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Combined Ultrasound and Microneedle for Enhanced Transdermal Delivery: Synergistic Mechanisms and Therapeutic Advances

Lingling Lei1,2,3,a, Wangrui Peng1,3,a, Jianfeng Cai4,* and Chao Lu5,*

1Key Laboratory of Medical Imaging Precision Theranostics and Radiation Protection, College of Hunan Province, The Affiliated Changsha Central Hospital, Hengyang Medical School, University of South China, Changsha, China

2Department of Medical Imaging, The Affiliated Changsha Central Hospital, Hengyang Medical School, University of South China, Changsha, China

3Institute of Medical Imaging, Hengyang Medical School, University of South China, Hengyang, China

4Department of Chemistry, University of South Florida, Tampa, Florida 33620, USA

5State Key Laboratory of Bioactive Molecules and Druggability Assessment, Guangdong Basic Research Center of Excellence for Natural Bioactive Molecules and Discovery of Innovative Drugs, College of Pharmacy, Jinan University, Guangzhou 511436, China

aThese authors contributed equally to this work.

*Correspondence to: Jianfeng Cai, E-mail: jianfengcai@usf.edu; Chao Lu, E-mail: chaolu@jnu.edu.cn

Received: May 15 2025; Revised: July 14 2025; Accepted: July 24 2025; Published Online: November 18 2025.

Cite this paper:

Lei L, Peng W, Cai J et al. Combined Ultrasound and Microneedle for Enhanced Transdermal Delivery: Synergistic Mechanisms and Therapeutic Advances. BIO Integration 2025; 1–21.

DOI: 10.15212/bioi-2025-0095. Available at: https://bio-integration.org/

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© 2025 The Authors. This is an open access article distributed under the terms of the Creative Commons Attribution License (https://creativecommons.org/licenses/by/4.0/). See https://bio-integration.org/copyright-and-permissions/

Abstract

Transdermal drug delivery (TDD) offers a non-invasive alternative to conventional administration routes, yet the efficacy of TDD is constrained by the impermeable stratum corneum (SC) of the skin, particularly for macromolecules exceeding 500 Da. While microneedle technology addresses this barrier by creating micro-scale channels, the limited penetration depth of microneedles restricts drug delivery to superficial epidermal layers. This review highlights the transformative potential of combining ultrasound and microneedles for enhanced TDD, which synergistically integrate physical disruption and energy-driven permeation enhancement. Unlike material-based or charge/magnetism-dependent strategies, ultrasound leverages multifactorial mechanisms (mechanical stress, cavitation, and thermal effects) to propel drugs through microneedle-generated pathways into deeper tissues. The ultrasound-microneedle (US-MN) system enables spatiotemporally controlled drug release with sonochemical/piezoelectric effects expanding applications in precision medicine. The underlying mechanisms, technological innovations, and clinical translation challenges of the US-MN have been critically evaluated herein, emphasizing the versatility for macromolecules and precision medicine. By bridging mechanisms with translational gaps, this work provides a roadmap for optimizing the US-MN platform, offering researchers actionable strategies to advance TDD for chronic diseases, vaccines, and targeted therapies.

Keywords

Drug delivery, microneedle, transdermal therapeutic strategies, ultrasound, ultrasound-assisted microneedle.

Introduction

Transdermal drug delivery (TDD), an innovative drug delivery system, has emerged as a promising alternative to oral administration and subcutaneous injection [1]. TDD enables non-invasive systemic drug absorption utilizing formulations, such as creams, gels, films, and patches by bypassing hepatic first-pass metabolism and gastrointestinal absorption [2, 3]. These advantages have facilitated the widespread application of TDD in chronic disease management, dermatologic therapies, and novel vaccine development [4, 5]. However, TDD faces persistent challenges due to the inherent barrier properties of the skin. As the largest organ in the human body, the skin exhibits a complex multilayered structure comprising the epidermis, dermis, and hypodermis [6]. The stratum corneum (SC), which is the outermost epidermal layer (10–20 μm thick), serves as a formidable barrier to the penetration of exogenous substances, particularly limiting the effective permeation of macromolecular drugs >500 Da in size, such as gene therapies, proteins, and antigens [7–9]. Overcoming this barrier remains a critical challenge in advancing TDD technologies.

To address this limitation, TDD strategies have evolved significantly. Initial approaches relied on drugs with inherent skin permeability, followed by the incorporation of chemical permeation enhancers to improve the delivery of small molecules. More recently, physical techniques, including thermal ablation, electroporation, ultrasound (US), microneedles (MNs), and combined strategies involving two or more of these techniques, have been developed to enhance the delivery of macromolecules and vaccines [6, 10]. Among the physical enhancement methods, MNs and US have garnered substantial attention in TDD research due to the minimally invasive nature, operational simplicity, and precise local delivery capabilities [11–13]. MNs facilitate drug permeation by creating micron-scale channels that bypass the SC barrier. In addition, US leverages multifactorial biophysical effects, including mechanical, cavitation, and thermal mechanisms, to enhance drug permeation [14].

However, the penetration depth of MNs is typically confined to the epidermal layer, thereby limiting effective drug delivery to deeper dermal or subcutaneous tissues [15, 16]. This limitation stems from the extracellular matrix (ECM), which functions as an additional barrier layer that compromises the efficiency of TDD. Even when drugs successfully traverse the SC, the dense ECM persists as a secondary barrier, hindering access to deeper pathologic regions or subsequent permeation into vascular or lymphatic systems [17].

To overcome this constraint, a strategy of integrating external energy-assisted mechanisms has been explored to augment drug permeation. For example, magnetic field-driven drug targeting and iontophoresis (combining MNs with electroosmotic flow) enable directional transport of magnetic or charged drug molecules, respectively [18, 19]. The requirement for specific drug properties (e.g., magnetism or charge) limits broader applicability. Notably, the combination of US and MNs has shown exceptional potential in achieving effective drug delivery to deeper tissues. US enables drugs traversing MN-created channels to actively diffuse into deeper tissues without requiring drug modifications [13, 14]. Moreover, the combination of US and MNs effectively addresses a critical limitation of US-mediated TDD—its characteristically delayed therapeutic onset. A seminal study demonstrated that standalone US application requires a 2h post-irradiation period to achieve significant lidocaine permeation enhancement. In stark contrast, US-MN co-administration accelerated permeation kinetics by >20-fold compared to US alone, establishing immediate therapeutic efficacy [20]. To date, the combination of US and MNs has been successfully applied across diverse diseases. Unlike conventional synergistic approaches based on mechanistic superposition, the US-MN platform achieves superior TDD efficiency via a relay-enhanced hierarchical permeation mechanism that synergistically combines MN-mediated SC disruption with US-driven active diffusion into deeper tissues [21, 22]. Furthermore, the spatiotemporal precision of US allows on-demand drug release from MN reservoirs. Indeed, the combination could accelerate release kinetics to establish early therapeutic windows and inhibit disease progression [23, 24]. The sonochemical and piezoelectric effects of US further expand the therapeutic application in disease management, offering novel opportunities for precision medicine [25] (Figure 1).

