Transdermal theranostics

Skin offers an easily accessible site for drug administration as well as for health signal monitoring, with non‐invasiveness or minimal‐invasiveness, convenience, and good patient compliance. Transdermal theranostics promises for personalized, home‐based, and long‐term management of chronic diseases, and is expected to change the landscape of healthcare profoundly. In this article, we review the recent advances in transdermal drug delivery, diagnosis based on sweat and skin interstitial fluid, and wearable devices. The advantages, limitations, and commercialization of these emerging techniques are comparatively discussed.


INTRODUCTION
Human skin, the largest organ, accounts for ≈16% of total body weight and covers a surface area of ≈2 m 2 . [1] It is composed of epidermis, dermis, and hypodermis. Depending on the anatomical site, age, sex, and body mass index, thickness of these three tissue layers varies, roughly ≈0.1 mm, 1.5-2 mm, and 7-20 mm, respectively. [2] The epidermis, the outermost layer of skin, is comprised of keratinized, stratified squamous epithelium, and consists of five layers, from the most superficial layer of stratum corneum (SC) to stratum lucidum, stratum granulosum, stratum spinosum, and stratum germinativum (also called stratum basale). [3] The dry and keratinized SC layer is the major physical barrier between the body and environment. The ≈20 µm thick SC layer is made of keratin-rich corneocytes embedded within the lamellarstructured lipid matrix. In the stratum spinosum, dendritic cells (also called Langerhans cells) are present and can function like macrophages to fight against invading microbes. [4] These cells can also act as antigen-presenting cells for both CD4 + and CD8 + T cells. Hence, skin dendritic cells are a delivery (TDD) has several advantages over conventional systemic administration methods. [9] It circumvents the limitations suffered by oral administration, such as gastrointestinal tract absorption and degradation, hepatic first-pass metabolism. In contrast to parenteral drug administration using conventional hypodermic needles, TDD is usually painless, non-invasive or minimum-invasive, self-administrable, and thus ensures good patient compliance. In addition, TDD approach can provide sustained drug release without high dosage shock and with high bioavailability. Consequently, therapeutic efficacy is improved and toxicity is reduced. Overcoming SC barrier is the main challenge for TDD. Various drug formulations and delivery methods have been developed.
Skin tissue is a volume conductor that allows bioelectricity originated from muscle, heart, and nerve to be measured on the skin surface. Various wearable electronic sensors have been developed to monitor bioelectric signals such as electroencephalogram (EEG) and electromyogram (EMG) signals. [10] The skin interstitial fluid and sweat are regarded as the alternatives to blood for biomarker detection. Sweat, which is a hypotonic fluid secreted from eccrine glands (≈500-700 mL/day), contains various biochemical markers, such as electrolytes (eg, potassium, sodium, and chloride ions), metabolites (eg, glucose, lactate, and pyruvate), proteins, amino acids, and xenobiotics (such as drug molecules). [11] They are correlated with diseases or pathological conditions. Therefore, sweat renders a convenient source for non-invasive and continuous monitoring of biochemical signals. Another information-rich bio-fluid present in the skin is the cutaneous interstitial fluid (ISF). Recent studies have shown that ISF has similar profiles of small molecules, metabolites, proteins, and RNAs as blood. [12] It contains biomarkers that are not only comparable to blood (eg, exosomes, memory T cells), but also highly relevant for skin disorders (eg, melanoma). [13] This review summarizes the emerging technologies on transdermal drug delivery, and transdermal diagnosis based on biomarker detection from sweat and ISF as well as bio-signal detection by wearable devices.

PERMEATION ENHANCERS
Skin, particularly the brick and mortar like structure of SC layer, is the primary barrier for TDD. [14] Only a few molecules like low molecular weight (<500 Da) lipophilic ones can pass through the intercellular lipid matrix of SC. [15] Chemical permeation enhancer (CPE) improves drug permeation into the skin through several mechanisms, such as inserting amphiphilic molecules into the SC lipid bilayers to disrupt their highly ordered molecular packing, extracting lipids to create nanoscale defects in lipid bilayers, improving the fluidity of lipid bilayers, or increasing the thermodynamic activ-ities of drug molecules. [16] Many CPEs have been developed, including fatty acids, alcohols, glycols, terpenes, sulphoxides, tensides, laurocapram, pyrrolidones, surfactants, urea, etc. [17] CPEs have advantages, including flexible formulation chemistry and easy application even on large surface area. However, it is challenging for CPEs to deliver hydrophilic macromolecules. In addition, CPEs often do not selectively act on SC lipids, therefore causing skin irritation or injury. [18] Karande et al proposed that safe and effective enhancers should enhance Ch2 symmetric stretching of the SC bilayer lipids, without altering the amide I band or denaturation of SC proteins. [19] The study suggested that compounds with long and saturated carbon tails or multiple aromatic rings are desirable. Some studies have demonstrated that synergistic combinations of CPEs (SCOPE) can enhance potency, safety, and macromolecular permeability. [20] In vitro skin impedance guided high-throughput (INSIGHT) screening was carried out using the known CPEs. [20b] Among ≈5000 SCOPE formulations based on ≈500 binary paring of individual CPEs with different chemical compositions and concentration, the study found that the combinations of N-lauroyl sarcosine: sorbitan monolaurate and Na laureth sulfate: phenyl piperazine can increase the flux of macromolecules (heparin, leutinizing hormone releasing hormone, oligonucleotides, leuprolide acetate) by 50-100-fold, without inducing skin irritation (tested on hairless rats). Exploration of new CPEs and their combination remains an active area of transdermal research.
Some peptides can act as permeation enhancers for TDD. A study by Chen et al reported that a short synthetic peptide (ACSSSPSKHCG), identified by screening through in vivo phage-displayed peptide libraries, facilitated transdermal insulin delivery through the intact skin, particularly via hair follicles. [21] A natural pore-forming peptide, magainin, was also found to be effective in increasing skin permeability, particularly when being synergistically combined with a surfactant enhancer. [22] It was shown that the combination of magainin and N-lauroyl sarcosine (NLS) in 50% ethanol increased human skin permeability by 47-fold, while NLS-ethanol only increased 15-fold. [22a] Recently, several cell-penetrating peptides (CPPs) have been proposed as novel transdermal penetration enhancers. [23] The water soluble, short cationic or amphipathic CPPs can carry electrostatically or covalently bound macromolecules across cellular membranes, without relying on any receptors or causing membrane damage. Rothbard et al reported that arginine oligomers conjugated to cyclosporin A led to increased topical skin absorption in mouse and human skin. [24] Several other peptide CPPs (eg, Tat peptide, Pep-1 peptide, penetratin) have been demonstrated for transdermal delivery of small molecule drugs, proteins, peptides, nucleotides, and nanoparticles (NPs). [25] There are several review articles on CPPs [23,25b] and CPEs. [17a,b] F I G U R E 1 Micro and nano delivery systems for transdermal drug delivery, [36] Copyright 2017 Elsevier. Reprinted with permission from Elsevier [36]
As microemulsion contains a variety of components, and the effect of each component is concentration dependent, the choice and optimal concentration of each component need to be carefully chosen and validated for each drug formulation. Using microemulsion, several studies demonstrated transdermal delivery of various therapeutic compounds, including tacrolimus, retinoic acid, 5-fluorouracil, triptolide, ascorbic acid, lidocaine, prilocaine hydrochloride, and many nonsteroidal anti-inflammatory drugs. [30][31] These studies showed that microemulsion formulations could provide higher drug permeation (≈2-3 times) and bioavailability when comparing with the conventional gel or ointment. Antifungal and antibacterial agents have been encapsulated into microemulsions for topical application in order to increase skin permeation and bioavailability, and subside the side effects associated with nontarget accumulation. For examples, microemulsion containing amphotericin B for invasive fungal infection, microemulsion containing itraconazole and clotrimazole for sporotrichosis. [32] Using microemulsion droplet, transdermal delivery of collagen and hyaluronic acids (HA) were also reported recently. [33] Simple and low-cost preparation, high solubilization for both hydrophilic and hydrophobic drugs, and improved drug permeation make microemulsion attractive for TDD. [30][31] However, local skin irritation remains as a risk for usage of microemulsions. And most microemulsions have very low viscosity, which restricts their application. Recent studies have introduced hydrogel to formulate microemulsions in order to increase viscosity, stability, spreadability, and drug loading capacity, as well as to improve skin tolerability and permeability. [34] Various polymers have been used as gelling agent to formulate microemulsion hydrogel, including natural polymers (eg, collagen, gelatin, HA, various gums, and polysaccharides), semi-synthetic polymers (eg, carboxymethyl cellulose, CMC) and synthetic polymers (eg, carbopols, poloxamers, polyacrylamide, polyvinyl alcohol). [34,35]

NANOCARRIERS
To improve skin impermeability, supramolecular structures that function as both carrier and permeation enhancer have been devised [36] (Figure 1). Liposomal vesicles are one of them and well-studied for TDD, because these vesicles not only increase skin permeability, but also are able to incorporate hydrophilic and lipophilic drugs within their aqueous cores. [37] Liposomes contain one or more phospholipid bilayers. The lipid composition determines the thermodynamic properties of liposomes, which in turn defines their interaction with the skin. Normally, liposomal bilayers are rigid and in gel state, which can even lead to reduced drug permeation. [38] Phospholipid bilayers can be made more flexible and deformable, and liposomes in such a liquid crystalline state can enhance transdermal drug penetration. [39] Through proper formulation, highly deformable and stable liposomal vesicles have been synthesized. For instance, invasomes that contain small amounts of ethanol and terpenes or terpene mixtures have been demonstrated with flexible and fluidic soy-phosphatidylcholine-based phospholipid bilayers. [40] By fluidizing and disturbing SC lipid bilayers, invasomes offer effective transdermal delivery of both hydrophilic and lipophilic drugs. Another type of highly transformable liposomes is called transfersomes, which exhibit ultra-flexibility (several times more elastic than other vesicular systems). [41] Transfersomes are usually composed of phospholipids and surfactants (such as Tween80, Span60), and can rapidly penetrate skin through the intercellular lipid pathway. Transfersomes have been utilized to deliver proteins, peptides and DNAs. Another novel liposomal vesicle is called ethosomes because of their high percentage of ethanol. [42] Ethanol compromises the lipid packing as well as decreases the transition temperature of phospholipid bilayers, making ethosomes highly deformable and penetrable comparing to the conventional liposomes. Niosomes are formed by self-assembling of one or more nonionic surfactants (such as sorbitan esters, polyoxyethylene alkyl ether), with or without lipids (such as cholesterol, soy-phosphatidylcholine). [43] Niosomal surfactants can disrupt SC lipids, allowing niosomes to deliver a variety of drugs into skin. These biocompatible and biodegradable molecules form liposome-like vesicles as effective drug carriers for TDD. However, colloidal and chemical instability, and high cost limit their usage in real applications.
