECs, designed to enhance long-term drug retention and enable on-demand release by NIR laser irradiation. These DECs were successfully fabricated through three sequential processes: DE generation using microfluidics, encapsulation by solvent evaporation, and one-pot PDA coating. By systematically optimizing polymer concentration and inner phase flow rate in microfluidics, we achieved an 85.2% DEC yield, while ensuring the formation of a dense PCL shell for improved structural integrity and reduced passive leakage. Additionally, by fine-tuning DA concentration and coating time, we reduced passive leakage over 8 days by ~ 20%, from 74% in uncoated DECs to 56% in PDA-coated DECs, demonstrating the effectiveness of PDA coating as a diffusion barrier. Furthermore, PDA coating enabled programmable, light-triggered drug release through localized heating under NIR laser exposure. Notably, the high photothermal conversion efficiency of PDA facilitated DEC rupture even in water and hydrogels, despite their high specific heat capacities, demonstrating its feasibility in diverse biological environments. These findings establish PDA-coated DECs as a versatile, dual-function platform, combining long-term drug retention with light-responsive drug release. Moreover, this capability could further expand its potential for a wide range of biomedical applications, including cancer and gene therapy, orthopedic and intraocular implants, and diabetes management, by enabling control over the rupture temperature through the use of polymers with varying melting points, and by reducing the risk of thermal damage through minimization of the laser spot size.

Materials and methods

Materials

The inner phase solution was prepared using DI water as a solvent, 1% (w/v) poly(vinyl alcohol) (363170, Sigma-Aldrich) as a surfactant, 3% (w/v) glycerol (G7757, Sigma-Aldrich) as a humectant, and 2.5 mg/mL (w/v) fluorescein isothiocyanate–dextran (FD40S, FITC-dextran, Sigma-Aldrich) as a fluorescent dye and drug model. The outer phase and collecting solutions had the same composition as the inner phase, excluding FITC-dextran. The middle phase solution consisted of polycaprolactone (PCL, 440744, Sigma-Aldrich) dissolved in dichloromethane (DCM, D65100, Sigma-Aldrich). A 10 mM Tris–HCl buffer solution (pH 8.5), used for PDA coating, was prepared using Tris base (252859, Sigma-Aldrich) and hydrochloric acid (HCl, H1026, SAMCHUN). Dopamine hydrochloride (DA, H8502) and agarose (A9539) were purchased from Sigma-Aldrich. All materials were used as received without further purification.

Custom-built microfluidic chip

DEs were generated using a custom-built sequential co-flow capillary microfluidic chip (Supplementary Fig. S1a). To fabricate the microfluidic chip, a stainless-steel needle tip (30 GA, 7018178, Nordson), bent at a 90° angle, was inserted near the end of a polyvinyl chloride (PVC) tubing (4 cm pre-cut, ACF00001, Tygon), allowing the needle tip to extend outside the PVC tubing. A highly hydrophobic polytetrafluoroethylene (PTFE) tubing (1.5 cm pre-cut, AWG28TW, Tef Cap) was then partially inserted (approximately one-third of its length) between the needle tip and the PVC tubing so that the end of needle tip was positioned at the middle of PTFE tubing. The PTFE tubing was then inserted through a side hole, created near the end of a polyethylene (PE) tubing (5 cm pre-cut, BD 427440, Intramedic), ensuring the PTFE tubing to extended beyond the PE tubing. Finally, a hydrophilic square glass capillary tube (8100, Vitrocom) was partially inserted between the PTFE tubing and the PE tubing, enclosing the PTFE tubing within the glass tubing. All junctions were completely sealed with epoxy glue (1365868, Loctite) to prevent fluid leakage.

Double emulsion (DE) generation

The microfluidic chip was connected to three syringes, delivering the inner, middle, and outer phases through the needle tip, PVC tubing, and PE tubing, respectively (Supplementary Fig. S1b). The flow rates of each phase were precisely controlled using syringe pumps (NE-300, New Era Pump Systems). In this study, DEs were generated by maintaining the middle (0.2 mL/min) and outer (1.0 mL/min) phase flow rates, while varying the inner phase flow rate from 0.05 to 0.30 mL/min. The generated DEs were stored in collecting solution, and their morphology (inner and outer diameters) was analyzed using an optical microscope (Supplementary Fig. S2) and image analysis software, Image J.

Double emulsion capsule (DEC) yield calculation

DEC yield was calculated using an image-based analysis method. An overall image of the generated DEs, stored in the collecting solution, was captured using a digital camera (EOS 200D II, Canon) under blue light to enhance the visibility of FITC-dextran fluorescence. The DEs were then converted into DECs through solvent evaporation, and an additional image was taken to facilitate direct comparison of their numbers before and after the evaporation process. Image J software was used to quantify the number of DEs and DECs in the acquired images. Fluorescence signals were isolated by applying an intensity threshold, and the Watershed function was employed to separate clustered DEs and DECs. Finally, the number of DEs and DECs was counted using the Analyze Particle function (size: 2000 ~ infinity, circularity: 0.20 ~ 1.00), and the yield was determined by calculating the ratio of DEs to DECs.

One-pot polydopamine (PDA) coating

DECs and test samples were coated with PDA using a one-pot coating method35. All samples were thoroughly rinsed three times with DI water to remove any residual impurities. The cleaned samples were then immersed in a coating solution consisting of dopamine hydrochloride (DA, 0.2, 2, or 20 mg/mL) and Tris–HCl buffer (50 mL). The solution was stirred at 450 rpm at room temperature for the designated coating time (1, 2, or 3 h). After coating, the samples were rinsed three times with DI water to eliminate excess PDA residues. The coated samples were then individually stored in DI water at room temperature. All samples were used within 3 days of preparation.

