Pharmaceutical Development and Technology

ISSN: 1083-7450 (Print) 1097-9867 (Online) Journal homepage: http://www.tandfonline.com/loi/iphd20

Recent progress in transdermal sonophoresis Kevin Ita To cite this article: Kevin Ita (2015): Recent progress in transdermal sonophoresis, Pharmaceutical Development and Technology, DOI: 10.3109/10837450.2015.1116566 To link to this article: http://dx.doi.org/10.3109/10837450.2015.1116566

Published online: 25 Nov 2015.

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Date: 04 December 2015, At: 03:25

http://informahealthcare.com/phd ISSN: 1083-7450 (print), 1097-9867 (electronic) Pharm Dev Technol, Early Online: 1–9 ! 2015 Taylor & Francis. DOI: 10.3109/10837450.2015.1116566

REVIEW ARTICLE

Recent progress in transdermal sonophoresis Kevin Ita

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College of Pharmacy, Touro University, Vallejo, CA, USA

Abstract

Keywords

Transdermal drug administration has a number of advantages that cannot be leveraged for therapeutic benefits because of the robust barrier provided by the stratum corneum. One of the promising techniques for circumventing the stratum corneum is sonophoresis – the use of ultrasound for facilitating transdermal drug delivery. In this review, the mechanisms underlying sonophoresis and the utilization of the technique for transdermal delivery are discussed. The challenges of this mode of drug administration have also been highlighted and insight from a number of toxicological studies is described.

Acoustic streaming, cavitation, drug delivery, sonophoresis, transdermal

Introduction The skin (Figure 1) is the largest organ in the body and can be used for systemic drug administration. It is the primary interface between the body and the environment. The skin accounts for 15% of the total bodyweight, with a surface area of 1.5–2.0 m2 1. This organ provides an effective barrier against microbial pathogens and other deleterious physical and chemical factors2. Because of the large surface area, it is the focus of significant research efforts as a route of drug administration. However, a major limitation of using the skin as a transdermal drug delivery route is that only few medications can penetrate the stratum corneum (SC) (the outermost layer of the epidermis) in therapeutic quantities3. Drug transport through the skin is predominantly by diffusion through the SC, viable epidermis and the dermis4. The amount of drug that can cross the skin into systemic circulation is influenced by factors such as the integrity of the skin barrier, the physicochemical properties of the penetrant and the nature of excipients used5. Transdermal drug delivery systems currently used in clinical practice are formulated using approximately 19 active pharmaceutical ingredients6. Drug delivery scientists are working assiduously to expand the range of compounds that can be delivered through the transdermal route. In addition, there has been a rapid expansion in the number of biotechnology products. These products are mainly peptides, proteins, small interfering RNA, genes and oligonucleotides (ODNs). Because of the thermal and enzymatic lability of some of these compounds, transdermal drug delivery research efforts are currently being expanded to take into consideration the molecular structure and stability of these compounds. Taken together, these developments have led to the development of several strategies to deliver medications across the skin. These techniques include microneedles (MNs)7–9, iontophoresis10,11, chemical penetration enhancers12–14 and sonophoresis15–18.

History Received 16 September 2015 Revised 28 October 2015 Accepted 29 October 2015 Published online 25 November 2015

Sonophoresis refers to the use of ultrasound for transdermal drug delivery enhancement16,17,19–23. Based on the frequency used, this technique can be classified into low-, intermediate- and high-frequency sonophoresis24,25. This review will examine the recent advances that have been made in this area and some of the challenges that are inherent with the use of ultrasound for percutaneous transport enhancement. Mechanisms The experimental setup for sonophoresis26 is shown in Figure 2. It shows ultrasound source, the horn as well as the coupling medium. The skin is usually placed between the transducer and a mechanical support. The coupling medium provides a distance between the skin and the horn. There is substantial evidence in the literature that the application of ultrasound can facilitate transdermal drug delivery19,26–29. Table 1 shows selected reports in the literature using ultrasound for transdermal delivery enhancement. Despite these reports, the precise mechanisms through which ultrasound works are poorly understood. It has been postulated that when ultrasound is applied to the skin, it has the capacity to increase skin permeability through a variety of mechanisms including acoustic streaming17,30, rectified diffusion22, cavitation20,22,24 and the cellular-level effect31. Increased skin permeability leads to facilitated percutaneous penetration of drugs and biologicals. Ultimately, it is the capacity of sonophoresis to enhance drug delivery in a safe and efficient manner that is of interest in clinical practice. On the other hand, it is also important to understand the underlying mechanisms of ultrasound-mediated transdermal drug delivery. A better understanding of these mechanisms will enable scientists design and develop transdermal drug delivery systems for a broad range of drugs and biotechnology products. Acoustic streaming

Address for correspondence: Kevin Ita, PhD, College of Pharmacy, Touro University, Vallejo, CA 94592, USA. Tel: +1 707 638 5994. E-mail: [email protected]

The term streaming generally refers to the flow caused by a force which acts upon a fluid and it is normally dominated by its fluctuating velocity components32. In the specific case of sound waves interacting with fluids, bulk movement caused by

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Figure 1. Anatomy of the skin (reproduced with permission from reference 4).

