Driving Drugs Deeper: Force‐Driven Microneedle Systems for Active Therapy and Their Road Towards Clinical Implementation

ABSTRACT

Transdermal delivery of macromolecules and biologics is constrained by the barrier function of skin. First‐generation microneedle arrays address this issue by penetrating the stratum corneum; however, their passive, diffusion‐based release limits delivery depth, kinetics, and adaptability. In this review, we discuss force‐driven microneedles (F‐MNs) as a shift from passive permeation to active propulsion. F‐MNs integrate mechanisms that generate localized physical driving forces to rapidly and controllably drive therapeutic payloads into tissue, achieving millimeter‐scale penetration within seconds to minutes. We categorize four principal modalities based on their force‐generation mechanism: gas‐driven microneedles, cavitation‐driven microneedles, magnetically actuated microneedles, and iontophoresis‐enhanced microneedles. For each class, underlying mechanisms, material choices, safety considerations, and translational challenges are presented. We further discuss how artificial intelligence can guide F‐MNs' design, manufacturing and closed‐loop theranostic control. We also highlight the unique fabrication requirements of F‐MNs, including multi‐material integration, spatial compartmentalization of force‐generating components, stimulus‐responsive architectures, and device‐level quality control. By framing F‐MNs around their force‐generation mechanisms and translational barriers, this review outlines design rules and clinical application spaces for deep, rapid, and controllable microneedle‐mediated therapy across transdermal, mucosal, gastrointestinal, wound, and tumor‐related delivery settings.

Force‐driven microneedles are presented as active therapeutic microdevices that overcome the diffusion limits of conventional patches. By integrating gas generation, magnetic actuation, cavitation, or iontophoresis, these systems drive payloads deeper into tissue, enabling faster, more controlled delivery and highlighting key design principles and translational challenges for clinical implementation.f

1. Introduction

For centuries, humans have sought effective ways of administering therapeutics into the body. Ancient civilizations relied on simple routes, including the oral ingestion of herbal potions, inhalation of medicinal smokes, and topical application of salves [1]. Historical records show that drug delivery through the skin began with early topical therapeutic practices, where medicinal formulations were applied to the body to treat disease [2, 3]. Nature further inspired more direct delivery strategies, as venomous snakes and bees demonstrate how bioactive molecules can be introduced beneath the skin for rapid action [4]. These concepts eventually influenced the development of modern injectable systems [5]. However, conventional injections are still associated with pain, needle anxiety, and the need for trained personnel, motivating the development of gentler and more patient‐friendly delivery platforms [5, 6]. Acupuncture offers another ancient, minimally invasive route for influencing physiology through the skin [7]. By inserting fine needles into specific cutaneous points, practitioners modulate neural, vascular, and immune pathways to relieve pain and restore function without introducing exogenous chemicals. Beyond its therapeutic effects, acupuncture demonstrated that targeted penetration of the stratum corneum can safely trigger systemic responses, an insight that presaged modern intradermal approaches such as microneedles and other gentle transdermal technologies [7, 8].

In the late 20th century, the skin once again became a portal for systemic therapy, this time utilizing a minimally invasive approach. Transdermal patches were developed to deliver drugs across the skin without needles, enabling controlled systemic administration through the skin barrier [9, 10]. In 1979, the first transdermal patch delivering scopolamine to prevent motion sickness was approved by the Food and Drug Administration (FDA), demonstrating that an adhesive patch applied to the skin could steadily release a drug into the bloodstream over several days [9]. Over the following decades, transdermal patches became a mainstream option for certain small‐molecule drugs (e.g., nicotine for smoking cessation, hormonal therapy, and analgesics). Yet the skin's protective barrier (i.e., the stratum corneum) permits only small, lipophilic molecules to penetrate effectively; larger biomolecules or hydrophilic drugs are typically blocked. In other words, traditional patches work for a limited number of medications because the skin, while convenient, is highly impermeable to most compounds. These limitations highlighted the need for strategies that could bypass or transiently breach the stratum corneum while preserving minimal invasiveness [10, 11, 12, 13].

To penetrate the skin's protective barrier while maintaining minimal invasiveness, microneedle patches were introduced [13]. A microneedle patch contains an array of dozens or hundreds of small needles, each only a few hundred microns in length, which can painlessly puncture the skin's stratum corneum layer without reaching deep enough to hit nerves. The concept was first proposed in the 1970s (patented in 1976 by Gerstel and Place), although the technology to fabricate these small needles lagged behind the idea. By the late 1990s, advancements in microfabrication turned the concept of microneedle prototypes into a reality [14, 15]. In 1998, it was shown that solid silicon microneedles could successfully pierce the skin and deliver a model drug (a fluorescent compound, calcein) into the tissue. Once the feasibility was shown, the field rapidly expanded [13]. Multiple types of microneedle arrays were developed in the 2000s, for example, hollow microneedles that allow injection of liquid formulations (first patented in 1996, demonstrated by 2000) [16], drug‐coated microneedles that carry dry drug on their surfaces (introduced in 2004) [17], dissolving microneedles made of biodegradable polymers that fully dissolve to release their payload (first reported in 2006) [18, 19], and swelling hydrogel microneedles that absorb interstitial fluid and swell to gradually dispense drugs (invented in 2012) [12, 14]. Through these innovations, microneedle patches emerged as an exciting drug delivery platform that could bypass the stratum corneum barrier by mechanically creating micro‐scale pathways into the skin. Equally important, they promised a painless and easy‐to‐use experience for patients, often described as simple as “applying a Band‐Aid” in contrast to conventional hypodermic injections [11, 12, 15]. The first‐in‐human evaluation of a dissolvable microneedle patch for influenza vaccination, a randomized, partly blinded, placebo‐controlled Phase 1 study conducted by the Emory/Georgia Tech team (Prausnitz laboratory), was reported in The Lancet in 2017 and demonstrated that microneedle‐based delivery was safe, well‐tolerated, immunogenic, and acceptable to participants [20]. Figure 1a depicts the historical development trajectory of microneedle‐based drug delivery.

FIGURE 1.

(a) Historical roadmap of microneedle drug delivery evolution. (b) Conceptual overview of force‐driven microneedles (F‐MNs) for deep drug delivery. Conventional dissolving microneedles (left) are limited to passive diffusion, achieving shallow penetration (less than $600\mu \mathrm{m}$). F‐MNs overcome this barrier using active propulsion: Gas‐expansion microneedles use internal chemical reactions to generate propulsive gases (900–$1200\mu \mathrm{m}$ penetration); magnetic‐actuation microneedles use an external magnetic field to drive magnetic particles loaded with therapeutics deep in tissue (650–$850\mu \mathrm{m}$ penetration); cavitation‐driven microneedles use ultrasound to create shockwaves from collapsing microbubbles (750–$1000\mu \mathrm{m}$ penetration); and an iontophoresis‐integrated microneedle patch in which a gentle direct current propels therapeutics through the micro‐channels, extending delivery far beyond the diffusion‐limited depth achieved by passive microneedles (600–$750\mu \mathrm{m}$ penetration). All images were generated by BioRender.

Microneedle technology is especially appealing for modern biopharmaceuticals, such as mRNA vaccines, peptide biologics, and gene therapies, that are difficult or impossible to deliver orally [21, 22]. Transdermal delivery via microneedle patches offers several advantages for these therapeutics [23, 24, 25, 26]: it (i) bypasses enzymatic degradation in the gastrointestinal tract and first‐pass hepatic metabolism; (ii) places drug directly within the viable epidermis and dermis, tissue layers rich in antigen‐presenting cells and dense lymphatic networks that can improve targeting for vaccines and other immunomodulators; (iii) creates transient micro‐channels that accommodate large, hydrophilic molecules and even nanoparticles that otherwise cannot cross the stratum corneum; and (iv) enables minimally invasive, needle‐free, potentially self‐administered dosing that can improve acceptance, adherence, and opportunities for controlled or prolonged release at the application site. Despite these advantages, first‐generation passive microneedle patches exhibit several important limitations [1, 10, 27, 28]:

• Passive, diffusion‐limited transport: Drug embedded in dissolving or coated microneedles typically diffuses into the surrounding tissue solely down a concentration gradient, without any active driving force.

• Shallow penetration and localization: Payloads often remain confined to the epidermis or upper dermis, with effective penetration depths of only a few hundred micrometers, limiting access to deeper tissue targets or systemic circulation.

• Slower onset and reduced bioavailability: Compared with subcutaneous injection, passive patches can exhibit delayed absorption and lower bioavailability, particularly for large biomolecules that require rapid or extensive distribution.

• Safety–comfort trade‐off with needle length: Simply increasing microneedle length (e.g., $\sim$ 1 mm) risks contacting nociceptors and blood vessels, undermining the goal of a pain‐free, minimally invasive device.

• Dose‐loading constraints: The finite tip and surface volume available in coated or dissolving formats can limit the deliverable dose, especially for high‐dose macromolecules.

• Heterogeneous deposition: Variability in skin properties (e.g., hydration, elasticity) and patch placement can lead to inconsistent tissue exposure and inter‐application variability.

These challenges have driven the development of “active” microneedle delivery systems that add a driving force (e.g., pressure, electrotransport, acoustofluidics) to propel payloads deeper into tissue.

1.1. Force‐Driven Microneedle Concept

Force‐driven microneedles (F‐MNs) shift drug delivery from passive diffusion toward active transport mechanism. In these systems, each microneedle is engineered not only as a static drug depot but also as a microscopic patch that generates localized mechanical forces (such as pressure, convection, and electro‐osmotic flow) to “push” therapeutic agents into deeper skin layers and even into circulation [29, 30, 31]. This active propulsion can deliver payloads on the order of millimeters deep within seconds to minutes, exceeding the typical $\sim$ 500 $\mu \mathrm{m}$ penetration of passive microneedle diffusion. By integrating energy sources or reactive agents into microneedles, F‐MNs overcome the fundamental transport resistance of the skin [32]. The result is faster uptake, greater penetration depth, and a more uniform drug distribution in the tissue compared to passive patches. This approach decouples transdermal delivery from Fickian diffusion kinetics, instead utilizing controlled forces to drive molecules – akin to an “active pump” at the microscale. Such control is especially important for potent biologics that require precise dosing and localization; by modulating force, F‐MNs can potentially tune the amount of drug delivered at specific depths or time points, thereby improving therapeutic indices [29, 33]. Thus, F‐MNs can mitigate the six limitations previously described for passive microneedle patches.

Functionally, F‐MNs can be classified into two broad forms according to the target of the applied force. In the first form, the microneedle structure itself is mechanically driven into tissue, thereby improving insertion, localization, or access to otherwise difficult anatomical sites. This strategy is most evident in magnetically actuated microneedles and microneedle‐robot systems, where external fields guide, orient, vibrate, or push the needle‐bearing structure into tissue [34, 35]. In the second form, the microneedle primarily serves as an entry interface or drug reservoir, while the therapeutic cargo is subsequently driven beyond the needle tip by an additional force such as gas pressure, cavitation‐induced convection, or electrotransport [29, 30, 36]. This distinction clarifies that F‐MNs may act either by actively enhancing tissue insertion or by actively propelling the payload after insertion. Building on this functional classification, the systems reviewed here are further grouped according to their primary mechanism of force generation. It is also important to distinguish whether the applied force functions only as a delivery aid or whether it also contributes directly to therapy. In some F‐MNs, the force mainly enhances insertion, penetration depth, or cargo transport, while the therapeutic effect is provided by the delivered drug. In other systems, the force‐generating component itself has therapeutic activity, creating a dual‐function platform. For example, gases such as ${\mathrm{H}}_2$ and ${\mathrm{O}}_2$ can act not only as propulsive agents but also as bioactive modulators of oxidative stress, inflammation, hypoxia, or tissue repair [29, 37, 38]. Similarly, electrically driven microneedles can enhance drug transport through electrophoresis and electroosmosis, while the applied current may also provide electrotherapeutic effects such as antibacterial activity, biofilm disruption, or stimulation of tissue repair [30, 39, 40]. Therefore, F‐MNs can be further classified into force‐assisted delivery systems and dual‐function force‐therapeutic systems.

In this review, we categorize F‐MNs into four main types: (1) Gas‐driven microneedles (GDMNs), which produce gas (e.g., hydrogen (${\mathrm{H}}_2$) [29], oxygen (${\mathrm{O}}_2$) [41], carbon dioxide (${\mathrm{CO}}_2$) [42] or carbon monoxide (CO) [43]) via in situ chemical reactions to drive drug flow; (2) Magnetic‐actuation microneedles (MAMNs) [36, 44], which incorporate magnetic materials or microrobots that move or vibrate under external magnetic fields; (3) Cavitation‐driven microneedles (CDMNs), which harness ultrasound‐triggered microbubble expansion for propulsive fluid motion [45]; and (4) iontophoresis‐enhanced microneedles (IEMNs) [30], which combine microneedles with electrical currents (iontophoresis) to actively push charged drugs into tissue (Figure 1b). Because several F‐MN systems combine multiple physical and therapeutic effects, this review classifies them according to the dominant mechanism responsible for active insertion or post‐insertion cargo transport. Thus, gas‐driven, cavitation‐driven, magnetic‐actuation, and iontophoresis‐enhanced systems are discussed as mechanism‐dominant categories rather than strictly separate functions. The main distinctions among these mechanism‐dominant categories are summarized in Table 1. By examining these categories in detail, including their working principles, materials, performance, safety, and translational progress, this review provides a comprehensive analysis of how force‐driven strategies are broadening microneedle‐mediated therapy beyond conventional passive transdermal delivery toward deeper, faster, and more controllable treatment across diverse anatomical sites. We further discuss future directions, including how the synergy between artificial intelligence (AI) and smart design and F‐MNs could optimize personalized therapy and accelerate clinical translation.

TABLE 1.

Comparison of major force‐driven microneedle mechanisms.