Dendrimers, the highly branched and star-shaped macromolecules, were proposed as a novel drug carrier and permeation enhancer, because of excellent monodispersity and solubility, a high degree of surface functionalities, copious internal cavities for drug loading, and strong interactions with phospholipids. [44] They can be synthesized by stepwise addition of branches in each generation around the core. Drugs can be loaded in dendrimers via hydrophobic encapsulation, ionic interaction, or direct conjugation. Dendrimers improve drug permeation by extracting lipids to create nanopores in SC lipid bilayers. [45] Different dendrimers, including poly(amidoamine) (PAMAM), poly(propylene imine) (PPI), or poly(L-lysine) dendrimer, have been demonstrated for transdermal delivery of various small molecule drugs, including indomethacin, ketoprofen, diclofenac, 5-fluorouracil, tamsulosin, riboflavin, etc, without causing skin irritation or injury. [46] Dendrimers are however not biodegradable. Cell toxicity associated with dendrimers are well documented, particularly the higher generation dendrimers and dendrimers with positive charges on the surface. [47] To solve the biocompatibility issue, dendrimers have grafted with different chemical moieties, including PEG, carbohydrates, and acetyl groups. [48] Biodegradable and biocompatible dendrimers have been recently synthesized for controlled and targeted drug delivery, for examples, polyester dendrimers, polyacetal dendrimers, and peptide dendrimers. [49] A recent study revealed that arginine-terminated peptide dendrimers (carrying varying charges, 4+, 8+, and 16+) could considerably increase transdermal permeation of ketoprofen into mouse skin, compared to diffusion. [50] Different types of nanostructured lipid carriers are explored for TDD. [16a] For example, solid lipid NPs that are composed of a crystalline solid core with high melting point lipid (eg, glycerides, waxes, etc) and coating with lipid monolayers or surfactants. [36] Such biodegradable and biocompatible nanocarriers can improve solubility, avoid degradation, facilitate skin penetration, and control release profile of both hydrophilic and hydrophobic drugs. Studies have demonstrated transdermal delivery of various drug-loaded solid lipid NPs, including meloxicam, [51] acyclovir, [52] diclofenac, [53] triptolide, [54] and Ivermection. [55] Other colloidal lipid nanocarriers including cubosomes and hexosomes are also explored for TDD. [56] For example, topical application of cubosomes loaded with capsaicin showed similar skin permeation, but increased skin retention of drug, in comparison to the commercially available cream. [57]

SONOPHORESIS
Sonophoresis technique uses ultrasound waves to temporarily increase skin permeability up to several hours, through acoustic cavitation effect on skin. [58] The mechanical vibration waves produced from piezoelectric crystals can create cavitation, that is, nucleation as well as expansion, contraction, or distortion of gaseous bubbles, in a liquid medium during acoustic pressure cycles. [59] In stable cavitation state, microstreaming caused by continuous oscillation of small bubbles in response to relatively low acoustic pressure can generate hydrodynamic shear stress. [60] In contrast, inertial cavitation occurs when oscillating bubbles expand beyond the equilibrium radius and eventually collapse, depending on the acoustic pressure intensity and bubbles size. [61] It then generates shock waves (when symmetrical bubbles collapse) or microjets (when asymmetrical bubbles collapse). [61] Given that bubbles are difficult to oscillate and grow in the densely packed skin tissue, liquid medium, such as hydrogel, is usually applied between the transducer and skin. The mechanism of acoustic cavitation on the skin layers, particularly SC, involves in several aspects, including structural modification or disruption of SC lipid bilayers and formation of microchannels. [62] Cavitational ultrasound is superior to conventional non-cavitational ultrasound, because of its ability to focus energy at the target site without causing tissue heating and injury. [63] Sonophoresis-assisted TDD has been reported, using various ultrasound parameters, such as frequency (eg, 20 kHz to 3 MHz), intensity (eg, 100 mW cm 2 to 7-8 W cm 2 ), sonication time (eg, 1 min to 4 h), for a wide range of therapeutic compounds, including caffeine, diclofenac, fentanyl, heparin, insulin, ketoprofen, steroids, salicylic acids, and lidocaine. [58,63a] For example, sonophoresis (20 kHz, 5 min) was employed to enhance diclofenac permeation across EpiDerm 3D-skin by fivefold. [64] Similarly, sonophoresis (20 kHz) was shown to enhance transdermal delivery of rivastigmine by 3.1-fold, generating approximately three times enhancement of plasma drug concentration in pigs, comparing to topical application. [65] It was also found that sonophoresis (40 kHz over 60 min) enhanced transdermal delivery of insulin, and thereby provided a much better glucose lowering effect on rats. [66] Recently, a double-blind, randomized clinical study found that patients with acute low back pain treated with ultrasound-assisted delivery of diclofenac and thiocolchioside showed good therapeutic outcome. [67] Studies have shown that dual-frequency sonophoresis could provide better TDD performance without producing thermal effect, due to the enhancement of cavitation activity. [68] For example, application of dual-frequency sonophoresis (20 and 800 kHz) enhanced the permeability of sinomenine hydrochloride into porcine skin, as compared to a single frequency sonophoresis. [69] Dual-frequency sonophoresis uses a low-frequency transducer perpendicular to the surface together with a high-frequency transducer at a 90 • angle to it. [68] Combined usage of sonophoresis with other TDD techniques like microneedles (MNs) and electroporation has also been explored. It was found that combination of sonophoresis (20 kHz for 2 min) and MN application dramatically enhanced the permeation of hydrophilic macromolecule (FITC-dextran) into porcine skin, comparing to MNs application alone. [70] Combination of sonophoresis, MN, and electroporation (pulse voltages of 50-300 V) provided the best skin permeability, specifically, approximately two-, four, 35-, seven, and 100-fold higher than MN plus electroporation, MN plus sonophoresis, electroporation plus sonophoresis, MN alone, and sonophoresis alone, respectively. [71] Similarly, a synergistic effect of sonophoresis (20 kHz, pulsed for 1.5 min) and stainless steel MN roller was demonstrated on the transder-mal delivery of siRNAs and protein (ovalbumin) into porcine skin, with approximately fivefold enhancement as compared to MNs alone, and ≈7-15-fold as compared to the negative control. [72] Likewise, simultaneous treatment of sonophoresis (280 and 350 kHz) and iontophoresis (200 µA cm 2 ) significantly enhanced penetration of hydrophilic drug (glutamic acid) into porcine skin by 2.4-fold. [73] Recently, Liu et al studied the combined application of sonophoresis and iontophoresis on penetration of lipid-soluble drug oxaprozin using surface-enhanced Raman spectroscopy. [74] It was found that combined application of sonophoresis and iontophoresis or Azone (a chemical enhancer) synergistically improved drug penetration by 6.1-and 8.4-fold, respectively.
Sonophoresis-assisted transdermal deliveries of nanoparticles (NPs) has been investigated. In a recent study, almost 10-fold permeation enhancement was obtained when ketoprofen-A16 dendrimer complex (an arginine-terminated peptide dendrimers having 16+ charges) and sonophoresis were applied simultaneously on mouse skin, comparing to sonophoresis alone (20 kHz, pulse for 30 min). [75] Permeation enhancement was even greater (≈26-fold) when dendrimer complex was pretreated for 2 h before sonophoresis was applied. In contrast, sonophoresis alone (20 kHz, continuous for 2 min) failed to enhance liposome penetration into porcine skin, and even lowered the penetration of solid lipid NPs and niosomes, likely due to the breakage of those lipid nanocarriers by the sonication energy. [75] The potential use of ultrasound contrast agents (UCA) in sonophoresis is also examined for efficient sonophoresisassisted TDD, because actively induced cavitation with lowintensity ultrasound (<1 MPa) can efficiently disrupt SC lipid bilayers and form aqueous channels in SC. [76] The UCA are small gas-filled microbubbles (≈3 µm) typically used for ultrasound imaging. It was reported that the presence of UCAs largely increased the cavitation activity of sonophoresis, and skin permeability. [76b] Using the commercially available UCAs (Definity and SonoVue; the lipid-coated microbubbles filled with octafluoropropane and sulphur hexafluoride gas, respectively), a concentration of 1:1000 UCAs was found to yield the maximum enhancement of sonophoresisassisted porcine skin permeation of glycerol and ferulic acid (1.1 and 2.4 MHz, pulsed for 10-30 min), that is, 3.1-fold higher than sonophoresis without UCAs, and 7.5fold higher than negative control (no sonophoresis). Different types of engineered microbubbles have been synthesized to augment sonophoresis-assisted TDD, for examples, airfilled albumin-shelled microbubbles, [77] air-filled lysozymeshelled microbubbles, [78] and nanocups that are cup-shaped polymeric NPs with gas bubbles trapped in their cavities. [79] Recently, perfluorohexane-based liquid-core cavitation seed was utilized as microbubbles, because they can sink by gravity and thus position near the skin surface for better cavitation activity. [80] It was found that sonophoresis with such cavitation seeds could increase transdermal delivery of 150-kDa dextran by 6.4-fold, with a penetration depth of 500 µm, compared to sonophoresis without seeds (1 MHz, pulsed for 20 min).

IONTOPHORESIS AND ELECTROPORATION
Iontophoresis method uses a continuous low-voltage electric current to provide an electrical driving force for transport of molecules across SC. [81] Charged molecules can be driven into the skin via electrophoresis or electrostatic repulsion between drug-loaded electrode and counter electrode, while uncharged or weakly charged molecules are moved via electro-osmosis because of liquid movement induced by an applied potential. Iontophoresis-assisted TDD depends on many factors, such as drug charges, molecular weight and concentration, electrode configuration and shape, current density and duration, etc. [81,82] An advantage of iontophoresis is that it can provide programmed drug delivery. Because lowdensity current (usually a direct current ≤ 1 mA/cm −2 ) used in iontophoresis cannot affect SC layer itself, this non-invasive method is usually limited to small molecules up to a few thousand Daltons. The maximum delivery rate is restricted by the electric current tolerable for the patients without producing skin irritation and pain. Iontophoresis devices usually require high power and repeated battery charges, which makes them difficult for miniaturization and continuous monitoring. Highly conductive carbon nanotubes (CNT) have been used as the active element for highly efficient electrophoretic pumping of drug at low power, with variable and programmable delivery rates, and fast fluid flow. [83] It was shown that a transdermal nicotine delivery device based on CNT membranes was able to switch between high and low fluxes of delivery with minimal skin irritation. [84] Recently, self-powered iontophoretic TDD systems have been developed. Ogawa et al developed organic iontophoresis patch with built-in enzymatic biofuel cells (Figure 2A). [85] The patch consists of two enzyme-modified carbon fabrics (CF) (fructose dehydrogenase-modified CF anode and a bilirubin oxidase-modified O 2 -diffusion CF cathode), connected with a conducting polymer-based internal resistor. The hydrogel films containing biofuel fructose and drug molecules are placed directly under the anode and cathode. The patch, when mounted on skin, can generate a transdermal ionic current to produce osmotic flow from anode to cathode, thereby administrating drug into skin. Ouyang et al developed on-demand TDD system driven by a triboelectric nanogenerator (TENG) as the motion sensor and energy harvester that can convert biomechanical motions into electricity for iontophoresis ( Figure 2B). [86] The TENG basically consists of a poly(tetrafluoroethylene) (PTFE) film sandwiched between

F I G U R E 2
Iontophoretic transdermal drug delivery systems. (A), Organic iontophoresis patch with built-in biofuel cell, [85] Copyright 2015 WILEY-VCH. (B), Self-powered, on-demand iontophoretic transdermal system driven by triboelectric nanogenerator, [86] Copyright 2019 Elsevier. Reprinted with permission from WILEY-VCH [85] and Elsevier [86] two radial-arrayed copper films, one is rotatory and the other is static. The iontophoresis drug patch contains screen-printed drug-loaded electrode and counter electrode, and polypyrrolebased drug reservoir film. It was shown that manually rotating the TENG (30-40 rpm) for 1.5 min can release drug from the patch at 3 µg/cm 2 . Similarly, Wu et al developed a wearable TDD system driven and regulated by biomechanical motions. [87] The system contains a PTFE and aluminiumbased TENG that is electrically connected to a hydrogel-based soft patch with side by side electrodes designed to enable TDD.
