Gratieri T, Kalia YN (2011) New technologies for topical delivery and their application in aesthetic dermatology. Kos Med 32: 165 – 175.

1. Introduction

In the past, cosmetics have been defined solely as a category of products intended to be applied to the human body for cleansing, beautifying, promoting attractiveness or altering appearance. Thus, evaluation of efficacy was essentially a subjective affair. The recent addition of active ingredients to “cosmetic” formulations has made more objective evaluations of their efficacy possible. However, the term “cosmetic” is no longer considered adequate and terms such as “quasi-drug”, “cosmeceutical” and “active cosmetic” have been introduced to denote prescription items or active ingredients that addressed skin appearance issues [1]. For the purpose of this review we use the terms “cosmetic” and “cosmeceutical” interchangeably.

 

The active ingredients present in cosmetic formulations exert their action in the viable epidermis or dermis. To do this, they must penetrate the outermost layer of the skin, the stratum corneum (SC) – whose thickness varies significantly in different regions of the body from ~20 µm in the forearm to ~200 µm in the palm [2] – and which is the principal barrier against the entry of microorganisms and other foreign substances. It is composed of dead cells, called corneocytes that are surrounded by a lipid extracellular matrix. The lipid-rich intercellular space and high degree of tortuosity combine to limit transepidermal water loss – at the same time they restrict the transport of many cosmeceuticals into and across the skin. Therefore, several permeation enhancement technologies have been developed to overcome the SC barrier and so increase delivery.

 

Such enhancement technologies can be divided into chemical or physical methods (Fig. 1). In this scheme, chemical methods include all modifications performed on the active ingredient or the formulation. They can be done to increase stability and/or the capacity to cross biological barriers such as the SC [3, 4]. For example, topical corticosteroids are usually applied as mono- or di-ester derivatives or acetonides in order to increase lipophilicity [5]. Similarly, esterification of the carboxylic acid group in naproxen has been shown to increase dermal permeation [6]. In addition, concentration and degree of saturation of the active ingredient in the vehicle (measures of the thermodynamic activity) are parameters that can be modified to enhance skin penetration. Encapsulation processes are also an alternative to increasing the amount of substance in the formulation (i.e., its loading) [7]. Many carriers using particulate and vesicular systems are employed in cosmetic preparations not only to increase permeation but also to improve general stability, compatibility with other ingredients and sensorial feeling – e.g., microvesicles, nano/microemulsions [8], liposomes [9], polymeric nano/microparticles [10], solid lipid nanoparticles [11, 12], nanostructured lipid carriers [13, 14].

Moreover, formulation components can be used to alter SC barrier function transiently; lipids, surfactants, esters or alcohols can act as penetration enhancers to (i) disrupt or fluidize the lipid lamellae and increase diffusion of the substance, (ii) improve partitioning of the active ingredient into the SC by increasing either its thermodynamic activity in the formulation or its solubility in the membrane [15, 16].

However, when chemical enhancement methods are insufficient, other artifices may be used, the so-called physical methods of absorption enhancement. Their mechanisms of action can be very different: some act on the skin by removing the SC to enhance passive diffusion, others like iontophoresis, act on the molecule by using a second driving force, in addition to the concentration gradient, to increase penetration of active ingredients. This review will focus on the main physical technologies developed for promoting topical and transdermal delivery and how these new technologies can be applied in the field of aesthetic dermatology.

 

2. Microdermabrasion

Microdermabrasion or particle resurfacing is an FDA-approved cosmetic procedure that was developed in the 1980s to treat signs of photo-aging (fine lines, wrinkles, hyper-pigmentation), tattoos, and superficial scars [17]. It is a process that uses an abrasive component, usually Al₂O₃ or NaCl crystals and negative pressure for superficial peeling of the outer surface of the skin. The extent of peeling is determined mainly by the crystal flow rate and exposure time [18]; this enables precise control of skin ablation and reduces the risk of deep tissue damage due to operator error [19 – 21]. The microdermabrasion device works by placing a handpiece on the skin. Upon tip occlusion, the crystals flow into the inlet port and abrade the skin. Simultaneous vacuum suction is used to collect used crystals and skin debris (Fig. 2). This closed-loop system prevents cross-contamination between patients and exposure of medical personnel [22].