F‐MN category	Classification criterion	Operating principle	Advantages	Limitations	Applications	Translational challenges
GDMNs	In situ gas formation is the main source of pressure‐driven or convective transport; the gas may also be therapeutic.	Tissue fluid or external stimuli trigger ${\mathrm{H}}_2$, ${\mathrm{O}}_2$, ${\mathrm{CO}}_2$, CO, NO, or ${\mathrm{H}}_2\mathrm{S}$ generation, causing bubble expansion, local pressure, and deeper payload transport [29, 37, 38, 46, 47].	Hardware‐free active delivery; combines propulsion with gas therapy; relevant to hypoxic, inflamed, infected, or tumor tissues.	Gas kinetics and dose must be controlled; precursors may alter pH/osmolarity or compatibility; CO, NO, and ${\mathrm{H}}_2\mathrm{S}$ have narrow therapeutic windows.	Local inflammatory disease, chronic wounds, photodynamic therapy, and accessible tumors.	Translation requires control of gas dose, precursor stability, packaging, repeated‐use safety, and pharmacokinetics of bioactive gases.
MAMNs	Magnetic field is the primary external input for insertion, motion, navigation, heating, or release.	Magnetic particles/substrates respond to static, rotating, alternating, or gradient fields to enable actuation, microrobot navigation, detachable tips, magnetothermal heating, or triggered release [34, 35, 44].	Wireless and spatially controllable; enables navigation, localized heating, triggered delivery, and access to difficult tissue interfaces.	Requires magnets/coils and field alignment; penetration and heating vary with tissue depth; magnetic materials must be biocompatible, degradable, or retrievable.	Deep wounds, biofilm‐associated infections, gastrointestinal delivery, and internal tissue interfaces.	Translation requires field‐control hardware, user‐safe actuation, tracking, magnetic‐material safety, and reproducible dose placement.
CDMNs	Acoustic energy dominates bubble oscillation, expansion, or collapse near MN‐generated pathways.	Ultrasound activates microbubbles, nanobubbles, gas pockets, or cavitation nuclei, producing microstreaming, shockwaves, and transient mechanical disruption to enhance transport [36].	Rapid and localized delivery; externally tunable by acoustic parameters; valuable in dense tissues where diffusion is limited.	Cavitation must remain within safe limits; coupling, bubble distribution, and exposure are variable; requires ultrasound hardware and monitoring.	Tumor delivery, dense dermal lesions, and localized drug penetration.	Translation requires standardized acoustic parameters, cavitation monitoring, safety windows, and device integration.
IEMNs	Electrical current is the dominant driver of molecular transport through MN‐created skin pathways.	Low current drives electrophoresis and electroosmosis of charged or polar therapeutics, while MNs reduce barrier thickness and transport distance [30].	Precise on/off dosing; current‐ and time‐controlled delivery; suited to charged drugs, peptides, and wearable closed‐loop therapy.	Mainly limited to electrotransport‐compatible molecules; electrodes and power are required; risks include pH drift, electrode reactions, Joule heating, and irritation.	Antibiotic delivery, wound treatment, pain management, insulin or peptide delivery, and wearable therapy.	Translation requires current‐density safety, electrode stability, reproducible skin impedance, battery integration, and repeated‐use tolerability.

1.2. Clinical Motivation

The advent of F‐MNs is largely motivated by clinical demands that cannot be met by passive microneedle patches or conventional topical and injectable approaches [48]. One major motivation for developing F‐MNs is the delivery of therapeutics that have poor permeability into the skin but are ideally desired to be administered outside of hospitals and clinics [49]. For example, the success of mRNA vaccines (such as COVID‐19 vaccines) has stimulated interest in utilizing microneedle patches for self‐vaccination [50, 51, 52]. However, intradermal or transdermal delivery of mRNA requires penetration to deeper immune‐cell‐rich layers to trigger a robust response, a requirement that passive microneedles fall short of [50, 53]. In contrast, F‐MNs are being investigated as a way to enhance the deposition of mRNA nanoparticles into deeper dermal or subdermal compartments, mimicking the effect of an intramuscular syringe injection [28, 54, 55]. Similarly, other therapeutics based on peptides and proteins (insulin, growth factors, monoclonal antibodies) ought to be delivered to the systemic circulation or to deeper layers of tissues [56].

Another important motivation for developing F‐MNs is their utility in treating deep skin lesions, chronic wounds, and conditions such as psoriasis and eczema that involve inflammation throughout the dermis [29]. While topical dressings only address the surface and systemic drugs cause side effects, F‐MNs can deliver anti‐inflammatory therapeutics (e.g., cytokine inhibitors) across the full thickness of a psoriatic plaque and limit the side effects [29, 57]. In fact, recent studies indicate that hydrogen gas released from microneedles can alleviate imiquimod‐induced psoriasis‐like inflammation in skin by reducing oxidative stress in deeper layers [29]. Chronic wounds (e.g., diabetic foot ulcers) also exemplify the need for active delivery. These wounds often suffer from hypoxia and biofilm infections in the wound bed (several millimeters below the surface) [30]. Simply spreading growth factor gel or antibiotics on the wound is ineffective because they do not penetrate the necrotic tissue. F‐MNs could both breach the devitalized tissue and actively perfuse the wound bed with oxygen and drugs. For instance, oxygen‐generating microneedle arrays have been shown to create localized hyperoxic conditions that accelerate diabetic wound healing and promote tissue granulation. Infected wounds can benefit from iontophoretic microneedles, which improve antibiotic delivery and can electrically disrupt biofilms, achieving deeper tissue bacterial reductions unattainable by passive methods.

Another motivation behind developing F‐MNs is their potential in cancer therapy, particularly for treating tumors accessible on the skin or just below it. Melanoma and cutaneous lymphoma are examples of such cancers that are typically treated with systemic drugs that cause toxicity [58]. Although direct intratumoral injection is an option, it is painful and inadequate in uniformly distributing the drugs [59]. Microneedle patches can be applied to such tumors, but their passive diffusion‐based mechanisms may limit the delivery of drugs only to the periphery of tumors [59, 60]. Active approaches like cavitation or magnetic agitation can propel drugs throughout the tumor mass [61]. Similarly, a photothermal gas‐generating microneedle patch has been developed to treat melanoma by simultaneously heating the tumor and releasing CO gas to enhance the sensitivity of cancer cells to chemotherapy, resulting in better tumor regression in a mouse model [61]. These examples underscore the clinical promise of combining microneedles with force‐mediated delivery to tackle diseases that require deep, targeted, and timely drug deposition. From the patient's perspective, standard microneedles offer several benefits, such as reduced pain, the potential for self‐administration, and the elimination of hazardous sharps, while also expanding the range of drugs that can be delivered to deeper layers of the skin. By enabling the transdermal delivery of medications that are currently administered by injection, such as insulin, cytokine inhibitors, and monoclonal antibodies, these technologies could improve patient compliance and quality of life, especially for individuals with chronic diseases.

1.3. Design Considerations for Local Versus Systemic F‐MN Delivery

A key design principle for F‐MNs is whether the intended therapeutic effect is local or systemic. For local therapy, the main objective is to concentrate the payload within a defined diseased tissue region while minimizing systemic exposure. This design logic is relevant to inflammatory skin diseases, chronic wounds, infected wounds, and accessible tumors, where therapeutic benefit depends on achieving sufficient drug or stimulus intensity within the lesion rather than maximizing blood exposure [29, 30, 57, 58]. Therefore, locally acting F‐MNs should prioritize controlled deposition depth, lateral distribution within the lesion, tissue residence time, and spatial confinement of both the drug and the force‐generating stimulus. In these systems, stronger propulsion is not always advantageous; excessive force may drive payload away from the target region or increase systemic leakage. Instead, the force magnitude, activation time, microneedle length, array density, and matrix degradation rate should be matched to the thickness, stiffness, hydration state, and clearance profile of the target tissue [29, 59, 60]. For example, wound‐ or tumor‐directed F‐MNs may benefit from depot‐forming matrices, localized gas or electrical stimulation, and activation strategies that distribute the therapeutic agent throughout the diseased microenvironment while limiting off‐target diffusion.

In contrast, F‐MNs intended for systemic transdermal delivery are designed to move therapeutic cargo beyond the local skin depot and into dermal microcirculation, lymphatic drainage, or deeper tissue compartments. This design logic is particularly important for vaccines, insulin, peptides, proteins, monoclonal antibodies, and other biologics that require reproducible systemic or immune exposure [23, 24, 25, 26, 28, 54, 55, 56]. For these applications, the main design priorities are dose accuracy, reproducible skin insertion, controlled release rate, efficient transport across the viable epidermis and dermis, and predictable pharmacokinetic or immunological outcomes. Unlike local systems, where retention at the application site is often desirable, systemic F‐MNs should reduce uncontrolled local trapping, degradation, or heterogeneous deposition. Their force‐generating mechanism should therefore be tuned to enhance flux toward vascular or lymphatic uptake while maintaining patient comfort and avoiding tissue damage. This distinction indicates that F‐MN design should not be optimized only for maximum penetration depth, but rather for the intended therapeutic destination: localized tissue retention for regional therapy or reproducible absorption for systemic delivery.

2. Gas‐Driven and Gas‐Therapeutic Microneedles

Gas‐generating microneedle systems can be broadly divided into two functional categories: gas‐driven microneedles (GDMNs) and gas‐therapeutic microneedles (GTMNs). In GDMNs, the generated gas is primarily used as a physical driving force to enhance convective transport of conventional therapeutic payloads [29]. In GTMNs, the generated gas itself also serves as a therapeutic agent. For example, ${\mathrm{H}}_2$ can reduce oxidative stress and inflammation, ${\mathrm{O}}_2$ can relieve tissue hypoxia and support oxygen‐dependent therapies, and gasotransmitters such as NO, CO, and ${\mathrm{H}}_2\mathrm{S}$ can modulate immune, vascular, or tumor microenvironmental responses [37, 38]. Recent designs increasingly combine these two roles, enabling a single microneedle patch to provide both gas‐propelled delivery and local gas therapy, as illustrated in Figures 2 and 3.

FIGURE 2.

Gas‐generating microneedles using chemical reactions for deep drug delivery and gas therapy. Using ${\mathrm{H}}_2$ (a,b). (a) Hydrogen‐generating microneedle patch for psoriasis treatment using live E. aerogenes bacteria as internal ${\mathrm{H}}_2$ microreactors. (b) Fluorescence images of drug penetration at 0, 4, 8, and 24 h post‐application across depths from 0 to 900 $\mu \mathrm{m}$, showing ${\mathrm{H}}_2$ propulsion enables penetration up to $\sim 1$ mm, far deeper than passive diffusion. Using ${\mathrm{O}}_2$ (c–j). (c) Schematic of oxygen‐propelled dissolving microneedles loaded with sodium percarbonate (SPC) and photosensitizer Ce6. (d) Photographs of microneedle arrays with ${\mathrm{O}}_2$ bubbles visible after activation. (e,f) Fluorescence and elemental mapping. (g) Force–displacement curve. (h) Simulated intratumoral oxygen distribution with and without ${\mathrm{O}}_2$ generation and representative fluorescence images showing deeper Ce6 penetration in SPC+Ce6(+) groups. (i) Tumor growth curves and quantitative analysis. (j) H&E and Ki‐67 immunohistochemistry of tumor sections. Using CO (k–p). (k) Carbon monoxide‐generating microneedle system. (l) Treatment schedule indicating tumor induction, microneedle application, light irradiation, and data collection. (m) Tumor volume curves. (n) Photographs of excised tumors from each treatment arm. (o) Quantitative comparison of tumor weights confirming superior efficacy of CO‐augmented chemotherapy. (p) Body‐weight monitoring. Panels a and b adapted with permission from Ref. [29], Copyright [2024], Nature Publishing Group UK, London. Panels c–j adapted with permission from Ref. [46], Copyright [2022], Wiley Online Library. Panels k–p adapted with permission from Ref. [47], Copyright [2024], American Chemical Society.

FIGURE 3.

Gas‐therapeutic microneedle platforms co‐delivering therapeutic gas and drug and their integration into multi‐step cancer therapy. (a) Schematic of a dissolvable GTMN patch for rivastigmine (RVT) delivery incorporating citric acid (CA), ${\mathrm{K}}_2\mathrm{CO}$ , and PVA‐based RVT particulates. (b) Gas‐triggered drug release and permeation profiles showing rapid initial payload deployment followed by sustained delivery in gas‐generating microneedles compared with non‐gas controls. (c) Confocal microscopy at early and late time points demonstrating deeper and broader intradermal penetration of fluorescent probes with gas‐generating microneedles. (d) in vivo near‐infrared imaging. (e) Schematic timeline of a subcutaneous tumor model integrating intravenous injection of cancer cells. (f) Tumor volume curves showing superior suppression in gas‐enabled composite microneedle groups (PA‐CMNs) relative to controls. (g) Representative H&E, Ki‐67, and TUNEL staining of tumor sections. (h) Quantitative analysis of necrotic area, Ki‐67 cells, and ${\mathrm{TUNEL}}^{+}$ cells corroborating enhanced tumor cell killing in gas‐enabled cohorts. (i) Histology of skin at 30 min and 24 h post‐patch application showing mild, transient microdisruption and near‐complete architectural recovery, highlighting the local biocompatibility and repeatability of GTMN application. Panels a–c adapted with permission from Ref. [80], Copyright [2024], Elsevier. Panels d–i adapted with permission from Ref. [81], Copyright [2021], American Chemical Society. 

2.1. Design Strategy and Operating Principles

GDMNs convert chemical energy into mechanical work by generating bubbles in situ within the micro‐channels and pores created by the needles. Reactive components embedded in the microneedle matrix (or in internal microchambers) contact interstitial fluid and rapidly produce gases such as ${\mathrm{CO}}_2$, ${\mathrm{O}}_2$, or ${\mathrm{H}}_2$ [29, 62]. The expanding gas acts as an internal micro‐piston that drives therapeutic fluids and particles away from the needle tips and into the surrounding tissue matrix, analogous to a syringe plunger [29]. This convective transport pushes payloads millimeters beyond the needle length [29, 62, 63], enhancing cumulative permeation (e.g.,$\sim$2‐fold increase over passive microneedles) and accelerating onset of action [29, 47] while retaining the minimal invasiveness and patch‐based simplicity of microneedles.