On the other hand, electroporation technique uses high voltage pulses to temporarily disrupt lipid bilayer and subsequently create aqueous pores in SC. [88] Although high voltage are applied for only milliseconds and electropores (≈1 nm) are created within sub-microseconds, those pores are long-lived and metastable, and can persist from seconds to hours, during which drug molecules can successfully pass through SC. [89] Small molecule drug (<1 kDa) can be transported through the epidermis via electroporation. Chemical enhancers (eg, urea) are usually needed for electroporative transport of larger molecules, particularly macromolecules with negative charges. [90] The presence of lipid (eg, anionic phospholipids) can enhance the electroporative transport of molecules up to 10 kDa. [91] Because the electrical resistance of SC is 5-25 kV/cm 2 and the electrical breakdown potential is ≈70-100 V, the electroporation methods usually use ≈100-1500 V. [92] Although small and macromolecules (proteins, peptides, and nucleotides) can be delivered into skin, the possible pain, muscle stimulation, and cell death as well as the need of complex device make electroporation-based TDD largely limited at the research stage. [93]

MICRONEEDLES FOR TDD
TDD methods like chemical enhancers, iontophoresis, and electroporation can only improve skin permeation for certain drug molecules to some extent via creation of microscopic pores in SC, usually temporarily and reversibly. Using these methods, only 10-20% of the loaded drug in the topical cream can be delivered into skin, giving poor bioavailability. [8b,94] Hypodermal needle injection can rapidly deliver ≈100% of drugs into the skin. But it is invasive and painful with risk of infection. Thus, it requires skill or clinic visit and has poor patient compliance. Besides it cannot provide sustained and controlled drug delivery. Microneedle (MN)-based TDD approach can overcome those limitations. MNs, with micrometer sized tip, can penetrate SC barrier in a minimally invasive and pain-free manner, and deliver loaded drugs directly into skin, offering an effective and patient-friendly approach. [94] MNs are developed in various dimensions depending on materials used and application purposes, usually in a range of 100-1500 µm in length, 50-400 µm in base width, and 5-50 µm in tip diameter. Various types of MNs are fabricated for TDD, including solid, porous, coated, hollow, dissolving, and hydrogel forming MNs. [95] Solid MNs simply create microchannels in SC and dermis through which drug molecules can diffuse into the skin layers, and thus increasing the drug permeation for a certain period. [96] Solid MNs can enhance skin permeability without causing injury and infection as microchannels can heal quickly. It requires two-step application process, that is, MN application followed by topical drug application. Hence, delivery efficiency still depends on passive diffusion of drug compounds that is difficult for large weight molecules. Metals, silicon, and some polymers (eg, poly(methyl methacrylate), PMMA; polylactic acid, PLA) are mainly used to fabricate solid MNs. [97] Metals, particularly stainless steel, titanium (Ti), nickel, palladium, palladium-cobalt alloys, are the most used materials to make solid MNs, because of good mechanical properties and biocompatibility. [98] Dry etching process is commonly used to fabricate solid silicon MNs. Potassium hydroxide (KOH) based wet etching process is also used to fabricate silicon MN with tunable high-pitch ratio. [98b] Recently, Narayanan et al used tetramethyl ammonium hydroxide (TMAH) etching process to fabricate silicon MNs that are hard enough to penetrate skin. It was shown that gold (Au) coating on silicon MNs using metal sputtering technique improved the mechanical strength and biocompatibility. [99] Solid MNs can be coated with drugs or drug containing layer, whose dissolution leads to rapid release of drugs into the skin layers. [100] The coating thickness and MN size determine the loading capacity of drugs. Proteins, peptides, nucleotides, and virus can be easily coated on solid MNs using different coating methods, including dip coating, spray coating, gas-jet dry coating, electrohydrodynamic atomization (EHDA)-based coating, and direct inkjet printing. [101] For example, MNs can be simply dipped into a drug-containing solution, withdrawn, and dried. However, the drying process is slow and create uneven coating. [102] Gas-jet drying coating method involves application of gas-jet on the coating solution suspended on the MNs, creating uniform coating layer on the MN surface. [103] Viscosity and surface tension are the two key factors for a coating solution to determine the coating thickness and uniformity. Spray coating involves atomization of the coating solution into microdroplets that are deposited and adhered to the MN surface, and the coalescence of microdroplets forms a uniform coating. [104] EHDA-based coating method uses an electrical field to produce the atomized microdroplets. [105] In addition to the viscosity and surface tension, flow rate, electrical conductivity, angle, and distance between nozzle and MN surface are important parameters for such coating processes. Piezoelectric inkjet printing uses piezoelectric crystals to induce vibrations that create F I G U R E 3 Microneedles for transdermal theranostics. (A) 3D-printed MNs coated with insulin. [107] (B) 3D-printed polymeric MNs. [108] (C) Porous polymer coatings on metal MN. [109] (D) Solid MN coated with influenza vaccine. [112] (E) Titanium porous MN. [118] (F) Porous polymeric MN. [120] (G) 3D-printed hollow MNs. [127] (H) Built-in active dissolving MN loaded with magnesium microparticles. [137] (I) 4D-printed polymeric MN with backward-facing barbs. [146] (J) Dissolving PLGA MN. [139] (K) Circular obelisk-type multilayer MN. [147] (L) Long hollow MN for transdermal diagnostics. [121b] (M) Hydrogel coated MN for transdermal diagnostics. [148] [107,112,120] The Royal Society of Chemistry, [108] WILEY-VCH, [137,146,139] American Chemical Society [148] a pressure pulse in the ink chamber, forcing microdroplets ejected from the nozzle. [106] It allows controlled distribution of microdroplets onto MN surface, but limits to the usage of low viscosity coating solution. Recently, inkjet printing coating of insulin formulation was applied on the 3D-printed resin-based MNs [107] ( Figure 3A). MNs were fabricated using a Form 2 stereolithography printer by Formlabs followed by ultraviolet (UV) light-induced photo-polymerization, while mannitol, trehalose, and xylitol were used as insulin carriers in order to preserve its activity. Luzuriaga et al fabricated 3D-printed polylactic acid (PLA)-based MNs, with the tip sizes in the range of 1-55 µm, using fused deposition modeling 3D printing method, followed by chemical etching and drug coating ( Figure 3B). [108] Drug delivery efficiency of polyvinyl alcohol (PVA) coated PLA-MNs reached ≈90% (sulforhodamine B). [101b] Porous polymer coating on metal MNs was also explored for enhanced TDD ( Figure 3C). [109] Polylactic-co-glycolic acid (PLGA)-based coating solution containing gelatin and PVA was used to create a porous coating layer on MNs, and it was found that drug delivery increased concomitantly with increased porosity. Le et al fabricated individually coated MN arrays for co-delivery of multiple compounds with different properties. [110] MNs were individually decorated with poly(vinyl pyrrolidone) (PVP) containing different molecules, proteins, and nanoparticles. Beak et al coated poly(L-lactide) (PLLA) MNs with an anesthetic agent (lidocaine) for rapid release. [111] Shin et al coated zymosan (a glycan derived from Saccharomyces cerevisiae) and/or poly(I:C) (polyinosinic-polycytidylic acid) together with whole inactivated influenza vaccine on stainless-steel MNs for vaccination ( Figure 3D). [112] It was found that addition of zymosan to the vaccine antigen significantly improved stronger IgG and hemagglutination inhibition (HI) titers in BALB/c mice, suggesting zymosan as an immunostimulant. In a study, DNA vaccine was loaded in PLGA-based cationic NPs that in turn were coated on stainless-steel MNs (PVP and trehalose as coating solution). [113] It was found that BALB/c mice treated with MN-delivery of polyplex encapsulated with pH1N1 plasmid showed a greater humoral immune response (ie, higher IgG titer) compared to those delivered with intramuscular injection, suggesting the advantage of skin immunization based on intradermal delivery of vaccines. Duong et al also reported a rapid delivery of polyplex-based DNA vaccine (pDNA, pTarget-Ig-A -Fc), using a polycarbonate MN array coated with polyelectrolyte multilayers, which were formed via assembly of heparin and charge reversal pH-responsive copolymer composed of oligo(sulfamethazine)-b-poly(ethylene glycol)-b-poly(amino urethane)). [114] The charge reversal copolymer (isoelectric point at pH 7.4) allows the electrostatic repulsion between heparin and polymer in skin, triggering the release of loaded DNA vaccines, and producing robust immune responses in BALB/c mice (ie, producing higher IgG titer, comparing to those produced in subcutaneous administration). More recently, MN-delivery of lipid-coated cisplatin NPs via its CMC-based coating layer was reported for efficient cancer therapy. [115] Using a xenografted head and neck tumor model, MN-delivery of cisplatin-loaded pH-responsive lipid NPs largely inhibited the tumor growth, without producing systemic side effects. Solid MNs for cutaneous immunotherapy was also explored to prevent airway allergies. [116] Specifically, mice pre-treated with MN-delivery of Der p1, a house dust mite allergen, into skin had similar protective effects against development of Der p1-induced airway allergy, compared to those received with subcutaneous injection.
Porous MNs can be loaded with both dry and liquid drug formulation. Metals, silicon, ceramics, and some polymers (eg, PLA) have been used to fabricate porous MNs. [117] Recently, a titanium porous MN array (≈30.1% porosity and pore size of 1.3 µm) was fabricated using a metal injection molding method for transdermal delivery of drug liquid into skin ( Figure 3E). [118] Cahill et al used the stainless steel to create a porous MN patch (≈36% porosity with ≈2 µm in pore diameter) using a series of fabrication processes that involve hot embossing, sintering, and subsequent electropolishing. [119] The drug delivery efficacy of such porous MNs was threefold higher than that of topical administration. Using a modified hot embossing method, Li et al developed a porous PLGA MNs, called as gradient porous microneedle array (GPMA), with ≈20.1% porosity mostly at the MN tip (Figure 3F). [120] The authors showed that GPMA could effectively deliver insulin to lower blood glucose level in diabetes rats.
Hollow MNs (HMN) have also be devised for continuous delivery of liquid drug formulation. Hollow MNs can be a standalone device or integrated with an external drug reservoir (eg, syringe or micro-pump). They are suitable for a rapid bolus injection of drugs flowing through the microchannels of HMNs. HMNs can be fabricated from metal, silicon, and glass. [121] The size of microchannel determines the flow rate and mechanical strength of HMNs. A MN with an external drug reservoir can provide an unlimited dose of drugs. It, however, requires complex fabrication process and has risk of drug clogging. Schipper et al fabricated HMNs from fused-silica capillaries by wet etching with hydrofluoric acid. [122] It was demonstrated that mice with microinjection of inactivated polio vaccine produced similar IgG and virus-neutralizing antibody titers as compared to intramuscular immunization. De Groot et al also used fused-silica HMNs to deliver ovalbumin-loaded PLGA-NPs to elicit T cell-mediated immunity against bacterial infection. [123] It was shown that such immunization provides protection against a recombinant strain of the intracellular bacterium, ovalbuminsecreting Listeria monocytogenes. Similarly, Du et al investigated the intradermal delivery of ovalbumin loaded NPs (PLGA NPs, liposomes, mesoporous silica NPs, gelatin NPs) using such fused-silica HMNs, and found that PLGA NPs and liposomes induce stronger IgG2a responses than silica and gelatin NPs, attributable to the sustained release of antigen. [124] Maaden et al devised the digitally controlled HMN injection system (DC-hMN-iSystem) based on fusedsilica HMNs, and demonstrated that intradermal administration of therapeutic cancer vaccine (HPV E743-63 synthetic long peptide containing liposomes) could more efficiently induce cytotoxic and T-helper responses in mice, as compared to intradermal immunization. [125] Niu et al recently demonstrated the intradermal delivery of vaccine-loaded immunostimulatory NPs, containing ovalbumin with imiquimod and monophosphoryl Lipid A, using a stainless-steel based hollow micro-structured transdermal system (hMTS; from 3 M company). [126] MN delivery of such NPs demonstrated faster kinetics of antibody affinity maturation and a higher IgG2a antibody response in rats, in comparison to intramuscular or intradermal delivery. A 3D-printed microfluidic-enabled HMN device was recently fabricated using a Form 2 stereolithography printer (from Formlabs) and class IIa biocompatible resin (a mixture of methacrylic acid esters and photoinitiators) for TDD ( Figure 3G). [127] Twelve of such devices can be created in 2.5 h in a single print. With the built-in hydrodynamic mixing capability, the device enables homogeneous mixing of multiple fluids under different flow rates, and thus is suitable for combinational drug therapy.