 

It has been suggested that the inflammatory response induced by damaging the SC can stimulate fibroblast activity along with new dermal collagen deposition, resulting in the positive cosmetic results seen after microdermabrasion treatment [23, 24]. The potential of microdermabrasion for increased skin delivery has been demonstrated for active cosmetic ingredients such as vitamin C [25] and 5-aminolaevulinic acid (ALA), which is used in photodynamic therapy [26] for the treatment of non-melanoma skin cancer [27, 28], actinic keratoses [29] and photoaging [30, 31]. Efficacy depends on the absorption of the drug precursor (ALA) and its conversion via the heme cycle into protoporphyrin IX (PpIX), a photosensitizing agent which, upon irradiation with a light source, forms reactive oxygen species that destroy target cells [28]. ALA is a hydrophilic molecule and a zwitterion at physiological pH that is poorly absorbed by the lipophilic SC [32].

 

It has been shown in a pig skin model in vitro that microdermabrasion for various treatment periods resulted in a 5- to 15-fold enhancement in the skin permeation of ALA. A linear relationship was observed between drug flux and (i) treatment duration and (ii) the etched thickness of the SC [26]. Enhanced skin delivery of ALA was also demonstrated after microdermabrasion using nude mice in vivo. Confocal laser scanning microscopy (CLSM) of microdermabrasion-treated skin revealed intense red fluorescence of ALA-transformed protoporphyrin (PpIX) in the epidermis and upper dermis [3].

 

The use of microdermabrasion for delivery of active ingredients is still an expanding field. A better understanding of the relationships between microdermabrasion parameters and the effects on SC barrier function in different locations of the body will be useful to achieve controlled penetration enhancement while minimizing the risk of damage to deeper tissues.

 

3. Microporation technologies

3.1. Microneedles

Microneedle arrays typically consist of a plurality of microprojections attached to a support. They were proposed as a hybrid to overcome the limitations of hypodermic needles and conventional patch systems (Fig. 3). The microneedles are inserted in to the skin crossing the whole SC in order to open channels through which small or large molecules and even nanoparticles can penetrate [34 – 36].

 

Silicon microneedles have been proposed as a physical method to enhance ALA delivery in to the skin [37]. Puncturing excised murine skin with 6×7 arrays of microneedles 270 μm in height, with a diameter of 240 μm at the base and an interspacing of 750 μm led to a significant increase in the transdermal delivery of ALA released from a bioadhesive patch containing 19 mg ALA per cm²; studies with porcine skin showed that the ALA was localized in the upper regions of the excised tissue [37]. Skin pretreatment with biodegradable polymeric microneedles has been shown to be safe for human use. Microneedles with a length of 600 μm and a width of 300 μm with an interspacing of 300 μm in arrays with 121 microneedles per cm² enabled a reduction of ALA concentration (i.e., loading) in the formulation and the application time for treatment [38]. Microneedles have been studied for the transdermal delivery of peptides and proteins, e.g., desmopressin [39, 40], insulin [41, 42] and human growth hormone [40] and, in particular, vaccines [43 – 46].

 

One of the drawbacks of microneedle patches was the difficulty in applying them over large areas due to skin deformation under the patch. To address this issue, microneedle rollers (Fig. 4) that were originally developed to cause microdamage to the skin and to induce collagen production, started to be used for drug delivery [47, 48]. Recently, a microparticulate melanoma cancer vaccine (~ 2 µm) was delivered in mice using such technology. In order to study the efficacy of vaccination, the mice were challenged with the live S-91 tumour cells and the transdermally-vaccinated mice showed no measurable tumour growth when observed for 35 days after tumour injection [43].

 

Microneedle rollers have also been applied in the field of aesthetic dermatology. Full-face photorejuvenation treatment performed in 21 patients with 630 nm light and broadband pulsed light after multiple passes with a microneedle roller and 1 hour ALA incubation was shown to be well tolerated and produced satisfactory results. The needles were 300 µm in length and 108 µm in width and resulted in confluent mild erythema without bleeding [30]. Another study showed that application of a polymeric microneedle-roller (17 microneedles per array bar, square pillar type geometry, 500 µm length with sharp ends) to shaved rat skin increased ascorbic acid permeation 10.54-fold and resulted in faster hair growth as compared to the control (no microneedle-roller treatment) [49].

 

Thus, microneedle rollers seem to be a simple, easy-to-use and cost effective alternative for the enhancement of active ingredients. They are available in dimensions from 100 µm to more than 1 mm in length [50]. However, it is important to note that needles with lengths > 500 µm have been shown to cause bleeding [51] which may discourage its use. Moreover, although numerous studies have been reported, the majority have been done using murine or rat skin and it should be noted that their mechanical perforation is much easier and requires less force than human skin.