Figure 1 Schematic illustration of the ultrasound-combined microneedle systems for enhanced transdermal drug delivery.

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This review comprehensively examined the US-MN combination for enhanced TDD, focusing on the synergistic mechanisms and technological innovations. Current challenges in clinical translation were critically evaluated and future directions for this synergistic platform are proposed. By introducing a hierarchical permeation framework, this work aimed to provide new perspectives for advancing efficient drug delivery and precision therapeutics in TDD.

MNs: an emerging powerful platform for TDD

Advances in TDD

The advent of the first commercially approved transdermal patch (hyoscine hydrobromide adhesive matrix) in 1979 constituted a pivotal advancement in TDD and catalyzed paradigm shifts in non-invasive pharmaceutical administration [26]. TDD facilitates controlled permeation of active pharmaceutical ingredients through the SC barrier by functioning as a non-invasive delivery modality with subsequent sequential diffusion across the epidermal, dermal, and hypodermal layers. Systemic bioavailability is achieved via capillary uptake into the peripheral circulation, enabling both localized therapeutic targeting and systemic pharmacologic effects, while circumventing gastrointestinal degradation and hepatic first-pass metabolism [5, 7]. TDD exhibits superior pharmacokinetic advantages compared to conventional oral or parenteral routes, including the following: (1) sustained zero-order release kinetics minimizing plasma concentration fluctuations; (2) elimination of gastrointestinal irritation and drug-enzyme interactions; (3) enhanced patient adherence through simplified self-administration protocols; and (4) programmable dose titration via patch surface area modulation [27, 28]. These attributes collectively position TDD as a transformative platform for chronic disease management and biologics delivery with growing adoption across neurology, endocrinology, and oncology therapeutics.

The technological evolution of TDD has progressed through four distinct generations, each marked by paradigm-shifting innovations [29]. First-generation TDD systems relied on passive transdermal transport of lipophilic small molecules but was constrained by inherent limitations in delivering macromolecular or polar therapeutics [30]. Second- and third-generation TDD systems addressed these barriers through reversible disruption of the SC or the application of additional driving forces, such as modifying drug properties or using physical methods to create micropores in the SC, thereby enhancing drug delivery efficiency [31]. The current fourth generation TDD systems integrate bioresponsive electronics with wearable platforms, enabling closed-loop drug release control through real-time biomarker monitoring [30]. These advances have led to an increase in the scope and precision of TDD systems.

Current research in TDD primarily focuses on second- and third-generation systems with an aim to improve skin permeability using a range of strategies, including iontophoresis (electro-osmosis), electroporation (using brief high-voltage pulses to create temporary aqueous channels in the skin) [32], thermal ablation (utilizing heat to enhance skin permeability), lasers (which melt the SC), chemical permeation enhancers, and nanocarriers. Among these strategies, chemical permeation enhancers function by altering the conformation of proteins, lipid fluidity, or the solvent properties of the SC, thereby improving drug permeation. These enhancers have been extensively used in the treatment of conditions, such as Alzheimer’s disease and migraines [33, 34]. Nanocarriers, such as solid lipid nanoparticles (SLNs), nanostructured lipid carriers, and liposomes, benefit from a lipidic composition and interactions with the lipid layers of skin, which results in enhanced drug loading capacity and improved permeability [7, 35]. In addition, an emerging platform – MNs which create micron-scale pores in the SC to enhance drug permeation [7], is gaining increasing attention due to its simplicity, low invasiveness, and transdermal ability [36] (Figure 2).

Figure 2 Advances in transdermal drug delivery.

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MN-mediated TDD

MNs, an innovative method for TDD, were first conceptualized in the 1970s and officially introduced in 1998 [37]. These systems use micrometer-scale needles to create temporary microchannels in the SC, facilitating efficient drug delivery into the dermal layer. This approach improves drug bioavailability and avoids metabolic processes in the digestive system, significantly reducing patient discomfort compared to traditional injections [38–40].

Types of MN systems

MN systems are categorized into the following five MN types: solid; hollow; coated; dissolving; and hydrogel. Each MN type has distinct mechanisms and materials, making MNs suitable for diverse diagnostic and therapeutic applications (Figure 3). Solid MNs, made from materials (such as stainless steel, silicon, or titanium) create reversible microchannels in the skin, which promotes passive drug diffusion [41]. However, the short duration of action, limited to 24 h due to the natural healing process of the skin, restricts the use of solid MNs for prolonged drug delivery [42]. Coated MNs, in which the drug is applied to the needle surface, provide rapid drug release upon skin penetration [43], although limitations related to coating capacity and uniformity are an issue. Hollow MNs, with microfluidic channels, allow continuous drug delivery and have advanced to clinical trials for vaccine delivery and TDD applications. However, the fragile structure and complex manufacturing process of hollow MNs increase the risk of breakage and limit reliability [44].

Figure 3 Modes of action for different types of microneedles.

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Dissolving MNs, which are fabricated from biodegradable polymers through techniques, such as micro-molding, stretch lithography, and 3D printing, have had a significant reduction in cost and an expanding market, which have substantially enhanced patient accessibility [45]. The ability to dissolve within the skin ensures painless administration and precise drug delivery, making dissolving MNs particularly effective for treating deep-seated tumors, including melanoma and breast cancer [46, 47]. Hydrogel MNs represent cutting-edge medical technology and are typically constructed from cross-linked hydrogels, like gelatin methacrylate, hyaluronic acid methacrylate, and PVA-dextran. Upon insertion, these MNs expand, facilitating targeted drug delivery [48, 49]. By modifying the chemical structure of hydrogels, responsive variants can be developed for the treatment of specific diseases [50]. The engineering precision of these MNs allows for remarkable adaptability in therapeutic applications.

