Current Advancements in Transdermal Biosensing and Targeted Drug Delivery

Abstract

In this manuscript, recent advancements in the area of minimally-invasive transdermal biosensing and drug delivery are reviewed. The administration of therapeutic entities through the skin is complicated by the stratum corneum layer, which serves as a barrier to entry and retards bioavailability. A variety of strategies have been adopted for the enhancement of transdermal permeation for drug delivery and biosensing of various substances. Physical techniques such as iontophoresis, reverse iontophoresis, electroporation, and microneedles offer (a) electrical amplification for transdermal sensing of biomolecules and (b) transport of amphiphilic drug molecules to the targeted site in a minimally invasive manner. Iontophoretic delivery involves the application of low currents to the skin as well as the migration of polarized and neutral molecules across it. Transdermal biosensing via microneedles has emerged as a novel approach to replace hypodermic needles. In addition, microneedles have facilitated minimally invasive detection of analytes in body fluids. This review considers recent innovations in the structure and performance of transdermal systems.

Keywords: Iontophoresis; drug delivery; electroporation; fluorescent biosensors; luminescent sensors; microfabrication; microfluidics; microneedle; transdermal biosensing.

Figure 1

Schematic illustration showing the pathways of topical and transdermal delivery, including electrically assisted delivery by iontophoresis, electroporation, and electroincorporation [7].

Figure 2

Reverse iontophoretic extraction flux rates of valproate as a function of time and subdermal concentration (a) in the range 209–1050 μM, and (b) in the range 21–104.5 μM. Reverse iontophoretic extraction flux rates of valproate and glutamate as a function of time relative to their values at 24 h. (c) Ratio of the reverse iontophoretic extraction flux rates of valproate and glutamate as a function of time and subdermal valproate concentration (range 21–104.5 μM). (d) Each data point represents the mean ± standard deviation (n = 6) [17].

Figure 3

(a) An iontophoretic drug delivery system comprising donor and receptor compartments along with a current source and controller. D⁺: cationic drug; M⁺: biological cations; X⁻: biological anions. (b) Vyteris Inc. LidoSite^TM topical system [18].

Figure 4

Glucose electrode inserted in subcutaneous tissue. Glucose diffuses from the intravasal compartment (G₁) into interstitial compartment (G₂0; it is then taken up by cells if insulin is present [19].

Figure 5

(a) Apparent extraction of glucose by reverse iontophoresis in 2 h. Reverse iontophoretic extraction of (b) titrated water and (c) ¹⁴C-labeledethanol in 2 h [20].

Figure 6

Tattoo-based platform for noninvasive glucose sensing. (1) Schematic of the printable iontophoretic-sensing system displaying the tattoo-based paper (purple), Ag/AgCl electrodes (silver), Prussian Blue electrodes (black), transparent insulating layer (green), and hydrogel layer (blue). (2) Photograph of a glucose iontophoretic sensing tattoo device applied to a human subject. (3) Schematic of the time frame of a typical on-body study and the different processes involved in each phase. (A) Chronoamperometric response of the tattoo-based glucose sensor to increasing glucose concentrations from 0 μM (dashed) to 100 μM (plot “l”) in buffer in 10 μM increments. (B) Interference study in the presence of 50 μM glucose (plot “a”), followed by subsequent 10 μM additions of ascorbic acid (plot “b”), uric acid (plot “c”), and acetaminophen (plot “d”). Potential step to −0.1 V (vs. Ag/AgCl). Medium was phosphate-buffer with 133 mM NaCl (pH 7) [21].

Figure 7

Molecular imprinting of cyclodextrins as receptors of nanometer-scaled templates. (a) The cyclodextrin is cross-linked by diisocyanate, in dimethylsulfoxide, in the presence of the template. (b) The vinyl monomer of cyclodextrin is copolymerized with methylenebisacrylamide, in water, in the presence of the template [22].

Figure 8

Construction steps for the planar configuration of the screen-printed amperometric transducer. (a) PETE support material; (b) printing of conducting silver basal track; (c) printing of insulation layer; (d) printing of Ag/AgCl pads, the central circular pad is the iontophoresis electrode for reverse iontophoresis while the other pad is the reference electrode; (e) printing of graphite pads, which serve as the working and counter electrodes [24].

