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Publicly Available Published by De Gruyter July 27, 2013

Laser ablation-enhanced transdermal drug delivery

Laserablations-verstärkte transdermale Medikamentenverabreichung

  • Yajuan Li , Liangran Guo and Wei Lu EMAIL logo

Abstract:

Transdermal delivery offers an excellent route for drug and vaccine administration. Nonetheless, the lipid-rich outer stratum corneum layer of the skin presents a critical challenge to drug penetration. Laser ablation perforates epidermis through selective photothermolysis, making skin more permeable to hydrophilic and macromolecular drugs such as peptides, proteins, and genes. This review summarizes recent applications to laser ablation-enhanced transdermal delivery. Needle- and pain-free transcutaneous drug delivery via laser ablation provides an alternative approach to achieve local or systemic therapeutics.

Zusammenfassung:

Die transdermale Applikation bietet eine hervorragende Möglichkeit für die Verabreichung von Medikamenten und Impfstoffen, stellt aber aufgrund der lipidreichen äußeren Hornschicht der Haut eine besondere Herausforderung dar. Mittels Laserablation kann die Epidermis durch selektive Photothermolyse perforiert werden, was die Haut durchlässig für hydrophile und makromolekulare Substanzen wie Peptide, Proteine und Gene macht. Der vorliegende Review-Artikel fasst die neuesten Anwendungen der Laserablations-verstärkten Verabreichung von Wirkstoffen zusammen. Die nadel- und schmerzfreie transkutane Prozedur mittels Laserablation bietet einen alternativen Ansatz zur Verabreichung lokaler oder systemischer Therapeutika.

1 Introduction

The development of transdermal drug delivery systems (TDDS) is attractive as skin is the largest organ. TDDS have distinct advantages over oral administration or injections since they directly deliver drugs into the skin or even the systemic circulation, avoiding first-pass clearance of liver thus enhancing bioavailability. TDDS provide sustained and steady-state pharmacokinetics, therefore decreasing administration frequency and improving patient compliance. Further, TDDS avoid the limitation of injections such as pain, accidental needle-sticks, and possible side effects due to transiently high plasma drug concentration [1–3].

However, skin presents a natural barrier to protect the body from xenobiotic substances. It forms multilayers in the epidermis, which include stratum corneum (SC), stratum lucidum, stratum granulosum, stratum pinosum and stratum basale from topical toward dermis. The SC is the outermost layer and consists of dead keratinocytes or corneocytes intercalated with lipids [4]. This 10- to 20-μm thick layer is the formidable barrier preventing most drug molecules from permeation. Only a lipophilic drug with molecular weight (MW) <500 Daltons (Da) is able to penetrate the skin barrier, such as clonidine, fentanyl, and lidocaine [3, 5].

A variety of methods have been tried to enhance the permeability of the SC. Chemical enhancers promote the drug penetration through the SC by disrupting the highly ordered bilayer structures of the intracellular lipids in the SC [3]. Conventional chemical enhancers such as Azone (1-dodecylazacycloheptan-2-one) as well as newly developed biochemical enhancers like peptides are of interest [6, 7]. However, chemical enhancement has been shown to have little impact on delivery of hydrophilic drugs and macromolecules and to cause irritation to living cells in the deeper skin [3]. On the other hand, physical enhancement techniques including mechanical and thermal approaches have been used to make micrometer dimensions of disruptions to SC structures. These micro-scale disruptions create channels of sufficient dimensions for passage of macromolecules. The thermal ablation activated by microheaters [8], radio frequency [9–11], superheated steam ejectate [12] or laser [13–17] is a non-invasive technique to selectively remove small portions of the SC. These perforations are temporary, since the layers of the SC are continually replaced through the natural process of desquamation [8]. Some physical enhancement technologies have been investigated in clinical trials of TDDS such as BA058 transdermal microneedle patch [18], transdermal basal insulin patch with microporation [19], teriparatide acetate TDDS [20], and electroacupuncture for opioid detoxification [21].