 

3.2. Laser Ablation

Laser light is monochromatic (of one wavelength), coherent (the peaks and troughs occur synchronously or in phase) and collimated (the light rays are parallel and do not diverge like conventional light) [52]. When it is applied to the skin it is absorbed and the energy is converted to heat. It has been used to induce epidermal and dermal remodelling in various dermatological applications including resurfacing of rhytides, scars and photodamage [53-55]. It has been reported that during laser treatment, the thermal effects account for visible collagen shrinkage and immediate skin tightening. In the longer term, it leads to the formation of granulation tissue and new collagen (fibroplasia) [56]. Although the exact mechanisms for cosmetic improvement have not yet been completely determined, one hypothesis for the effects of the erbium:yttrium-aluminum-garnet (Er:YAG) laser, is that Reactive Oxygen Species (ROS) may be produced which, in relatively low concentrations, stimulate signal transduction processes for transcription factor activation, gene expression, muscle contraction and fibroblast growth, thus playing a key role in collagen and extracellular matrix formation [57].

 

Besides being commonly used in dermatological applications, lasers can ablate the SC for enhanced transdermal delivery. The basic principle is that the laser, emitting light at a selected wavelength, excites specific molecules in the skin; their superheating within the space of a few microseconds and subsequent explosive evaporation lead to the formation of pores [58-60]. The Er:YAG laser is currently the most widely investigated laser for transdermal applications. It emits light at 2940 nm, corresponding to the excitation wavelength of water, which enables the laser to ablate skin with minimal residual thermal damage, thereby reducing the risks of post-inflammatory hyperpigmentation [54]. Studies have shown that Er:YAG ablation (beam diameter 7 mm) effectively enhanced delivery of several molecules, including nalbuphine, morphine, buprenorphine [61], lidocaine [58], indomethacin [62], methotrexate [63], 5- fluorouracil [64], vitamin C [65, 25] and ALA [26, 66]. Laser treatment followed by skin vaccination with a lysozyme antigen in mice in vivo produced a 3-fold increase in antibody levels in the serum [67]. The delivery of siRNA was also increased (2.4- to 10.2-fold) after Er:YAG laser pre-treatment in nude mouse skin in vitro when compared with the non-treated group [68].

 

The effect of four physical enhancement methods on ALA delivery was compared using pig skin in vitro; the increases in transport following Er:YAG laser ablation, iontophoresis, microdermabrasion and electroporation were 243-, 15-, 5- to 15- and 2-fold, respectively [26]. The effect of Er:YAG laser ablation on ALA delivery was also tested in mice in vivo; laser treatment produced a higher accumulation of PpIX within superficial skin and subcutaneous tumours as compared to those of the non-treated group (Student t-test, p < 0.05). The enhancement ratios observed with laser treated skin ranged from 1.7- to 4.9-fold as compared to the control depending on the parameters used. The barrier properties of the skin irradiated by the laser had completely recovered within 3 days [69]. Higher efficacy and improved aesthetic results after treatment of recurring nodular basal cell carcinomas have been obtained using a combination of PDT with a ALA methyl ester and Er:YAG laser ablation [70] and more recently laser radiation was also shown to facilitate transdermal delivery of ALA without resulting in ulceration or scarring and enabled reduction of the ALA dose and shortened the duration of PDT therapy [66].

 

For permeabilizing the skin, the control of the laser power output – continuous or pulsed – and laser fluence (energy delivered per unit area) may ensure minimum tissue damage (Fig. 5) while specifically localizing the ablation area to the upper epidermis, lower epidermis or even the dermis [58, 71]. These techniques show great potential for indications where the active ingredient needs to exert its effect in a specific layer of the skin. More recently, fractional ablation, which results in microscopic treatment zones, that are rapidly re-epithelialised by the surrounding undamaged tissue, has been used to reduce the risk of infection and scarring [72, 73]. In some devices, a grid is used to split the laser beam into “microbeams” and manual orientation of the headpiece is required to make multiple passes. In some ways, these approaches are better suited to permeabilizing the skin to active ingredients than for improving the appearance – i.e., removing rhytides and repairing photodamage – since the mild to moderate alterations in the skin mean that the latter require multiple treatment sessions [74]. The recently developed P.L.E.A.S.E.^® fractional ablation system uses a scanner to generate the pore array meaning that the focus and energy used to create each micropore is identical – the pores are created sequentially with their number and depth being user-defined – pores can be “drilled” into the superficial or deep epidermis or even reach into the deep dermis by increasing the number of pulses used to create the pore [75]. The presence of the scanner reduces inhomogeneity and the risk of operator error. It has been used to increase delivery of lidocaine and diclofenac into the skin [76, 58].