2.1.1. Hydrogen‐Based Systems

Hydrogen is an attractive fuel because its generation is highly exergonic, its solubility and diffusivity favor rapid clearance, and it has emerging roles as an antioxidant and anti‐inflammatory agent [64]. Chemically, ${\mathrm{H}}_2$ can be produced from water‐reactive hydrides (e.g., sodium hydride (NaH)) [65], hydrolysable boron hydrides (e.g., ammonia borane) [66], or biodegradable metals such as magnesium [67], each offering distinct trade‐offs between reaction rate, byproducts, and storage stability. Biologically, fermentative bacteria encapsulated within microneedles can act as living microreactors that continuously produce ${\mathrm{H}}_2$ from endogenous substrates [68, 69].

Hydrogen‐generating microneedles exemplify the dual propulsion–therapy paradigm. One approach uses NaH embedded within the needle base; upon insertion into moist mucosa, NaH reacts with water to release ${\mathrm{H}}_2$, generating sufficient thrust to self‐drive the microneedles through nasal mucus and then dissolve to release a corticosteroid, improving outcomes in allergic rhinitis models [70]. A more biologically inspired design is shown in the “Using ${\mathrm{H}}_2$” panels of Figure 2. Here, dissolving microneedles encapsulate Enterobacter aerogenes as living microreactors that ferment glucose to ${\mathrm{H}}_2$ [29]. The schematic in Figure 2a illustrates bacterial ${\mathrm{H}}_2$ evolution within the needles, propelling the vitamin D derivative calcipotriol across psoriatic epidermis and dermis and attenuating cytokine circuits such as IL‐17/IL‐22. Depth‐resolved fluorescence images in Figure 2b show drug penetration up to $\sim 1$ mm over 24 h, representing more than a two‐fold increase compared with passive diffusion [29]. In addition to propulsion, molecular hydrogen exerts antioxidant and anti‐inflammatory effects, scavenging reactive oxygen species and further dampening local inflammation in both psoriasis and allergic rhinitis models.

Beyond these examples, ${\mathrm{H}}_2$ can be produced from boron hydrides (e.g., ammonia borane encapsulated in mesoporous silica for slow hydrolysis) [71] or from biodegradable metals such as magnesium (Mg + 2${\mathrm{H}}_2\mathrm{O}$ $\to$ Mg(OH) + ${\mathrm{H}}_2$), although these chemistries require careful control of reaction kinetics and byproduct handling. Acid‐degradable metal–organic frameworks (MOFs) co‐loaded into microneedles provide another GTMN strategy: in acidic tumors, such MOFs release both ${\mathrm{H}}_2$ gas and doxorubicin [95]. The generated ${\mathrm{H}}_2$ repolarizes tumor‐associated macrophages toward an M1 phenotype, downregulates metastasis‐related enzymes (e.g., MMP‐2), inhibits P‐glycoprotein and ATP production in resistant cells, and protects normal tissues by scavenging radicals, thereby potentiating the cytotoxic action of doxorubicin [73, 74, 75].

2.1.2. Oxygen‐Based Systems

Another common strategy is generating oxygen gas via decomposition of peroxides. Enzymatic ${\mathrm{O}}_2$ generation can be achieved by incorporating catalase to rapidly convert endogenous hydrogen peroxide (${\mathrm{H}}_2\mathrm{O}$ ) into ${\mathrm{O}}_2$ and water [46, 76, 77]. Alternatively, solid peroxides like calcium peroxide (${\mathrm{CaO}}_2$) can serve as self‐contained ${\mathrm{O}}_2$ sources. Upon exposure to aqueous fluids, ${\mathrm{CaO}}_2$ nanoparticles hydrolyze to yield ${\mathrm{O}}_2$ (and antimicrobial ${\mathrm{H}}_2\mathrm{O}$ ) directly at the target site [33]. In a recent design for diabetic wound healing, microneedles loaded with ${\mathrm{CaO}}_2$ were used to bypass the diffusion limits of skin and bacterial biofilms, delivering ${\mathrm{O}}_2$ into hypoxic wound tissue while simultaneously releasing bactericidal ${\mathrm{H}}_2\mathrm{O}$ . Similarly, dissolving microneedles loaded with sodium percarbonate (SPC) have been applied to solid tumors to enhance photodynamic therapy by generating oxygen microbubbles in situ [46]. The rapid decomposition of SPC upon insertion produces ${\mathrm{O}}_2$ gas that forms vigorous fluid convection currents in the interstitial fluid, actively transporting a co‐delivered photosensitizer (chlorin e6) deeper into tumor tissue (Figures 2c–f) [46]. This active delivery increased drug penetration depth (demonstrated in hydrogel and ex vivo skin models) and reduced tumor hypoxia during photodynamic therapy, leading to an increase in cancer cell death rates from $30\%$ to $99\%$ in vitro and markedly improved tumor suppression in vivo (Figure 2h,i). These examples illustrate how on‐demand ${\mathrm{O}}_2$ generation can both physically enhance drug transport and chemically augment oxygen‐dependent therapies, and decrease tumor proliferation (Figure 2j).

Other microneedle designs employ ${\mathrm{CaO}}_2$ nanoparticles to deliver ${\mathrm{O}}_2$ into diabetic wounds, simultaneously oxygenating ischemic tissue and exerting antibacterial effects, or catalase ${\mathrm{H}}_2\mathrm{O}$ and solid peroxide systems to generate hyperoxic niches in tumors [78]. In squamous cell carcinoma models, such ${\mathrm{O}}_2$‐releasing microneedles suppress HIF‐1$\alpha$, increase singlet‐oxygen generation during photodynamic therapy (PDT), and approximately double treatment efficacy relative to non‐oxygenating controls [32, 33, 76].

2.1.3. Carbon Dioxide and Carbon Monoxide‐Based Systems

${\mathrm{CO}}_2$ can serve as a precursor for generating therapeutic CO [79]. In the “Using CO” panels of Figure 2, microneedles incorporated with an effervescent tartaric acid/carbonate pair and polyethylene glycol (PEG)‐coated copper sulfide (CuS) nanostars are demonstrated. Upon insertion, ${\mathrm{CO}}_2$ is generated; during near‐infrared (NIR) irradiation, the photothermal CuS catalyst converts part of this ${\mathrm{CO}}_2$ into CO [47]. The schematic in Figure 2k outlines three stages: insertion and swelling with ${\mathrm{CO}}_2$ evolution, light‐triggered photocatalysis to produce CO, and removal of the intact array to terminate exposure and retrieve the catalyst. The treatment scheme in Figure 2l shows tumor induction, repeated microneedle application, irradiation, and end‐point analysis. Tumor growth curves (Figure 2m) and photographs of excised tumors (Figure 2n) demonstrate that CO‐generating microneedles combined with cisplatin (${\mathrm{eMN}}_{\text{CO}}$ + ${\mathrm{L}}^{+}$ + CDDP) yield the strongest tumor suppression. Quantified tumor weights (Figure 2o) confirm this synergy, while stable body weights (Figure 2p) indicate the absence of overt systemic toxicity [47].

Carbonate–acid effervescent systems are commonly employed as ${\mathrm{CO}}_2$ sources in GDMNs [82]. At the intracellular level, pH‐responsive formulations encapsulate sodium bicarbonate and doxorubicin in poly(lactic‐co‐glycolic acid) (PLGA) hollow particles. Acidification within endosomes and lysosomes induces ${\mathrm{CO}}_2$ bubble generation, which mechanically disrupts these compartments, thereby releasing doxorubicin into the cytosol and circumventing P‐glycoprotein–mediated efflux [37]. At the tissue level, the GTMN architecture in Figure 3a illustrates a conceptually similar effervescent engine. Microneedles are co‐loaded with citric acid (CA), potassium carbonate (${\mathrm{K}}_2\mathrm{CO}$ ), and rivastigmine (RVT) within a polyvinyl alcohol (PVA) matrix [37]. A superficial “valve” layer of free RVT provides an early bolus, while deeper RVT particulates act as a depot. Upon insertion, interstitial fluid dissolves CA and ${\mathrm{K}}_2\mathrm{CO}$ , generating ${\mathrm{CO}}_2$ microbubbles that push the freely soluble RVT rapidly across the dermis, while capillary forces and matrix erosion sustain delivery from the depot [37, 83]. The release and permeation curves in Figure 3b show a characteristic biphasic profile in gas‐generating microneedles (F@ICG‐P@RhB‐GMN), with both a steeper initial phase and higher overall permeation than non‐gas controls (F@ICG‐P@RhB‐MN) [80]. Confocal images (Figure 3c) demonstrate deeper and broader intradermal distribution of dual fluorophores for the gas‐generating formulation, and in vivo imaging (Figure 3d) confirms wider tissue coverage [81].

2.1.4. Gas‐Therapeutic Immunotherapy and Microenvironment Remodeling

Gas‐therapeutic platforms can also reshape the tumor microenvironment to support immunotherapy [38]. Nitric oxide (NO)‐releasing nanoparticles, for example, have been engineered as self‐propelled nanomotors whose NO output normalizes tumor vasculature, degrades extracellular matrix, and increases cytotoxic T‐cell infiltration from $\sim 2\%$ to $\sim 28\%$ in vivo [81]. When combined with adoptive T‐cell transfer, these systems yield improved tumor control relative to static nanoparticles. Figures 3e–h illustrate how such principles can be embedded in multi‐step onco‐immunotherapy regimens using microneedles [81]. The timeline in Figure 3e outlines subcutaneous tumor engraftment, intravenous administration of nanomotors or adjuvants, T‐cell infusion, and subsequent application of gas‐capable composite microneedles [81, 84]. Tumor volume curves (Figure 3f) show the most pronounced growth suppression in gas‐enabled composite microneedle groups (PA‐CMNs) compared with static microneedle (S‐CMN) and control cohorts. Representative H&E, Ki‐67, and TUNEL staining (Figure 3g) reveals reduced proliferation and increased apoptosis/necrosis in tumors treated with gas‐therapeutic microneedles, and quantitative analysis (Figure 3h) confirms larger necrotic areas, fewer cycling cells, and more TUNEL‐positive cells [81]. Dermal histology in Figure 3i (30 min and 24 h post‐patch) shows that ${\mathrm{CO}}_2$‐generating GTMNs induce only mild, transient microdisruption, with near‐complete restoration of skin architecture by 24 h and no evidence of destructive cavitation [81]. This supports the feasibility of repeated application in chronic regimens. Across all systems, microneedle geometry (length, tip shape, bore size, and array spacing) and internal layout (e.g., core–shell structures and microchambers) strongly influence gas accumulation, pressure, and flow paths [47, 85]. Denser arrays or longer needles confine gas to smaller volumes, increasing pressure but also the risk of discomfort, whereas tailored venting and spacing can steer flow toward desired targets [86]. Early computational models coupling Fickian diffusion and Darcy‐type interstitial flow are beginning to capture how gas production, tissue poroelasticity, and device architecture jointly determine spatiotemporal profiles of both drugs and gases, offering a route to rational optimization.

2.2. Safety Considerations

Biocompatibility hinges on both the structural matrix and the gas‐generating components [71]. Most devices rely on water‐soluble, clinically familiar polymers such as PVA [87], polyvinylpyrrolidone (PVP) [73], and hyaluronic acid [88], which dissolve or degrade into benign products and typically elicit only minor, mosquito‐bite‐like irritation. Gas precursors are frequently benign inorganic salts: calcium carbonate (${\mathrm{CaCO}}_3$), calcium peroxide (${\mathrm{CaO}}_2$), and SPC decompose into calcium ions, water, soda ash, and ${\mathrm{O}}_2$/${\mathrm{CO}}_2$, all of which tissues can manage in moderate quantities [89, 90]. In ${\mathrm{CaO}}_2$‐based wound patches, formulation strategies are used to prevent premature peroxide decomposition; at the point of use, oxygen and controllable ${\mathrm{H}}_2\mathrm{O}$ release promote bacterial clearance and tissue oxygenation without enduring toxicity [47, 79, 91]. SPC, which is widely used as an emergency oxygen source, also exhibits similarly favorable behavior in PDT microneedle applications [92].

Highly reactive hydrides such as NaH demand stricter containment, as they react vigorously with water and generate alkaline sodium hydroxide (NaOH) [93]. Designs therefore confine NaH to the needle base, shield it with polymer layers until full insertion, and embed it within hydrogels that rapidly swell and quench the reaction [46, 70]. In intranasal ${\mathrm{H}}_2$‐generating patches, such strategies prevented observable mucosal injury in animal models while maintaining propulsion. Future systems are likely to favor more physiologic fuels, such as glucose‐oxidase–driven gas generation or bicarbonate–acid reactions where possible. Also, local perturbations in pH or osmolarity are mitigated by buffering excipients and by the strong buffering capacity of blood and interstitial fluid. Across the platforms underlying Figure 2 and Figure 3, short‐term in vivo studies have reported no adverse histopathology at the application sites. However, the intrinsic toxicity profiles of different gases vary widely. ${\mathrm{H}}_2$ is largely inert at physiological conditions and has even been evaluated as an inhaled therapeutic, with minimal reported toxicity [42]. The small quantities generated by microneedle patches are rapidly dissolved and cleared, suggesting a low risk of embolism or systemic poisoning under the tested dosing conditions. ${\mathrm{O}}_2$ can cause oxidative injury at high partial pressures, but in microneedle systems the generated amounts are modest and transient, aimed primarily at relieving hypoxia or driving convective flow; available data indicate no oxidative damage and, instead, net benefit via correction of pathological hypoxia [92, 94]. By contrast, endogenous gasotransmitters such as CO, NO, and ${\mathrm{H}}_2\mathrm{S}$ have narrow therapeutic windows. CO binds hemoglobin and mitochondrial cytochromes with high affinity; NO and ${\mathrm{H}}_2\mathrm{S}$ modulate vascular tone and redox signaling [79, 88]. GTMNs using these gases therefore employ stringent spatial and temporal control: pH‐ or ${\mathrm{H}}_2\mathrm{O}$ ‐responsive donors, light‐triggered catalysis, and removable arrays that terminate exposure once treatment is complete, as exemplified by the CO‐generating microneedles in Figure 2k–p. For chronic indications, formal dose‐escalation and pharmacokinetic studies will be required to define safe exposure ranges and monitoring strategies (e.g., carboxyhemoglobin measurements for CO).