Solid and hollow MNs are commonly used for diagnostic purposes in clinical settings, such as interstitial fluid extraction, although material brittleness and biocompatibility are potentially problematic [51]. In contrast, polymer-based dissolving and hydrogel MNs excel in therapeutic applications. Dissolving MNs hold promise for transdermal cell delivery with substantial potential in innovative applications, such as vaccine delivery and chronic disease management [52]. Hydrogels mimic the ECM environment, making hydrogels invaluable in cell culture studies [53]. The microchannel architecture of hydrogels also enables the absorption of interstitial fluid, which offers the potential for biomarker detection and real-time health monitoring [54, 55]. These advances highlight the progress of MN technology towards safer, more adaptable, and clinically versatile applications in precision medicine [37, 47].

Development of MN-mediated TDD

MNs, as a transformative biomedical technology, have shown considerable promise in TDD and disease treatment. However, conventional smooth-tapered MNs face clinical translation barriers due to insufficient mechanical adaptability and limited drug-controlled release capabilities [56, 57]. Recent breakthroughs in smart-responsive systems and bioinspired design strategies offer promising solutions to overcome these challenges. Innovations in MN systems, by leveraging stimuli-responsive mechanisms and biologically inspired architectures, are progressively addressing the limitations of traditional MNs, significantly enhancing drug delivery efficiency and therapeutic outcomes. Moreover, wearable, wireless, and closed-loop MN systems provide patients with more convenient, efficient, and personalized treatment options (Figure 4).

Figure 4 Development of new microneedle osmotic promotion systems for therapeutics.

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The integration of smart-responsive systems with MN technology has enabled the development of stimuli-responsive MNs, creating a powerful platform for spatiotemporally controlled drug delivery. Stimuli-responsive MNs enable spatiotemporal drug delivery by recognizing endogenous or exogenous signals to undergo structural transformations, such as degradation, swelling, rupture, or collapse [58]. Exogenous triggers include US, light, electricity, force, magnetic fields, and temperature. For example, a magnetically responsive MN robot guided by external magnetic fields efficiently delivers macromolecular drugs through intestinal barriers, demonstrating blood glucose regulation in preclinical models [59]. Hydrothermal-responsive MNs (HRMAM), which are composed of polycaprolactone, release docetaxel upon exposure to water and generate localized heat for tumor ablation. This system has been reported to achieve 75.11% and 72.29% tumor inhibition in melanoma and breast cancer models, combining hydrothermal therapy with chemotherapy for deep penetration [60]. Drug delivery based on light-responsive MNs can be categorized into photothermal and photochemical responses [16]. Photothermal MNs use photo-absorbers, such as polydopamine, to convert light energy into heat, triggering responsive stimuli through thermosensitive materials [61]. In contrast, photochemical MNs absorb light energy to break the photosensitive covalent bonds in the MN platform, thereby triggering the release of encapsulated drugs [62]. The design of endogenous stimulus-responsive MNs is primarily based on the physiologic changes in disease sites; commonly studied endogenous stimuli include pH [63], glucose [64, 65], and enzymes [66]. Several reviews have summarized the development of these stimulatory probes, so the progress of stimuli-responsive MNs will not be a focus of the current review.

Multilayer MNs offer unique advantages to meet the need for sequential drug delivery. A Christmas tree-like MN patch encapsulating fluorouracil, leucovorin, irinotecan, and oxaliplatin (FOLFIRINOX) chemotherapeutics mimics clinical dosing schedules through layer-by-layer fabrication. The patch adheres to moist tumor surfaces and releases drugs in batches, enhancing spatiotemporal control [67, 68]. It has been shown that electro-stimulated MN patches, which enable rapid drug penetration (>1.0 mm depth within 5 min) under microcurrent stimulation, while activating immune responses against deep fungal infections [69]. Preserving therapeutic efficacy throughout the drug delivery process is of paramount importance. Frozen MNs (cryoMNs) have emerged as a breakthrough in immune cell delivery, enabling the intradermal transfer of viable dendritic cells (DCs), while preserving cellular integrity and inducing antigen-specific immune responses. It has shown that cryoMNs loaded with ovalbumin-pulsed DCs significantly delay tumor progression in melanoma models, outperforming both intravenous and subcutaneous administration methods. This finding highlights the promising potential for use in cancer vaccine applications [52, 70].

Bioinspired innovations have significantly improved the mechanical adaptability of MN systems. For example, a hydrogel MN system combining polydopamine and poly(ethylene glycol) diacrylate-sodium alginate tips mimics octopus suckers, achieving strong adhesion in both dry and wet environments [71]. Similarly, angle-tunable MNs inspired by cat tongue barbs enable dynamic wound closure under high tension [72]. Crab claw- and shark skin-inspired MNs with dual-anchor mechanisms further reduce insertion resistance in joint wounds [73]. These biomimetic designs provide new perspectives for clinical translation of TDD systems. Pomegranate-inspired dual-chamber MNs significantly increase drug payload capacity for drug delivery enhancement, which co-loads free and particle-bound rivastigmine, increasing drug capacity by 40% and sustaining cognitive recovery in Alzheimer’s disease models [74]. Bio-inspired designs also enable multifunctional applications. Specifically, sea cucumber-inspired MNs incorporating piezoelectric zinc oxide nanoparticles and conductive graphene oxide generate microcurrents to promote neural repair, while reducing muscle atrophy [75]. Stem cell niche-inspired MN arrays promote angiogenesis in diabetic wounds by enhancing adipose-derived stem cell proliferation and VEGF secretion [76].

The growing demand for precision and personalized medicine has driven advances in wearable MN-based sensors. Closed-loop wearable systems utilize MN arrays to detect interstitial fluid biomarkers in real time and deliver therapeutics, demonstrating exceptional potential in diabetes management. A self-powered wearable MN patch incorporating a triboelectric nanogenerator (F-TENG) converts mechanical motion into electrical energy to drive deep tissue delivery of drug-loaded nanoparticles. This system enables rapid drug release in acidic tumor microenvironments and achieves superior chemo-photodynamic synergy in melanoma models compared to conventional MNs [77, 78]. While these innovations have addressed traditional limitations in tissue adhesion, drug penetration, and therapeutic precision, further improvements in treatment efficacy and patient compliance are essential to accelerate clinical translation.