 

3.3. Thermal/radiofrequency ablation

Another approach to microporate the skin is to apply heat (temperatures > 100 ºC) for short time intervals using an array of micro-scale heaters; thus, as with laser fractional ablation, only a small area of the SC is ablated within the treatment area [77]. The pores have definite size and shape depending on the geometry of the ablative elements. The depth is normally dependent on the temperature and duration of exposure to the heat source. Selective thermal ablation was used to create micropores of 30 µm diameter and 70 µm depth without causing necrosis in surrounding tissue [78]. Another study reported the formation of elliptical micropores with 80 µm width, 300 µm length and 40–50 µm depth [79].

 

Recent studies have shown that the changes in the SC after exposure to heat depend more on the temperature than on the duration of heating; longer application times cause greater heating of neighbouring tissues resulting in collateral damage and pain. It has been hypothesized that after heating the SC for 100 ms, 1 s and 5 s to (i) ~100–150°C, the small enhancement observed is due to disordering of the SC lipid structure; (ii) ~150-250°C, the one to two orders of magnitude enhancement is attributed to disruption of the SC keratin network structure and (iii) >300°C results in several orders of magnitude enhancement attributed to decomposition and vaporization of keratin to create micron-scale holes in the SC. Longer exposure times only led to greater increases in skin permeability at temperatures of 150-250°C [80]. Although it may seem impractical to heat the skin to such high temperatures, short heating periods are employed (< 1 s) which, in principle, restricts the ablation just to the SC in a minimally invasive manner [80]. The increase in temperature of micro-scale heating elements can be achieved using resistive electrodes or radiofrequency energy (Fig. 6).

 

Systems using resistive electrodes and radiofrequency energy technologies include, respectively, the PassPort^TM by Altea Therapeutics (Tucker, GA, USA) [81] and the ViaDor^TM by TransPharma Medical [82]. Both systems are reported to be safe and minimally invasive. Current clinical studies are reported to be underway for delivery of drugs including calcitonin, exenatide and hPTH [81, 83], a peptide fragment of parathyroid hormone (PTH). The projected cost of these devices and the possibility of home use might lead to the further development of the technologies and their application in cosmetology, especially for newer high molecular active ingredients used in cosmetic formulations.

4. Iontophoresis

Iontophoresis is a painless and non-invasive technique that involves application of a mild electric current (no more than 0.5 mA cm²) to enhance the penetration of hydrosoluble, ionized molecules into and through tissues [84, 85]. Two different mechanisms for permeation enhancement have been proposed, electromigration (EM) and electroosmosis (EO). EM occurs due to the movement of ions under the influence of the applied electric field, whereas EO occurs because at physiological pH the skin is negatively charged and acts as a cation-selective ion-exchange membrane generating a convective solvent flow in the anode-to-cathode direction. The practical consequence of EO is that it contributes to the permeation of cations but opposes the movement of anions. Furthermore, under physiological conditions, it also enables the transport of neutral molecules from the anode into the body (Fig. 7) [86].

Iontophoresis has been studied for several dermatological and cosmetic applications and the molecules studied include, ascorbic acid [87, 88], acyclovir [89], ALA [90-92], botulinum toxin [93, 94], several anti-cancer drugs as 5-fluorouracil [95, 96], 6-mercaptopurine [96], doxorubicin [97], methotrexate [98, 99], lidocaine.

An advantage of iontophoresis for ALA delivery was that therapeutic levels of protoporphyrin IX (PpIX) were achieved in the epidermis in 10 min or less even when applying low ALA concentrations in the formulation [90] – this represents a very short application time in comparison with the time required to achieve therapeutic levels without physical enhancement [100]. As ALA is expensive and degrades rapidly (via a second-order reaction), reducing the required dose is also a considerable advantage. Transport of ALA across the skin was 4-fold higher with iontophoresis relative to the passive application of a DMSO formulation [101]. A strategy to further improve iontophoretic delivery of ALA that was tested involved the use of ALA ester prodrugs, which have a net positive charge under physiological conditions and may be transported via electromigration [101-104]. Iontophoresis of a homologous series of 5-ALA esters (methyl, ethyl, butyl, hexyl and octyl) was evaluated in vitro. A more than 50-fold enhancement of the methyl ester compared to the zwitterionic parent ALA was observed, an effect that gradually decreased with increasing aliphatic chain length. Another iontophoretic approach to improve PDT is to iontophorese the photosensitizing agent directly as demonstrated for the meso-tetra-(N-methylpiridinium-4-yl)-porphyrin (TMPyP) and meso-tetra-[4-sulfonatophenyl]-porphyrin (TPPS4), charged porphyrin derivatives [105, 106].