Dissolving MNs are made of biocompatible and biodegradable polymers, which encapsulate drug molecules. [128] The major advantage of dissolving MNs is that it can control the releasing kinetics by tailoring the degradation rate of polymers, and thus improve therapeutic efficacy. As MNs dissolve completely inside the skin, the application process is simple, that is, pressing and removal of supporting substrate, thus ensuring good patient compliance. However, they often suffer from the drawbacks such as non-ideal mechanical properties and limited loading capacity. The fabrication typically involves a simple molding process in which drug-blended polymer solution is filled into a poly(dimethylsiloxane) (PDMS)-based mold and dried. PDMS-based molds can be simply prepared by pouring sylgard polymer over a stainlesssteel master structure followed by curing. Alternatively, using a laser micro-machining method, PDMS molds with different geometries can also be directly fabricated from a piece of PDMS. [129] Various water-soluble and biocompatible polymers have been used to fabricate fast dissolving MNs, for examples, PVA and PVP which can give strong mechanical strength for skin penetration. Esser et al employed PVA-MNs loaded with unadjuvanted tetanus toxoid for protective immune response in pregnancy. [130] It was found that MN delivery of tetanus toxoid to pregnant mice not only provided better survival rate (up to 100%) to tetanus toxin, but also delivered superior immune response to their new-born neonate mice, comparing to intramuscular vaccination. Littauer et al studied the stable incorporation of granulocyte-macrophage colony stimulating factor (GM-CSF), an immunomodulatory cytokine known for its increased immunogenicity, into PVA-MNs to improve vaccine-induced protective immunity against influenza. [131] It was found that MN delivery of GM-CSF adjuvanted egg-grown-subunit monovalent-influenza vaccine into the skin induced robust and long-lived antibody responses, leading to rapid lung viral clearance against H1N1 influenza virus in mice. Zhu et al evaluated pertussis toxin (PT) vaccine delivery in mice using PVA-MNs. [132] It was demonstrated that MN delivery of PT together with CpG oligodeoxynucleotide as an adjuvant achieved superior immune response in mice over subcutaneous injection (ie, higher serum levels of PT-specific IgGs and cytokines). Recently, Lee et al developed a tearable MN system, comprising of an array of tearable PVA-MNs (PVA with sucrose at a ratio of 14:1) and micropillars. [133] Exerting a vertical force on the micro-pillars can rapidly separate MN tips from the base and insert into the porcine skin. Lahiji et al encapsulated a hydrophobic drug (valproic acid) into carboxymethyl cellulose (CMC) based MNs to induce hair regrowth. [134] It was found that dermal micro-wounds created by MN application, together with delivered valproic acid, could stimulate wound re-epithelialization signals involved in hair follicle regrowth. Similarly, Kim et al packed the powder of a lipophilic drug (finasteride) into the cavity of CMC-based MN to treat androgenetic alopecia. [135] It was shown that application of such MNs together with topical application of diffusion enhancer (containing ethanol, propylene glycol, and water) yielded higher efficacy in promoting hair growth in mice, compared with topical gel application. For better solubilization of hydrophobic drugs in hydrophilic polymer, Guo et al utilized nanostructured lipid carriers (NLCs) to encapsulate hydrophobic anti-inflammatory molecule aconitine (ACO-NLCs). [136] A significant inhibitory effect on the paw swelling and inflammation in a rat model of adjuvantinduced arthritis was observed using ACO-NLCs-loaded PVP-MN array. Recently, active MN delivery platform was realized for deeper and faster TDD, by loading magnesium microparticles into PVP-MNs ( Figure 3H). [137] Reaction between magnesium microparticles and ISF can lead to rapid explosive-like production of hydrogen bubbles, providing the force to breach dermal tissue and thus enhance drug delivery. Using a B16F10 mouse melanoma model, active MN delivery of anti-CTLA-4 (a checkpoint inhibitor drug) resulted in significantly improved immune response and survival rate.
Hyaluronic acid (HA) is a non-sulfated glycosaminoglycan distributed abundantly in human connective tissues including skin. [138] As the natural water-soluble polymer and having a robust mechanical property and fast dissolution rate, HA is a suitable matrix material for dissolving MNs. Recently, fast dissolving MN array made from small molecular weight HA (miniHA, 3-10 KDa) was utilized for transdermal delivery of anti-obesity compounds to subcutaneous fat tissue. [139] 3-adrenoceptor agonist and thyroid hormone T3 quickly released from HA-MNs can diffuse to and accumulate in the underlying subcutaneous fat, whereby promoting white fat browning and energy consumption. Localized fat browning and reduction was observed in high-fat diet induced mice treated with such anti-obesity MN patches. HA-MN patch was also utilized for 5-aminolevulinic acid (ALA)-based photodynamic therapy for cancer treatment. [140] It was found that MN treatment for subcutaneous tumor-bearing BALB/c nude mice showed better efficacy than the injection by hypodermic needle. HA-MN array loaded with live attenuated Bacille Calmette-Guerin bacillus (BCG) vaccine powder was developed for painless and lesion-free vaccination against tuberculosis, with efficacy comparable to intradermal immunization. [141] Choi et al demonstrated transdermal delivery of canine influenza vaccine (derived from canine influenza H3N2 virus) using tip-separable HA-MN (HA MNs on polycaprolactone substrate), which produced two-times higher HI antibodies in guinea pigs than intramuscular injections. [142] Both vaccination groups showed lower viral titers in guinea pigs challenged with influenza virus 2 weeks after the second vaccination. Du et al demonstrated in mice that methotrexate-loaded HA-MNs alleviated psoriasis-like imiquimod-induced epidermis thickening. [143] HA-MN delivery of ovalbumin as the model antigen was studied in mice bearing ovalbumin-expressing EG7 tumor. [144] It was found that transdermal delivery of ovalbumin effectively increased the ovalbumin-specific CD8 + and CD4 + T cell populations and induced ovalbuminspecific cytotoxic T cell response against the tumor. Recently, hydroxypropyl--cyclodextrin was incorporated into HA to improve mechanical strength and hydrophobic drug-loading capacity of MNs for hypertrophic scar therapy. [145] It was shown that intradermal delivery of triamcinolone acetonide using such MNs could achieve better drug delivery into the hypertrophic scar, and thus enhanced the therapeutic effects.
Crosslinked polymers have been used to fabricate hydrogel forming MNs for sustained drug release. Zhang et al showed that MNs made of crosslinked methacrylated-HA (MeHA) enabled sustained drug release for approximately 3 days. [149] Localized fat browning and reduction was observed in diet-induced obese mice using MeHA-MN patch carrying rosiglitazone-loaded NPs. Than et al have demonstrated a kind of double-layered (DL) MN with a fast dissolving HA core and a slowly dissolving MeHA shell for ocular drug delivery. [150] The biphasic drug releasing enabled by DL-MN array provides not only an initial bolus dose (within 5 min) to quickly reach the therapeutic level, but also sustained drug release (≈3 days) to maintain therapeutic effect for a much longer period, thus leading to much enhanced therapeutic efficacy. In another study, HA (≈250 kDa) was crosslinked with 1,4-butanediol diglycidyl ether (BDDE) before filling into the MN mould. [151] In contrast to HA MN that released 100% of loaded FITC-dextran within 2 h in PBS, crosslinked HA MN released relatively slowly (≈60% at 2 h). Similarly, it was demonstrated that FITC-dextran loaded in diamine-crosslinked HA MNs can retain in mouse skin up to 5 days. [152] Poly(ethylene glycol) diacrylate (PEGDA) is a biocompatible polymer and can be crosslinked under UV exposure. Gap junction inhibitor peptides were loaded in PEGDA-MNs during MN swelling in aqueous solutions for keloid scar treatment. [153] Recently, PEGDA-MNs with bioinspired backward-facing curved barbs, called 4D-printed MNs, were fabricated for enhanced tissue adhesion and sustained TDD ( Figure 3I). [146] The MNs were fabricated by a digital light processing, micro-stereolithography 3D printing technique, while backward-facing barbs on the MNs were created by desolvation-induced deformation utilizing cross-linking density gradient in a photocurable polymer PEGDA. Migdadi et al integrated a lyophilized drug reservoir layer with hydrogel forming MNs made of crosslinked poly(methylvinylether-co-maleic acid). [154] Such MN array can deliver ≈50 mg of metformin within 24 h in porcine skin. Chitosan-based hydrogels also have attracted a lot of attention for controlled drug delivery since chitosan is a natural cationic copolymer that can be degraded by lysozyme in humans. [155] Recently, Zan et al developed an antimicrobial MN patch made of chitosan-poly(ethylenimine) (PEI) copolymer for treating deep cutaneous fungal infection. [156] The MNs swelled immediately after skin insertion and released loaded cargo over approximately 3 days in mice. The synergistic actions from the antimicrobial activities of both chitosan-PEI and the encapsulated antifungal agent (amphotericin B) offered outstanding effectiveness.
Sustained drug delivery avoids frequent drug administration, drastic fluctuation of drug level, and hence gives better patient compliance, enhanced therapeutic effect, and reduced side effects. [157] MNs can be modified for sustained drug release system. Given the known biocompatibility, biodegradability, and slow-degradation of PLGA, a PLGA-MN patch was utilized for sustained delivery of a browning agent -3 adrenoceptor agonist into subcutaneous fat ( Figure 3J). [139] Immediately after insertion, PLGA-MNs can detach from HA-substrate within 2 min and remain in skin for ≈2 weeks. It was shown that treatment with such a fat-browning MN patch largely inhibited diet-induced body weight gain by ≈15% as compared with the untreated obese mice. Kim et al fabricated PLGA-PVP multilayered MNs with circular obelisk and beveled-circular obelisk geometries, using micro-milling and spray deposition methods, for controlled TDD (Figure 3K). [147] Obelisk-shaped MNs are mechanically stronger than pyramid-shaped MNs. Recently, Li et al fabricated a patch consisting of an array of rapidly separable PLGA-MNs loaded with a contraceptive hormone -levonorgestrel (LNG). [158] Bubble structures between MNs and the patch backing allow MNs to penetrate skin within 5 s under a small compressive shear force. Its application in rats could maintain plasma concentrations of LNG above the human therapeutic level for 1 month.