Laser ablation enhancement belongs to a physical approach that utilizes laser to perforate or remove the SC barrier in order to enhance the drug penetration. Water and pigments in the skin absorb the laser light energy and transform it into heat to achieve thermolysis of the skin. The heating duration must be controlled within microseconds in order to avoid heat propagation to deeper tissues [22]. The laser ablation approach enables precise control of depth of skin permeation, having the potentials for percutaneous delivery of biomacromolecules such as peptides, proteins, vaccines, and DNAs [15]. The present review focuses on recent progresses of laser ablation-enhanced TDDS.

2 Direct laser ablation enhancement

Although many types of lasers with a broad wavelength range (193–10,600 nm) are available in clinical practice such as ruby laser, neodymium:yttrium-aluminum-garnet (Nd:YAG) laser, alexandrite laser, carbon dioxide (CO2) laser and erbium:yttrium-aluminum-garnet (Er:YAG) laser (Table 1), only a few have been applied to transdermal delivery so far. Pulsed CO2 and Er:YAG laser are in common use for SC ablation [27]. The ruby laser (694 nm) and the alexandrite laser (755 nm) belong to near-infrared (NIR) lasers (650–900 nm). The NIR light causes little tissue absorption or minimal thermal effect [28], which is not sufficient to remove the SC. By contrast, the wavelengths of the CO2 and Er:YAG laser are 10,600 nm and 2940 nm, respectively. Both lasers directly induce heating and microporation of the skin through water excitation and explosive evaporation from the epidermis. However, the wavelength of the mid-infrared Er:YAG light matches a principal absorption wavelength for water molecules [13]. Compared with the CO2 laser, the Er:YAG laser is about 15 times better absorbed in skin [27]. Therefore, the Er:YAG laser has a much higher ablation efficacy and a lower ablation threshold [29]. The Er:YAG laser shows the reduced thermal damage even in deeper crater holes in comparison with the pulsed CO2 laser [27, 29]. These favorable properties make the Er:YAG laser an ideal light source not only for skin surgery but also for enhanced transdermal drug delivery. A comparison of three sources of laser, the ruby, CO2 and Er:YAG laser, on the skin permeability for 5-fluorouracil (5-FU) showed that the ruby laser only moderately enhanced the drug flux [23]. The Er:YAG laser with fluences at 0.8–1.4 J/cm2 enhanced the flux of 5-FU by 53–133 times than untreated skin. The CO2 laser increased penetration of 5-FU by 36–41 times under the fluences of 4.0 and 7.0 J/cm2 with certain thermal effects [23].

Table 1

Light sources for thermal ablation-enhanced transdermal drug delivery.

Light sourceWavelength (nm)Pulse durationRole of the light in TDDSOther characteristics
Traditional Er:YAG laser2940250–400 μsEpidermal ablation; dermal removal [23]One beam; spot ablation
Fractional Er:YAG laser294010–300 μsFractional epidermal removal [15, 24]Microbeams; shorter pulse; less damage to epidermis; fractional photothermolysis
Short pulse CO2 laser10,60050 msEpidermal removal; dermal thermal injury [23]Ablation; vaporization
Nd:YAG in tandem with Ti:sapphire laser690–95015 nsPhotothermal ablation [25]Photothermal conversion by HCuSNPs
Xenon bulb750–1000Continuous lightPhotothermal ablation [26]Surface plasmon resonance of gold nanorods

Laser-induced thermal ablation heats the skin to hundreds of degrees for very short periods of time (micro- to milliseconds) to disrupt the SC [3]. The extent of structure alteration of the SC is proportional to the temperature locally elevated, i.e., (i) disordering of SC lipid structure by temperature between 100°C and 150°C, (ii) disruption of SC keratin network structure by temperature between 150°C and 250°C, and (iii) decomposition and vaporization of keratin to create micron-scale holes in the SC by temperature above 300°C [22]. Correspondingly, skin permeability was increased from a few fold to three orders of magnitude [22]. For thermal ablation-enhanced TDDS, high energy of laser with pulse duration less than microseconds is required because it generates limited or negligible heat transfer to surrounding tissue [13–16]. The microsecond-pulsed laser steepens the temperature gradient across the SC. The skin surface is extremely hot but not the viable epidermis and deeper skin tissues [12]. This technique referred to as “cold ablation”, thereby, largely eliminates side effects and vastly improves safety.