Figure 9

Construction of microneedle assembled screen-printed electrode (SPE)-based amperometric sensor for hydrogen peroxide. (A) Electrode assembly connected to a dedicated electrochemical detector; (B) Microneedle assembled SPE in three electrode configuration; (C) Electrode connector fixed to SPE: WE = working electrode, RE = reference electrode and CE = counter electrode) [26].

Figure 10

Construction of a microneedle assembly-based screen-printed electrode (SPE)-based potentiometric potassium ion sensor [27].

Figure 11

Size-dependent absorption and emission spectra of CdSe semiconductor nanocrystals with various sizes. (a) Size-dependent luminescence color and (b) schematic presentation of size, color, and emission wavelength of CdSe–ZnS QDs. (c) Absorption (solid lines) and emission (broken lines) spectra of CdSe QDs with various sizes [33].

Figure 12

(a) Schematic illustrations of on-demand transdermal drug delivery using near-infrared (NIR) light-responsive microneedles (MNs), composed of polycaprolactone and NIR absorbers as well as silica-coated lanthanum hexaboride (LaB₆-SiO₂) nanostructures. (b) Temperature changes of R6G-loaded microneedles after continuous exposure to 808 nm NIR laser light at an output power of 7 W cm⁻². (c–h) Infrared thermal images of microneedles after continuous irradiation for 10, 30, 90, 120, 240, and 300 s. The insets in (b) show the light-triggered melting behavior of microneedles [37].

Figure 13

Fabrication process of hollow silicon out-of-plane microneedles [41].

Figure 14

Schematic showing steps (a–f) used for assembly of the microneedle device. Images of microneedle array and cadaveric porcine skin after microneedle insertion. (g) Optical micrograph showing delivery of trypan blue into microneedle-fabricated pores within cadaveric porcine skin scale bar 1 = mm. (h) Optical micrograph showing hollow microneedles before insertion into cadaveric porcine skin. (i) Optical micrograph showing hollow microneedles after insertion into cadaveric porcine skin [42].

Figure 15

MTT viability data for cells grown on e-shell 200 acrylate-based polymer and glass cover slip. (a) MTT viability of human epidermal keratinocytes grown on e-Shell 200 acrylate-based polymer compared to glass cover slip. A and B denote statistical differences p < 0.05 between the polymer and the control. (b) MTT viability of human dermal fibroblasts grown on e-Shell 200 acrylate-based polymer compared to glass coverslip. A and B denote statistical differences p < 0.05 between the polymer and the control [42].

Figure 16

Breath and transdermal data for (a) Subject 1 and (b) Subject 2. (c) Two phase design of the data analysis system (left) and the blood/lung/breath analyzer and blood/skin/TAS systems (right) [44].

Figure 17

(a) Rat on its back; 1-mm thick water-tight stand-off between the shaved abdomen and array. Reservoir inside the stand-off filled with saline. US (ISPTP = 100 mW/cm²) applied to exposed group for 20 min; US not applied to the control. (b) In presence of glucose oxidase, glucose becomes gluconic acid via glucono-δ-lactone: hydrogen peroxide is generated as byproduct. Electrochemical biosensor uses the signal caused by oxidation of hydrogen peroxide. One glucose molecule generates an oxygen molecule, a water molecule, and two electrons in the reaction. (c) Voltage between working electrode and reference electrode, 0.7 V, applied from potentiostat controlled by computer. At the same time, current between two electrodes is determined by potentiostat and recorded in the computer [50].

Figure 18

The cymbal transducer made of piezoelectric material lead zirconate-titanate (PZT)-4 operated at a frequency of 20 kHz. The cymbal disk was placed between two titanium caps with air cavities beneath the caps, which convert the radial oscillations of the disk into flexure motions of the caps. Graph of the change in blood glucose over the 90-min transdermal insulin delivery experiment duration, comparing both controls (insulin only and ultrasound only) to the insulin and ultrasound results generated by the 2 × 2 array (insulin and ultrasound 2 × 2) and 3 × 3 array (insulin and ultrasound 3 × 3) [51].

Figure 19

(a) A single B-FIT device cell with all masking layers. (b-(i)) Cross-section of the released device, (b-(ii)) SEM micrographs of SU-8 dry etching using masking layer. (c) Fully functional B-FIT pictured on human arm [53].