Clinical research involving MN-mediated TDD

MNs have rapidly advanced into clinical applications since the formal introduction in 1998. The first clinical trial using MNs explored use in embryo-assisted hatching [79]. Extensive human studies have focused on MN delivery performance and therapeutic applications since 2008 with cosmetic indications making up approximately 50% of these trials. Radiofrequency (RF) microneedling, which combines MN penetration with targeted RF energy, has become a key technique in aesthetic medicine. Insulated needles deliver RF energy to the dermis, stimulating neocollagenesis and tissue remodeling. Initial studies in 2009 showed RF-induced elastin and collagen production with later trials confirming efficacy in skin rejuvenation and acne scar treatment [80].

Recent advances have explored comparative and combined approaches with lasers. Fractional Er:Glass lasers and RF MNs have similar effectiveness in treating facial acne scars, both of which offer good safety profiles [81]. Synergistic effects have been noted when combining RF MNs with non-ablative 1927-nm thulium fiber lasers (TFLs) for wrinkle reduction [82]. RF MNs also outperform non-ablative 1565-nm fractional lasers (NAFLs) in treating nasolabial fold laxity, although RF MNs may cause more discomfort, requiring protocol adjustments for optimal results [83, 84].

In addition to aesthetics, MN applications are expanding into medical fields, such as alopecia areata, psoriasis, verruca vulgaris, and periorbital aging, offering new therapeutic approaches [85–88], For example, the dexmedetomidine hydrochloride MN patch has received clinical approval and is now authorized for use during preoperative sedation in pediatric patients (CXHL2400157). MNs have also revolutionized vaccination strategies. Early-phase trials with influenza vaccines have shown that low-dose intradermal MN delivery achieved immunogenicity equivalent to full-dose intramuscular injections in 180 adults [89]. This success has extended to rabies and poliovirus vaccines [90, 91]. A 2024 Lancet trial confirmed the role of MNs in pediatric immunization, reporting mild local reactions and no serious adverse events from measles-rubella vaccine delivery via MNs patches, proving safety in children [92]. Additionally, MN-based clinical trials for primary axillary hyperhidrosis, diabetes management, dry eye disease, and localized anesthesia highlight MN versatility, as shown in Table 1.

Table 1 Summary of MN-Based Clinical Trials for Primary Axillary Hyperhidrosis, Diabetes Management, Dry Eye Disease, and Anesthesia

Disease Application Year Delivery/Monitoring MN Key Finding Ref
Intraoperative pain Anesthesia 2022 Lidocaine Roller Microneedle Reduced anesthesia onset time and improved pain control with no adverse effects [136]
IV Cannulation pain Anesthesia 2022 Lidocaine MicronJet600 (NanoPass Technologies Ltd, Israel) 11-fold pain reduction compared to no pretreatment [137]
Intraoperative pain Anesthesia 2016 Lidocaine + Prilocaine Fractional Microneedling Significantly lower pain scores vs. topical anesthesia alone (P < 0.05) [138]
Dental injection pain Anesthesia 2021 Lidocaine Microneedle Patch Reduced pain during dental injections and improved patient comfort [139]
Diabetes Glucose control 2011, 2012, 2015 Insulin 34-gauge 1.5-mm steel microneedle Faster insulin absorption/onset and reduced glycemic variability over 72 hours. [140–142]
Diabetes Glucose monitoring 2014 Glucose Hollow microneedle array 15% MARD accuracy and 72-hour wear feasibility [143]
Diabetes Glucose control 2013 Insulin Single hollow microneedle 22-minute faster onset and reduced insertion pain (P = 0.005) [144]
Primary axillary hyperhidrosis Sweat secretion control 2015, 2020, 2024 N/A Fractional Microneedle Radiofrequency (FMR) Botulinum toxin-A showed higher efficacy and lower pain. Both treatments significantly reduced HDSS scores. Improved HDSS vs. sham (P < 0.05) [145–147]
Dry eye disease Tear secretion control 2022, 2024 N/A Biodegradable Microneedle Acupuncture (BMA) Equivalent efficacy to intradermal acupuncture (improved OSDI, VAS, QoL, and tear secretion; P < 0.05) with no adverse events, demonstrating clinical safety/efficacy for future trials [148, 149]

US-mediated TDD

US, a form of mechanical wave energy, offers a non-invasive, cost-effective strategy for drug delivery with the distinct advantage of providing precise spatiotemporal control. Moreover, the deep tissue penetration capability of US enables new possibilities for disease diagnosis and treatment. US energy attenuates via reflection, refraction, scattering, and absorption during tissue propagation. Penetration depth, which is defined as the distance where intensity decays to 1/e (~37%) of the initial value, is inversely correlated with frequency [93, 94]. US reaches a 3–5 cm depth at 1 MHz, which substantially exceeds optical methods [visible light: < 1 mm; NIR: ~1 cm] [95]. This advantage stems from lower absorption in aqueous and biological media versus light waves, permitting non-invasive precision energy delivery to deep targets.

Mechanisms underlying US-mediated TDD

The initial combination of ultrasound with hydrocortisone ointment for enhanced drug delivery in polyarthritis treatment was pioneered in the 1950s [21]. This US-based method was subsequently extended to the treatment of bone and joint diseases as well as bursitis [96]. Extensive research has demonstrated that US enhances skin permeability and promotes efficient TDD through multiple mechanisms, including thermal effects, cavitation effects, and acoustic streaming effects [97]. Each mechanism has a distinct role in enhancing drug diffusion. The thermal effect of US primarily increases the kinetic energy of drug molecules, thereby enhancing diffusion across the skin. In many studies this thermal effect has been combined with iontophoresis (known as sonophoresis), significantly improving TDD [13]. However, some studies have suggested that the US intensity used in sonophoresis is relatively low with no significant increase in temperature. Therefore, the thermal effect may not be a major contributor to enhancing drug diffusion [98].