Table 1. Ultrasound-enhanced transdermal delivery of selected compounds.

Figure 2. Application of ultrasound for transdermal drug delivery (reproduced with permission from reference 26).

momentum transfer from the acoustic field to the fluid is present33. Acoustic streaming is a type of macroscopic flow caused by a sound field, which exists in viscous fluid17,33,34. It has been theorized that acoustic streaming is engendered by ultrasound reflection, nonlinearity and attenuation17,33. Nonlinear losses in momentum of intense sound cause solenoidal mean mass flow, which are characteristic of multi-dimensional flows33. AS is also caused by the non-uniformity of the acoustic field32. Azagury et al. observed that streaming velocities developed during AS may lead to shear stresses which may affect tissues including the skin17. Interestingly, Baker et al. subdivides acoustic streaming into bulk and microstreaming35. The authors define bulk streaming as fluid flow caused by unidirectional propagation of ultrasound beam while microstreaming is defined as vortices

Compound

Authors and references

Lidocaine Insulin Penbutolol Ketoprofen Tetanus toxoid Heparin Caffeine Heparin Dextran Bovine serum albumin Anti-sense oligonucleotides FITC-dextran Diclofenac Insulin Glucose and mannitol

Nayak et al.55 Han and Das15 Ita and Popova18 Herwadkar et al.26 Tezel et al.23 Lanke et al.28 Pires-de-Campos50 Mitragotri and J. Kost29 Schoellhammer et al.47 Han and Das15 Tezel et al.27 Wolloch and Kost53 Huang et al.56 Boucaud et al.54 Merino et al.19

of flow proximate to an ‘‘oscillating source’’35. During the process of microstreaming, there is unidirectional flow of fluid in response to bubble dynamics in an acoustic field. The microstreaming process is linked to fluid properties such as acoustic viscosity and density, as well as ultrasound properties such as transducer aperture size, pressure amplitude, temporal average intensity and frequency21. Microstreaming can generate high velocity gradients and hydrodynamic shear stresses17,21. Many investigators have reported that the predominant mechanisms by which skin permeability is increased by US are (1) acoustic streaming, arising from attenuation of US as it propagates through a medium and (2) microstreaming resulting directly from cavitation activity of gas bubbles within a medium.

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Furthermore, acoustic radiation force can also act upon drug or vaccine particles if they are of comparable dimensions to, or larger than, the wavelength of the US beam, pushing them through a medium and thereby increasing particle permeation36. Such radiation forces will not be significant when using nano- or micro-sized particles in a kiloHertz US field, but will begin to dominate in the presence of MHz-range US36. Although, there is some understanding of the fundamental biophysical processes, it is still challenging to predict the extent of drug delivery enhancement produced by any given application of low frequency ultrasound. This is due to the fact that the extent of cavitational disordering of the SC depends upon numerous factors such as transducer geometry, transducer to skin distance, as well as the availability and distribution of dissolved gas37.