Gas‐generating microneedles can decrease inflammatory burden by lowering required drug doses and by leveraging the intrinsic antioxidant properties of gases such as ${\mathrm{H}}_2$ [95]. Nevertheless, foreign components, including metal catalysts (e.g., CuS nanostars), MOF scaffolds, enzymes, and especially encapsulated bacteria, carry risks of immunogenicity, granuloma formation, or infection. Removable, non‐dissolving backings that retrieve catalysts after use, preferential use of GRAS nanozymes, and engineered containment or self‐destruction of microbial payloads are all important design features for clinical translation. The bacterial ${\mathrm{H}}_2$‐generating system in Figures 2a and b underscores the need for robust sterility assurance and fail‐safe strategies to prevent unintended dissemination [29]. Histological analyses in GTMN‐treated skin (Figure 3i) demonstrate that, when materials and doses are appropriately chosen, gas events induce only transient, self‐limiting tissue responses. However, long‐term studies under repeated application will be needed to exclude subtle fibrosis or microvascular remodeling [81].

Finally, GDMNs and GTMNs will be regulated as combination products integrating a device, an active pharmaceutical ingredient, and a gas‐generating formulation. Regulatory dossiers must therefore address device safety, chemical and structural stability of gas precursors, pharmacology and toxicology of the generated gases, and any biologics (e.g., antibodies, cells, microbes) included in the system. For gases with narrow safety margins, regulators are likely to require stringent control of precursor loading, robust manufacturing quality systems, and mitigation plans for worst‐case scenarios such as accidental over‐activation.

2.3. Clinical Outlook

The performance of gas‐generating microneedles is inherently sensitive to local tissue mechanics and fluidics. Skin thickness, collagen content, and the presence of fibrosis or psoriatic plaques, as well as tumor extracellular matrix architecture and interstitial pressure, all influence how gas propagates and where drugs are deposited [68, 96]. The deep penetration seen in controlled models (e.g., $\sim 1$ mm for the ${\mathrm{H}}_2$ system in Figure 2b, or the broad Ce6 distribution in Figure 2h) may therefore vary in the clinic. Patient behavior adds further variability: incomplete patch application, inadequate dwell time, or mis‐timed external activation (e.g., NIR irradiation) can all degrade performance [97, 98, 99]. Training, ergonomic design, and potentially built‐in tactile or visual indicators of proper deployment will be important.

Fine control over gas release is central to both efficacy and safety. Spontaneous reactions initiated solely upon contact with tissue fluid can produce uncontrolled gas release; if gas evolves too quickly, it may cause pain or forcefully expel the payload, whereas if it evolves too slowly, it may not generate sufficient convective flow to be effective [63, 87]. Compartmentalized architectures (core–shell needles and segregated drug and fuel chambers as in Figure 3a) and externally triggerable chemistries (light‐, pH‐, or ${\mathrm{H}}_2\mathrm{O}$ ‐responsive donors) provide avenues for better control [37, 100].

Manufacturing adds an additional layer of complexity. Microneedle arrays that integrate multiple layers, catalysts, enzymes, or microbes are more challenging to scale than simple dissolving microneedles. Large‐scale manufacturing will likely require roll‐to‐roll or other high‐throughput molding approaches, automated filling of microcompartments, and in‐line assays to verify gas precursor loading and mechanical integrity. Reactive components (peroxides, hydrides, gasotransmitter donors) may demand specialized packaging and cold‐chain logistics, with implications for cost and global deployment [99].

Most existing data derive from single or short‐course treatments. Chronic indications, such as repeated GTMN administration for neurodegenerative disease, chronic inflammatory skin disease, or long‐term cancer immunotherapy, raise additional questions. Repeated gas pulses could, in principle, induce cumulative microvascular or stromal changes, or adaptive shifts in hypoxia‐ and stress‐signaling pathways. Chronic low‐level exposure to gasotransmitters with narrow safety windows (NO, CO, ${\mathrm{H}}_2\mathrm{S}$) must be carefully evaluated. Regulators will expect long‐term toxicity, immunogenicity, and carcinogenicity studies in appropriate models, especially for systems incorporating living microbes or novel nanomaterials [32, 33]. For widespread adoption, gas‐generating microneedles must demonstrate clear advantages over existing therapies in clinically meaningful end points, faster wound closure, superior tumor control, improved symptom management, or reduced dosing frequency, rather than merely incremental improvements in pharmacokinetic metrics. Devices requiring external hardware (e.g., lasers) or complex handling will initially be restricted to specialized centers unless integrated, user‐friendly designs can be developed. Cost‐effectiveness analyses will be important, particularly for multi‐component GTMNs whose manufacturing costs may exceed those of standard injectables or simple microneedles [83, 96].

Nonetheless, the preclinical evidence summarized in Figures 2 and 3 suggests that gas‐generating microneedles can uniquely combine minimally invasive administration, tunable convective transport, and local gas therapy. Continued advances in materials science, modeling, device engineering, and regulatory science will determine the pace at which these platforms progress from experimental patches to clinically adopted tools.

3. Magnetic‐Actuation Microneedles (MAMNs): Wireless Mechanical Force

Magnetic actuation is considered a complementary route in which deep delivery is rendered by wireless mechanical work rather than just a chemical propulsion [34]. In MAMNs, ferromagnetic or superparamagnetic materials are embedded in microneedles' tips or substrates so that an external field aligns, translates, heats, or decouples the structures to place payloads precisely within resistant tissues and hollow organs [36, 41, 101]. Our discussion in this section is not restricted to transdermal patches and includes magnetically controlled needle‐robot systems that operate inside the body. Although needles can be used to deliver therapy to deep‐seated targets, localizing them in hard‐to‐reach regions generally requires an active control system [102]. Magnetic‐actuation platforms are particularly suited to this task, as they enable wireless, precise guidance of needles within complex anatomical environments.

3.1. Design Strategy and Operating Principles

Field–material coupling is engineered at two distinct levels; heat generation at magnetically lossy tips and force transmission through magnetized backings [34, 36]. In a bilayer hyaluronic acid patch, iron oxide (${\mathrm{Fe}}_3{\mathrm{O}}_4$) nanoparticles concentrated at the tips were used to convert an alternating electromagnetic field into heat that is confined to the distal micron‐scale volume; the base remains comparatively cool, permitting sterilization of deep biofilm without collateral burning, and subsequent release of selenium nanoparticles regulates redox balance and angiogenesis in the wound bed (Figure 4a) [44]. The field source is provided by a disk‐shaped coil whose geometry yields minimal attenuation through tissue and concentrates electromagnetic loss in the needle tip, an effect verified by finite‐element analysis and by the spatial pattern of tip heating. These features enable penetration through scab, hyperkeratotic crusts, and mature biofilm, followed by magneto‐thermal ablation and staged nanoparticle delivery to promote regeneration [44, 90, 103]. Another example is an orally administered microneedle robot platform designed to deliver macromolecular drugs to the gastrointestinal tract [34]. These microneedle robots consist of three functional components: a drug‐loaded tip (fabricated from a biodegradable hydrogel such as gelatin methacryloyl), a separable middle link, and a magnetic base containing encapsulated neodymium‐iron‐boron microparticles. The components are UV‐polymerized in sequence, forming a composite microneedle with 650 $\mu \mathrm{m}$ in height mounted on a 3 mm magnetic disk. Encapsulation in a pH‐sensitive enteric capsule protects the device through the stomach, then releases it in the small intestine. Once deployed, an external magnetic field is used to navigate and orient the microneedle robot toward the intestinal wall and drive its needle tip into the tissue (Figure 4b–e) [34]. Controlled magnetic manipulation in ex vivo intestines demonstrated that the devices can actively move, align, and insert into gut tissue under field guidance. Upon tissue penetration, the middle link rapidly dissolves in intestinal fluids, detaching the drug‐loaded tip from the magnetic substrate. The embedded tips remain in the mucosal tissue and gradually release their drug payload, while the magnetic substrates are excreted. This approach achieved over 90% successful tissue penetration and tip detachment in intestinal tissue (Figure 4f). In a pig model of diabetes, magnetically guided oral microneedles loaded with insulin effectively delivered the hormone across the intestinal wall, achieving a reduction in glucose levels that standard oral formulations cannot accomplish (Figure 4h–k). These results highlight the potential of magnetic force not only to increase penetration depth but also to enable delivery to internal organs via non‐invasive routes [34, 44, 104, 105].

FIGURE 4.

Magnetic‐actuation microneedles (MAMNs) enable wireless, on‐demand drug delivery and therapy (a) Schematic of magnetically driven microneedle arrays: an external rotating/alternating magnetic field concentrates forces/heat at needle tips to (left) kill deep‐tissue bacteria by magneto‐thermal sterilization and (right) trigger nanoparticle/drug release that promotes regeneration. (b, c) Concept of a detachable drug‐loaded tip on a magnetic substrate: field‐guided penetration, tip–substrate separation, and controlled cargo release inside tissue. (d) Composite microneedles (polyvinyl alcohol + mesoporous iron oxide NPs + minoxidil) illustrating field‐mediated mass transport to dermis. (e) Rotation sequence (0–360) showing reversible bending of the needle array under a rotating magnet. (f, g) Performance metrics: penetration depth and bending distance versus substrate thickness and magnetization time. (h–i) Ex vivo tissue insertion and lateral steering demonstrated by switching magnetic polarity (arrows track tip motion). (j) Field‐strength dependent penetration and histology confirming minimal acute damage. (k–l) In vivo outcomes: improved tissue recovery and enhanced intradermal delivery (fluorescence) with MAMNs versus control. Panel a adapted with permission from Ref. [44], Copyright [2022], John Wiley Library. Panels b–i adapted with permission from Ref. [34], Copyright [2021], John Wiley Library.

Synergistic multi‐stimuli microneedle systems combine triggers like heat and magnetic fields with therapeutic payloads to treat challenging deep‐tissue conditions [106, 107, 108]. An exemplary platform is a bilayer microneedle patch designed for deep infected wound therapy, which integrates ferromagnetic nanoparticles and antioxidant agents into a degradable hyaluronic acid needle matrix. The tips of these needles are loaded with ${\mathrm{Fe}}_3{\mathrm{O}}_4$ nanoparticles, while the base contains micelle‐stabilized selenium nanoparticles (SeNPs) [34, 109]. When the patch is applied to a wound and exposed to an alternating electromagnetic field from a disk‐shaped coil (Disk‐ZVS device), the magnetic nanoparticles ${\mathrm{Fe}}_3{\mathrm{O}}_4$ at the tips of the needle heated by inductive magnetothermal conversion [34]. Importantly, the specially designed Disk‐ZVS field penetrates tissue with negligible attenuation and concentrates the electromagnetic energy at the microneedle tips. This focused heating can effectively sterilize deep infections, even beneath thick eschar or biofilms, by raising the temperature locally to kill bacteria. Meanwhile, the lower portion of the microneedles remains relatively cool, protecting healthy surrounding skin. As the hyaluronic acid matrix gradually dissolves in the enzyme‐rich wound environment, the embedded SeNPs are released [34]. Selenium is a potent reactive oxygen species (ROS) scavenger and pro‐angiogenic agent; its release mitigates oxidative stress from inflammation and stimulates new blood vessel formation to accelerate healing. Thus, this magnetically responsive microneedle patch provides a combination of deep‐targeted hyperthermia (to disinfect and disrupt biofilms) and biochemical therapy (antioxidant, pro‐healing effects), exemplifying how multiple stimuli can be harnessed to extend microneedle efficacy into deep or pathological tissues that were previously difficult to treat [34, 35, 44, 110]. To ensure a strong magnetic response, high‐permeability or high‐saturation magnetization materials are used. Common choices are iron oxide, known for its biocompatibility and use in FDA‐approved contrast agents; nickel, which requires a protective coating to prevent nickel ion release due to its lack of biocompatibility when left in the body; and neodymium‐iron‐boron microparticles, recognized for their strong magnetic properties, with a necessary precaution to prevent corrosion. Iron‐platinum or iron‐cobalt alloys can also be embedded to generate a strong effect, but are more exotic [43, 111, 112]. The shape of the magnetic material matters, too: aligned nanowires vs. random powder can respond differently to directional fields. Needles themselves can be made from a magnetic polymer composite that is cured in a magnetic field to align particles, giving a stronger actuation in a preferred direction [36, 101, 113, 114]. One reported design used screw‐like or spiral microneedles containing magnetic material. Under a rotating field, these could act like screws drilling into tissue to increase penetration depth beyond what initial insertion achieved. After insertion, a rotation could pull the needle deeper (or drive it like a screw). This is similar to the principles used in magnetically assisted capsule endoscopy, where rotating magnets induce movement, now adapted to a patch setting.

3.2. Safety Considerations

MAMNs generally use field strengths that are considered low‐risk when appropriately controlled, although safety depends on exposure conditions, device geometry, and patient‐specific factors. Static magnetic fields (like from a handheld magnet) cause no known tissue damage [115]. Time‐varying fields used for actuation are typically low frequency (for mechanical movement on the order of Hz to kHz), which induces negligible eddy currents in tissue (thus minimal heating from the field itself) [116]. However, if induction heating is the goal (for thermal therapy), higher frequencies (e.g., 100 kHz to MHz) and field strengths are used, which can heat magnetic particles that are intended but must be controlled to avoid overheating tissue beyond therapeutic levels [117, 118]. In Wei et al.'s [119] study on wound therapy, they ensured that the temperature stayed in a safe range ( 43 for hyperthermia) suitable for killing bacteria but not causing burns.