In recent years the cavitation effect has been identified as the primary mechanism underlying US-mediated TDD. Cavitation occurs when US induces compression and expansion of liquid within tissues, reducing local pressure below the vapor pressure and causing bubble formation. Cavitation can be categorized into stable cavitation and inertial cavitation. Stable cavitation involves the oscillation of bubbles at lower intensities, causing periodic contraction and expansion, while inertial cavitation is characterized by rapid bubble expansion and collapse with release of significant energy [99]. Both types of cavitation are influenced by US intensity. Studies have elucidated the role of stable cavitation in TDD Krasovitski et al. suggested that US at intensities below the cavitation threshold induces stable cavitation and leads to the formation of pores in the bilayer membrane, which facilitates material transport [100]. Stable cavitation can also increase drug diffusion by generating microflows. The choice of US frequency has a critical role in this mechanism. It has baren shown that the size of cavitation bubbles is inversely proportional to US frequency [10, 101]. High-frequency US can enhance the diffusion rate of salicylic acid by up to 4-fold under the same treatment conditions, while low-frequency US has minimal effects on diffusion [102]. This finding is due to the smaller bubble size under high-frequency US, which allows for more efficient penetration and diffusion within the skin. We hypothesize that freely diffusing bubbles enhance drug permeation through the skin by interacting with drug molecules and facilitating penetration. However, low-frequency US has also been shown to have significant efficacy in enhancing drug delivery, particularly for the transdermal delivery of insulin in diabetic rats [13]. In fact, low-frequency US has been shown to increase the permeation of ethanol and sucrose up to 1000-fold compared to therapeutic US [1 MHz] [103]. The dominant mechanism in low-frequency US is inertial cavitation, in which larger bubbles rapidly collapse in the coupling medium and generate microjets that disrupt the SC, thereby enhancing skin permeability and drug delivery efficiency [104]. Further research by Ueda et al. demonstrated that lower US frequencies produce stronger cavitation effects [105]. Additionally, the size of the pores formed on the skin surface after US treatment correlates with the frequency of the applied US with lower frequencies resulting in larger pore sizes [106] (Figure 5).

Figure 5 Mechanistic contrast between low- and high-frequency US synergized with microneedles.

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Advances in US-mediated TDD

Preclinical studies have substantiated the efficacy of US in enhancing TDD. US has been shown to improve the transdermal delivery of small molecules, such as quantum dots and gold nanoparticles [107, 108]. However, larger molecules, such as peptides and proteins, encounter significant challenges in penetrating the skin barrier due to size. Notably, Park et al. successfully demonstrated the transdermal delivery of insulin using a piezoelectric transducer, showing that this non-invasive method yields a higher effective transdermal delivery of insulin and offering a promising alternative for insulin delivery in diabetic patients [109]. Technological advances have also addressed the limitations of traditional single-frequency US. Recent studies have indicated that dual-frequency US, which combines high and low frequencies, enhances the cavitation effect and leads to a substantial increase in drug delivery efficiency [110, 111]. For example, Polat et al. used a dual-frequency approach with 20 kHz low- and 1 MHz high-frequency US, which resulted in a 3.81- and 13.6-fold increase in the delivery efficiency of inulin and glucose, respectively, across porcine skin [106].

US-mediated transdermal drug delivery has made notable progress in clinical settings. The first clinical applications focused on local anesthesia and in 2004 the FDA approved the SonoPrep® device, which uses US to enhance the skin permeability of local anesthetics, such as lidocaine [13]. This device has been shown to effectively reduce the onset time and enhance the anesthetic effects of lidocaine. The US devices commonly used in clinical trials and the corresponding parameters are summarized in Table 2. US has also facilitated the penetration of dendrimer molecules into human skin. Specifically, dendrimers were detected in the dermis following 10 min of US treatment, whereas the penetration under passive diffusion was minimal [112]. These findings underscore the significant potential of US in clinical TDD.

Table 2 Ultrasound Devices and Parameters Commonly used in Clinical Trials

Device Frequency Intensity (W/cm2) Duty Cycle Participants (n) Drug Delivered Gov Identifier Year Ref
Langevin-type transducer (Transderma) 36 kHz, US1: 2.72 (SPTA), US2: 3.50 (SPTA) US1: 2:5 s
US2: 3:5 s		 10 Histamine N/A 2010 [150]
Langevin-type transducer (Transderma) 36 kHz, 2.72 (SPTA) 2:5 s 15 Betamethasone 17-valerate N/A 2010 [3]
Sonopuls 992 (Enraf Nonius) 1 MHz, 1.0 (SPTA) Continuous 22 Halcinonide N/A 2013 [151]
Sonopuls 992 (Enraf Nonius) 1 MHz, 1.0 (SPTA) Pulsed (1:4) vs. Continuous 93 Lidocaine NCT01404468 2012 [152]
NAVA-01 (Noah Tongzhou) 40 kHz, 300 mW/cm2 (peak), 100 mW/cm2 (pulse avg) Pulsed 110 CM gel NCC2015Z-01 2019 [153]
Carewear LFUS patch 45 kHz, 0.075–0.09 (TA) Continuous 40 Dexamethasone sodium phosphate N/A 2020 [154]
Chattanooga Vectra Genisys Therapy System 1 MHz, 1.5 W/cm2 50% Duty Cycle 40 Dexamethasone sodium phosphate N/A 2020 [154]
NAVA-01 (Noah Tongzhou) 40 kHz, 300 mW/cm2 (peak), 100 mW/cm2 (pulse avg) Pulsed 110 Chinese medicine formula N/A 2020 [155]

While commercial sonicators have proven effective for US-mediated TDD, the ultrasonic probe or converter from a commercial sonicator can weigh ≥ 1 kg. Therefore, developing lightweight, portable US devices is critical for advancing clinical translation. To address this need, Smith et al. proposed using low-profile, lightweight lanthanum-modified lead zirconate titanate (PZT) array transducers (total weight < 22 g) for transdermal delivery. The miniaturized device generated sufficient acoustic intensity to achieve efficient insulin delivery in hyperglycemic rat and rabbit models [113, 114]. Similarly, PZT transducers have been evaluated for efficacy in enhancing insulin delivery in pigs [109]. In addition, the development of wearable US devices has further increased the clinical applicability of US-mediated drug delivery. A stretchable electronic facial mask (SEFM) was developed to synergize sonophoresis with topical drug masks, integrating real-time sensors and actuators for personalized facial healthcare [115] (Figure 6). At the same time Yu et al. engineered a conformable US patch (cUSP) comprised of 4 PZT elements embedded in a polydimethylsiloxane (PDMS) substrate (5 cm diameter). This wearable platform enhanced transdermal delivery of cosmetic agents (e.g., niacinamide) by 26.2-fold in porcine skin models, outperforming MNs and traditional methods through synergistic reduction of skin impedance and enhanced permeation [10]. Despite device miniaturization, Han et al. developed nanobubble ultrasonic coupling agents to address the limitations of conventional ultrasonic coupling agents in TDD. By amplifying US-induced cavitation effects, these agents disrupt the dual barriers of skin (SC and epidermis), enabling stepwise drug penetration into the dermis and achieving deep, non-invasive TDD [22].