Passive leakage study

Since FITC-dextran, encapsulated in the DECs, is a fluorescent dye, passive leakage was estimated by measuring the fluorescence intensity of the surrounding solution over time. To investigate passive leakage for 8 days, 20 DECs were placed in DI water, with the total mass adjusted to 2 g. The leakage of FITC-dextran was quantified by periodically collecting the surrounding solution and measuring its fluorescence intensity using a microplate reader (SpectraMax Mini, Molecular Devices). The fluorophore concentration in the solution was then determined using a calibration curve, which correlates fluorophore concentration with fluorescence intensity (Supplementary Fig. S5). To establish the calibration curve, seven solutions with predetermined fluorophore concentrations (0, 500, 1000, 1500, 2000, 2500 and 3000 mg/L) were prepared, and their intensities were measured using the plate reader. After obtaining all results over the 8-day period, all DECs were ruptured to release the remaining FITC-dextran, and the concentration was measured to determine the total concentration. Passive leakage was then calculated as the ratio of the FITC-dextran concentration released over time to the total concentration.

Shell thickness measurement

To measure the shell thickness of DECs and PDA-coated DECs, DEC samples were cut in half and thoroughly rinsed three times with DI water to remove any residual contaminants. The samples were then air-dried overnight to ensure complete dehydration. Cross-sectional images were obtained using a scanning electron microscope (SEM, JSM-IT300, JEOL), and the shell thickness was measured using Image J. Multiple regions of each DEC sample were analyzed to determine the average shell thickness.

Rectangular film sample preparation

A glass mold with a confined casting area (28 mm × 26 mm × 1 mm) was fabricated by placing 1 mm thick spacers between two glass slides. The mold was then filled with a middle phase solution containing 6% PCL and dried overnight in a fume hood. To create PDA-coated samples, the dried PCL samples were coated using the one-pot PDA coating method under optimized coating conditions (2 mg/mL DA solution for 2 h), and subsequently dried in a fume hood for 24 h. The dried samples were then cut into rectangular film specimens (25 mm × 23 mm × 0.2 mm). These rectangular film samples were used for Fourier-transform infrared spectroscopy (FT-IR) analysis, contact angle measurement, and rupture studies.

Fourier-transform infrared spectroscopy (FT-IR) analysis

FT-IR spectra of uncoated and coated rectangular films were obtained using a Fourier-transform infrared spectrometer (Spectrum 3, PerkinElmer) equipped with an attenuated total reflectance (ATR) accessory. Transmittance data were recorded over a wavenumber range of 4000–560 cm−1 with a resolution of 4 cm−1, averaging 32 scans per measurement.

Contact angle measurement

The contact angles of DI water on the surface of rectangular films were measured using a contact angle measurement platform (L2004A1, Ossila). A glass slide was precisely positioned horizontally on the platform to minimize tilting effects. The rectangular film sample was then placed on the glass slide with a 0.5 mL of DI water, applied between them to ensure a flat sample surface. The samples were then air-dried for 24 h to achieve firm attachment to the glass substrate. A 3 µL of water droplet was carefully deposited on the sample surface, and photo images of the droplet were captured at 30-s intervals for 3 min using a digital camera. Contact angles were then measured using Image J software with the contact angle plugin. The average contact angle values were determined from three independent measurements.

Absorption measurement

Absorption spectra of uncoated and coated PCL samples were obtained using a microplate reader (SpectraMax Mini, Molecular Devices). Disk samples were prepared by dispensing 0.15 mL of a middle phase solution containing 6% PCL into a 96-well plate and allowing them to dry in a fume hood for 24 h. To create PDA-coated disk samples, the dried PCL samples were coated using the one-pot PDA coating method under optimized coating conditions (2 mg/mL DA solution for 2 h), and subsequently dried in a fume hood for another 24 h. Absorption spectra were recorded at 1 nm intervals over a wavelength range of 300–900 nm using the microplate reader.

Custom-built laser system

A custom-built laser system was built with the following major components: a linear stage (LTS300/M, Thorlabs), a focusing lens (LA1740, Thorlabs), a NIR laser (1W, 808 nm, LCDRS-SHINE), a thermal imaging camera (D384A, Guide Sensmart), and a switchable (chamber-type and film-type) sample holder (Fig. 5a). The system is controlled by a custom-written LabVIEW (National Instruments) script.

Rupture study

To study the photothermal effect of PDA, the maximum temperatures of uncoated and coated rectangular film samples under NIR laser irradiation were monitored using the custom-built laser system. The rectangular films were placed on the film-type sample holder in the system, and the NIR laser was focused on their surfaces. The thermal imaging camera in the system provides surface temperature of the samples during laser irradiation. The samples reached their maximum temperatures within 5 s of laser exposure.

On-demand drug release

The on-demand drug release of PDA-coated DEC was demonstrated using the custom-built laser system with the chamber-type sample holder. To create an agarose gel as a soft tissue model, 0.5% agarose was dissolved in DI water while simultaneously heated at 200 °C under stirring at 600 rpm for 20 min. A 2 mL of the heated agarose solution was then dispensed into the chamber and partially solidified by cooling at 4 °C for 5 min. The PDA-coated DEC was embedded into the gel, followed by an additional 5-min cooling step in a refrigerator for complete gelation. The chamber was then positioned in the laser system, and the laser was focused on the DEC surface for 2 min to induce DEC rupture. Drug release was monitored using a digital camera, and blue light was applied to enhance visualization.