In physically enhanced TDDS, the controllable depth and wound area of skin perforation by the laser ablation should be well considered. Based on the clinical data from microneedle and thermal ablation-enhanced transdermal delivery, micron-scale defects in the SC are well tolerated by patients as long as there is no significant damage to living cells in the viable epidermis and dermis [3]. To solve this issue, a laser microporation technology called P.L.E.A.S.E.® (Precise Laser Epidermal System; Pantec Biosolutions AG, Liechtenstein) has been developed by using a diode-pumped fractional Er:YAG laser (Figure 1A) [14, 15]. Instead of conventional Er:YAG laser in clinics that ablates a 7-mm spot on the skin, the system generates a matrix of identical micropores with 100–150 μm wide of each (Figure 1B). Since the concentrated laser beam is divided into microbeams, P.L.E.A.S.E.® efficiently and fractionally ablates skin with less damage (Figure 1C) [14]. In addition, the pulse duration of the fractional laser from P.L.E.A.S.E.® is shorter than conventional Er:YAG laser to ensure the localization of heat transfer to the skin surface without allowing heat to propagate to the viable tissues below. This technology is patient-friendly since it is programmed to precisely control the number of micropores in unit area and depth of micropores based on the laser fluence [15].

Figure 1 P.L.E.A.S.E.® technology. (A) The photograph of the hand-held device. (B) Formation of a micropore array in the skin surface using the device. (C) Hematoxylin and eosin staining of micropores created in porcine ear skin after laser microporation using the device at fluences of 4.53 J/cm2, 22.65 J/cm2 or 135.9 J/cm2 (from left to right). From reference [15] (© Elsevier; reprinted with permission).
Figure 1

P.L.E.A.S.E.® technology. (A) The photograph of the hand-held device. (B) Formation of a micropore array in the skin surface using the device. (C) Hematoxylin and eosin staining of micropores created in porcine ear skin after laser microporation using the device at fluences of 4.53 J/cm2, 22.65 J/cm2 or 135.9 J/cm2 (from left to right). From reference [15] (© Elsevier; reprinted with permission).

3 Photothermal nanoparticle-mediated laser ablation enhancement

The development of nanotechnology brings a breakthrough in the limited application of NIR lasers in TDDS. Gold nanostructures such as nanoshells [30], nanorods [31], nanocages [32, 33], and hollow nanospheres [34] possess unique optical properties due to strong and tunable surface plasmon resonance (SPR). They can be synthesized to specifically absorb NIR light and convert photo energy into thermal energy to raise the temperature of the surrounding tissue [30, 35]. Nanoparticles with the property of such photothermal coupling effect are called photothermal nanoparticles. Gold photothermal nanoparticles can be applied to photothermal ablation therapy of tumor cells [36–39], as well as the NIR laser-controlled drug release [40–44]. The absorbance of NIR light is desirable because it causes minimal thermal injury to normal tissues with optimal light penetration [28, 45]. Recently, a surfactant/protein/gold nanorod complex has been applied to transdermal delivery of proteins [26]. The solid-in-oil dispersion system has been formulated through incorporation of gold nanorods as the photothermal ablation enhancer to disrupt the skin barrier. This approach effectively enhances the protein permeation through the skin in vitro and induces an immune response in vivo [26]. In this application, instead of pulsed laser, a xenon lamp that required high light power (6 W/cm2) and long duration of light exposure (20 min) has been used to ablate the SC [26]. Therefore, the heat propagation to the deeper tissue could be a major concern.

Semiconductor copper sulfide (CuS) nanoparticles (CuSNPs) are a new class of photothermal nanoparticles that provide an alternative to gold analogs. Compared to gold, CuS is much less expensive [46]. Irradiated with an NIR laser, CuSNPs generate heat for photothermal destruction of tumor cells [46–49]. Hollow CuSNPs (HCuSNPs) have been utilized for photothermal ablation-enhanced transdermal drug delivery [25]. A nanosecond-pulsed Nd:YAG laser in tandem with a Ti:sapphire laser (900 nm) has been used to induce rapid heating of the nanoparticles and instantaneous heat conduction. Such a type of laser with nanosecond pulse duration provides focused thermal ablation of the SC and minimizes skin heat accumulation. The average temperature of the irradiated skin area only increases to ∼40–50°C. The depth of skin perforation can be precisely controlled by adjusting the laser power. The skin disruption by HCuSNPs-mediated photothermal ablation significantly increases the permeability of macromolecule drugs, providing effective percutaneous delivery [25].