One safety aspect is ensuring that the microneedles or any magnetic fragments do not break off and remain in the body [120]. Iron oxide is relatively safe (similar to MRI contrast, eventually broken down by body), but materials like nickel or cobalt could cause local inflammation or systemic toxicity if significant amounts dissolve [121, 122]. To mitigate this, magnetic particles are often coated with biocompatible layers (silica, polymer, gold). Also, the quantities used in a patch are small. Nonetheless, thorough biocompatibility testing is needed, particularly for chronic or repeated use. The iron oxide nanoparticles injected for MRI eventually end up in the liver and spleen and can cause some transient changes but are generally safe; locally in the skin, a small deposit might just be cleared by macrophages over time [104, 123]. Magnetic fields exert forces not just on the patch but potentially on any other metal objects in or on the patient [124]. If the patient has a pacemaker or metal implant near the patch, caution is required, a strong magnet could perturb a pacemaker's function or cause a metal implant to move slightly (though typical actuating fields are not very strong) [125]. Patients with implanted electronics would have to consult with their doctor before using a magnetic patch system. Fortunately, actuation can be performed using moderate field strengths (ranging from 0.001 to 0.1 Tesla) which diminish with distance. Unlike MRI, which operates at 1.5–3 Tesla over the entire body, these fields are localized. If an alternating electromagnetic coil is used (for induction heating or vibrations), there might be some induced currents within the body [48, 117, 120]. However, at the frequencies and intensities used for a patch, these currents would be below the threshold that leads to peripheral nerve stimulation. Nevertheless, it is important for designers to ensure adherence to electromagnetic exposure guidelines, as stipulated by IEC standards [126].

Other potential safety considerations for magnetically actuated microneedle patches primarily relate to external magnetic fields, mechanical motion, and thermal effects, but can be effectively mitigated through design and usage guidelines [117]. Strong external magnets (e.g., MRI systems or industrial magnets) could potentially interact with magnetic components in the patch, causing movement or heating; however, this scenario is unlikely in unsupervised settings, as patients are routinely instructed–and packaging should clearly warn to remove such patches prior to MRI or avoid strong magnetic environments [122, 127]. Mechanically, excessive or uncontrolled vibration or rotation of microneedles could risk tissue tearing or pain, but reported systems typically employ small, controlled oscillations, with histological analyses showing only slightly enlarged micro‐channels without significant damage, emphasizing the importance of limiting motion amplitude and rotational speed [127, 128, 129]. Thermal side effects are a concern in magnetothermal applications, where local heating must be precisely regulated (e.g., via feedback control) to maintain temperatures within a mild hyperthermia range (43) to avoid burns [130]. Practically, magnetic actuation offers advantages such as deeper penetration and the ability to operate through clothing, but requires proper alignment of magnets or coils for efficient actuation and clear user instructions to ensure timely removal of permanent magnets to prevent overtreatment, premature drug depletion, or unnecessary tissue stress [122, 131].

3.3. Clinical Outlook

Several translational considerations warrant careful evaluation. First, field logistics in intestinal applications remain non‐trivial. Microrobots and magnetic substrates rely on finely tuned static and gradient fields, with propulsion speed and capture distance dictated by magnetization protocol, substrate geometry, and tissue conditions (Figure 4f,g) [34]. These dependencies will translate into inter‐patient variability due to differences in bowel motility, wall thickness, and luminal contents, necessitating robust controllers, safety margins, and patient‐specific planning to ensure reliable insertion depth and lesion coverage. Second, dose governance under motion is complex. Once tips detach, they elute cargo as peristalsis resumes; achieving uniform distribution along extended lesions will require coordinated navigation and controlled release prior to separation, supported by longitudinal tracking (Figure 4h,i) and, ultimately, feedback mechanisms to avoid clustering or under‐dosed segments [34, 101]. In dermal systems, magnetothermal regimens must balance bactericidal or adjuvant hyperthermia with host comfort and safety; while tip‐focused heating can confine energy deposition, clinical translation will demand real‐time thermometry, conservative temperature ceilings, and controlled duty cycles to prevent microcavitation or cumulative thermal injury (Figure 4a). Third, manufacturing and robustness present non‐trivial challenges: magnetized substrates must preserve remanence through fabrication, sterilization, and storage yet resist inadvertent demagnetization in clinical environments; degradable interlayers must exhibit consistent kinetics across physiologic variations in pH and enzymatic activity to ensure timely tip–substrate separation. Fourth, regulatory classification will be driven by combined device–drug behavior. The intestinal platform functions as an ingestible, externally actuated, transient implant, implicating device and combination‐product pathways and requiring rigorous evidence of safe transit, complete excretion, absence of obstruction, and preserved mucosal integrity, as supported by histology under both weak and strong field regimens (Figure 4j) [34]. For wound systems, long‐term electromagnetic exposure and thermal dose metrics must align with standards for chronically applied therapeutic devices.

Beyond these constraints, magnetic microneedle platforms offer distinctive advantages for controlled and responsive therapy. Their on‐demand actuation enables pulsatile or adaptive dosing; for example, a “smart bandage” with magnetic hydrogel microneedles that deliver antibiotics in short, scheduled bursts via a wearable coil, with actuation frequency increased in response to sensor‐detected infection for closed‐loop treatment. Similarly, magnetothermal microneedle patches could localize hyperthermia and chemotherapy to accessible tumors (e.g., chest wall recurrences, cutaneous metastases) [128], leveraging precedents from magnetic nanoparticle hyperthermia while reducing systemic exposure. Integration with MRI‐based guidance for steering magnetic microrobots remains speculative but conceptually extends existing MRI‐driven catheter and particle navigation paradigms to microneedle‐delivered agents [128].

A key advantage of magnetic actuation is depth reach: unlike optical or electrical triggers, low‐frequency magnetic fields can penetrate tissue, enabling post‐injection repositioning or concentration of magnetic depots toward deeper targets, in line with ongoing efforts in magnetic targeting of intravenously administered nanoparticles [36, 44]. From the patient's perspective, magnetic actuation is silent and minimally perceptible, supporting adherence. External magnets or coils can be embedded in compact, wearable formats, such as wraps or bandages, that activate patches on demand (e.g., around an arthritic joint during pain flares). Overall, magnetic actuation provides a wireless layer of mechanical and thermal control for deep, localized delivery, effectively complementing gas‐mediated and passive approaches. Recent demonstrations have shown that field‐directed penetration, rotation, and separation can be successfully orchestrated in vivo while maintaining tissue architecture integrity. Future advancements will rely on field‐aware dosing algorithms, scalable and stable magneto‐fabrication processes, integrated sensing for feedback control, and clinically applicable hardware that maintains benchtop precision across various patient anatomies and use cases [117].

4. Cavitation‐Driven Microneedles (CDMNs)

Ultrasound‐driven cavitation refers to the formation, oscillation, and collapse of gas bubbles generated when alternating compression and rarefaction cycles of an acoustic field impose sufficiently negative pressure to nucleate vapor cavities [87]. During oscillation, cavitation bubbles induce microstreaming and shear forces; during collapse, they generate highly localized mechanical impulses, including shockwaves and microjets [132, 133]. These mechanical effects transiently disrupt the ordered lipid‐protein architecture of the stratum corneum and enhance convective transport within deeper dermal layers. Because microneedles already bypass the superficial barrier by physically breaching the stratum corneum, coupling them with cavitation provides a synergistic mechanism in which bubbles act as microscale actuators that further propel drug payloads away from the needle tips and into the tissue matrix [34].

4.1. Design Strategy and Operating Principles

Ultrasound‐mediated cavitation forms the physical basis of CDMNs. Cavitation arises when alternating compressional and rarefactional phases of an acoustic wave generate sufficiently negative pressure to nucleate vapor cavities from dissolved gases. Once formed, these gas nuclei can undergo either stable oscillation or violent inertial collapse, producing localized hydrodynamic forces including microstreaming, microjets, shockwaves, and transient pore formation. Such forces increase permeability in both the stratum corneum and deeper dermal compartments by perturbing lipid lamellae, loosening intercellular junctions, and enhancing convective mass transport. Microneedles provide a unique substrate for cavitation because they physically bypass the outermost barrier layer and deliver acoustic energy directly into hydrated tissue microdomains where cavitation is most efficient [79, 133].

To unify the diverse CDMNs platform, we categorize activation mechanisms according to the mode of bubble generation: (1) chemically induced cavitation; (2) nanobubble‐assisted cavitation; (3) electrochemically induced cavitation; and (4) mechanically induced cavitation using integrated acoustic transducers. These four categories capture differences in bubble formation, spatial localization of cavitation, and the resulting drug propulsion dynamics.

Chemically driven CDMNs embed gas‐generating reactants within biodegradable microneedle matrices, enabling in situ ${\mathrm{CO}}_2$ generation upon contact with interstitial fluid. As shown in Figure 5a,b, gas‐reactive compounds are encapsulated beneath a diffusion‐modulating PEG/PLGA shell that delays hydration and controls bubble nucleation kinetics [134]. Upon insertion, interstitial fluid diffuses into the core, triggering acid‐base reactions that release ${\mathrm{CO}}_2$ microbubbles. Under ultrasound exposure, these bubbles serve as cavitation nuclei whose oscillation amplifies interfacial shear stress and generates vortex flow around the needle tips. This chemical–acoustic coupling enhances the penetration of both hydrophilic and hydrophobic payloads, evident from the deeper methylene‐blue (MB) fluorescence profiles (Figure 5d–f) and the increased reactive oxygen species generation during photodynamic therapy (Figure 5g,h). The resulting intratumoral accumulation leads to enhanced tumor regression without systemic toxicity (Figure 5i,j). Related strategies employing effervescent reactions for microneedle propulsion have been demonstrated in other gas‐generating microneedle systems for transdermal insulin delivery, chemotherapeutic enhancement, and oxygenating therapies [38, 135, 136], underscoring the scalability of chemical cavitation.

FIGURE 5.

Cavitation‐driven microneedle (CDMN) systems for ultrasound‐enhanced transdermal delivery. (a) Schematic illustration of the fabrication of bubble‐generating MNs using PEG/PLGA coatings and the mechanism of ultrasound‐induced cavitation. (b) SEM images of fabricated MNs showing uniform cone geometry and smooth surface morphology (scale bar: 200 $\mu \mathrm{m}$). (c) Histological and 3D‐projected sections of gelatin gel and ex vivo porcine skin. (d) Experimental timeline for 4T1 tumor–bearing mice. (e) Confocal fluorescence images of tumor sections showing intracellular MB distribution after different microneedle treatments (DAPI: blue; MB: red). (f, g) Quantification of MB fluorescence intensity and penetration depth. (h, i) Singlet Oxygen Sensor Green (SOSG) imaging and quantification of reactive oxygen species (ROS) production following PDT. (j) Tumor growth curves over 14 days. (k) Mouse body‐weight trajectories showing no treatment‐related systemic toxicity. (l) Micrographs showing the ZnO‐nanowire–modified microneedle surface and the emergence of microbubbles during electrochemical stimulation. (m) Quantification of bubble size distribution. (n) Time‐resolved bubble count. Panels a, b, and d–k adapted with permission from Ref. [134], Copyright [2021], American Chemical Society. Panel c adapted with permission from Ref. [45], Copyright [2022], American Chemical Society. Panels l‐n adapted with permission from Ref. [61] Copyright [2019], Wiley Online Library.

Nanobubble‐assisted CDMNs introduce preformed lipid‐stabilized gas nuclei onto the microneedle surfaces to lower the cavitation threshold and provide predictable bubble dynamics. The microneedles shown in Figure 5c utilize a layer‐by‐layer assembly of phospholipids (e.g., 1,2‐dipalmitoyl‐sn‐glycero‐3‐phosphoglycerol (DPPG) and cholesterol) that encapsulate gas pockets within nanobubble shells [45]. Upon ultrasound exposure, these nanobubbles undergo controlled oscillation and collapse, releasing the associated drugs and generating microstreaming that enhances diffusion through ex vivo skin models. Unlike chemical systems, nanobubble coatings act as a modular surface interface that can be added onto diverse microneedle substrates without altering core mechanical properties. Similar nanobubble‐enhanced acoustic systems have been reported for antibody delivery, anti‐inflammatory therapies, and nanoparticle transport through dense tissues [69, 133], highlighting their versatility.

Electrochemical CDMNs generate bubbles through voltage‐driven electrolysis at the needle‐tissue interface. As illustrated in Figure 5l–n, vertically aligned zinc oxide (ZnO) nanowires on silicon MNs catalyze the decomposition of interstitial fluid to produce gas bubbles upon application of low‐voltage stimulation [61]. Ultrasound then drives these electrolysis‐generated bubbles into oscillation and collapse, producing high‐energy mechanical impulses that disrupt tumor extracellular matrices and increase intracellular uptake of chemotherapeutics such as paclitaxel. Because bubble production is electrochemically tunable, these systems offer precise temporal control of cavitation activity. Analogous electrolysis‐based microneedle systems have been reported using gold, platinum, and iron‐oxide catalysts [137, 138], demonstrating broad material compatibility for electrochemical cavitation.

In mechanically actuated CDMNs, cavitation arises from direct delivery of ultrasound through the microneedle structure rather than from gas‐generating chemistries. In the hollow microneedle arrays fabricated via high‐resolution 3D printing [139], a piezoelectric transducer mounted beneath the microneedle patch transmits MHz‐range ultrasound through the needle shafts, focusing acoustic energy at the tip‐tissue interface. This configuration promotes bubble nucleation from dissolved tissue gases and generates intense acoustic streaming within the hollow lumen, enabling programmable liquid drug extraction and dispensing. Similarly, the wearable acoustic microneedle platform [140] uses digitally modulated acoustic fields to create tip‐localized vortex flows that actively pump liquid formulations from a drug reservoir into tissue. These mechanically driven systems support multiple delivery modes (single‐pulse, batch, and sustained release) and eliminate the need for reactive chemistries. Additional forms of mechanical cavitation have been demonstrated using surface acoustic wave (SAW) devices [141], which create nanoscale oscillations and localized microstreaming without large bubble formation, enabling delivery of macromolecules up to $\sim$2000 kDa.

