transdermal drug delivery with a pressure wave

www.elsevier.com/locate/addr

Advanced Drug Delivery Reviews 56 (2004) 559–579

Transdermal drug delivery with a pressure wave

Apostolos G. Doukasa,*, Nikiforos Kolliasb

aDepartment of Dermatology, Wellman Laboratories of Photomedicine, Massachusetts General Hospital, Harvard Medical School,

Boston, MA 02114, USAbJohnson & Johnson Consumer and Personal Products Worldwide, 199 Grandview Road, Skillman, NJ 08558, USA

Received 9 September 2003; accepted 13 October 2003

Abstract

Pressure waves, which are generated by intense laser radiation, can permeabilize the stratum corneum (SC) as well as the cell

membrane. These pressure waves are compression waves and thus exclude biological effects induced by cavitation. Their

amplitude is in the hundreds of atmospheres (bar) while the duration is in the range of nanoseconds to a few microseconds. The

pressure waves interact with cells and tissue in ways that are probably different from those of ultrasound. Furthermore, the

interactions of the pressure waves with tissue are specific and depend on their characteristics, such as peak pressure, rise time

and duration. A single pressure wave is sufficient to permeabilize the SC and allow the transport of macromolecules into the

epidermis and dermis. In addition, drugs delivered into the epidermis can enter the vasculature and produce a systemic effect.

For example, insulin delivered by pressure waves resulted in reducing the blood glucose level over many hours. The application

of pressure waves does not cause any pain or discomfort and the barrier function of the SC always recovers.

D 2004 Elsevier B.V. All rights reserved.

Keywords: Photomechanical waves; Shock waves; Stratum corneum barrier; Stress waves; Transdermal delivery

Contents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 560

2. Generation and propagation of pressure waves . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 561

3. Biological effects of pressure waves . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 564

4. Experimental arrangement . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 565

5. Transdermal drug delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 567

6. The synergy of pressure waves and sodium lauryl sulfate . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 571

7. The mechanism of the permeabilization of the stratum corneum . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 573

8. Examples of transdermal drug delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 575

8.1. Delivery of allergens. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 575

8.2. Systemic delivery of insulin . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 575

9. Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 576

Acknowledgements . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 577

References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 577

0169-409X/$ - see front matter D 2004 Elsevier B.V. All rights reserved.

doi:10.1016/j.addr.2003.10.031

* Corresponding author.

E-mail address: [email protected] (A.G. Doukas).

A.G. Doukas, N. Kollias / Advanced Drug Delivery Reviews 56 (2004) 559–579560

1. Introduction
penetration depends on anatomical site, age, sex, skin
In clinical drug therapies, topical application allows

localized drug delivery to the site of interest. This

enhances the therapeutic effect of the drug while

minimizing systemic side effects. Furthermore, topical

application of drugs bypasses systemic deactivation or

degradation and minimizes gastrointestinal incompat-

ibility and potential toxicological risk.

The stratum corneum (SC) of the skin is an effective

barrier to molecular transport. It is composed of

corneocytes (f 30 Am by 0.5–0.8 Am thick cells)

which are filled with keratin and lack nuclei and

cytoplasmic organelles. There are 10–50 cell layers

in the human SC with an intercellular spacing of the

order of 20 nm [1]. The intercellular volume is 5–21%

of the SC volume [2]. The intercellular regions are

composed mainly of neutral lipids, which originate

from the membrane-coating granules in the stratum

granulosum [3]. These granules are 0.15–0.5 Am by

0.3–0.7 Am and are composed of stacks of 7.5–8 nm

disks thick in a triple-layered membrane [3]. They fuse

with the plasma membrane and release their contents

into the intercellular matrix as the cells in the stratum

granulosum progress into the SC. Elias [4] has pro-

posed a heterogeneous two-compartment model of the

SC which attributes a special role to the intercellular

lipids in the regulation of the SC barrier function.