To optimize drug dosing and therapeutic efficacy based on specific pathophysiological conditions, MN patches capable of stimuli-responsive drug delivery have been developed. For instance, a smart insulin MN path was developed using hypoxia as the bio-stimulus to regulate insulin release ( Figure 4A). [159] The glucose oxidase (GOx) loaded glucose-responsive vesicles were formed by self-assembly of HA modified with hypoxia-sensitive hydrophobic group (2-nitroimidazole). Based on the body glycemic condition, local hypoxic environment due to enzymatic conversion of glucose to gluconic acid by GOx can trigger disassembly of the vesicles, and subsequently insulin release. Such patch could regulate blood glucose level within 30 min and maintain a normoglycemic state for ≈4 h in diabetic mice. Ye et al fabricated a MN patch containing synthetic glucose-signal amplifiers and pancreatic cells for cell-based glucose responsive insulin delivery ( Figure 4B). [160] Disassembling of glucose-responsive vesicles due to high glucose level leads to the release of -amylase and glucoamylase. The subsequent enzymatic reactions convert -amylose into glucose, which in turn activates pancreatic cells to release insulin. Such cell-based MN patch showed better glucose control for a prolonged period (≈10 h) in diabetic mice. Based on GOxcatalyzed glucose oxidation and hydrogen peroxide (H 2 O 2 ) production, Hu et al integrated H 2 O 2 -responsive polymeric vesicles (PVs) into MN array as a self-regulated smart insulin patch. [161] The PVs were formed from self-assembly of block copolymer incorporated with PEG and phenylboronic ester (PBE)-conjugated polyserine. Production of H 2 O 2 from glucose oxidation leads to degradation of hydrophobic PBE, triggering rapid dissociation of PVs and consequently release of loaded insulin. In vivo testing indicated that a single patch could regulate blood glucose level in diabetic mice. Yu et al utilized both hypoxia and H 2 O 2 as the endogenous stimuli to make a more sensitive smart insulin patch. [162] The dual-sensitive PVs were synthesized using a diblock copolymer consisting of PEG and polyserine modified with 2-nitroimidazole via a thioether moiety. In the local hypoxic environment with abundant H 2 O 2 due to GOx -mediated glucose oxidation, the H 2 O 2 -sensitive thioether moiety renders the polymer more hydrophilic as it is converted into a sulfone by H 2 O 2 . Together with the hypoxia-sensitive 2-nitroimidazole moiety, these chemical changes can trigger rapidly dissociation of PVs and thus release of insulin. It was shown that such MN patch could regulate blood glucose level for ≈10 h in diabetic mice. Wang et al also designed a MN patch for self-regulated insulin delivery, without the risk of hypoglycemia and inflammation caused by the generated

F I G U R E 4 MN array patches capable of stimuli-responsive drug delivery. (A) The MN patch loaded with hypoxia-sensitive vesicles for
glucose-responsive insulin delivery, [159] Copyright 2015 National Academy of Sciences. (B) The MN patch integrated with pancreatic cells and synthetic glucose-signal amplifiers for glucose-responsive insulin delivery, [160] Copyright 2016 WILEY-VCH. (C) Core-shell MNs gel for self-regulated insulin delivery , [163] Copyright 2018 American Chemical Society. (D) Glucose-responsive insulin patch for the regulation of blood glucose, [165] Copyright 2020 Springer Nature. Reprinted with permission from WILEY-VCH, [160] American Chemical Society, [163] Springer Nature [165] H 2 O 2 ( Figure 4C). [163] The core component of the core-shell MNs in this patch is composed of PVA matrix crosslinked with an H 2 O 2 -labile small linker, and GOx-encapsulated acrylated nanogel that can generate H 2 O 2 to trigger insulin release. The crosslinked-PVA shell is embedded with catalase-encapsulated nanogel to serve as an active strainer for scavenging excessive H 2 O 2 . Charge-switchable polymeric complex was also developed as glucose-responsive insulin delivery vehicle for either subcutaneous injection or transdermal MN application. [164] In the presence of a hyperglycemic state, glucose binding of glucose-responsive phenylboronic acid (PBA) converts the charge of the polymeric moiety from positive to negative, thereby enabling release of insulin from the complex. Recently, a MN patch bearing a glucose-responsive insulin delivery system was developed and demonstrated in insulin-deficient diabetic mice and minipigs ( Figure 4D). [165] Under hyperglycemic conditions, glucose binding with PBA units within the MN polymeric matrix can reversibly form glucose-boronate complexes that in turn induces MN swelling. The subsequent weakening of the electrostatic interactions between the negatively charged insulin and polymers leads to rapid release of insulin. Such patch with a size of ≈5 cm 2 was able to effectively regulate plasma glucose levels for >20 h in minipigs (>25 kg).

MICRONEEDLES FOR TRANSDERMAL DIAGNOSTICS
Microneedle (MN)-based transdermal diagnostics have gained increasing attention due to its ability to sample skin bio-fluids for point-of-care testing (POCT) and continuously monitoring of various biomarkers, in a minimally invasive and painless manner. [166] MNs can effectively extract blood or interstitial fluid (ISF) via negative pressure, capillary action, or material absorption. Hollow, porous, solid, and hydrogel forming MNs have been designed for extraction of skin biofluids. Long hollow MNs (HMN) >1500 µm in length based on silicon, metals, or polymers have been deployed for blood sampling, as they are long enough to reach dermal

F I G U R E 5 MNs for transdermal biofluids collection. (A)
A self-powered one-touch hollow MN-based blood extraction system, [168] Copyright 2015 The Royal Society of Chemistry. (B) A MN-based device for one-step painless collection of capillary blood samples, [169] Copyright 2018 Springer Nature. (C) A microfluidic chip integrated with porous MNs for collection of skin ISF, [170] Copyright 2019 Springer Nature. (D) A Swellable MN patch for rapid extraction of skin ISF, [171] Copyright 2017 WILEY-VCH. Reprinted with permission from The Royal Society of Chemistry, [168] Springer Nature, [169,170] WILEY-VCH [171] capillaries located at 300-1800 µm underneath the skin ( Figure 3L). [121b,167] Recently, Li et al devised an MN-based self-powered blood extraction system, in which a PVP-sealed long HMN was attached to a PDMS-based pre-vacuum actuator( Figure 5A). [168] Once the device is pressed on skin, a stainless-steel HMN penetrates, detaches the PVP-sealer, and extracts blood into the vacuum-chamber via negative pressure with a flow rate of ≈7.8 µL/s. Blicharz et al developed a one-step painless blood extraction device, containing a long solid MN array, a microfluidic system, and a vacuum chamber ( Figure 5B). [169] With a single push of actuation button, a stainless-steel MN array can be rapidly deployed and retracted, creating skin punctures through which blood is withdrawn into the device by negative pressure (≈100 µL in 3 min). Shorter MNs (<1000 µm) are usually used to extract ISF, which not only acts as an alternative biomarker source to blood but also provides a good applicability in continuous monitoring of biomarkers. Strambini et al. developed a densely packed silicon-dioxide-based HMN array (up to 1 × 10 6 MNs per cm 2 ) that can extract ISF up to 1 µL/s through simple capillary action. As shown by Miller et al, the concentric holder of HMN can exert local mechanical pressure at skin surrounding the HMN insertion point, and thus squeeze skin ISF flow into the HMN [13a] . By this way, an array of five HMNs can extract up to 16 µL of ISF from humans within 1-2 h. A porous MN that contains an interconnected network of microchannels is also capable of rapid ISF extraction by capillary action. [172] To extract ISF, Liu et al developed a porous poly(glycidyl methacrylate)-based MN array (6 × 6, ∼1 µm pore-diameter) by photo-polymerization of acrylate monomers in the presence of PEG porogen. [173] Ceramicbased porous MNs, with the pore diameter ∼1 µm, were also reported for ISF extraction. [174] Recently, Takeuchi et al developed a microfluidic system connected with a porous PDMS-MN array to enable a direct analysis of extracted ISF ( Figure 5C). [170] The fast-dissolving HA was prefilled in porous PDMS-MNs to reinforce their mechanical strength.
Hydrogel-forming crosslinked polymers can rapidly absorb nearby fluid and swell without dissolution. [175] Chang et al demonstrated the use of crosslinked methacrylated-HA (MeHA) to rapidly extract ISF for off-line analyses of glucose and cholesterol ( Figure 5D). [171] A patch containing 10 × 10 MeHA-MNs can extract ≈1.4 and ≈2.3 µL of ISF in mice within 1 and 10 min, respectively, without need of any external devices. Similarly, Zhu et al demonstrated that a MN patch (11 × 11 MNs) made by crosslinked methacrylatedgelatin (GelMA) can collect ≈2.5 µL (≈2.5 mg) of ISF from a rat within 10 min. [176] A hydrogel MN patch (10 × 10 MNs) made of PVA and chitosan was recently fabricated by repeated freezing and thawing method without using any cross-linking agent. [177] Such MN patch can extract ≈1.25 mg ISF from a rabbit within 10 min. Mandal et al coated the Ca 2+ crosslinked alginate as the hydrogel forming layer on solid PLLA-MNs for sampling of ISF and cells to monitor skin-resident immunity. [178] With specific antigen and adjuvant being loaded in the alginate layer, such MN is also capable of ISF sampling in order to capture lymphocytes. Recently, Sulaiman et al used hydrogel coated solid MNs for ISF extraction and detection of nucleic acid biomarkers (Figure 3M). [148] It was demonstrated that an array of PLLA-MNs (77 MNs) coated with an alginate-peptide nucleic acid hybrid material can extract ≈6.5 µL ISF in 2 min, and enable detection of specific nucleic acid biomarkers either on the patch itself or in solution after light-triggered release from the hydrogel. As an alternative to hydrogel coating, Kolluru et al utilized the filter paper to absorb ISF for subsequent offline analysis. [179] A patch containing nine stainless-steel MNs with a strip of filter paper adhered to the patch backing can extract >2 µL ISF within 1 min.
Recently, researchers have integrated MNs with a microchip platform containing microfluidic channels and electrochemical electrodes to function as a miniaturized diagnostic analyzer for POCT. For example, Miller et al developed an HMN-based microchip platform containing eight immuno-modified electrode transducers for myoglobin and troponin detection in ISF ( Figure 6A). [180] Based on the electrochemical immunoassay, it can detect specific proteins in the range of 100-1000 ppb (parts-per-billion). Ranamukhaarachchi et al utilized the inner lumen of HMN as the immunoassay microreactor. [181] Together with an integrated optofluidic device and immuno-modified HMN, this biosensor only requires <1 nL ISF to measure vancomycin concentration, with low limit of detection (LoD) of <100 nM. Integration of HMN with a paper-based colorimetric sensor provides a low cost and simplified transdermal diagnostic system, without the need of optical or electronic components. [182] Such HMN-based colorimetric sensor was employed for detection of glucose and cholesterol in ISF ( Figure 6B). [182] Finger pressing of the PDMS-chamber pushes HMNs into skin, and when the finger is released, negative pressure exerted by the deformable chamber extracts blood into the paper-based multiplex sensor for colorimetric detection. The whole process only needs 3 min. Similarly, Nicholas et al assembled the paper-based colorimetric sensor on the backside of a HMN for detection of glucose in ISF. [183] Such MN patch can discriminate glucose concentrations between 0 and 10 mM within 30 s with naked eye.
Apart from the single-use disposable MN devices for POCT, MN devices for continuous monitoring of biomarkers in ISF have been demonstrated too. For example, Mohan et al devised a MN-based electrochemical biosensor for continuous assessment of alcohol in ISF. [184] The device consists of three HMNs integrated with platinum (Pt) and sliver/silver chloride (Ag/AgCl) wires within their apertures. As the alcohol oxidase-immobilized working electrode is covered by a size-exclusion poly(o-phenylenediamine) (PPD) layer and a negatively charged Nafion film that could eliminate electroactive interferents (eg, uric and ascorbic acids), such a device displays an interference-free linear detection of ethanol from 0 to 80 mM. Mishra et al reported an MN-based electrochemical biosensor for continuous detection of nerve agents in ISF, using three carbon-paste packed HMN-based electrodes. [185] As the working electrode is immobilized with organophosphorus hydrolase (OPH) and nafion layer, the biosensor offers a linear detection of OP agent (methyl paraoxon) over a range of 20-180 µM. Recently, Ribet et al reported the continuous monitoring of glucose in ISF using a single HMN integrated with a miniaturized sensing probe consisting of three electrochemical electrodes ( Figure 6C). [186] Using the GO-based enzymatic electrochemical detection method, this device can track glycemia (0-14 mM) over time with ≈10 min delay.