Across chemical, nanobubble‐assisted, electrochemical, and mechanically induced systems, the unifying mechanism of CDMNs lies in converting ultrasound energy into localized hydrodynamic forces that overcome the diffusion limits of passive microneedles. As shown throughout Figure 5, cavitation transforms MNs from static conduits into active micro‐pumps capable of driving therapeutics hundreds of micrometers to millimeters into tissue within seconds. The spatial confinement of cavitation near microneedle tips may improve local delivery efficiency while aiming to limit off‐target effects under controlled acoustic conditions, providing a route for deep, controlled, and minimally invasive drug administration.

4.2. Safety Considerations

Ensuring the safe operation of CDMN requires careful control of the acoustic field, cavitation thresholds, and thermal load imparted to tissue. Cavitation activity increases nonlinearly with acoustic pressure; thus, operating in regions where bubble oscillation remains stable is critical for minimizing irreversible damage. At typical CDMN driving conditions (20–100 kHz or 0.5–3 MHz), stable cavitation dominates when the acoustic intensity is below $1.0\mathrm{W}{\text{cm}}^{-2}$ and the mechanical index (MI) remains under 0.8, generating reversible microstreaming that perturbs but does not rupture cellular structures [132, 133]. In contrast, inertial collapse appears when MI exceeds 1.5, producing microjets capable of breaching vascular endothelium. Studies using the chemically activated ${\mathrm{CO}}_2$ platform (Figures 5a and b, d–k) consistently operated below these thresholds, and no histological disruption was observed under pulsed ultrasound (1.5 W ${\mathrm{cm}}^{-2}$, 10% duty), corresponding to a time‐averaged intensity of $\sim$0.15 W ${\mathrm{cm}}^{-2}$. Similarly, nanobubble‐assisted cavitation (Figure 5c) is initiated at even lower pressures due to the high compressibility of lipid‐coated bubbles, enabling safe operation at intensities where no thermal‐ or ultrasound‐induced histological damage was reported under the applied ultrasound conditions. These quantitative boundaries ensure that CDMN platforms can leverage acoustic enhancement without crossing into regimes associated with microvascular rupture or thermal injury [133, 135].

Material composition represents an additional safety dimension. Biodegradable matrices such as PVP, PEG, and PLGA used in chemically activated systems are FDA‐cleared for injectable use and degrade into non‐toxic by‐products. Nanobubble coatings composed of DPPG and cholesterol integrate lipids commonly used in clinical liposomal formulations, ensuring compatibility with cutaneous tissue. Electrochemical CDMN platforms (Figure 5l–n) introduce ZnO nanowires as catalytic elements; however, safety assessments showed that complete blood counts and serum biochemical parameters remained comparable to controls, and H&E staining showed no necrosis at the probe insertion sites. Mechanically induced platforms using integrated piezoelectric transducers [139, 140] emphasize localized acoustic confinement to the microneedle–tissue interface. Infrared thermal measurements showed that the microneedle device with an embedded ultrasonic transducer maintained microneedle and skin‐surface temperatures near 30, while the internal piezoelectric element reached $\sim {48}^{\circ }\mathrm{C}$ under 110 kHz, 2 W ${\mathrm{cm}}^{-2}$ operation without evidence of heat‐induced tissue damage [139]. In contrast, the wearable 34 kHz acoustic microneedle patch reported by Wu et al. produced a localized skin temperature rise of $\sim$7–8 under high‐voltage actuation, yet these exposures remained within short‐term safety limits and did not result in histological abnormalities [140]. Repeated actuation studies further show that polymeric microneedle shafts maintain structural integrity under cyclic vibration, preventing tip fracture, a common failure mode in mechanically excited systems.

From a regulatory perspective, reported CDMN demonstrations operate within the sonophoresis regime (1–3.5 W ${\mathrm{cm}}^{-2}$), and under pulsed excitation their time‐averaged intensities can fall within the diagnostic ultrasound safety limits established by the FDA, namely the spatial‐peak temporal‐average intensity (ISPTA $\le$ 720 mW ${\mathrm{cm}}^{-2}$; MI $\le$ 1.9) [142]. These thresholds define the permitted thermal and mechanical output for clinical diagnostic systems and provide a useful benchmark for evaluating emerging microneedle‐integrated ultrasonic devices. Longitudinal studies tracking inflammatory markers (IL‐1$\beta$, TNF‐$\alpha$), barrier function (transepidermal water loss), and histopathology indicate full recovery of micro‐channels within 24–48 h without fibrosis or chronic inflammation. The lack of systemic toxicity, rapid resolution of transient erythema, and minimal thermal load position CDMN as one of the safest classes among active transdermal enhancement techniques. Collectively, these data support the clinical viability of CDMN, provided that future devices incorporate closed‐loop acoustic control and real‐time monitoring to avoid inadvertent transition into inertial cavitation regimes [139].

4.3. Clinical Outlook

The ability of CDMN systems to deliver payloads rapidly, deeply, and with spatial precision broadens their therapeutic scope across oncology, dermatology, vaccination, infection control, and regenerative medicine. Chemically driven platforms [134] have already demonstrated clinically meaningful enhancements in intratumoral penetration of methylene blue for photodynamic therapy, achieving deeper distribution within poorly perfused tumor cores where passive microneedles fail. Nanobubble‐enhanced cavitation [45] may similarly enable high‐efficiency delivery of peptide vaccines, siRNA constructs, and hydrophilic biologics that require rapid deposition into the viable epidermis and upper dermis. Electrochemical cavitation [61] has shown potential for chemotherapeutic delivery into stiff, fibrotic tumor matrices by generating ultrasound‐responsive micro‐channels that transiently widen interstitial pathways. This mechanism is especially promising for desmoplastic cancers such as pancreatic ductal adenocarcinoma and certain sarcomas, where diffusion barriers limit therapeutic access.

Wearable mechanically induced CDMN systems [139, 140] open opportunities for ambulatory and closed‐loop drug administration. Their digitally programmable acoustic fields allow fine control of dose timing, depth, and rate, enabling transdermal insulin bolusing, analgesic titration, or local immunomodulation without traditional infusion pumps. Hollow microneedle reservoirs integrated with on‐patch electronics further permit delivery of liquid biologics such as monoclonal antibodies or growth factors, circumventing the payload limitations of dissolving microneedles. In wound‐care applications, cavitation‐enhanced transport could deposit antimicrobial peptides or reactive oxygen species within hypoxic or biofilm‐rich wound beds, accelerating re‐epithelialization and reducing bacterial burden. For regenerative medicine, cavitation‐assisted delivery of stem‐cell‐derived exosomes, cytokines, or pro‐angiogenic factors may overcome stromal diffusion limits and promote controlled tissue remodeling.

The path to clinical translation will depend on scalable manufacturing, consistent acoustic coupling, and standardized exposure metrics. Advances in high‐resolution three‐dimensional (3D) printing enable production of low‐cost, mechanically compliant patches feasible for scalable production. Regulatory adoption is facilitated by the fact that CDMN operates under existing diagnostic ultrasound frameworks, simplifying preclinical safety requirements. Early‐phase human trials will likely focus on dermatologic oncology, chronic wound care, and vaccine delivery, where cavitation‐enhanced penetration addresses clear clinical unmet needs. Hybrid systems integrating ultrasound with optical therapy, electrochemical sensing, or feedback‐controlled dosing are emerging directions with the potential to transform CDMN into multifunctional theranostic platforms.

5. Iontophoresis‐Enhanced Microneedles (IEMNs)

Iontophoresis‐enhanced microneedles (IEMNs) merge electric‐field‐driven transport with the micro‐channel‐creating capability of microneedles to achieve active, precisely dosed transdermal delivery [39, 40]. By superimposing a mild current across hydrated skin, iontophoresis adds electrophoretic and electroosmotic flux components to passive diffusion, greatly improving penetration of hydrophilic or charged macromolecules that normally traverse the stratum corneum poorly [143, 144, 145]. Electrically driven microneedle systems can be separated into transport‐dominant and electrotherapeutic designs. In transport‐dominant IEMNs, the applied current mainly serves to increase drug flux through electrophoresis and electroosmotic flow. In electrotherapeutic systems, however, the electrical stimulus may also contribute biologically by disrupting bacterial biofilms, modulating local electrochemical conditions, or stimulating wound‐repair processes while simultaneously enhancing drug penetration [30, 39, 40]. This dual role makes electrically driven microneedles particularly attractive for infected wounds, chronic ulcers, and other conditions where both deeper drug delivery and local bioelectric therapy may be beneficial.

5.1. Design Strategy and Operating Principles

Field‐responsive microneedles exploit materials that combine mechanical fidelity with ionic conductivity and electronic compatibility. A PVP matrix loaded with naloxone established that polymeric microneedles doped with buffer electrolytes can sustain transient current densities without structural collapse [39]. Incorporation of silver–silver chloride (Ag/AgCl) contacts within the patch base created a distributed anodal surface capable of delivering a six‐fold higher drug flux within 1 h compared to passive diffusion. The mechanical dissolution of the PVP network shortened the lag time of diffusion from tens of minutes to five, allowing a rapid pharmacokinetic onset suitable for emergency administration. Such electrically augmented dissolving patches demonstrated that the polymer phase not only supports ion migration but also modulates electroosmotic flow through internal hydration gradients, establishing a tunable electrochemical permeability [39].

Beyond dissolving architectures, permanent or reusable geometries integrate conductive scaffolds. A porous array was fabricated from crosslinked glycidyl methacrylate and triethylene glycol dimethacrylate, creating a micro‐porosity gradient that stores liquid drug within interconnected voids [30]. Under applied potential, these channels act as distributed conduits for electrophoretic transport, producing nearly complete penetration efficiency in ex vivo tissue while maintaining structural integrity under repeated compression cycles. The conductive path is reinforced by a hydrogel interlayer that serves simultaneously as an ionic reservoir and as a compliant contact for uniform current distribution. The inclusion of Ag/AgCl electrodes on flexible printed‐circuit substrates further transforms the patch into a self‐contained electrochemical transducer, capable of both driving drug flux and recording current–time profiles indicative of skin impedance or analyte concentration.

The iontophoresis‐microneedle array patch (IMAP) powered by a smartphone has demonstrated an electronically addressable platform [146]. The IMAP integrates a polymethyl‐methaccalrylate microneedle matrix, a porous drug reservoir, and a miniature constant‐current circuit powered by a mobile phone. By coupling press‐and‐release mechanical actuation with electronically modulated iontophoresis, the system implements the triphasic administration sequence of penetration, diffusion, and iontophoresis. The microneedle array first breaches the stratum corneum, then retracts, leaving transient aqueous microholes through which charged nanovesicles are driven under controlled current. The precision of this electronically stabilized current source allows linear modulation of delivery rate with current amplitude, confirming direct proportionality predicted by Faraday's law of electrolysis. Such architectures exemplify the convergence of microelectronics and polymer mechanics into wearable transdermal circuits [146].

To directly demonstrate the synergy between microneedle‐created micro‐channels and electric‐field–driven transport, Peng et al. developed carbon nanotube (CNT)‐enhanced, glucose‐responsive hydrogel microneedles capable of forming a transient layer that modulates insulin release under hyperglycemic conditions [147]. Their platform (Figure 6a,b) shows how conductive hydrogels facilitate both ion mobility and electroosmotic solvent flow during iontophoresis. High‐molecular‐weight FITC‐dextran experiments (Figure 6c–e) further revealed that electroosmotic flow dominates long‐range transport for larger macromolecules, validating a key mechanistic advantage of IEMNs: microneedles create short diffusional paths while iontophoresis supplies size‐independent convective enhancement.

FIGURE 6.

Iontophoresis‐enhanced microneedle (IEMN) systems for electrically driven transdermal drug delivery. (a, b) CNT‐enhanced glucose‐responsive microneedles for iontophoretic insulin delivery. Panels illustrate the hydrogel‐CNT network, formation of the glucose‐dependent skin layer, and iontophoresis‐assisted transdermal transport. (c–e) FITC‐dextran studies: iontophoresis enhances macromolecular delivery compared to passive permeation, with electroosmotic flow dominating transport for larger dextrans. (f) The spatial concentration distributions during transdermal delivery reveal progressive OVA permeation across seven conditions (control, 1 iontophoresis, MN, MN/0.5 iontophoresis, MN/1 iontophoresis, MN/1.5 iontophoresis, MN/2 iontophoresis). Iontophoresis drives OVA along the applied electric field, while microneedle‐created micro‐channels enhance diffusion. Panels a‐e adapted with permission from Ref. [147], Copyright [2024], American Chemical Society. Panel f adapted with permission from Ref. [148], Copyright [2023], Nature Publishing Group, London.

This mechanistic framework was expanded by mapping spatial concentration distributions of ovalbumin (OVA) under microneedle‐only, iontophoresis‐only, and combined microneedle and iontophoresis conditions [148]. As shown in Figure 6f, the deepest and most uniform penetration occurred with simultaneous microneedle insertion and iontophoresis, demonstrating that diffusion through aqueous micro‐channels and electric‐field‐driven transport are additive rather than redundant. These findings confirm that IEMNs uniquely enable transdermal delivery of hydrophilic proteins that otherwise show negligible passive permeation.

Beyond polymeric and rigid structures, hydrogel and conductive polymer microneedles have redefined iontophoretic sensing [149, 150]. Under iontophoretic stimulation, charged solutes move electrophoretically while the negatively charged skin and hydrated polymer network generate electroosmotic solvent flow; passive diffusion contributes the remainder [30, 151].

Recognizing these concurrent transport modes has guided the design of hydrogel and conductive‐polymer microneedles that maintain high water content, low interfacial resistance, and efficient ion–electron transport during actuation [152, 153]. Hydrogel microneedles composed of hyaluronic acid or methacrylated derivatives permit reversible swelling that adapts to electric‐field‐induced solvent flow, maintaining low interfacial resistance. When doped with conducting polymers such as poly(3,4‐ethylenedioxythiophene)‐polystyrene sulfonate (PEDOT:PSS), these matrices achieve mixed ionic–electronic conductivity, allowing simultaneous actuation and sensing. The hydrated polymer lattice supports electrochemical redox cycling for in situ detection of biomarkers (glucose, lactate, or electrolytes) while serving as an electrode–electrolyte continuum for iontophoretic pumping [69]. Porous and hollow microneedles extend this function to macromolecular analytes by minimizing diffusional path length and generating localized electroosmotic vortices that enhance convective transport.