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Rectified diffusion When a sinusoidal pressure field of a particular frequency and amplitude is imposed on a liquid with dissolved gas, cavitation bubbles form in the presence of nucleation sites38. The resulting inhomogeneity in the liquid in the form of microscopic gas bubbles has been postulated to become nucleation sites38. If the imposed pressure field is beyond a certain value, the tiny gas bubbles repeatedly expand and compress and dissolved gas in the liquid flows into the gas bubbles by rectified diffusion38. During the positive pressure half cycle of the sinusoidal sound field, the bubble is compressed and the gas diffuses out of the bubble into the surrounding liquid22,39. During the negative pressure half cycle of the pressure field, the bubble expands and the gas flows from the liquid into the bubble38. Rectified diffusion in fact consists of two effects. The first effect is an ‘‘area’’ effect. During bubble expansion, the surface of the expanded bubble is much greater than that of the compressed bubble. Therefore, the amount of gas which enters the bubble during its expansion is higher than the amount of gas leaving the bubble during its compression and so the bubble gains a significant quantity of gas over many cycles. The second effect is the ‘‘shell’’ effect. The gas diffusion is controlled by a thick diffusion layer or a shell that forms in the liquid surrounding the bubble38. When the bubble expands, the shell becomes thinner and the difference in gas concentration increases. In this case, the flow rate of gas to the bubble also increases. When the bubble is compressed, the shell is thicker as a result the concentration gradient decreases. The combination of the ‘‘area effect’’ and the ‘‘shell effect’’ results in a quantity of gas being pumped into the bubble in each acoustic cycle. These effects are opposed by the normal tendency of a bubble existing in a gassy liquid to dissolve. Because of surface tension, the pressure inside the bubble is higher than that in the liquid immediately adjacent to the bubble. This pressure which is termed Laplace pressure, given by the expression, Pg – Pl¼2/R, where  is the surface tension and R is the radius of the bubble, can be large for very small bubbles38. Also, in the process of rectified diffusion, the bubbles generated by cavitation are pushed by the so called Bjerknes force under acoustic pressure gradient to move downwards. Bjerknes force is created under the acoustic field which directly increases the diffusion rate of the drug molecules. It can then push the bubbles forward, thereby, increasing the diffusion rate40. The foregoing indicates that there is a threshold acoustic pressure amplitude, above which a bubble of a given size will grow, and below which it will tend to dissolve. This is true for a distinct frequency of the sound field. The dissolved gas concentration and ultrasound frequency affects the threshold pressure38. Cavitation As can be seen from the above description, there is a relationship between rectified diffusion and cavitation. Indeed, rectified

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diffusion can be viewed as an integral part of cavitation. Cavitation may be defined as the process by which a liquid is pulled apart when it is acted upon by a force in excess of its tensile strength, causing the formation of voids in the system22. Stable cavitation refers to the periodic oscillation of bubbles over several acoustic cycles, whereas inertial cavitation is the dramatic expansion of bubbles during the rarefactional ultrasonic halfcycle, followed by their violent collapse during the compressional half-cycle under the effect of the inertia of the surrounding medium36. Inertial cavitation is usually identified by the broadband acoustic response that it generates, whereas stable cavitation can be identified by the harmonics and subharmonics of the fundamental US frequency it creates36. Acoustic cavitation involves the creation and oscillation of gas bubbles in a liquid41. In the low pressure portions of an ultrasound wave, dissolved gas and vaporized liquid can form gas bubbles. These bubbles alternately contract and expand in size, oscillating in response to high- and low-pressure portions of the ultrasound wave41. Acoustic cavitation occurs due to the nucleation of small gaseous cavities during acoustic pressure cycles21. The probability of cavitation occurring is closely related to ultrasound frequency as well as bubble size and shape. Since cavitation nuclei in biologic environments are random in type, size, type or shape, it is challenging to predict cavitation21. In stable cavitation, there is continuous oscillation of bubbles around the equilibrium radius in response to relatively lower acoustic pressures in an acoustic field21. During inertial cavitation, there is violent growth and collapse of bubbles which can occur within a period of a single cycle or a several cycles and this process depends on acoustic pressure as well bubble size and frequency21. The inertial cavitation threshold corresponds to the minimum acoustic pressure amplitude required to induce fast growth and collapse of cavitation bubbles36. The growth of cavitation bubbles becomes increasingly challenging with increasing US frequency because of the reduced timescale available for bubble growth and so the threshold pressure for occurrence of inertial cavitation is directly linked to US frequency36. In the course of inertial cavitation, lower frequencies give bubbles more time to grow in the expansion cycle and this consequently leads to a more violent collapse during the compression cycle21. A spherical collapse of bubbles yields high pressure cores that discharge shock waves with amplitudes exceeding 10 kbar while an aspherical collapse of bubbles near boundaries such as the SC yields microjets with diameters of about one-tenth of the bubble diameter and velocities of approximately 100 m/s21,42. Microjets are also capable of physically entering into the SC and facilitating SC permeability. These processes can lead to either the permeabilization of the SC due to shock waves, disruption of the microstructure of the SC due to the impact of microjets on its surface or transport of microjets into the SC42. A schematic sketching of cavitation21 is shown in Figure 3. Major bubble collapse mechanisms17 which include shock waves, and microjets are depicted in Figure 4. The disruption of a target exposed to such a pressure wave may also occur through shear strain, relative particle displacement, compressive failure or tensile stress21. Microjet distortion due to bubble collapse depends on the surface encountered by the bubble. If the surface is larger than the resonant size of the bubble (radius of 1–150 mm at about 5 MHz–20 kHz), the resulting collapse will be in the form of a microjet21,42. Shock waves generated by inertial cavitation can cause structural alterations in the surrounding keratinocyte–lipid interface regions, resulting in the creation of diffusion channels through which drugs may be potentially delivered42. Furthermore, impact pressure of the microjet on the skin surface may enhance SC permeability by disrupting SC lipid bilayers21,42. In the context of sonophoresis, increased skin permeability can be achieved when cavitation