MN-based wearable analytical sensors have been developed for real-time on-body analysis of metabolites and electrolytes in ISF. For example, a wearable device containing a MN-based electrochemical sensing platform was used to detect tyrosinase enzyme, a biomarker for melanoma. [187] The catechol-coated carbon-paste packed HMN array (3 × 3 MNs) as the three-electrode system is attached with a flexible printed circuit board for real-time data analysis and wireless transmission. In the presence of tyrosinase in skin tissue, the immobilized catechol is rapidly converted to benzoquinone that is detected amperometrically, with a current signal proportional to the tyrosinase level (0-2.5 mg/mL). Goud et al also reported a wearable electrochemical MN sensor for continuous monitoring of levodopa for Parkinson's disease management. [188] Rapid detection of levodopa in ISF (50 to 250 µM, with 50 µM increments) was demonstrated using a dual-sensing platform, based on square-wave voltammetry and chronoamperometry with unmodified and tyrosinasemodified carbon-paste HMN electrodes, respectively. Parrilla et al reported a wearable all-solid-state MN patch for potentiometric detection of potassium (K) ions in ISF (Figure 6D). [189] Stainless steel MN was sequentially coated with carbon, thin film of functionalized multi-walled CNTs (f-MWCNTs) and K + -selective membrane to act as a working electrode, before being embedded into a flexible PDMS substrate together with an Ag/AgCl reference MN electrode.
MNs can be functionalized with specific antibody or antigen molecules to act as effective immunosensing platform for capturing and detection of protein biomarkers in ISF. Muller et al exploited the surface-modified silicon MN array for selective detection of dengue virus nonstructural protein 1 (NS1) with a LoD of 8 µg/mL. [190] Lee et al reported the capture of circulating plasmodium falciparum biomarker -histidine rich protein-2 (HRP2) using a skin patch equipped with antibodies-coated silicon MNs. [191] Recently, Coffey et al developed highly packed, Au-coated silicon MN array (eg, 30 000 MN/cm 2 ) for rapid and selective biomarker capturing in skin ISF. [192] It was demonstrated that such high-surfacearea MN patch coated with FluoSpheres/methylcellulose layer could rapidly capture antigen-specific IgG from mice within 30 s. To detect multiple antigens simultaneously, Ng et al developed a multiplex immunodiagnostic device. [193] Different antibodies were immobilized on hexamethylenediamine (HMDA)-modified PLA-MNs. Using a blotting method in conjunction with densitometry analysis, such device can simultaneously detect interleukin (IL)-1 and IL-6 from skin, with LoD at <10 pg/mL. A recent study by Zhang et al demonstrated the multiplex detection of ISF biomarkers using MNs encoded with photonic crystal (PhC) barcodes. [194] Different antibody probes were immobilized on the silica colloidal crystal bead (SCCB) and hydrogel hybrids to form PhC balls that were then loaded in PEG-PEGDA MNs. With the addition of corresponding fluorescent probes in vitro to form sandwich immunocomplexes, the relative content of specifically enriched biomarkers (tumor necrosis factor , IL-1b, and IL-6) can be measured by the fluorescence intensity of the barcodes.
Apart from the immunosensing platform, MNs can also be modified to function as the built-in electrochemical sensors for detection of biomarkers in skin ISF. Chen et al constructed an implantable MN-based glucose sensor by layerby-layer deposition of Au/Pt NPs, GO-entrapped polyaniline nanofibers, and porous poly(vinylidene difluoride) (PVDF)-Nafion layers on stainless MNs ( Figure 6E). [195] As Pt NPs were also incorporated into the conductive polyaniline layer, electroactive surface areas as well as electrocatalytic activities between enzymes and electrodes were largely enhanced. This MN-based sensor has fast response (< 30s), an LOD of 0.1 mM and a wide range of glucose detection (0-20 mM). Bollella et al devised a MN-based lactate sensor by functionalizing Au-coated polycarbonate MNs with three layers, including electrodeposited gold-multiwalled carbon nanotubes, electro-polymerized redox mediator-methylene blue (MB), and immobilized enzyme lactate oxidase (Figure 6F). [196] Such biosensor (combining four patches of 4 × 4 MN array, three as functionalized working electrodes and one as silver-coated reference electrode) can achieve interference-free lactate detection with high sensitivity (1473 µA cm −2 mM −1 ), an LOD of 2.4 µM and a wide range of 10-200 µM. Recently, Gowers et al developed a MN biosensor for continuous monitoring of -lactam antibiotics in vivo. [197] Chrome and Au-coated polycarbonate MNs were decorated with multiple layers containing insulating lacquer, iridium oxide, -lactamase hydrogel, and PEI. The potentiometric sensing is based on the pH-sensitive iridium oxide that responds to local pH change resulting from -lactam hydrolysis by -lactamase. This biosensor comprising three of 4 × 4 working MNs and one of 4 × 4 silver-coated MN array achieved a low LOD of 6.8 µM. A MN-based -lactam biosensor with an LOD of ≈0.17 mg/L was recently evaluated in healthy volunteers for monitoring of phenoxymethyl penicillin, a commonly used antibiotic in clinics. [198] It was found that the drug pharmacokinetic (PK) profiles were similar between MN biosensing and microdialysis method.
MN-based electrochemical sensors have been developed to detect pH and biomolecules such as nitric oxide (NO) and H 2 O 2 in skin tissue for real-time monitoring of important pathophysiological conditions. For example, Tang et al modified the acupuncture needle by depositing with Au NPs, poly(ethylenedioxy thiophene) (PEDOT) and iron-porphyrin functionalized graphene composite layers for real-time monitoring of NO in acupoints. [200] The good catalytic activities of iron-porphyrin and the excellent conductivity of graphene allow the biosensor to achieve an LOD down to 3.2 nM. Hegarty et al designed a transdermal pH sensor, in which polystyrene MNs were loaded with carbon NPs/fibers or Ag NPs to serve as working and reference electrodes, respectively. [201] Mani et al also developed a MN-based sensor for direct, label-free, and real-time pH sensing using a pH-sensitive amphoteric material zinc oxide (ZnO). [202] The sensor comprising the working electrode of ZnO-coated tungsten MN and the Ag/AgCl reference electrode can measure a range of pH between 2 and 9, with a high sensitivity of −46 mV/pH. Liu et al utilized ZnO nanowires to construct MN biosensor for detecting H 2 O 2 in skin tissue. [203] The water-dissolvable PVP was further coated on the ZnO-coated stainless steel/resin MN array (9 × 9) to protect the superficial ZnO NWs during MN penetration, and thus increased the sensitivity by threefold (1.12 mA cm −2 mM −1 ). Similarly, Jin et al exploited PVP coating to protect the nanostructures (nanohybrid consisting of reduced graphene oxide, rGO, and Pt NPs) integrated on the surface of stainless steel MNs. [204] The patch containing 10 MN working electrodes (PVP-coated Pt/rGO-integrated MNs) can measure H 2 O 2 in situ and in real time, with high sensitivity (≈0.1 mA/mM) over the range of 0-20 mM.
The continuous and long-term recording of biopotentials, such as electrocardiography (ECG), electromyography (EMG), and electroencephalogram (EEG) is of great importance to monitor various pathophysiological conditions. [10] As MNs can directly insert into the conductive layers of the skin epidermis in a painless manner, MN-based electrodes can offer advantages over conventional non-invasive wet and dry electrodes, specifically, elimination of the need for gel application or skin preparation to decrease electrode-skin interface impedance, and motion artifacts arising from electrode-skin attachment. Chen et al coated Ti/Au film on magnetizationinduced self-assembled MN array (6 × 6) for biopotential recording (EMG, ECG). [205] Wang et al devised a silicon MN electrode array on flexible parylene-based substrate to allow long-term recording of EEG without motion artifacts, with a low impedance density of 7.5 KΩ cm 2 @10 Hz (Figure 6G). [199] Ren et al developed a novel magneto-rheological drawing lithography method to cost-effectively fabricate a flexible MN electrode array (6 × 6 Ti/Au film-coated MNs) on polyimide substrate for the wearable biopotential monitoring (EMG, EEG, ECG). [206] Kim et al further proposed that a flexible MN electrode array can be integrated to a flexible printed circuit board (FPCB) for wireless transmission of recorded EMG signal. [207] 9 WEARABLE SWEAT SENSORS Sweat shows great promise to examine pathophysiological important analytes, because it can be monitored noninvasively and continuously through a wearable platform. [11b] Sweat can also be generated non-invasively through local chemical stimulation at different body locations, and thus serve as on-demand bio-fluid for in situ analysis. However, poor blood-to-sweat biomarker correlations as well as bloodto-sweat lag time often make sweat sensors difficult to infer the true biomarker concentration. Hence, in vivo blood-sweat correlation validation is required to verify clinical usefulness of the sweat sensor and sweat-based biomarker monitoring. Several sweat-based biosensors for healthcare monitoring based on different mechanisms have shown promise. Among them, colorimetric sensing method is simple and easy to F I G U R E 7 Wearable sweat sensors. (A) A microfluidic device for colorimetric sensing of sweat, [208] Copyright 2016 American Association for the Advancement of Science. (B) A microfluidic device for fluorometric sensing of sweat, [209] Copyright 2016 The Royal Society of Chemistry. (C) A microfluidic-based colorimetric sensor for detection of sweat glucose, [210] Copyright 2019 American Chemical Society. (D) A microfluidic-based colorimetric sensor for detection of sweat creatinine and urea, [211] Copyright 2019 The Royal Society of Chemistry. (E) A textile-based colorimetric sensor for detection of sweat pH and lactate, [212] Copyright 2019 Elsevier. (F) A microfluidic device for sweat analysis in aquatic settings, [213] Copyright 2019 American Association for the Advancement of Science. (G) Fully integrated wearable sensor for multiplexed sweat analysis, [214] Copyright 2016 Springer Nature. (H) A soft microfluidic device for continuous monitoring of sweat glucose and lactate levels, [215] Copyright 2017 American Chemical Society. (I) Roll-to-roll gravure printed electrochemical sensors used in wearable devices, [217] Copyright 2018 American Chemical Society. (J) A graphene-based electrochemical device with thermoresponsive MNs for diabetes monitoring and therapy, [218] Copyright 2016 Springer Nature. (K) An electrochemically enhanced iontophoresis interface wearable platform for sweat analysis, [220] Copyright 2017 National Academy of Sciences. (L) A wearable biosensor that enables simultaneous monitoring of sweat and ISF, [222] Copyright 2018 WILEY-VCH. (M) A wearable sweat cortisol sensor based on molecularly selective organic electrochemical transistors, [223] Copyright 2018 American Association for the Advancement of Science. (N) A wearable sensing platform for simultaneous electrochemical, colorimetric, and volumetric analysis of sweat, [226] Copyright 2019 American Association for the Advancement of Science. (O) A textile sensor patch for real-time and multiplex sweat analysis [227] Copyright 2019 American Association for the Advancement of Science. Reprinted with permission from American Association for the Advancement of Science, [208] The Royal Society of Chemistry, [209,211] American Chemical Society, [210,215,217] Elsevier, [212] Springer Nature, [214,218] National Academy of Sciences [220] analyze sweat biochemistry with naked eye or a smartphone. For example, Koh et al developed a soft and flexible PDMS-based microfluidic device for colorimetric sensing of sweat ( Figure 7A). [208] The device is composed of three subsystems, (a) an adhesive layer with openings for sweat collection, (b) microfluidic channels and reservoirs filled with color-responsive reagents for sweat analysis, and (c) a magnetic loop antenna with near-field communication (NFC) electronics for interfacing to external wireless devices. Four different paper-based colorimetric assays are resided in the reservoirs for simultaneous detection of glucose, Cl, lactate and pH, with the resolution of 0.1 mM, 0.2 mM, 0.3 mM, and 0.5 pH units, respectively, based on 1% change in the R channel of the RGB images. Sekine et al applied a fluorometric sensing approach for in situ detection of sweat Na, Cl, and Zinc ions using ions-responsive fluorescence probes and a smartphone-based imaging analyzer ( Figure 7B). [209] The PDMS-based soft, flexible, and skin-interfaced microfluidic device contains three main layers, specifically, an adhesive layer with openings for sweat collection, microfluidic channel layer that routes sweat to micro-reservoirs filled with fluorescent reagents, and a detachable light-shielding layer to prevent light exposure prior to the readout process. The central flower-shaped microchannel allows measurements of total sweat loss and rate. Based on the fluorescence intensities produced from the reactions between the probes and ions, the device is able to measure 20-60 mM Na, 5-100 mM Cl, and 1-20 µM Zinc. Similarly, Xiao et al developed a simple, wearable microfluidic-based colorimetric sensor for detection of sweat glucose ( Figure 7C). [210] The PDMS-based device consists of a center cavity for sweat sampling, from which five microfluidic channels are branched out to detection microchambers filled with enzymes and reagents. This sensor can detect a linear range of sweat glucose between 0.1 and 0.5 mM, with an LOD of 0.03 mM. Zhang et al recently reported a soft, microfluidic-based wearable colorimetric sensor for sweat capture and detection of creatinine and urea, which are relevant to kidney disorders ( Figure 7D). [211] The PDMS-based circular pad consists of (a) an adhesive layer with openings for sweat collection, (b) microfluidic layer with a central serpentine microchannel for measuring sweat rate and loss, and three other microchannels in conjunction with micro-reservoirs containing paper-based chemical assays and enzymes, and (c) a capping layer and a transparent polyester adhesive film containing color reference markers. Based on RGB values from the captured images, the device can determine the sweat pH (5-7), creatinine concentration (0-500 µM), and urea concentration (0-250 mM).