The combination of these materials and mechanisms converts the microneedle patch from a passive sampling membrane into an electro‐mechanically coupled transducer. Each applied current pulse simultaneously drives molecular motion and yields a measurable electrical signature that correlates with tissue conductivity, hydration, or analyte concentration.

5.2. Safety Considerations

Electrical augmentation in microneedle systems must operate strictly within low current‐density regimes to remain clinically relevant [154]. Most iontophoretic microneedle platforms are intentionally designed to run at $\le 0.5\mathrm{mA}{\mathrm{cm}}^{-2}$, as higher currents increase the risk of irritation, thermal rise, and electrochemical instability and therefore fall outside acceptable safety limits for human use [30, 155]. At these low currents, the applied potential across hydrated skin is typically kept below $10\mathrm{V}$, and no irreversible epidermal changes or barrier disruption have been reported. Histological assessments of porous and hydrogel‐based arrays likewise show negligible irritation or hypersensitivity after repeated stimulation [30, 155].

Because iontophoretic performance deteriorates rapidly when electrode polarization, pH drift, or gas evolution occurs, stable electrochemical interfaces are essential; noble or passivated materials (Ag/AgCl, gold, carbon) remain preferred to avoid metal dissolution. Modern iontophoretic microneedle patches, including smartphone‐controlled and wearable current drivers, incorporate charge‐balancing or biphasic driving schemes to ensure the net delivered charge per cycle remains near zero [156], suppressing skin polarization and maintaining physiological pH at the electrode‐tissue interface.

Thermal regulation is inherent to safety management [157]. Joule heating in hydrated tissue is minimal at the milliwatt level, but localized hot spots can occur at sharp electrode edges or within resistive polymer regions [158]. Finite‐element simulations and infrared thermography in recent prototypes indicated temperature elevations below $2^{\circ }\mathrm{C}$ under continuous operation [69], well below thresholds for protein denaturation or nociceptive stimulation. Nonetheless, integration of micro‐thermistors or impedance‐based feedback loops is advisable for future diagnostic implementations, allowing automatic current modulation upon detecting rising temperature or altered skin resistance.

Chemical compatibility and skin physiology govern chronic‐use safety. The electrolytes and polymers must remain non‐cytotoxic and non‐sensitizing; hydrogels cross‐linked from PVA, PLGA, or hyaluronic acid generally fulfill these requirements. Because iontophoresis transiently alters transepidermal water loss and barrier function, recovery intervals are essential between successive operations. Longitudinal animal studies using insulin and methotrexate models reported complete restoration of barrier integrity within 24 h, validating reversible permeability modulation [31].

From an electrical engineering perspective, compliance with IEC and ISO standards on skin‐contact current density and electromagnetic exposure (less than 10 mA RMS total body current) ensures that IEMNs remain within safe operating envelopes.

5.3. Clinical Outlook

Iontophoresis sustains controlled infusion of therapeutic agents while simultaneously measuring impedance or electrochemical feedback to assess tissue response [159, 160]. The methotrexate platform exemplified this duality: pre‐pored human skin under anodal stimulation exhibited similar transdermal flux in psoriatic and healthy tissues yet markedly different intradermal accumulation, demonstrating that electrical modulation can compensate for disease‐related barrier heterogeneity while revealing pathology‐dependent uptake profiles [31]. Such electrodiagnostic signatures could in principle guide adaptive dosing by quantifying local resistance or capacitance as surrogates for skin inflammation, edema, or fibrosis, enabling personalized adjustment of current amplitude and treatment duration in real time.

Clinical translation will depend on miniaturized, self‐regulating electronics and standardized dosing protocols. Smartphone‐integrated current controllers, as pioneered in iontophoresis‐integrated insulin patches, already demonstrate patient‐specific programming of current amplitude, duration, and waveform using low‐voltage constant‐current drivers [146].

Wearable vaccine and insulin delivery patches further show that iontophoresis hardware can be packaged into wrist‐ or patch‐mounted modules with simple user interfaces, rechargeable batteries, and single‐button operation. The convergence of flexible printed electronics, conductive microneedles, and wireless telemetry will streamline treatment by enabling programmable, closed‐loop dosing with real‐time feedback on current stability and skin impedance, could support more consistent transdermal delivery in outpatient or home‐care settings after further validation. Coupling these systems to mobile health platforms also opens a path to remote supervision, where clinicians can review usage logs, adjust dosing regimens, and impose safety cut‐offs without requiring in‐person visits.

From a regulatory perspective, IEMNs' delivery platforms align with established safety precedents of both transdermal electrical stimulators and dissolving or hydrogel‐forming microneedle patches. Their classification as combination products mandates dual evaluation of device reliability and biochemical compatibility, yet the individual subsystems‐low‐voltage constant‐current circuitry, Ag/AgCl or carbon electrodes, and biodegradable polymers already possess extensive clinical acceptance across glucose monitors, TENS devices, and microneedle‐based vaccines.

As manufacturing advances toward scalable roll‐to‐roll imprinting of microarrays and integrated flexible circuits, the cost per unit is expected to fall toward that of disposable electrochemical test strips, bringing consistent, programmable drug delivery into routine outpatient and home‐care settings. Robust quality control for microneedle geometry, array uniformity, polymer conductivity, and current‐source stability will be essential to ensure batch‐to‐batch reproducibility and to satisfy regulatory expectations for long‐term reliability.

Future clinical deployment will also require attention to population‐specific and long‐term use considerations. Patients with chronic inflammatory skin disease, diabetes, peripheral vascular disease, or neuropathy may exhibit altered skin impedance, hydration, and sensation, necessitating conservative current ceilings and automated safety interlocks [39, 149]. Repeated application over weeks to months will need to be evaluated for cumulative irritation, changes in barrier function, and potential sensitization to electrode or adhesive components.

At the same time, the ability to electronically tune flux makes IEMNs attractive for drugs with narrow therapeutic windows, including methotrexate, insulin, naloxone, and emerging biologics, by decoupling dose from a fixed patch loading and enabling titration based on pharmacokinetic response or symptom relief. Compared with thermal, acoustic, or jet‐based enhancement methods, iontophoretic microneedles offer the distinct advantage of low‐temperature, electrically programmable transport using familiar materials and form factors, positioning them as a promising route for chronic, self‐administered transdermal therapies.

Through judicious control of material conductivity, current density and waveform, and integrated signal feedback, iontophoretic microneedle systems extend the paradigm of microneedle technology from purely mechanical penetration to dynamic electronic communication with the body, transforming skin from a passive barrier into an actively interrogated and actuated therapeutic interface.

6. Fabrication Strategies and Manufacturing Considerations for F‐MNs

The fabrication of F‐MNs differs fundamentally from that of conventional passive microneedles. Conventional microneedles are typically designed to create transient microchannels in the stratum corneum and then release their payload through coating dissolution, polymer degradation, swelling, or passive diffusion [11, 13, 14, 15]. Their fabrication therefore, mainly focuses on controlling needle geometry, tip sharpness, mechanical strength, drug loading, dissolution kinetics, and biocompatibility. Established approaches include silicon or metal microfabrication for solid and hollow microneedles, dip‐coating or spray‐coating for coated microneedles, micromolding and casting for dissolving polymeric microneedles, and hydrogel crosslinking for swelling microneedles [12, 16, 17, 18]. In contrast, F‐MNs must be fabricated not only as skin‐penetrating structures or drug reservoirs, but also as active microdevices capable of generating, transmitting, or coupling external forces to enhance tissue insertion or drive payload transport beyond passive diffusion.

A key fabrication requirement for F‐MNs is the integration of force‐generating or force‐responsive components within the microneedle architecture. As discussed, these components may include gas precursors, enzymes, catalysts, magnetic particles, acoustic cavitation nuclei, conductive hydrogels, electrodes, or external hardware interfaces [29, 30, 34, 36]. As a result, F‐MN fabrication usually involves multi‐material and multi‐layer designs rather than single‐phase polymer casting. The location of each functional component is critical: force generators must be positioned where they can produce useful propulsion or actuation, while drugs must remain stable and spatially accessible for controlled release. Therefore, F‐MNs often require compartmentalized tips, core–shell structures, detachable components, porous reservoirs, conductive backing layers, or integrated circuits. This manufacturing complexity is one of the major features distinguishing F‐MNs from conventional microneedle patches.

For gas‐driven and gas‐therapeutic microneedles, fabrication must balance mechanical strength, chemical reactivity, gas‐generation kinetics, and drug stability. Reactive components such as carbonates, peroxides, hydrides, gas donors, catalysts, or living microreactors must be incorporated into the microneedle matrix without premature activation during processing or storage [29, 37, 46, 47]. This typically requires spatial separation of reagents, moisture‐protective packaging, and careful selection of polymer matrices such as PVA, PVP, hyaluronic acid, or other dissolving or swelling polymers [73, 87, 88]. In some systems, the gas precursor is embedded in the needle base or internal microchambers so that interstitial fluid triggers gas evolution only after insertion [29, 62, 70]. In others, acid–carbonate or peroxide‐based reactions are distributed within the needle body to generate ${\mathrm{CO}}_2$ or ${\mathrm{O}}_2$ microbubbles that push the payload deeper into tissue [37, 46, 81]. These designs require control over precursor particle size, loading density, reagent distribution, polymer hydration rate, and gas escape pathways, because excessive gas generation may cause tissue irritation or payload expulsion, whereas insufficient gas generation may fail to produce meaningful convective transport.

Magnetic‐actuation microneedles require a different fabrication logic. Their performance depends on the amount, distribution, and magnetic responsiveness of embedded ferromagnetic or superparamagnetic materials [34, 35, 44]. Magnetic particles such as ${\mathrm{Fe}}_3\mathrm{O}$ , neodymium‐iron‐boron microparticles, nickel, or other high‐magnetization fillers must be incorporated without compromising tip sharpness, insertion strength, or biocompatibility [34, 43, 44, 111, 112]. Fabrication approaches include casting magnetic particles into polymer matrices, concentrating magnetic nanoparticles at the needle tips, aligning anisotropic fillers under an external magnetic field during curing, or assembling detachable drug‐loaded tips onto magnetic substrates [36, 101, 113, 114]. For magnetothermal designs, magnetic fillers must be localized where heating is desired, usually near the tip, to maximize antibacterial or therapeutic effects while avoiding excessive heating of surrounding tissue [44, 109]. For microneedle‐robot systems, fabrication is even more demanding because the device must combine a drug‐loaded biodegradable tip, a dissolvable or separable linker, and a magnetic base that can be guided and later excreted or retrieved [34]. Thus, magnetic F‐MNs require simultaneous optimization of magnetic torque or force, mechanical penetration, payload release, and post‐treatment safety.

Cavitation‐driven microneedles are fabricated to couple microneedle‐created microchannels with ultrasound‐induced hydrodynamic forces. In these systems, the microneedle surface or reservoir may be modified with bubble‐generating coatings, gas‐stabilizing particles, nanobubbles, microbubbles, or acoustic interfaces [45, 61, 134]. For example, PEG/PLGA coatings can be used to generate or stabilize cavitation nuclei near the microneedle tip, while electrochemical systems can use catalytic surfaces such as ZnO nanowires or noble‐metal electrodes to produce gas bubbles before ultrasound activation [61, 134, 137, 138]. More advanced mechanically induced systems integrate hollow microneedles with piezoelectric transducers or acoustic reservoirs, enabling ultrasound to be transmitted through the needle shaft and focused near the tip–tissue interface [139, 140, 141]. Fabrication of CDMNs therefore requires control over needle geometry, acoustic coupling, surface roughness, bubble nucleation sites, coating adhesion, and mechanical durability under repeated ultrasound exposure. Unlike passive microneedles, where surface smoothness is often desirable for reproducible insertion, CDMNs may intentionally engineer surface heterogeneity or gas pockets to promote controlled cavitation.

Iontophoresis‐enhanced microneedles require integration of mechanical microneedle arrays with electrochemical and electronic components. Their fabrication involves conductive or ionically conductive materials, drug reservoirs, flexible electrodes, hydrogel interfaces, insulating layers, and current‐control circuits [30, 39, 40, 146]. Polymeric microneedles may be doped with buffer electrolytes to sustain current flow, while Ag/AgCl electrodes, conductive hydrogels, porous reservoirs, or flexible printed‐circuit substrates are incorporated to distribute current uniformly across the patch [30, 39, 146]. In such systems, fabrication must prevent short‐circuiting, electrode delamination, uncontrolled pH shifts, Joule heating, and nonuniform current density at the skin interface. The choice of hydrogel is particularly important because it can act simultaneously as a drug reservoir, ionic conductor, skin‐contact layer, and mechanical cushion [146, 147, 148]. Therefore, IEMN manufacturing requires co‐design of the microneedle geometry, electrical pathway, drug formulation, and wearable power source.

Across all F‐MN categories, several manufacturing challenges are shared. First, the device must maintain sufficient mechanical strength for skin insertion despite the inclusion of functional fillers, pores, reservoirs, reactive agents, or electronic layers. Second, the force‐generating component must remain stable during fabrication, sterilization, storage, and handling, but become active only under the intended physiological or external trigger. Third, drug stability must be preserved in the presence of reactive gases, catalysts, magnetic particles, ultrasound exposure, electric fields, or local pH changes. Fourth, the fabrication process must provide batch‐to‐batch reproducibility in needle geometry, payload distribution, force‐generator loading, and activation kinetics. Finally, clinical translation will require scalable manufacturing methods, such as high‐throughput micromolding, roll‐to‐roll processing, automated reagent filling, additive manufacturing, and in‐line quality control to verify mechanical strength, dose uniformity, electrical continuity, magnetic response, acoustic performance, or gas‐generation rate [83, 96, 99, 161, 162]. These requirements position F‐MNs not simply as modified microneedle patches, but as integrated therapeutic microdevices whose fabrication must combine materials engineering, microfabrication, pharmaceutics, and device‐manufacturing principles.