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Figure 3. Schematic sketching of cavitation (reproduced with permission from reference 21).

desired permeability enhancement regardless of the choice of frequency, although the necessary energy density is higher at higher frequencies. Low frequencies (approximately 20 kHz) induced localized transport compared to a more dispersed effect seen with higher frequencies (approximately 58.9 kHz). Increase in sonophoretic transport also correlates with increasing intensity although two thresholds, both a lower cutoff intensity and upper decoupling intensity, may be present37. Cavitation is the term used to describe the formation and oscillation of microbubbles in a coupling medium36. These microbubbles are capable of collapsing on the surface of skin (SC) and increasing skin permeability and transdermal flux enhancement36,42. Tezel and Mitragotri recently described a cavitation model for sonophoresis42: "¼

Figure 4. Major bubble collapse mechanisms: (a) shock waves, and (b) micro jets(reproduced with permission from reference 16).

bubbles interact with and modify the skin barrier by inducing, dilating and connecting defects to form regions of increased permeability within the SC. However, the precise mechanisms, specifically the location(s) and type(s) of cavitation, responsible for skin permeabilization may be dictated by the frequency of ultrasound used for sonophoresis24. From the mechanistic perspective, it is known that acoustically-induced cavitation results in the formation of intercellular lipid channels and defects in the SC. These perturbations allow rapid diffusion of solutes into the skin37. The extent of enhancement is determined by four main sound wave variables – frequency, intensity, duty cycle and duration. Usually, transdermal drug delivery is facilitated by decreasing ultrasound frequency and this is due to the fact that ultrasound induces proportionally more cavitational activity at lower frequencies37. Tezel et al. used conductivity as an indirect measure of skin permeability and showed that for each frequency (in the range of 19.6–93.4 kHz), there exists a threshold intensity below which no detectable conductivity enhancement was observed43. The threshold intensity increased with frequency. The authors hypothesized that it was feasible to achieve the

P c

ð1Þ

where " is strain, P is shock wave amplitude,  and c are density and speed of sound within an element respectively. Equation (1) was modified to Equation (2)42:     d 1 rmin 3 p 2   ð2Þ dt 3 "c c2 3 Equation 2 establishes the relationship between cavitation bubble collapse and the rate of SC disruption42. "c is the critical strain which the authors gave as 0.02 while , which is the number of collapse events per unit volume per unit time depends on ultrasound intensity42. The value of rmin , determined through numerical simulations, is assumed to be about 1 mm42.  can also be determined by fitting Equation (2) to experimentally measured values of d=dt. Thermal effect The thermal effect of US on the skin results from transfer and conversion of mechanical energy produced by the vibration of a piezoelectric crystal in the sonophoresis probe19. When ultrasound passes through a medium, energy is partially absorbed.

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DOI: 10.3109/10837450.2015.1116566