Three possible pathways, transappendageal, trans-

cellular, and intercellular have been suggested for

molecular transport through the SC [5]. The trans-

appendageal pathway is primarily through hair fol-

licles. However, the transappendageal skin transport

in humans is limited by the small surface area avail-

able. The fractional area of hair follicles relative to the

skin area is 10� 2–10� 5 [6]. The transcellular path-

way requires the substrates to travel through the

corneocytes while the intercellular pathway is via

the extracellular matrix between the corneocytes.

For intercellular skin transport, hydrophilic substrates

are rate limited by the lipid environment of the

intercellular matrix of the SC [7]. On the other hand,

lipophilic substrates partition into the intercellular

lipids of the SC. However, the rate-limiting step is

the partition into the epidermis, which is practically an

aqueous environment. Molecular transport through

the skin has been described by a solubility–diffusion

model [8] and a transfer free energy model [9]. Skin

care, hydration, and temperature as well as contact

with organic solvents or surfactants [10]. In addition,

the molecular weight (MW) of the substrate affects

percutaneous absorption. The diffusion through the

SC follows the expression D~(MW)� b where the

value of b varies between 0.3 and 0.6 [11].

Depending on the substrate, there may be several

orders of magnitude difference in the rate of transport

through the SC [12]. Nevertheless, even the most

rapidly penetrating drugs actually diffuse very slowly

through the SC [5]. Once a substrate has diffused into

the epidermis and dermis it can enter the vasculature,

thus producing a systemic effect. For transdermal

delivery to be effective, the drugs have to enter into

the viable skin in sufficient quantities to produce a

therapeutic effect. There are a number of ways, which

can be utilized to enhance transdermal delivery. For

example, the drug can be mixed with a formulation

that promotes the delivery of the drug through the SC

over a period of hours [13]. Occlusion has also been

used together with formulations to enhance skin

penetration [14]. However, the dependence of the

molecular penetration on hydration is unclear because

hydration changes several parameters at the same

time. For example, diffusivity and skin thickness

increase with water content while the partition coef-

ficient decreases. In general, the amount of drug

absorbed through the skin increases with longer

contact time, larger contact area on the surface of

the skin, higher drug concentration in the formulation

and occlusion. However, high substrate concentration

can cause saturation of the transdermal absorption

pathways and extended exposure can increase the risk

of allergic contact skin sensitivity [15]. Furthermore,

these techniques usually require prolonged contact

and do not work for all substances.

Transdermal drug delivery has been the subject of

extensive research [16]. In addition to vehicle formu-

lations and chemical enhancers [13,17], physical

methods such as the application of continuous elec-

trical current (iontophoresis) [18,19], electrical pulses

(electroporation) [20,21] and ultrasound (phonopho-

resis) [22–24] have been used to facilitate the trans-

port of a variety of molecules across the SC. Physical

methods have the advantage of decreased skin irritant/

allergic responses, as well as no interaction with the

drugs being delivered.

A.G. Doukas, N. Kollias / Advanced Drug Delivery Reviews 56 (2004) 559–579 561

Pressure waves (high amplitude pressure transi-

ents) generated by lasers is, perhaps, one of the latest

platforms for drug delivery. A pressure wave (PW)

was first shown to permeabilize the cell plasma

membrane and allow macromolecules to diffuse

through it into the cytoplasm [25]. Pressure waves

have been used to permeabilize the SC and facilitate

the transport of macromolecules into the viable skin

[26]. They have also been shown to facilitate drug

delivery into microbial biofilms [27]. Last, we have

recently demonstrated that PW can permeabilize the

nuclear envelope and facilitate the delivery of macro-

molecules into the cell nucleus [28]. The many

diverse applications of pressure waves exemplify the

potential of this platform for drug delivery in many

different biological systems.