Figure 6 Wearable ultrasound-mediated transdermal delivery systems and nanobubble ultrasonic coupling agents. A stretchable electronic facial mask (SEFM) with an ear-hook design conforming to facial contours. The device integrates stretchable island-bridge mesh circuits and piezoelectric arrays to generate ultrasound, which enhances transdermal drug delivery through acoustic cavitation and thermal effects. Figure 6 was reproduced from Ref. [115] with permission from American Chemical Society. Copyright 2025.

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These technological innovations have significantly improved the efficiency of drug delivery. Looking ahead, US-mediated TDD is likely to be integrated with other enhancement techniques, such as electroporation and electro-osmosis, to further optimize delivery efficiency. With continuous advances in device technology, US therapy holds promise as a universal and efficient method for non-invasive drug delivery, particularly in clinical applications.

Combining US and MNs for enhanced TDD

The combination of US and MNs (US-MN) provides an innovative solution to overcome the limitations of conventional TDD. Although MNs effectively penetrate the SC via minimally invasive channels (300–400 μm in depth), the standalone application is constrained by shallow penetration depth and suboptimal drug delivery efficiency to deeper tissues. As a critical modality for enhancing transdermal drug permeation, US solely exhibits limited efficiency in facilitating drug transport across the SC.

The synergistic effects of US-MN enhance therapeutic efficacy through three multidimensional mechanisms: 1) Enhanced drug migration (Microneedles create microchannels through the SC, reducing drug penetration time in US-mediated transdermal delivery. US further enhances drug migration through acoustic streaming and cavitation-induced microjets, improving permeation efficiency.) 2) Relay delivery to deep tissues (The penetrating capability of US enables mechanical forces and thermal effects generated in deeper tissues facilitate post-microchannel drug diffusion, continuously driven drugs to subcutaneous layers.) 3) Spatiotemporal control for intelligent drug release (The spatiotemporal precision of US imparts smart functionality to MNs, allowing on-demand drug release and optimized release kinetics) (Figure 7).

Figure 7 Illustration combining ultrasound and microneedle for enhanced transdermal delivery.

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The US-MN synergy transforms MNs from static carriers into an intelligent “relay delivery” system. Additionally, the MN-created microchannels enhance US-mediated therapeutic efficiency within controlled parameter ranges. This integrated platform demonstrates unprecedented versatility, optimizing treatment windows for local anesthesia and infection control, while offering novel strategies for historically challenging issues, such as bacterial antibiotic resistance and subcutaneous tumor therapy. By combining the anatomic precision of MNs with the dynamic controllability of US, this paradigm pioneers new pathways for effective transdermal delivery (Table 3).

Table 3 Summary of the Ultrasound-Assisted Microneedle Platforms for Enhanced Therapeutics

Synergy Type Mechanism Experimental/System Case Outcome Ref
Enhanced transdermal drug delivery Thermal effects, cavitation effects, and acoustic streaming enhance drug diffusion and permeation Hollow MNs + US improve calcein and BSA delivery; solid MNs pretreatment + US for macromolecules Delivery efficiency increased 9–12×; reduced transdermal time from > 2 hours to 7 minutes (lidocaine) [20, 21]
Enhanced transdermal drug delivery Cavitation accelerates MN dissolution to release antibacterial materials HA MNs loaded with sonocatalytic materials triggered by ultrasound 99.94% antibacterial efficiency (S. aureus) [120]
Applications expansion Piezoelectric materials convert ultrasound energy to electricity for macrophage polarization modulation Piezoelectric MNs for localized electrical stimulation Non-invasive immune cell phenotype regulation [25]
Spatiotemporally controlled delivery Temperature/ultrasound dual-responsive release; ultrasound-activated targeted delivery MSNs grafted with copolymers (temperature/ultrasound-responsive)
MOF-based MNs for CRISPRa-UCP1 delivery (ultrasound-triggered ROS)		 On-demand drug release; localized fat ablation and white adipose browning [23, 24]
Applications expansion Fenton/electrolytic reactions generate ROS/gas for combined therapy Riboflavin MNs + PZY2 electrolysis (shortens ocular axis); Schottky heterostructure MNs (sonodynamic therapy) Enhanced scleral collagen crosslinking; efficient C. acnes elimination [127, 128]
Applications expansion Ultrasound-driven cascade reactions (glucose oxidase → H2O2 → OH → CO release) TiO2–x MNs + MPA–CO prodrug Controlled CO release for wound healing [126]

Combining US and MNs for enhanced TDD

The combination of US and MN enables efficient TDD through multiple mechanisms, involving microchannels, thermal effects, cavitation effects, and acoustic streaming. US-induced acoustic streaming reduces boundary layer resistance in enhancing skin permeability via fluid flow, amplifying drug diffusion through MN-created microchannels that can reduce drug penetration time in SC [20]. Concurrently, cavitation effects generate microjet formation through microbubble (MB) collapse, synergistically improving skin permeability to assist MN-mediated drug penetration. It has been shown that US-enhanced convective flow increased drug permeation through hollow MNs, achieving 9- and 12-fold enhancements in calcein and bovine serum albumin (BSA) delivery efficiency, respectively, compared to passive diffusion [21].

The US and MN combination enables a “relay” for efficient dermal penetration in enhancing drug permeation capabilities. The synergistic effects of US and MNs on TDD are summarized in Table 4. MNs, typically designed at a length of ~1000 μm, exhibit reduced effective insertion depths as a micro-scale transdermal platform due to skin viscoelasticity and insertion force limitations. Nayak et al. reported that MNs, 1100 μm in length, had a penetration depth of only 300–400 μm in ex vivo pig ear skin [20]. Makvandi et al. used optical coherence tomography imaging to evaluate an in situ penetration depth of 850 μm using high MN array patches in ex vivo pig skin and found that the maximum depth only reached ~600 μm [116]. Notably, the peripheral MNs penetrated 27% less than central MNs and further reduction occurred near skin appendages [117].

Table 4 Synergistic Effects of Ultrasound and Microneedles on Transdermal Drug Delivery

Technology Penetration Depth Permeation Enhancement Clinical Applications Ref
Low-frequency US (20–100 kHz) >10 cm (in soft tissue at 10 kHz) Up to 1000-fold for ethanol/sucrose Betamethasone valerate cream (2010)
Histamine (2010)
Anesthetic creams (2004)
TCM formulations (2020)		 [150, 155–157]
High-frequency US (1–3 MHz) 3–5 cm (at 1 MHz) Up to 4-fold for salicylic acid Limited reports [95]
Microneedles 300–600 μm (< designed height) 74.7% payload delivery over 6 h Vaccines (2017, 2022)
Betamethasone (2025)
Retinoic acid (2013)
Insulin (2011, 2016)
Nalmefene (2022)
Diclofenac (2012)		 [141, 158–164]

Given that the epidermis ranges from 200–1400 μm and the dermis ranges from 400–2400 μm in thickness, MNs primarily facilitate traversal of the SC within a constrained depth. Current MN systems have moderately improved drug permeation compared to passive diffusion but the diffusion within the dermal layer remains suboptimal and needs further enhancement. The auxiliary role of US addresses this limitation through spatial relay and kinetic enhancement. Moreover, US-generated thermal effects increase molecular kinetic energy, augmenting MN-driven diffusion rates post-SC penetration (Figure 8).