4 Drugs applied to laser ablation-enhanced transdermal delivery

In comparison with chemical enhancers that only improve the transdermal delivery of small molecules, laser ablation enhancement makes micrometer dimensions of disruptions to the SC structures suitable for the passage of both small and macro molecules such as 5-FU [23, 50], lidocaine [14], diclofenac [16], human growth hormone (hGH) [25], antithymocyte globulin (ATG) [15], ovalbumin (OVA) [26], polypeptides [24], fluorescein isothiocyanate (FITC)-labled dextran (FD) [13], nalbuphine [51], vitamin C [52], 5-aminolevulinic acid (ALA) [53], genes [54], and stem cells [55] (Table 2).

Table 2

Drugs/compounds used for transdermal delivery by laser ablation.

Drug/compoundIndication/purposeMolecule weight (kDa)Enhanced permeability (fold)Laser source
5-FUAntitumor0.130Up to 429Er:YAG [23, 50]
ImiquimodImmune response modifier0.2403Up to 127Fractional Er:YAG [24]
LidocaineLocal anesthetic0.2343Up to 18Fractional Er:YAG [14]
DiclofenacNon-steroidal anti-inflammatory drug (NSAID)0.29615Up to 119Fractional Er:YAG [16]
Vitamin CModel hydrophilic drug0.176Up to 260Er:YAG [52]
Methotrexate (MTX)Psoriasis or rheumatoid arthritisUp to 80Er:YAG [56]
hGHGrowth hormone deficiency22>1000Nd:YAG tandem with Ti: sapphire [25]
Antithymocyte/ basiliximabImmunosuppressive antibodies155/144Up to 145/ N/AFractional Er:YAG [15]
OVAAntigen44N/AXenon light [26]
Beta-galactosidase (bGal)Antigen465N/AFractional Er:YAG [57]
Recombinant Phl p 5Grass pollen allergen38
Equine heart cytochrome cModel proteins12.4N/AFractional Er:YAG [58]
Urinary follicle stimulating hormone30
FITC-labeled bovine serum albumin70
PeptidesModel peptides0.716–2.864Up to140Er:YAG [59]
DextranModel hydrophilic macromolecule4–150>1Continuous wave fiber laser [17], Er:YAG [13], fractional Er:YAG [24]
Insulin (hexameric)Diabetes mellitus36N/AEr:YAG [13]
NalbuphineAnalgesic /hydrophilic model drug0.357Up to 194Er:YAG [51]
IndomethacinNSAID /lipophilic model drug0.357Up to 30
ALAAnti-tumor/photosensitizer0.131133Fractional Er:YAG [53]
Antisense oligonucleotidesTest model5–8Up to 30Er:YAG [54]
Plasmid DNAExpress green fluorescent protein4.7 k base pairsUp to 164
RNASmall interfering RNA9.266Up to 10Er:YAG [60]
Adipose-derived stem cellsWound healingN/AFractional Er:YAG [55]

Note: “N/A” represents no numerical fold reported.

Dextran, a hydrophilic macromolecular model drug, was used to evaluate the skin permeation. By using a laser with a fluence above 1.7 J/cm2, the transdermal transport of FDs with MWs ranging from 4.4 kDa to 77 kDa was significantly enhanced. The possible mechanism could be ablation of the SC layer, photomechanical stress on intercellular regions, and alterations of the morphology and arrangement of corneocytes by the Er:YAG laser. Further, the transdermal delivery of hexameric insulin was higher than that of 38-kDa FD, suggesting the potential of laser ablative transdermal delivery of insulin [13].

ATG and basiliximab, two marketed antibodies for the induction of immunosuppression, were studied with fractional Er:YAG laser [15]. The result showed that the increase of pore numbers and laser fluence promoted the transdermal permeation of the antibodies. Total delivery of ATG at 24 h after laser treatment (900 pores, at a fluence of 45.3 J/cm2) increased 82.8-fold over the control (untreated skin). Increasing laser fluence from 22.65 to 135.9 J/cm2 enhanced total ATG delivery from 1.70±0.65 to 8.70±1.55 μg/cm2, respectively. Similar penetration enhancement was observed in basiliximab. Moreover, the in vitro and in vivo result was well correlated in a mouse model [15].