 

In the pharmaceutical field, until very recently, it was thought that iontophoretic transport would be limited to peptides smaller than 1.5 kDa since electric mobility tends to decrease with molecular weight [107]. However, studies have shown that larger peptides and proteins can be successfully delivered through the skin, contradicting previous assumptions by showing that electromigration and not electroosmosis was the dominant transport mechanism for these macromolecules. Cytochrome c, a 12.4 kDa protein was successfully iontophoresed into and across porcine skin in vitro; electromigration accounted for ~90 % of total delivery [108]. Similarly, ribonuclease A, an RNA cleaving enzyme with a molecular weight of 13.6 kDa, was successfully iontophoresed through porcine and human skin in vitro and retained its enzymatic activity post-delivery. Again, electromigration was the major driving force accounting for > 80% of the total flux [83]. Both cytochrome c and ribonuclease A were iontophoresed at pH<pI (i.e. they were positively charged) and it was thought that iontophoretic transport of high molecular weight anions was not possible; however, it was recently shown that biologically active (anionic) ribonuclease T1 an 11.1 kDa protein could also be delivered by iontophoresis [109]. These results could open the door to completely new approaches for the non-invasive delivery of high molecular weight active ingredients into the skin.

 

It has been hypothesized (and shown in animal models) that appendageal structures in the skin (hair follicles and sweat glands) play a key role in iontophoretic transport [110]. Therefore, it is interesting to note reports describing the iontophoretic delivery of Botulinum toxin A. This is well known for its use to reduce the appearance of lines and wrinkles; however, it is also used to treat hyperhidrosis. The first report into the iontophoretic administration of Botulinum toxin A for use in axillary hyperhidrosis appeared in 1996 [111]. It is believed that the toxin works by inhibiting the release of acetylcholine at the neuromuscular junction and affecting the postganglionic sympathetic innervation of sweat glands [112]. In 2004, Botulinum toxin A iontophoresis was reported to have been successfully used to treat two patients [93]. This was followed by a pilot study performed in 8 patients who were refractory to conventional therapy [113]. An in vivo study found the toxin in hair roots, sebaceous glands and arrector pili muscle fibres of Wistar rats after only 10 min of iontophoresis [94]. Botulinum toxin A is a two-chain protein with a 100 kDa heavy chain joined by a disulphide bond to a 50 kDa light chain and has a pI of 6.06. Given its high molecular weight and the known transport iontophoretic transport pathways, it is probable that iontophoresis may lead to Botulinum toxin accumulation in hair follicles or sweat glands – but this may not be enough to encourage its iontophoretic application for more conventional cosmetic treatments. However, the results obtained so far merit further investigation on the potential of iontophoresis to increase the penetration of other peptides and smaller proteins used for rejuvenation. Moreover, further research on vehicles that could facilitate administration as well as guarantee drug stability is also required.

 

5. Electroporation

Electroporation or electropermeabilization is the transitory structural perturbation of lipid bilayer membranes due to the application of short high voltage electric pulses [114, 115]. Skin electroporation has been the subject of considerable research since its first report in 1993 [116]. Increased molecular transport after electroporation may occur due to different mechanisms, including electromigration during the pulses and enhanced diffusion through the treated skin during and after voltage application [114, 117, 118]. A contribution of electroosmosis is theoretically possible although the short duration of the pulses means that its role will be minor if not negligible [119]. In addition, thermal effects of electroporation may also play a role on skin permeabilization; however, the issue is still the subject of discussion [120-121].