Figure 8 Mechanisms underlying US-MN platforms for enhanced transdermal delivery.

Next follows the figure caption

The US-MN platforms have demonstrated enhanced therapeutic effects in various applications. Combining MD pretreatment with US accelerated topical lidocaine anesthesia delivery 23-fold versus passive diffusion, achieving therapeutic concentrations in 7 min versus 30 min for US alone. Moreover, MN-created channels are a prerequisite for US enhancement because US alone fails to overcome SC barriers [20]. Therefore, to realize these synergistic effects in US-MN-based TDD, MNs are typically applied before US treatment. Wu et al. demonstrated that a 5-min MN pretreatment followed by a 10-min US application increased the finasteride penetration depth to 5.01 mm, which was 1.89-fold greater than MNs alone with cavitation-induced dermal fissures confirming the “relay” role of US in deep tissue diffusion [118]. It has been shown that US alone concentrates acoustic energy at the skin surface, forming cavitation-induced pores that hinder deeper propagation. As a result, no porous structures are observed in the deeper dermis, underscoring the necessity of MN pretreatment. A piezoelectric-driven MN platform (PDMN), which is composed of MN arrays, piezoelectric ceramics, and copper electrodes, generates low-frequency US with peak acoustic intensity at the MN tips. This finding enables the formation of cavitation-induced pores from the skin surface down to the needle tips. Unlike US- or MN-only groups, which showed limited diffusion in either direction, the PDMN group achieved enhanced drug diffusion both vertically and horizontally [119].

The US cavitation effects assist hyaluronic acid MNs loaded with TiO2 nanoparticles to dissolve rapidly for open wound infections, enabling deep biofilm penetration. This approach achieved 99.94% antibacterial efficacy against Staphylococcus aureus, rapidly clearing biofilms and promoting healing. These results highlight the dual role of US in enhancing MN-driven drug dispersion and amplifying therapeutic outcomes [120].

Piezocatalytic therapy (PCT) is an emerging dynamic therapeutic strategy that relies on US-activated anisotropic piezocatalysts with built-in electric fields [121]. The efficient delivery of piezoelectric nanoparticles (PNPs) is critical for PCT efficacy. A drug penetration depth of ~4.6 mm was observed in an agarose hydrogel model at 30 min using US combined with dissolving MNs (DMNs) for PNP delivery, doubling the depth achieved in the control group (~2.3 mm). The total diffusion depth reached ~500 μm with US-DMN co-treatment in porcine skin experiments compared to < 400 μm for DMNs alone, confirming that US further enhances deep tissue drug diffusion [122]. US-mediated drug delivery also improves cellular drug uptake via sonoporation. Zandi et al. developed an integrated electrochemical stimulator on a microfabricated silicon MN array coated with zinc-oxide nanowires (ZnONWs) to generate localized MBs for enhanced sonophoresis. While US combined with free drug administration reduced viable cell counts to 80%, the addition of MNs caused a dramatic decrease to ~52%, indicating that MN-generated MBs significantly amplify the anticancer efficacy of US-mediated drug delivery [123].

Combining US and MNs for spatiotemporally controlled drug delivery

Physical stimuli, such as light and US, assist MNs in achieving intelligent, responsive drug release. US supports deeper penetration into skin layers compared to light, reaching deep-seated tumor lesions without causing significant tissue damage, while offering superior controllability [124]. To date, MN platforms assisted by US-responsive materials or gene carriers have enabled spatiotemporally controlled drug delivery and therapeutic applications in preclinical studies. As the core transdermal delivery vehicle, MNs strategically integrated with US demonstrate unparalleled potential for precise disease treatment through spatiotemporal coordination and controlled release.

Anirudhan et al. developed a MN-compatible system based on mesoporous silica nanoparticles (MSNs) grafted with copolymer chains of aminoethyl methacrylate and 2-tetrahydropyranyl methacrylate (2-THPMA), which exhibit dual temperature- and US-responsiveness [23]. The 2-THPMA component contains US-cleavable acetal bonds, transitioning from a hydrophobic to hydrophilic properties under US exposure. These MSNs adopt an open conformation below 4°C for maximal drug loading, collapsing at physiologic temperatures (e.g., skin temperature) to seal pores. US assistance triggers copolymer cleavage, enabling on-demand drug release while enhancing MN-mediated skin permeation.

Building on the unique capabilities of US, Li et al. engineered a MN-dominant platform for spatiotemporally controlled gene therapy targeting obesity. The system encapsulates CRISPRa-UCP1 plasmid-loaded acoustically sensitive nanosystems into hyaluronic acid MNs (MN-MP/aUCP1), creating an US-assisted weight-loss platform. MNs deliver MP/aUCP1 to the subcutaneous fat layer, while US irradiation assists in triggering plasmid release from MNs and enhancing intracellular transfection in adipocytes via sonoporation, a process in which US-induced MB cavitation generates transient pores in cell membranes. This MN-driven strategy, when augmented by US, significantly upregulates UCP1 expression to promote white adipose tissue browning. US simultaneously assists in precision-triggered localized reactive oxygen species (ROS) generation to eliminate excess adipocytes, synergistically improving anti-obesity efficacy [24].

Combining US and MNs for expanding therapeutic applications

The combination of US and MNs extends the therapeutic applications of transdermal delivery. While MNs excel in treating superficial conditions, bacterial-mediated diseases (e.g., infected wounds and acne) face challenges due to antibiotic resistance. To address this finding, MNs establish transdermal pathways for ROS-generating materials, the therapeutic efficacy of which is enhanced by US assistance. Xiang et al. demonstrated that MNs deliver metal-organic frameworks (MOFs) to the skin, while US assists in generating ROS to eradicate Cutibacterium acne for acne treatment [125]. Similarly, Cao et al. developed a MN platform loaded with a sonosensitizer (TiO2–x), glucose oxidase, and an ·OH-responsive CO prodrug (MPA–CO). US assists by triggering Fenton reactions to convert H2O2 into ·OH radicals, which subsequently release CO gas, a cascade mechanism enabling precise diabetic wound therapy [126].