7. AI Integration for F‐MNs and Deep Drug Delivery

Combining AI with F‐MN technology creates the foundation for smart, adaptive drug delivery [163]. AI can enhance nearly every aspect of microneedle design and application, ranging from optimizing materials and structures to real‐time personalizing delivery protocols and managing complex datasets generated by these sophisticated devices.

7.1. Design and Optimization

One of the first impacts of AI has been in the design phase of microneedles [162]. Traditional trial‐and‐error or purely physics‐based simulations can be extremely time‐consuming when exploring the vast design space (materials, geometry, array patterns, drug formulations, etc.). Machine learning models can be trained on simulation or experimental data to predict performance outcomes (e.g., penetration depth, drug release rate, mechanical strength) for new designs [164, 165]. For example, a Bayesian optimization approach combined with machine learning algorithms was used to minimize the insertion pain of microneedles by varying design parameters [165]. This model evaluated different shapes and lengths, learned from the results, and intelligently suggested new design iterations, resulting in a microneedle geometry that theoretical models predicted to cause less pain upon insertion. Similarly, a Bayesian machine learning algorithm was employed to optimize microneedle design for sampling interstitial fluid (for diagnostic purposes). The algorithm determined the optimal needle spacing, length, and insertion force to maximize fluid extraction rate [161, 166, 167]. These examples show how AI can help balance several competing parameters at once, making the process more efficient than relying on traditional trial‐and‐error methods. In the context of F‐MNs, AI‐driven design might, for instance, find the ideal ratio of reagents in a gas‐generating needle to achieve a target pressure profile while maintaining structural integrity or the optimal pattern of magnetic particle distribution in a microneedle to maximize deflection for a given field (as hinted by early work using machine learning to predict 3D printing outcomes of magnetic needles) [161, 162].

7.2. Adaptive Control and Personalization

Once a microneedle patch is applied, AI can play a role in real‐time control of the delivery process. Many force‐driven systems (like iontophoretic patches, magnetic actuation, or ultrasound‐driven patches) are interfaced with electronic controllers. Integrating a control algorithm that adapts to patient‐specific feedback enables real‐time adjustment of drug delivery [168]. For example, an AI‐powered insulin patch takes continuous glucose monitor readings and, via a trained ML model, decides when and how much to drive insulin release using iontophoresis or magnetic stirring, aiming to mimic a healthy pancreas's insulin response. These AI‐powered algorithms might be more effective than pre‐programmed protocols because they can account for non‐linearities and patient‐specific kinetics (perhaps learned from that patient's historical data) [169, 170]. In a proof‐of‐concept study, researchers employed reinforcement learning, an AI technique, to modulate the current in an iontophoretic delivery system, successfully maintaining the targeted drug plasma concentration despite variations in skin resistance. Over multiple uses, the system adapts to the skin's behavior (which may vary with factors such as hydration and temperature) and improves its ability to achieve the desired concentration without an overshoot. Another domain of personalization is adapting based on skin variability [171]. Human skin varies widely in thickness, conductivity, hydration, and elasticity. AI models can leverage data such as age, anatomical site, and real‐time sensor data from the patch (e.g., impedance, temperature) to predict optimal force‐driven delivery parameters for each case. As noted in one study, incorporating patient‐specific skin impedance into a machine learning model allowed more accurate prediction of drug permeation rates. According to these forecasts, a smart patch could make adjustments; for example, if it detects that the skin is drier and presents more resistance, the embedded AI algorithm might recommend slightly increasing the iontophoretic voltage or extending the ultrasound duration to offset the change, all while remaining within safe limits [172].

7.3. Integration of Sensors and Big Data

Next‐generation patches will likely incorporate built‐in sensors, for example, microneedle patches capable of measuring interstitial fluid biomarkers, local pH, or temperature. These platforms will generate rich, continuous data streams that can be analyzed by AI algorithms to inform therapeutic decisions and enable broader, proactive health monitoring. AI is useful in extracting patterns from large, multidimensional datasets, for instance, linking subtle changes in local skin temperature and impedance to impending patch‐adhesion failure or early signs of skin irritation [165, 173].

In clinical trials, aggregated data from hundreds of smart patch usage (including drug dose, patient characteristics, and clinical outcomes) can be mined by machine learning models to refine dosing algorithms and reveal new insights, such as identifying which patient subgroups benefit most from a force‐driven delivery strategy. For example, machine learning has already been used to identify effective nanozyme treatments for alopecia delivered via microneedles, essentially screening through candidate formulations by learning which properties correlate with improved hair growth [174]. The same concept can be extended to optimize transdermal delivery strategies. AI could, for instance, determine which combination of penetration‐enhancement modalities yields the best response in different populations: it might reveal that gas‐driven microneedles with a specific bubble‐size distribution are optimal for administering vaccinesfo in certain age groups, whereas other age groups respond equally well to cavitation microneedles operated under specific ultrasound parameters relationships that would be difficult to uncover without large‐scale data and advanced analytics [175].

7.4. Patient Monitoring and Adherence

AI can also enhance the overall user experience. A smart microneedle patch could be paired with a smartphone app that uses AI to coach the patient in real time. For example, if sensor data and delivery logs reveal suboptimal patterns, such as inconsistent patch placement or low hydration affecting performance, an AI‐based digital assistant could provide personalized guidance (“It looks like your skin may have been a bit dry, which can reduce delivery efficiency. Try moisturizing the area 10 min before applying the patch.”) [165]. By analyzing trends and contextual information, AI can distinguish an isolated outlier from a true issue in a way that simple rule‐based systems often cannot.

In addition, AI could contribute to safety by using pattern‐recognition algorithms to identify abnormal impedance or temperature signals that may indicate device malfunction or unsafe operating conditions. If a sensor reading indicates something unexpected, such as a sudden drop in skin resistance that could suggest patch detachment or bleeding, an AI algorithm trained to recognize such patterns could immediately alert the patient or automatically shut off an iontophoretic current to prevent harm [176].

7.5. Challenges and Future Directions for AI‐Integrated F‐MNs

Integrating AI requires rich, high‐quality data and fortunately, modern microneedle patches can generate plenty. However, protecting patient privacy is paramount, especially when dealing with continuous health data from smart patches. This will likely favor secure, privacy‐preserving implementations such as on‐device or edge AI, where most processing occurs on the patch or the paired smartphone [167]. A parallel challenge is regulatory: AI‐driven dosing must demonstrate consistent reliability and safety. This will require extensive training and validation on large, diverse datasets, as well as mechanisms to lock, verify, and monitor algorithms over time. Regulators are increasingly open to AI in medical devices, including frameworks for “adaptive” algorithms, so progress will likely be incremental, for example, closed‐loop insulin patches, as noted, or adaptive analgesic patches that learn a patient's pain patterns and adjust drug delivery accordingly [166].

Looking further ahead, AI‐designed microneedle systems may become so optimized that they enable capabilities currently out of reach, such as non‐invasive delivery of monoclonal antibody therapies or precisely orchestrated multi‐drug regimens from a single patch, with different agents released at different times following an AI‐generated schedule tuned to a patient's metabolism. AI could also determine when a patch is no longer functioning optimally and prompt replacement, for instance by inferring when microneedles have fully dissolved or skin pores have closed [164].

In summary, by intelligently coordinating the mechanical and chemical forces involved in drug delivery, AI transforms microneedle systems from passive depots into active, learning caregivers that adapt to individual physiology and clinical context. This fusion of smart algorithms with advanced biomaterials captures the essence of personalized medicine: treatments that continuously learn and improve for each patient. The early demonstrations highlighted here, from design optimization to personalized control algorithms, are still at an early stage. As datasets expand and algorithms mature, AI may help expand the capabilities in microneedle‐mediated therapy, enhancing efficacy, safety, and patient experience beyond what is achievable with human planning and static device designs alone. These advances, as illustrated in Figure 7, point toward microneedle patches that are not only minimally invasive but also intelligent, adaptable, and patient‐centric drug‐delivery platforms. Taken together, these concepts are summarized schematically in Figure 7, which illustrates the potential roles of AI across the entire microneedle‐based drug delivery pathway.

FIGURE 7.

Applications of AI in microneedle patch technology for enhanced therapeutic delivery. The figure summarizes potential applications of AI in microneedle design and operation. It covers smart adaptive delivery systems, AI‐driven microneedle design optimization, real‐time control mechanisms for drug release, data‐driven insights from collected patient data, precision manufacturing processes, and improved patient support and safety features. The figure was generated by BioRender.

8. Future Outlook and Conclusion

F‐MN technologies are expanding the capabilities of minimally invasive delivery, going beyond the traditional limits of passive microneedle and transdermal platforms. By actively harnessing chemical reactions, acoustic energy, magnetic fields, and electrical currents, these systems can overcome diffusion barriers, improve tissue access, and actively drive therapeutic payloads within skin, wounds, tumors, mucosal tissues, hollow organs, and gastrointestinal sites. The case studies and developments discussed in this review illustrate an exciting future in which painless microneedle‐based systems can administer complex biologics in minutes, smart devices can adapt to patient physiology in real time, and minimally invasive platforms can achieve therapeutic outcomes that once required conventional injections or more invasive procedures.

Before these innovations realize widespread clinical adoption, several challenges and steps remain to be addressed. Safety and regulatory approval are paramount. Each force modality must continue to be validated in larger animal models and human trials to ensure that the short‐term and long‐term effects (including any tissue alteration or immune responses) are acceptable. The early results are very encouraging. For example, minimal inflammation with GDMNs, reversible and mild hyperthermia effects with MAMNs, and low irritation scores with IEMNs. Regulatory authorities are expected to classify many of these as combination products (device + drug), requiring demonstration of the safety and efficacy of both the drug and the device components. Close collaborations between engineers, pharmacologists, and regulatory experts will be essential to streamline this path.

Manufacturing scalability is another consideration. Microneedle patches are being produced by techniques like molding, photolithography, and, more recently, 3D printing. Incorporating layers for magnetic particles, embedded electronics, or enzyme reagents introduces additional complexity. Nonetheless, the field is advancing: for example, 3D‐printing techniques have been improved to rapidly produce high‐resolution needle arrays, and the processes for layering (for multi‐layer needles) are becoming automated. It is plausible that dedicated fabrication lines will be established for different types of smart patches, much like how micro‐electro‐mechanical systems fabrication matured in the semiconductor industry. The path forward will involve continued interdisciplinary collaboration. Material scientists will deliver improved biocompatible materials (e.g., new catalysts or magnetic composites) that enhance performance; mechanical and electrical engineers will refine device designs for reliability and user comfort; and data scientists and clinicians will develop the algorithms and protocols to use these devices optimally in patient care. The cost per patch will need to be reasonable for broad use–materials like sugars, polymers, and even nanoparticles are not inherently expensive at scale, and the additional costs related to electronics are often minimal (small batteries, printed circuit boards, etc.). Gradually, streamlined mass production and increased competition are expected to reduce costs, rendering these advanced patches more affordable for both patients and healthcare systems.

The experience of patients using these devices is expected to be better than current standards. Many of the technologies mentioned are designed to be minimally invasive and easy to use. Instead of patients self‐administering injections several times a day or visiting a clinic for their shots, they could simply apply a patch the size of a postage stamp and let it do its job, possibly with the help of a smartphone app. This approach not only reduces pain and inconvenience but also paves the way for telemedicine and remote healthcare. For instance, a doctor could remotely adjust a patient's smart patch settings through a secure app, fine‐tuning the dosage based on data collected in the cloud, all without the patient needing to visit the clinic–thus establishing a responsive treatment cycle over long distances. Equally significant is the role of patient education and acceptance, as it is crucial for patients to place their trust in and comprehend these devices. Initial trials consistently demonstrate a strong patient preference for microneedle patches over traditional injections, suggesting that with appropriate guidance, their adoption is likely to be seamless.

The smart patches also hold the potential for combining therapies and novel treatment paradigms emerging from these tools. A patch could simultaneously deliver an antibiotic and an anti‐inflammatory gas like NO to an infected wound or administer a chemotherapy drug and an immunotherapy agent deep within a tumor, potentially aided by ultrasound. The ability to control delivery both spatially and temporally could foster synergy, for instance, initially releasing a drug that degrades the extracellular matrix, followed moments later by another drug that targets tumor cells once penetration is enhanced (this could be achieved with a timed ultrasound sequence or sequential iontophoretic pulses). Such orchestrated multi‐drug regimens are challenging with traditional administration methods but are distinctly feasible with a microengineered patch.

From a broader perspective, force‐driven microneedles offer a strategy to address some persistent global health challenges. Microneedle patches are already being vigorously researched for global vaccination campaigns because they can be shipped without cold chain (if stabilized in solid form), potentially applied without trained personnel, and are minimally invasive. Another area is chronic disease management in places with limited resources; a smart patch that releases medications could reduce the need for frequent clinical visits.

In conclusion, the integration of microneedle technology with active force‐generating strategies represents a shift in minimally invasive drug delivery and active therapy. Over the past few decades, microneedles have progressed from laboratory prototypes to clinically relevant platforms for selected applications, including cosmetic delivery and vaccine administration. By equipping microneedles with chemical, acoustic, magnetic, or electrical forces, their practical capabilities can be expanded beyond passive skin permeation to include deeper tissue access, localized therapy, controlled cargo transport, and treatment of anatomically complex targets. Each force modality offers distinct advantages, such as chemical gas generation providing self‐contained power and, in some cases, therapeutic gas release; ultrasound enables deep‐tissue activation; magnetic fields offer wireless and penetrating control; and iontophoresis enables precise dosing of charged therapeutics. Rather than competing, these modalities are complementary tools that future clinicians may select or combine depending on the therapeutic task. F‐MNs are therefore poised to transform microneedle‐mediated therapy by making treatments less invasive, more spatially precise, and more adaptable to individual patient needs. With continued progress, the vision of medicine delivered exactly where and when needed may become an everyday reality across diverse clinical settings.

Author Contributions

All authors contributed to the development of the content for this work, as well as to the writing and revision of the manuscript.

Conflicts of Interest

The authors declare no conflicts of interest.

Data Availability Statement

Data sharing is not applicable to this article as no datasets were generated or analysed during the current study.