Topical delivery of DNA and RNA were also enhanced by laser ablation [54, 60]. With Er:YAG treatment, in vitro permeation of antisense oligonucleotides (ASOs) increased 3–30-fold, depending on the laser fluence and the MW of ASO. In vivo results showed an enhanced expression of plasmid DNA in the epidermis and subcutis [54]. Besides, it was also found that the delivery rate of siRNA was raised by several times by the laser application [60].

Laser-enhanced transcutaneous protein delivery provided a non-invasive immunization method [15, 57, 59]. The laser-induced microporation allowed high levels of antigen uptake. Further, transdermal delivery of vaccine targeting the potent epidermal Langerhans and dermal dendritic cells induces a strong immune response at much lower doses than hypodermic injection [61]. Transcutaneous application of OVA via laser-generated micropores using the P.L.E.A.S.E.® device induced equal or higher immune responses compared to immunization by subcutaneous injection [57]. In addition, targeting different layers of the skin had the potential to bias different T-cell polarization patterns [57]. The laser ablation enhancement followed by transcutaneous immunization of lysozyme with 129 amino acids (14,307 Da) induced antigen-specific immunoglobulin IgG in the serum by 3-fold compared to the control without laser treatment [59].

In addition to delivering drug compounds, laser ablation-enhanced transdermal delivery of adipose-derived stem cells (ADSC) was explored for wound healings [55]. After fractional Er:YAG laser treatment, bromodeoxyuridine (BrdU)-labeled ADSC was applied to the laser treated areas. After 4 and 48 h, 12% and 5.5% of the stem cells were found in the pretreated tissue, respectively [55]. This encouraging result furthered the studies to optimize the technology for future clinical applications.

Because of high photothermal conversion effect, the gold nanoparticles were utilized to achieve thermal ablation of skin to enhance transdermal delivery of OVA [26]. In this study, a solid-in-oil dispersion was formulated to incorporate both the gold nanorods and the drug. Therefore, the nanodispersion exerted two modules upon NIR light irradiation, i.e., thermal ablation of the SC by the gold nanorods and enhancement of skin permeation of OVA. In vivo experiment showed significant increase of immune response for the gold nanorod-OVA solid-in-oil dispersion with NIR light treatment than other groups [26]. Another study investigated the use of HCuSNPs as photothermal ablation enhancers [25]. The permeability of hGH in skin applied with HCuSNPs plus NIR laser was increased by three orders of magnitude in comparison with that of the intact skin. In vivo study showed that transdermal delivery of hGH using the HCuSNP-mediated photothermal ablation technique reached an average bioavailability of 83% relative to that of the subcutaneous injection. The peak drug concentration through transdermal delivery was only one-third of that via subcutaneous delivery [25]. This is clinically beneficial because it reduces the risk of side effect related to high concentrations and controls the drug concentration at a relatively stable level.

5 Conclusion

In conclusion, this review has discussed recent progress in laser ablation technology to enhance transdermal drug delivery. The success of delivery relies on locally thermal ablation of the SC. By adjusting the laser fluence and exposure time, the depth of the microporation can be controlled without harming the deeper living tissues such as the dermis. The microchannels allow skin permeation of hydrophilic and macromolecular compounds. Particular interest has been shown in the development of the photothermal nanoparticles that mediate photothermal ablation of skin and deliver drug in a single setting. As a clean, needle-free and non-invasive approach, laser ablation enhancement technology shows great potential for the future.

Conflict of interest statement

The authors declare no conflict of interest


Corresponding author: Wei Lu, Department of Biomedical and Pharmaceutical Sciences, College of Pharmacy, The University of Rhode Island, 7 Greenhouse Road, Kingston, RI 02881, USA, e-mail:

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Erhalten: 2013-5-28
Revidiert: 2013-7-1
Angenommen: 2013-7-3
Online erschienen: 2013-7-27
Erschienen im Druck: 2013-11-1

©2013 by Walter de Gruyter Berlin Boston

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