MN systems further benefit from physical properties of US for advanced applications. Hu et al. integrated Schottky heterostructures into MNs, in which US assistance amplifies electron transfer during Fenton reactions, generating singlet oxygen for sonodynamic tumor therapy [127]. Wang et al. engineered MNs with US-assisted charge delivery in piezoelectric applications. Piezoelectric MNs generate electrical signals that polarize macrophages in the dermis when activated by US, offering a non-invasive strategy for lymphedema management [25].

MNs serve as the primary vehicle for sustained riboflavin delivery to strengthen the sclera for ocular disorders, like high myopia. Zhong et al. enhanced this approach by combining riboflavin-loaded MNs with US-electrolyzable PZY2 material. Specifically, US assists in triggering PZY2 electrolysis to generate gas bubbles that dynamically shorten the axial length and powering MN-emitted blue light to crosslink scleral collagen. This MN-driven, US-assisted system achieves dual therapeutic structural reinforcement and dynamic ocular remodeling, providing an innovative solution for myopia-related complications [128].

Challenges and future

The integration of US and MNs has overcome the limitations of conventional TDD in terms of efficiency, penetration depth, and controllability, offering a novel multidisciplinary solution for therapeutic delivery. This synergistic approach not only significantly enhances drug permeation across biological barriers but also expands the application scope of transdermal technologies to complex indications, such as metabolic disorders, infectious diseases, oncology, and ophthalmology through precise spatiotemporal control, thereby advancing transdermal systems from localized therapy to systemic intervention.

However, this combination faces multifaceted challenges in achieving clinical translation. From a drug delivery perspective, the loading characteristics of this platform constrain utility. Conventional approaches favor lipophilic drugs or nanocarriers for enhanced efficiency [129]. Although MNs accommodate diverse payloads (small/large molecules, nucleic acids, nanoparticles, extracellular vesicles, and cells) [130], the interstitial fluid-filled channels create an aqueous microenvironment favoring hydrophilic agents. Hydrophilic nanoparticles exhibit 2.6-fold higher 48h cumulative transdermal permeation than hydrophobic counterparts. Confocal laser scanning microscopy revealed that hydrophilic rhodamine B diffused to a depth of 190 μm, markedly deeper than hydrophobic fluorescein isothiocyanate (FITC), which was confined to 130 μm within the MN channels. This difference is attributed to the covalent binding of FITC to skin proteins [131]. Furthermore, physicochemical properties (e.g., smaller size and negative charge, which leverage the negative surface charge of skin) significantly enhance intradermal diffusion [132]. Mechanically, aligning US parameters (frequency and intensity) with MN geometry is critical to prevent structural compromise or premature drug release, while miniaturizing wearable hybrid devices demands sophisticated engineering to maintain performance within energy-efficient, portable frameworks [21]. The use of Lead-Zirconate-Titanate (PZT) in US transducers raises biocompatibility concerns, which present potential lead leaching risks with adverse implications for human health and environmental safety [133]. Additionally, residual organic solvents, such as dimethyl sulfoxide and chloroform, which are often used during MN fabrication, may cause local tissue irritation and raise concerns about long-term biocompatibility [134]. Although platforms, like the PDMN developed by Chen et al., demonstrate favorable safety profiles (MN channels close within 5 min with skin recovery complete within 30 min at a 114-kHz US frequency). However, the treatment area is limited to 5×5 mm, restricting the ability to cover large or multifocal lesions [119]. High manufacturing costs and limited scalability of precision-engineered systems further hinder widespread adoption [134, 135].

Future advances should prioritize intelligent systems developed through interdisciplinary collaboration, such as real-time biosensors (e.g., pH and glucose) for dynamic adjustment of US intensity and drug release with biodegradable wireless designs for painless chronic disease management. AI-driven personalization of MN geometry and US parameters based on individual skin biomechanics or disease profiles may optimize efficacy. Expanding beyond small molecules to mRNA vaccines and microbiome-modulating probiotics broadens the therapeutic scope, while sustainable manufacturing using 3D/4D printing addresses ecologic concerns. Collectively, these innovations aim to transform US-MN platforms into scalable, patient-centric solutions for precision medicine, bridging engineering, pharmacology, and clinical science to redefine next-generation drug delivery paradigms.

Data availability statement

Data sharing is not applicable to this article because no new data were created or analyzed.

Ethics statement

No direct interactions with human or animal subjects were involved. Therefore, ethical approval and informed consent were not required.

Author contributions

LL: Writing – review & editing, Writing – original draft, Conceptualization, Investigation, Methodology. WP: Writing – review & editing, Writing – original draft, Investigation, Conceptualization, Methodology. JC: Writing –review & editing, Supervision, Resources, Funding acquisition, Project administration. CL: Writing—review & editing, supervision, resources. All the authors read and approved the final manuscript.

Funding

This work was supported by the National Natural Science Foundation of China (Grant nos. 82272028, 81971621, and 82102087), the Key R&D Program of Hunan Province (Grant no. 2021SK2035), the Natural Science Foundation of Hunan Province (Grant nos. 2022JJ30039 and 2022JJ40392), the Natural Science Foundation of Guangdong Province (Grant no. 2021A1515011177), and the Project of Science and Technology Innovation of Hunan Province (Grant no. 2021SK51807).

Acknowledgments

Not applicable.

Conflict of interest

The authors declare that there are no conflicts of interest.

Graphical abstract

Next follows the graphical abstract

Highlights

  • Microneedles (MNs) facilitate transdermal drug delivery by creating microchannels through the stratum corneum barrier.
  • Ultrasound (US) enhances transdermal delivery through cavitation and thermal mechanisms, enabling deep tissue penetration.
  • US-MN integration enables hierarchical drug permeation via relay-enhanced synergy, facilitating modification-free macromolecule/vaccine delivery.

In brief

This review introduced the US-MN platform, which synergistically integrates MN-mediated SC disruption with US-driven active diffusion to overcome extracellular matrix barriers and enable deep tissue delivery. This hierarchical permeation facilitates rapid, modification-free transdermal delivery of macromolecules and vaccines with spatiotemporal control. Despite the therapeutic promise, clinical translation faces material biocompatibility, device miniaturization, and regulatory standardization hurdles.

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