Ahmadi et al. observes that ultrasound energy is attenuated when a wave propagates through a medium44. Attenuation is the result of a complex combination of absorption and scattering effects. Absorption in tissue leads to the transformation of sound energy into heat and is the main component of attenuation44. In the human body, ultrasound energy absorbed by tissue causes a local temperature increase that is dependent upon ultrasound frequency, intensity, area of the ultrasound beam, duration of exposure and the rate of heat removal by blood flow or conduction21. When a material possesses higher ultrasound absorption coefficients, such as bone, thermal effects are higher compared with muscle tissue, which has a lower absorption coefficient45. Transfer of thermal energy to the drug molecules increases their diffusion through a barrier, particularly if the barrier has already been permeabilized36. It is also possible that elevated temperature may influence drug diffusion coefficient. The increase in the skin temperature resulting from the absorbance of ultrasound energy may increase the skin permeability coefficient because of an increase in the permeant diffusion coefficient46. It was shown that a temperature increase of 10  C caused a two-fold increase in the estradiol skin permeability. Because the typical skin temperature increase in case of therapeutic sonophoresis is 7  C, it can be hypothesized that thermal effect may be insignificant as temperature alone cannot account for a 13-fold increase in estradiol skin permeability upon application of ultrasound46. The absorption of the sound in skin increases as the ultrasound frequency goes up, which means that the energy would be stored in skin rather than transmit through40. The rise in the temperature of the skin increases the kinetic energy of the drug molecules which may enhance drug diffusion rate. However, the intensity in the application of sonophoresis is usually low; therefore, temperature effect on the kinetic energy of the drug molecule and hence the drug permeability may not be significant40. However, depending on the exposure conditions, this mechanism can become detrimental to drug delivery as prolonged heating damages both tissue and drug36. On the positive side, especially when a drug is thermostable, temperature increase of the skin may be beneficial as it may increase drug diffusivity. With increase in molecule diffusivity, permeability is facilitated. Cellular-level effect The cellular-level model (called ‘‘bilayer sonophore’’), described by Krasovitski et al., examines the physics of bubble dynamics as well as cell biomechanics with special emphasis on the dynamic behavior of two lipid bilayer membrane leaflets31. According to this hypothesis, the cellular membrane is intrinsically capable of absorbing mechanical energy from the ultrasound field and transforming it into expansions and contractions of the intramembrane space31. It is also hypothesized that the maximum area strain is proportional to acoustic pressure amplitude and inversely proportional to the square root of the frequency31. The authors investigated the direct interaction between oscillating acoustic pressure and cellular bilayer membranes. The aim was to examine in one model the influence of both cavitational and noncavitational ultrasound effects on biological tissue. The authors hypothesize that the intramembrane hydrophobic space between two lipid monolayer leaflets expands and contracts periodically when exposed to ultrasound. The two leaflets are pulled apart when the acoustic negative pressure overcomes the molecular attractive forces between the two leaflets. During the negative pressure cycle, the surrounding tissue is also pushed backwards. Consequently, the intramembrane leaflets are pulled back together in the positive pressure cycle. The authors propose the term ‘‘bilayer sonophore’’ (BLS) to show that the bilayer membrane is capable (under appropriate conditions) of transforming the

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(millimeter wavelength) oscillating acoustic pressure wave into (nanometric and micrometric) intracellular deformations31. This cyclic expansion and contraction of the BLS may be responsible for stimulating cycles of stretch and release in the cell membranes and in the cytoskeleton, which could activate mechano-sensitive proteins and/or increase membrane permeability31. To explore the dynamic response of the bilayer sonophore in a living cell to US exposure, the authors created physical models which include molecular forces, bubble dynamics, and gas diffusion in and around a membrane bilayer. Using such models, they evaluated the BLS response to US for multiple parameters, including the size of the free membrane when the BLS is surrounded by water and the combined effect of acoustic pressure amplitude and frequency for a more realistic BLS bounded by a thin layer of tissue on the order of a size of a cell31. To test this model, the authors treated goldfish with continuous ultrasound exposures at 1 or 3 MHz, or to both frequencies given in succession, at spatialaveraged, temporal averaged intensities of up to 2.2 W cm2 (0.25 MPa) and for durations of up to 360 s31. A force-balance technique was used to calibrate the intensities of exposure and the authors used an imaging scanner to detect cavitation. Usually, an epithelium’s response to ultrasound exposures is the generation of cavities over a range of shapes (from round and elliptical to parallel or undulated slits) and sizes (from narrow cavities550 nm wide between two neighboring desmosomes to a few micrometers in width)31. In this experiment, there were many cavities around the outer membranes and between the cells. Using the approach described by Krasovitski et al.31, numerous biological effects caused by ultrasound can potentially be interpreted as progressive stages along a graded scale of induced phenomena that differ in maximum area strain due to different ultrasound exposure parameters, proximity to a free surface, mechanical tissue properties, or the presence of extracellular gas bubbles31.

Experimental methodology Low frequency ultrasound is usually generated from a sonicator18,26. In the pretreatment protocol, human, porcine, hairless mouse or other skin sample is pretreated with ultrasound before mounting on diffusion cells. A piece of porcine skin was placed under the sonicator horn with the epidermal surface of the skin facing the horn. The sonicator horn is then placed in the coupling medium. After ultrasound pretreatment, skin sample is then mounted on diffusion cells for permeation studies18. In another experimental approach, Rich et al. generated ultrasound by sending a sinusoidal signal from a generator through a switch to a radio-frequency amplifier24. This was used to power one of two source transducers. It is important to design sonophoresis experiments meticulously as several parameters can be affected by experimental techniques.