Figure 20

(a) Schematics illustrating the principles of the fluorescence affinity hollow fiber sensor. In the absence of glucose, fluorochrome-labeled Concanavalin A is bound to the fixed glucose residues inside the porous beads (A-i). The beads are colored with dyes that prevent the excitation light from inducing Con A to fluoresce, keeping the fluorescence emission at 520 nm. After diffusion of glucose through the hollow fiber membrane (molecular weight cutoff, 10 kDa), Con A is displaced from the beads and diffuses out of them. Fluorochrome-labeled Con A becomes exposed to excitation light, resulting in a strong increase in fluorescence (A-ii). (b) Spectra of fluorescence and absorption of the various chromophoric components of the bead-based affinity sensor. (1) Excitation spectrum of fluorescein, (2) emission spectrum of fluorescein, (3) absorption spectrum of dye-labeled Sephadex beads. (c) Light microscopy picture of a small section of the hollow fiber that was filled with dye-colored Sephadex G150 beads. A bead fraction having a bead diameter of less than 25 µm was used in this study; the beads were obtained by sieving the original material (mesh size of 25 µm) [58].

Figure 21

(a) Photomicrograph of the host response to a pHEMA membrane implanted for 3 weeks. (b) Dependence of equilibrium water content, q_w, for flat sheet pHEMA membranes at 37 °C on crosslinking ratio. Experiments were performed in triplicate at each crosslinking ratio. (c) Dependence of creatinine permeability at 37 °C in flat sheet pHEMA membranes on equilibrium water content. For reference, the permeability of creatinine in a 1000 MWCO cellulosic dialysis membrane is 9.0 × 10⁻⁷ cm²/s [64].

Figure 22

(A) Schematic illustration of the fluorescent hydrogel fiber designed for long-term in vivo glucose monitoring. The fiber can be injected into subcutaneous tissues. The implanted fiber remains at the implantation site for a long period and transmits fluorescent signals transdermally depending on blood glucose concentration. The implanted fiber can be easily removed from the implantation site after use. (B) Fluorescent hydrogel fibers in a glass vial with a 50% glucose solution. The fibers are excited by ultraviolet light. The fluorescent image indicates that the glucose-responsive monomer is immobilized within the hydrogel fibers. (C) Inflammation indices of the mouse ears with the fibers one month after implantation. Inflammation was evaluated based on reddening, swelling and scab formation for a month. The inflammation index was obtained by summing the reddening, swelling, and scab formation scores. If the ear skin showed any reddening, reddening scores 1 point; similarly, swelling and scab formation were each also scored 1 point. PEG-bonded PAM induced less inflammation than PAM only. (D) Numbers of mice showing transdermal transmission of fiber fluorescence [67].

Figure 23

(A) Fluorescent microbeads fabricated by a microfluidic system. (A) Schematic diagram of the AFFD. The pregel solution flows into the inner channel of the AFFD, and the silicone oil flows into the outer channel. The droplets in silicone oil are collected from the solution with TEMED at 37 °C; they turn into microbeads after gelation. (B) Glucose responsiveness of GF-beads. (i–iii) Fluorescent images of GF-beads at a glucose concentration of 0 mg·dL⁻¹, 250 mg·dL⁻¹, and 1000 mg·dL⁻¹, respectively, (iv) The fluorescence intensity of GF-beads (n = 19) changes according to the increase and decrease in glucose concentration. The overlapping curves are evidence of the reversible reaction between GF-beads and glucose. (C) The fluorescence intensity traces blood glucose concentration. The time lag mainly comes from the time response between the interstitial fluid glucose concentration and the blood glucose concentration [68].

Figure 24

(a) Complexation of DSPBA probes with aqueous free cyanide. (b) Ratiometric response of the probes in water with increasing cyanide concentrations. (c) Fluorescence emission of o-DSPBA in water with time when excited at 475 and 375 nm [70].

Figure 25

(a) FL intensity for functionalized InP QWs at different cyanide concentration. (b) Relative intensity for functionalized InP QWs for various metal anions at 50 pM concentration [71].

Figure 26

Number of hypoglyceamia events (A), (B) median number of hypoglyceamic events within 72 h per patient before (CGMS-1) and after (CGMS-2) insulin therapy corrections [92].