 

More than 20 years ago, the combination of electroporation with bleomycin-chemotherapy was shown to be highly effective in producing tumour regression in vivo [123] revealing the potency of electroporation in oncology. The technique has since been widely applied for local chemotherapy of cutaneous tumours [124-128]. Synergy of electroporation and PDT has also been demonstrated [129]. Electric pulses were used to increase the transdermal transport of ALA [130]; a more than two-fold enhancement of PpIX production was observed with electroporative delivery, when compared with PpIX production after passive delivery. Superposition of a DC potential (constant current 6V) over the pulses resulted in a nearly five-fold enhancement of PpIX production as compared to the levels achieved with passive delivery. Since ALA–PDT relies on the intracellular conversion of ALA to PpIX, it is particularly important that the drug penetrate into the cells. In fact, several studies have demonstrated that electroporation facilitated cellular uptake of a number of molecules both in vitro [131-133] and in vivo [134, 135]. Therefore, the additional enhancement of PpIX production (up to three-fold over passive delivery) may be due to the cell permeabilizing effect of the electric pulses [136, 130].

 

Even though there is strong evidence of higher skin permeation of photosensitizing agents with electroporation, the majority of studies are focused on oncological treatments rather than cosmetic applications [137-139]. This may be due to the fact that the immediate side effect observed for electroporation is pain followed by strong muscle contractions and in some cases, mild, local muscle fatigue [140], which may make the everyday use of the technique in aesthetics rather impractical. Moreover, the problems associated with pulse application may be accentuated since facial skin is more sensitive than other areas; a study into the electroporation-assisted topical delivery of ascorbic acid showed that the tolerance threshold of facial skin for a single pulse was 50 V for 2 ms whereas for the forearm it was 70 V for 10 ms [141]. It is possible that new developments with respect to pulse protocol and electrode design may overcome such drawbacks since the use of closely spaced microelectrodes was shown to be able to constrain the electric field within the SC avoiding electric field distribution to deeper tissues, which contain sensory and motor neurons [142].

 

6. Sonophoresis

Sonophoresis involves the application of ultrasound to promote drug permeation. Ultrasound is a longitudinal acoustic wave at frequencies greater than that of the audible human hearing range (≥ 20 kHz). Such waves cause oscillation of the particles in the propagation medium due to alternating zones of high pressure (compression) and low pressure (rarefaction). This oscillation between high and low pressure gives rise to cavitation – the principal mechanism responsible for ultrasonic enhancement of skin permeability [143]. Sonophoresis has been investigated at both low frequencies (20 – 100 kHz) and high frequencies (≥ 0.7 MHz); however, low frequency sonophoresis has been proven in the last 15-20 years to induce a greater perturbation in the SC [144]. The resulting skin permeability enhancement can be controlled by varying the application time and other ultrasound parameters [145, 146].

 

Recent research has demonstrated the feasibility of using low frequency sonophoresis to deliver hormones, proteins, vaccines, liposomes and other nanoparticles through the skin [144] and in the last few years it has been also used in cosmetology. Successful delivery of methylprednisolone and cyclosporin into the skin was reported in a study involving 30 patients affected by alopecia areata [147]. Using similar ultrasound parameters, 96 women with either melasma or solar lentigo were treated by applying a skin-lightening emulsion containing ascorbic, azelaic, and kojic acids; significant positive effects after sonophoretic treatment were seen [147]. High-frequency ultrasound together with a coupling gel containing skin-lightening agents (ascorbyl glucoside and niacinamide) has also been shown to reduce facial hyper-pigmented spots compared with placebo and skin-lightening gel alone after 4 weeks [148]. In vitro studies using porcine skin also demonstrated that d-panthenol skin penetration could be increased with high-frequency ultrasound [149].

 

Ultrasound applicators with appropriate geometries and power ratings are currently being designed [150]. It is possible to expect that, as for the other physical techniques for promoting skin penetration, the use of low frequency sonophoresis in clinics will be largely dependent on the availability of new devices that are easy to use and cost effective.

 

7. Conclusion

The dynamic nature of aesthetic dermatology and its openness towards innovation means that many different technologies are already present in the clinic and being used to improve skin appearance. The next step is to use them to deliver active ingredients into the skin from cosmetic formulations. The scientific literature demonstrates that several physical enhancement technologies are capable of improving cutaneous penetration of cosmeceuticals: the choice of which technology to use will depend on several factors. The physicochemical properties of the active ingredient, its formulation requirements and stability under the application conditions will play a key role. The development of new formulations may be necessary since the technologies may have varying compatibilities with existing vehicles. Device design, ease-of-use and cost will also be important and advances in the microelectronics industry that encourage miniaturization and cost reduction with high scale production will also be critical determinants of which technologies can find a place in the world of mainstream aesthetic dermatology.