Sonophoretic delivery of small molecules Transdermal drug delivery has numerous advantages including painlessness, non-invasiveness, improvement in patient compliance, the possibility of increased bioavailability and sustained drug delivery6. However, these advantages can only be realized if adequate and therapeutically relevant serum levels of drugs are achieved. For over 35 years, only approximately 20 active pharmaceutical ingredients have been formulated as transdermal drug delivery system. Recent reports in the literature have shown that transdermal sonophoresis may be a promising approach for facilitating transdermal drug delivery26,47. Herwadkar et al. studied the influence of low frequency sonophoresis on percutaneous penetration of ketoprofen26. The authors examined the effect of several factors such as horn distance, ultrasound application time, coupling medium and duty cycle on

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percutaneous penetration of ketoprofen. They found out that application of ultrasound significantly facilitated transcutaneous permeation of ketoprofen from 74.87 ± 5.27 mg/cm2 for passive delivery to 491.37 ± 48.78 mg/cm2 for ultrasound-mediated delivery26. Drug levels in skin layers also increased from 34.69 ± 7.25 mg following passive permeation to 212.62 ± 45.69 mg following the application of low frequency ultrasound26. Tezel and Mitragotri have previously shown that, during transdermal sonophoresis, the number of collapse events per unit volume per unit time, depends on the ultrasound intensity42. To determine the required ultrasound intensity for their experiment, Herwadkar et al. used a calorimetric method. Briefly, the investigators placed 50 g of water in an insulated beaker and ultrasound was activated for predetermined time periods using different duty cycles. A thermometer was used to measure the temperature before and after ultrasound activation. Ultrasound intensity was estimated using the formula26,48: m  water  cpwater dT ð3Þ I¼ dt A where I is the intensity of ultrasound (W/cm2), m is the mass of water (50 g), cp is the specific heat capacity of water (4.18 J/g  C), A is the surface area of the sonicator horn and dT=dt is the rate of change of temperature of water. For sonophoretic transport experiments, the authors used ultrasound with a frequency of 20 kHz and intensity of 6.9 W/cm2. Indocyanine green (ICG) is considered a good alternative to conventional photosensitizers such as aminolevulinic acid and methyl aminolevulinic acid49. This is because ICG possesses minimal side effects and has average therapeutic efficacy. Lim et al. investigated the transdermal sonophoretic transport of ICG using 50 kHz ultrasound for 30 s. Sonophoresis was combined with fractional radiofrequency technology. To detect and image ICG fluorescence, the Maestro system with a near-infrared 740: 10:950 filter was used. The average fluorescence index was measured by the Maestro system immediately after treatment. Higher fluorescence intensity showed greater percutaneous ICG uptake49. The authors found that ICG-induced fluorescence was significantly higher on fractional RF and sonophoresis-pretreated skin than the untreated area. There is also a plethora of reports in the literature mentioning the combination of sonophoresis with other penetration enhancement techniques such as iontophoresis or electroporation. Piresde-Campos et al. used continuous ultrasound of 3 MHz with an intensity of 0.2 W/cm2 to deliver caffeine across pig skin50. The authors monitored morphological changes after drug administration. The authors reported that caffeine treatment was effective only when associated with ultrasound therapy50; the combination resulted in a significant reduction in the thickness of the subcutaneous adipose tissue, as well as damage to the adipocytes, consequently decreasing the number of cells50. Calcein is a charged and hydrophilic molecule and poorly penetrates the skin when delivered passively. In their publication, Zorec et al.51 noted that sometimes, sonophoresis and sonoporation are used interchangeably in the scientific literature. The authors defined sonoporation as the creation of aqueous pathways through the SC as a result of ultrasound application while sonophoresis was described as the migration of delivered molecules through acoustic streaming and convection effects51. The measured peak to peak amplitude of the ultrasound pressure was 166 kPa, with no difference observed for positive or negative pressure. The walls of the bath were made from Plexiglas and lined with an ultrasound absorber. The device was operated in a continuous mode at 30 kHz frequency51. Zorec et al. showed a statistically significant enhancement of calcein delivery across porcine skin after 5 min of ultrasound application43. The investigators combined sonoporation

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with electroporation hoping to achieve synergistic effects, however, the results showed no obvious dramatic improvement over a single method used alone51.