A textile-based colorimetric sensor was developed for simultaneous detection of sweat pH and lactate, by simply depositing three layers onto a cotton fabric, including (a) chitosan, (b) CMC, and (c) mixture of enzyme molecules for lactate assay and indicator dyes for pH ( Figure 7E). [212] Cetyltrimethylammonium bromide (CTAB) surfactant is added into the mixture to prevent color fading for at least 60 min. The sensor can determine sweat pH (1)(2)(3)(4)(5)(6)(7)(8)(9)(10)(11)(12)(13)(14) and lactate level (0-25 mM), and thus has potential to be incorporated into sport apparels and accessories for real-time monitoring of physiological conditions during sports. Reeder et al demonstrated a waterproof, electronics-enabled, wearable microfluidic device capable of collecting and analyzing sweat ( Figure 7F). [213] The patch is made with an adhesive layer with a central opening for sweat collection, and a hydrophobic poly(styrene isoprene-styrene) (SIS)-based microfluidic layer that contains microchannel connecting from an inlet to a colorimetric reagent filled chamber and an outlet. The skin adhesive forms a watertight seal, forcing sweat from the inlet to flow into the microchannel and displacing the air in the channel into outer environment through the outlet. Water from the environment does not enter the outlet because of the small and hydrophobic channel, and the air pressure associated with air in the unfilled region of the channel. The integrated near-field communication chip within the SIS-layer allows determination of skin temperature and wireless communication to the external devices. This device allows measurement of local sweat loss, sweat rate, Cl − ions (10-100 mM), and skin temperature even during vigorous physical activity and swimming.
Electrochemical sensing is another method used for sweat monitoring. Gao et al developed a wearable system for multiplexed perspiration analysis ( Figure 7G). [214] Five different electrochemical sensors (2 enzymatic amperometric sensors, 2 ion-selective electrode sensors, and a resistance-based temperature sensor) are integrated on a flexible poly(ethylene terephthalate) (PET) substrate, called the flexible integrated sensing array (FISA), and connected with the FPCB for realtime data analysis, and wireless transmission during exercise. FISA can simultaneously measure sweat glucose (0-200 µM), lactate (0-30 mM), Na + (10-160 mM), K + (1-32 mM), and skin temperature (20-40 • C). Martin et al developed a soft, skin-mountable microfluidic device for sweat sampling and continuous, real-time monitoring of glucose and lactate levels ( Figure 7H). [215] The device contains an adhesive layer with openings for sweat collection, a PDMS layer with the microfluidic channels connecting between inlets, a detection reservoir, and an outlet, and the another PDMS layer with enzymatic amperometric electrode that can measure glucose (4-20 mM) and lactate (2-10 mM). The fresh sweat entering into inlets is driven to the detection reservoir, and then to the outlet, allowing continuous replenishment, and thus continuous monitoring of sweat. The device is further integrated with FPCB for wireless real-time data communication. Abellán-Llobregat et al developed a stretchable and screenprintable glucose sensor based on Pt-decorated graphite. [216] An electrode array including a Pt-graphite working electrode immobilized with glucose oxidase was printed on a polyurethane (PU) sheet. The sensor performs well for quantification of sweat glucose, with a linear range between 33 µM and 0.9 mM. Roll-to-roll (R2R) gravure printed electrochemical sensors were developed for high throughput, cost-effective fabrication of sensing components for wearable devices (Figure 7I). [217] The carbon and/or silver ink-based electrodes can be printed on a 150 m long, flexible PET substrate roll, and functionalized into different sensors for detecting pH, ions (Na + , K + ), metabolites (glucose), heavy metals (Cu 2+ ), and other small molecules (caffeine). The polyethylene resin-based insulating layer was coated to protect electrodes from crosstalk or shorting in aqueous environments.
Lee et al devised a wearable patch for sweat-based diabetes monitoring and feedback therapy ( Figure 7J). [218] The silicone-based stretchable device contains a serpentine bilayer of Au mesh and Au-doped graphene with high electrochemical activities. The device has three main modules: (a) sweat control (a sweat-uptake layer), (b) electrochemical sensing (humidity, glucose, pH, and tremor sensors), and (c) therapeutic components (MNs, a heater and a temperature sensor). The high glucose concentration detected by the sensor can trigger the embedded heater to dissolve the thermo-responsive MN array, and thus realize a feedback TDD of loaded antidiabetic drug metformin. A wearable tattoo-based iontophoretic-biosensing system was also developed to measure sweat alcohol for real-time monitoring of alcohol consumption. [219] The sensor contains two iontophoretic electrodes (anode and cathode) for transdermal delivery of pilocarpine to induce sweat, and three sensing electrodes (working, reference, and counter electrode) for amperometric detection of ethanol using alcohol oxidase and Prussian blue as the signal transducers. The sensor is coupled with FPCB to transmit data wirelessly in real-time. Emaminejad et al developed an electrochemically enhanced iontophoresis interface that can extract sufficient sweat volume without causing discomfort in patients ( Figure 7K). [220] The integration of sensing electrodes (2 ion-selective electrodes for Na + and Cl − and an amperometric glucose sensor) and iontophoresis electrodes on the same PET-substrate allows the induced sweat to be analyzed on-site immediately. By further integrating with a FPCB, this wearable device transmits the measurements to a smartphone in real-time. Similarly, based on iontophoresis-based sweat induction method, Hauke et al developed a wearable sweat sensor that continuously monitors ethanol concentrations at 25 s interval, with a linear range of 0.014-3.67 mM and LOD of 1.7 µM. [221] It was found that blood-to-sweat lag time was 2.3-11.4 min for signal appearance, and 19.32-34.44 min for overall pharmacokinetic profile. Recently, Kim et al devised a wearable biosensor that could simultaneously monitor sweat and ISF using a dual-iontophoresis system, that is, iontophoretic delivery of sweat-inducing pilocarpine at the anode compartment, alongside with iontophoretic extraction of ISF at the cathode compartment ( Figure 7L). [222] The screen-printed iontophoretic systems, coupled with electrochemical biosensors and wireless FPCB enables real-time measurement of sweat-alcohol and ISF-glucose in human subjects.
A wearable sweat cortisol sensor was developed based on molecularly selective organic electrochemical transistors (OECTs; Figure 7M). [223] The device consists of (a) a laser-patterned microcapillary channel array for sample acquisition, (b) a molecularly selective membrane (MSM) based on molecularly imprinted polymers (MIPs) as the artificial receptors for cortisol, (c) an electrochemical transducing layer based on PEDOT: poly(styrenesulfonate) (PSS)-based OECTs with a planar Ag/AgCl gate, and (d) a styrene-ethylene-butylene-styrene (SEBS) elastomer-based flexible protection layer. It achieves a log-linear response of cortisol concentrations between 0.01 and 10.0 µM. A wearable sweat band was also developed for in situ monitoring of drug levels, based on CNT/Nafion working electrode and distinct oxidation of target molecules. [224] The platform, consisting of a triple-electrode array patterned on a flexible PET sheet and interfaced with a FPCB, exhibits a linear differential pulse voltammetry (DPV) of methylxanthine drug (caffeine) between 0 and 40 µM. Another wearable sweat band was developed to monitor levodopa concentration in sweat with an LOD of 1 µM, based on Au nanodendrite modified Au/chromium (Cr) electrodes with immobilized enzyme molecules (tyrosinase). [225] A battery-free, skininterfaced, wearable sensing platform was also devised for simultaneous electrochemical, colorimetric, and volumetric analysis of sweat ( Figure 7N). [226] The copper-on-polyimide substrate is magnetically attached to the microfluidic patch that is embedded with electrochemical sensors and colorimetric assays. The combined electronic and microfluidic platform can simultaneously monitor sweat rate and loss, pH, lactate, glucose, and chloride, and transmit signals to a smartphone. A recent work by He et al showed a wearable textile sensor patch for real-time and multiplex sweat analysis ( Figure 7O), [227] based on silk fabric-derived intrinsically nitrogen (N)-doped carbon textile (SilkNCT). The device can detect multiple analytes, that is, 25-300 µM glucose and 5-35 mM lactate evaluated by enzymatic amperometry, 20-300 µM ascorbic acid and 2.5-115 µM uric acid by DPV, and 5-100 mM Na + and 1.25-40 mM K + ions by ion-selective electrodes. A wearable sensor entirely laser-engraved was recently developed for simultaneous sweat sampling, chemical sensing, and vital sign monitoring. [228] The sensor patch, which can be prepared at a large scale, consists of graphenebased electrochemical sensors for monitoring of uric acid (20-80 µM) and tyrosine (50-200 µM), physical sensors for monitoring temperature and respiration rate, and multi-inlet microfluidic channels for dynamic sweat sampling.