Sonophoretic transport of macromolecules Schoellhammer et al. used dual-frequency ultrasound (20 kHz and 1 MHz wavelengths) to enhance transdermal delivery of 4 kDa dextran in PBS. A statistically significant increase in the flux of 4 kDa dextran was observed by using 20 kHz + 1 MHz as opposed to 20 kHz alone. There was a statistical difference between the flux observed for samples treated with 20 kHz + 1 MHz for 6 or 8 min but no increase in flux with treatment time for the use of 20 kHz alone47. Interestingly, the enhancement in flux at both treatment times using 20 kHz + 1 MHz was greater than what would be expected based on the increase in localized transport regions (LTR) sizes alone at each treatment time. The dextran flux was 3.5- and 7.1-fold greater using 20 kHz + 1 MHz at 6 and 8 min, respectively, compared to the use of 20 kHz alone, while LTR area increased 2.5- and 6-folds, respectively47. A previous study by the same authors had shown that dual-frequency ultrasound, utilizing 20 kHz and 1 MHz wavelengths together led to greater transient cavitation events, as assessed by pitting data, and resulted in larger localized transport regions in vitro25. Han and Das pretreated pig skin with MNs so as to alter the skin property and then applied ultrasound field on the pretreated area to let the cavitation further increase skin permeability15. The authors chose bovine serum albumin as a model molecule since it is a common protein and has a molecular weight of over 60 000 Da. Another reason for choosing BSA was because the molecule is hydrophilic and its molecular size ranges between the sizes of small peptide molecules and large vaccines15. The authors treated porcine skin with both ultrasound and MNs. Results of the experiments showed that the permeability of BSA increased to 1 mm/s with the combination of 1.5 mm MNs patch and 15-W ultrasound output15. This increase in permeability coefficient is about 10 times higher than the permeability obtained when BSA diffuses passively15. In comparison, the maximum permeability following ultrasound treatment was 0.4 mm/s15. Sodium dodecyl sulfate (SDS) is a surfactant and skin irritant which is often present in sonophoresis coupling media at 1% (w/ v). When used in this manner, the compound is known to enhance skin permeability52. However, because of problems related to skin irritation, Dahlan et al. examined the feasibility of using lowfrequency ultrasound-assisted transcutaneous immunization in the absence of SDS52. A lower ultrasound duty cycle of 10% generated higher antibody titers than a duty cycle of 20%, also despite causing lower skin damage. The authors concluded that lack of correlation between skin damage and immune responses showed that enhancement of skin permeability to topically applied antigen was not the major mechanism through which sonophoresis facilitated skin immunization52. ODNs are known to selectively target gene expression and are useful in the management of pathological disorders such as herpes simplex, skin carcinoma, psoriasis and melanoma. However, the transcutaneous delivery of large molecular weight, negatively charged molecules across intact SC is limited27. Tezel et al. used low frequency ultrasound (20 kHz, 2.4 W/cm2) to deliver therapeutically significant quantities of anti-sense ODNs into the skin27. The ODNs used in the study correspond to secondgeneration chemistries containing the 20 -O-methoxyethyl sugar modification. Concentrations of ODNs in the superficial layers of the skin ranged from approximately 0.5% to 5% of the donor concentration after a 10-min application of ultrasound and ODN27. The authors used both pretreatment and simultaneous sonophoresis modes of application. In the pretreatment mode, a

Recent progress in transdermal sonophoresis

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DOI: 10.3109/10837450.2015.1116566

10-min simultaneous application of ultrasound (20 kHz and 2.4 W/cm2) was used to permeabilize the SC. At the end of ultrasound application, skin conductance increased by about 100fold and permeability of ultrasonically-treated skin to ODN was measured to be 4.5  105 cm/h compared to nearly undetectable values across non-treated skin27. A 10-min simultaneous application of ODN and ultrasound (20 kHz, 2.4 W/cm2) resulted in dermal accumulation of 3500 disintegration per second (dpm/ cm2) of ODN in the skin. With this efficiency, about 53 g/cm2 ODN can be delivered into skin in 10 min from a donor solution containing 100 mg/ml ODN27. Low-frequency ultrasound has been postulated to enhance skin permeability in a heterogeneous manner. Application of ultrasound affects discrete domains of the skin, producing highly permeable localized transport regions53. Wolloch and Kost investigated the contribution of two cavitational phenomena (shock waves and microjets) with the aim of examining whether diversion of the bubble collapse mode to shock waves, or microjets, may result in a more homogenous effect on skin, and enhanced transdermal flux53. Skin permeability was evaluated by measuring the transport rates of a fluorescent conjugated molecule — FITC-dextran. The results of the study showed that the contribution of microjets to skin permeability enhancement is significantly higher than shock waves53. Boucaud et al. investigated insulin delivery across hairless rat skin using 20 kHz ultrasound54. The authors examined the contribution of different parameters – intensity, application time and pulse length to percutaneous flux. Change in blood glucose levels of the animals was used to evaluate insulin transport. The results showed a threshold below which no detectable changes in blood glucose level was observed for each sonophoresis parameter. In the study, it was demonstrated that significant hypoglycemia resulted with the application of less than 15 min ultrasound and was similar to subcutaneous injection of 0.5 U of insulin54.