WEARABLE OPTICAL SENSORS AND ELECTRONIC SKINS
The recent progress in wearable sensors integrated with wireless data communication technologies (eg, near-field communication, NFC; radiofrequency identification, RFID; Bluetooth; ZigBee; resonant antenna; etc.) enables real-time, remote, and continuous monitoring of multifaceted physiological data. [229] Optical technologies based on organic lightemitting diodes (OLED), organic photodetector (OPD), optical fibers make wearable sensors possible to non-invasively detect various physiological data, such as pulse rate, blood pressure, and blood oxygen levels (SpO2). A classic example is photoplethysmography (PPG) sensor that can measure SpO2 based on difference in light absorption between oxyhemoglobin (HbO 2 ) and deoxy-hemoglobin (Hb), pulse rate and blood pressure based on the frequency and magnitude of the fluctuating pulsatile component of PPG signal. Recently, ultra-flexible and conformable reflective pulse oximeter was fabricated by integrating with three-color, highly efficient polymer LEDs (PLED) and OPDs ( Figure 8A). [230] The three laminated layers of optoelectronic skin contain green and F I G U R E 8 Wearable optical sensors and electronic skins. (A) Ultra-flexible organic photonic skin, [230] Copyright 2016 American Association for the Advancement of Science. (B) A miniaturized battery-free pulse oximeter, [231] Copyright 2017 WILEY-VCH. (C) A stretchable active matrix temperature sensor array for electronic skin, [234] Copyright 2016 WILEY-VCH. (D) A schematic diagram of a soft, skin-like battery-free sensor for full-body pressure and temperature mapping. [236] Copyright 2018 American Association for the Advancement of Science. (E) An integrated stretchable and self-healable electronic skin system, [237] Copyright 2018 Springer Nature. Reprinted with permission from WILEY-VCH, [231,234] American Association for the Advancement of Science, [236] Springer Nature [237] red PLEDs and an OPD, and can be laminated to skin using adhesive tape. The pulse oximeter is only ≈30 µm thick, and allows direct visualization of data through its on-skin seven-segment digital displays and color indicators. Kim et al developed a miniaturized battery-free wearable pulse oximeter with NFC capabilities ( Figure 8B). [231] The NFC bare die chip can harvest power wirelessly for operation of a microcontroller that then drives red and infrared LEDs. The backscattered light captured through a photodiode is then transmitted back to the NFC chip to amplify, digitize, and transmit wirelessly. Near infrared spectroscopy (NIRS) sensors use light in NIR region between 650 and 1000 nm that allows deep tissue penetration. However, NIRS sensors are relatively expensive and difficult to be miniaturized. A compact wireless hybrid system capable of simultaneously monitoring surface electromyography (sEMG) and NIRS signals was developed to monitor muscle activation as well as muscle perfusion and oxygenation during contraction. [232] Various electronic skins (e-skin) have been developed for continuous monitoring of vital signs such as heart rate, body temperature, and so on, in real-time with high comfort. [233] As e-skin needs to be well adhered on the skin seamlessly and comfortably regardless of locations, movements, and conditions, stretchable and self-healable capabilities are two important considerations. Various sensors and electronics can be embedded in the e-skin for health monitoring, wireless transmission and feedback communication. Recently, a type of electronic skin was designed based on a stretchable polyaniline nanofiber-based temperature sensor array with an active matrix consisting of single-walled CNT (SWCNT)-based thin film transistors (TFT), for continuous body temperature mapping ( Figure 8C). [234] The SWCNT-TFTs, LEDs, and temperature sensors implanted on PET films are mounted on soft and sticky Ecoflex substrate (for skin attachment and protection against applied strain), and electrically connected via embedded interconnections of Galinstan, a eutectic alloy liquid metal for maintenance of e-skin performance under bending and stretching. Such e-skin has good mechanical stability under biaxial stretching of 30% and exhibits a high resistance sensitivity of 1% • C −1 , with a response time of 1.8 s in the temperature range of 15-45 • C. Park et al. showed a selfpowered ultra-flexible device for ECG monitoring. [235] Integration of ultra-flexible organic photovoltaic power source with organic electrochemical transistors (OECTs) on a flexible parylene substrate enables the electronic sensor to monitor biological signals continuously without the need of external power connections. Han et al developed a thin, soft, skin-like battery-free sensor capable of continuously monitoring of skin temperature and pressure in real-time ( Figure 8D). [236] The multilayered PDMS-based device contains an NFC chip, a loop antenna, a pressure sensor (spiral-shaped monocrystalline silicon that has piezoresistive properties), and a temperature sensor (a resistance thermometer), and is powered by RF power supplies. Son et al developed a stretchable and selfhealable e-skin in which an elastic conductive nano-network (CNT and silver nanowires) is embedded in a self-healable polymer (PDMS -methylene bisphenylurea -isophorone bisurea; Figure 8E). [237] By integrating with a strain monitor, an ECG sensor, a light-emitting capacitor (LEC) array, and electronics, this multi-functional e-skin powered by a lithium-ion battery can collect, process, and wirelessly transmit data to the LEC display for visual interpretation.

DISCUSSION
Topical drug products applicable on skin are available in the market, in the forms of creams, gels, ointments, and sprays that contain drugs such as nonsteroidal anti-inflammatory drugs and antimicrobial agents. A number of commercial transdermal drug products have been developed to treat or manage motion sickness (Scopolamine), smoking addiction (Nicotine), angina pectoris (Nitroglycerin), hypertension (Clonidine), pain (Fentanyl, Buprenorphine, Lidocaine, Tetracaine), enuresis (Oxybutynin), menopausal symptoms (Estradiol, Levonorgestrel), depression (Selegiline), dementia or Alzheimer's disease (Rivastigmine), Parkinson's disease (Rotigotine), and attention-deficit hyperactivity disorder (Methylphenidate). [238] These transdermal patches are designed as a single or multi-layer drug-in-adhesive or drug-reservoir type. Chemical enhancers, (eg, propylene glycol, terpene) are commonly incorporated in the drug formulations. [239] Some novel chemical enhancers are currently under clinical development, for example, AzoneTSenhanced triamcinolone acetonide drug formulation to treat dermatoses. Given the necessity to enhance skin permeation and drug solubility, chemical enhancers shall continue to be used as important excipients. Nanocarriers like liposomal vesicles are also instrumental, because these nanocarriers not only enhance skin permeability, but also are able to carry both hydrophilic and lipophilic drugs into skin. Liposomal-based dermal products, such as Pevaryl Lipogel (econazole) and Dolaut gel (diclofenac), are marketed for deep skin infection, inflammation, alopecia, atopic dermatitis, and so on. Liposomal T4N5 topical lotion has been developed to deliver DNA repair enzyme -bacteriophage T4 endonuclease V to patients with Xeroderma pigmentosum. [240] Some liposomal-like formulations, such as transfersomes and ethosomes, are commercially available or in clinical trials. For examples, transfersome gel with ketoprofen for knee osteoarthritis just completed phase III trial; transfersome with terbinafine for onychomycosis is currently in phase III trial; ethosomal formulation of anthralin for psoriasis is currently in phase IV trial.
A number of iontophoresis-based devices are available in markets. Most of them are tap water iontophoresis devices to treat hyperhidrosis (eg, Hidrex, Idromed). Only a few are for TDD, for example, IontoPatch. Some iontophoretic transdermal patches have even been withdrawn from the market because of safety concerns relating to skin erythema, edema, and burn, for example, LidoSite (to deliver lidocaine and epinephrine for local anesthesia), Ionsys (to deliver fentanyl for postoperative pain management), and Zecuity (to deliver sumatriptan for migraine). Clearly, there is a need to improve safety and effectiveness of iontophoresis-based transdermal techniques. However, considering the ability of iontophoresis to simultaneously enhance skin permeability and provide transport driving force, it is still attracting active research, and some iontophoresis-based TDD systems are under clinical trials, for example, iontophoretic delivery of corticosteroids to treat psoriasis, and treprostinil to enhance wound healing. Iontophoretic delivery of gonadotropin-releasing hormone to treat female infertility has recently completed Phase-II clinical trial. Because iontophoresis itself does not significantly alter skin barrier, it is not suitable for transdermal delivery of macromolecules. Other physical permeation enhancement methods using electroporation or sonophoresis may be effective for transdermal delivery of macromolecules. However, those methods are limited by the need of sophisticated devices. SonoPrep, a bulky battery-powered sonophoresisbased skin permeation device for transdermal delivery of lidocaine and other macromolecules, was withdrawn due to poor market uptake. Clearly, commercially viable TDD systems shall be not only safe and effective, but also low-cost, convenient, and user-friendly.
Technologies, particularly thermal ablation and radiofrequency ablation, have been developed to physically disrupt SC barrier by creating micron-scale channels for delivering a wide range of therapeutics including macromolecules such as peptides and proteins. [241] Several companies have applied those technologies in their marketed TDD devices, for examples, ViaDor and Passport that are portable devices used for thermal ablation of the skin prior to the application of a transdermal patch. Synera is a commercial transdermal patch that utilizes a controlled heat-assisted drug delivery (CHADD) technology to enhance lidocaine and tetracaine delivery for local anesthesia. Some companies have developed laser ablation systems for effective ablation of SC. Most of them are used for skin rejuvenation for cosmetic and dermatological purposes. [242] Only a few have been applied to TDD. A portable device that uses a precise laser epidermal system called P.L.E.A.S.E. to create a micropore array in the skin is available in the market.
MN technology allows transdermal delivery of a wide range of drugs including macromolecules with the highest bioavailability. Some MN-based TDD systems are already commercialized, for examples, microinjection systems with influenza vaccine (eg, Fluzone, Soluvia, Intanza), hollow MN injection systems (eg, Micronjet, AdminPen), and dissolving MN patches (eg, DrugMAT and VaxMAT). Parathyroid hormone (PTH)-loaded dissolving MN patch (MicroCor) and PTH-coated titanium MN patch (Macroflux) have recently completed Phase II clinical trials for osteoporosis treatment, while Zolmitriptan-coated titanium MN patch for acute migraine treatment has completed Phase III trial. There are ongoing clinical trials for MN patches to treat diabetes, hypoglycemia, migraine, influenza, psoriasis, cutaneous lymphoma, and so on. [243] Given the advantages of MN technology, the market share of MN-based TDD systems likely will grow quickly. However, MN technology still faces challenges in loading capacity, low-cost high-throughput production, and drug compatibility with the fabrication process. In principle, it can be combined with other TDD methods such as iontophoresis for synergistic actions. Although still in the immature stage, MN-technology is also promising for minimally invasive transdermal diagnostic systems. A few MN devices for painless collection of capillary blood and skin ISF are in clinical development (eg, HemoLink, Renephra). A single-use portable capillary blood collection device (TAP) containing ≈30 MNs has recently received clearance from US FDA. Despite facing challenges, MN-biosensors offer new prospects for multiplex biomarker detection in skin ISF.
Commercial wearable devices are available to measure sweat biomarkers, temperature, electrodermal activity, blood pressure, EEG, EMG, and so on in a continuous and noninvasive manner. The US FDA recently cleared two new features of the AppleWatch, that is, monitoring heart rate and detecting irregular heart rhythm. A graphene-based wearable sensor (GraphWear) is under development for real-time tracking of sweat glucose and electrolytes, particularly for athletes. The bluetooth enabled wearable smart patch (Echo H2) is being developed for real-time monitoring hydration, lactic acid, and calories levels for athletes. The key challenge for wearable devices is biocompatible, stable, conformal, comfortable, durable, and intimate interface with the curved and moving human body. To meet this challenge, developing skinlike self-healable interfacing materials and flexible components for energy supply and data transmission is critical. In addition, integrating these devices with TDD systems are highly desirable for feedback therapy to enhance efficacy and reduce side-effects.
In summary, skin provides an easily accessible interface for transdermal drug delivery and diagnosis, with noninvasiveness or minimal-invasiveness, convenience, and good patient compliance. Recent advances in scientific and technology developments have brought transdermal theranostics to a new level for personalized, home-based, and long-term management of chronic diseases.

CONFLICT OF INTEREST
The authors declare no conflict of interest.