Sonophoretic delivery from formulations Nayak et al. examined the influence of MNs and sonophoresis pretreatment on transdermal delivery of lidocaine from a polymeric hydrogel formulation55. Different ratios of carboxymethylcellulose and gelatin (NaCMC/gel ranges 1:1.60–1:2.66) loaded with lidocaine were used in the formulations. The authors used Franz diffusion cells to measure lidocaine penetration across porcine skin after pretreatment with stainless steel MNs and 20 kHz sonophoresis for 5 and 10 min. A 4.8-fold increase in lidocaine flux following combined application was observed compared with separate pretreatments after 30 min55. Sonophoresis pretreatment alone did not have a significant impact on lidocaine delivery during the initial 2 h period. However, MN application increased permeability at a time of 0.5 h for up to 17 fold with an average up to 4 fold. The time required to reach therapeutic levels of lidocaine was decreased to less than 7 min55. Huang et al. used polyamidoamine dendrimer coupled with sonophoresis to facilitate transcutaneous penetration of diclofenac (DF)56. The authors selected coupling media concentration, ultrasound-application time, duty cycle, distance from probe to skin, and a third-generation polyamidoamine dendrimer concentration as independent variables. Plackett–Burman factorial design was used for variable optimization. DF gel without dendrimer and ultrasound treatment to skin showed 56.69 mg/cm2 cumulative drug permeated through the skin, while the DF–dendrimer gel without sonophoresis treatment showed 257.3 mg/cm2 cumulative drug permeated through the skin after 24 h56. However, when same gel was applied to sonophoresis-treated skin, there was a dramatic increase in permeation. The cumulative drug that permeated through the skin was 935.21 mg/cm2 56. It is worth noting that in

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this study, only the cumulative amount delivered was reported instead of transcutaneous flux values. This may have been due to the fact that the authors were examining topical rather than transdermal drug delivery.

Challenges of sonophoretic drug delivery One of the major challenges regarding transdermal sonophoresis is the development of low cost ultrasound devices that can facilitate efficient transdermal drug delivery57. Echo TherapeuticsÕ is now marketing the SonoprepÕ device as an ultrasound skin poration system that can be used to enhance the transdermal delivery of local anesthetics. It is the only low-frequency ultrasound device which is approved by the United States Food and Drug Administration for transdermal drug delivery. This device transmits ultrasound energy to the skin. The collapse of cavitation bubbles created by ultrasound can permeabilize the SC and facilitate percutaneous penetration. Another concern is with the safety of transdermal sonophoresis. Several decades ago, precisely in 1917, Paul Langevin observed the immediate death of fish swimming near to the acoustic beam44. Since then scientists have carried out many studies to determine the threshold safety levels for transdermal sonophoresis. Singer et al. studied the safety of low-frequency sonophoresis58. Low-frequency ultrasound was applied for 60 s to the clipped abdominal skin of three anesthetized adult mongrel dogs using an ultrasound device operating at a frequency of 20 kHz with a maximal energy output of 400 W58. The authors found that transdermal sonophoresis at low intensities seemed to be safe for facilitating topical drug delivery. The authors only observed minor urticarial reactions at low intensities. However, at higher intensities, thermal injuries were observed. With still higher intensities, confluent epidermal necrosis became apparent58. Ahmadi et al. observes that significant increase in temperature at the transducer surface and sustained exposure to high temperatures can have many harmful effects on the skin and underlying tissues including burns, epidermal detachment and necrosis of tissues44. As a way to deal with this problem, it has been suggested that the coupling medium be replaced from time to time22,44. It has also been observed that cavitation in the tissue layers beneath the skin and sub-cutaneous tissues at low-intensity low-frequency exposures has received less attention44. Toxicological studies at this level are significant because the penetration depth of ultrasound (being inversely proportional to the tissue attenuation coefficient) increases at lower frequencies44. Ahmadi et al.44 also made extrapolations from studies conducted by O’Brien and Zachary59 and suggested that the pressure threshold for human lung hemorrhage damage might be approximately 500 kPa. There is no doubt that as new agents are delivered to and through the skin using ultrasound, regulatory authorities will require more toxicology studies to ensure that patients benefit from this powerful technology while avoiding harmful side effects.

Conclusions Transdermal sonophoresis continues to attract the attention of scientists and clinicians. The technique, which is based mainly on cavitation events, is capable of increasing skin permeability and facilitating percutaneous penetration of drugs and macromolecules especially products of the biotechnological industry. In this paper, the focus has been on the most important advances made especially in the understanding of the mechanics of the process. Conventionally, cavitation has been the major mechanism studied. Progress made regarding the cellular level model has also been highlighted. From a clinical standpoint, the approval of Sonoprep

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K. Ita

by the FDA has also been emphasized and challenges arising from the use of ultrasound discussed. Overall, as scientists gain more understanding of the skin-ultrasound interactions, more efficient ways of delivering medications with sonophoresis will be developed for the benefit of patients.

Declaration of interest The author has no interests to declare.

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