The applications of PW for drug delivery rest on

extensive studies of the interactions of laser-induced

pressure waves with cells [29–33] and tissue [34–

39]. Shock waves generated by extracorporeal shock

wave lithotripters [40,41] and to a lesser extent

ultrasound [42] have also contributed to our under-

standing of the effects of waves on cells and tissues.

Fig. 1. (A) Typical ultrasound waveform and (B) a typical PW used

in transdermal drug delivery and its characteristics, (PP) the peak

pressure, (RT) the rise time defined as the time from 10% to 90% of

the peak pressure, the duration (FWHM) defined as the full width at

half maximum and (D) the decay defined as the time from 90% to

10% of the peak pressure.

2. Generation and propagation of pressure waves

Pressure waves, stress waves and shock waves are

terms that have been used to describe high amplitude

pressure transients generated by laser [43] extracor-

poreal shockwave lithotripters [40], shock wave tubes

[44], and energetic material [45]. In addition, the term

photomechanical waves has often been used for laser-

generated pressure waves. Pressure waves should be

distinguished from ultrasound, which has been exten-

sively used for transdermal and cytoplasmic molecular

delivery. Fig. 1 shows the waveforms of an ultrasound

wave and a PW. The amplitude of the ultrasound

oscillates equally between positive and negative pres-

sure. The negative pressure (tensile component) is

responsible for the cavitation associated with low-

frequency ultrasound. On the other hand, the PW is

mostly a unipolar compression wave (positive pres-

sure), which does not show a measurable tensile

component and thus exclude biological effects in-

duced by cavitation [42]. Fig. 2 shows the bandwidth

of the PW generated by Fourier Transform. The

bandwidth depends on the rise time, decay and

duration. The shorter the rise time the broader the

bandwidth will be.

It should be emphasized that the operating param-

eters (amplitude and duration) of the two types of

waves are very different. For example, the peak

pressure of the PW, used in our experiments, has been

between 300 and 1000 bar (1 bar = 105 Pa = 0.987

atm.) while the amplitude of the ultrasound is usually

between 1 and 5 bar. Furthermore, the duration of the

PW, depending on the application, has been between

100 ns and 10 As. On the other hand, typical applica-

Fig. 2. The bandwidth of the PW shown in Fig. 1B produced by

Fourier Transform. In this scale the bandwidth of an ultrasound

wave (e.g., 300 kHz) would be represented by the vertical line.


tions of ultrasound for transdermal delivery are of the

order of tens of minutes. So there are many orders of

magnitude difference both in the pressure amplitude

and duration. What may be more important is the rate

at which the pressure is applied onto the skin. The rise

time of a PW is between 10 and 200 ns which, for a

500 bar PW, produced a rate of pressure increase

between 50� 109 and 2.5� 109 bar s� 1, a rate of

increase of billion atmospheres per second. It is

reasonable to assume that the application of this kind

of physical force onto the skin can produce novel

effects.

The generation of PW by lasers was demonstrated

in 1964 [46], shortly after the invention of the Q-

switched laser. The interactions of lasers with matter

that can generate pressure waves has been extensively

reviewed in the literature [47–50]. Pressure waves

can be generated by one of the following mechanisms:

optical breakdown, ablation, and rapid heating of an

absorbing medium (thermoelastic generation) [43].

These three modes of PW generation allow the

investigation of interactions of cells and tissue with

PW under a variety of conditions. It should be pointed

out that PW is only one of the phenomena under the

general category of photomechanical effects. Other

photomechanical effects include plasma formation,

cavitation, microstreaming and formation of jets.

The diversity of the phenomena involved during the

application of short pulsed lasers has hindered the

understanding of the PW effects and potential thera-

peutic uses until it was technologically possible to

separate and control these phenomena independently.

Ablation is a reliable method for generating PW

with consistent characteristics [51–53]. In ablation,

the laser radiation causes decomposition of the target

material into small fragments, which move away from

the surface of the target at supersonic speed [54].

Although the amount of material ejected is small, a

PW of high amplitude can be generated by the

imparted recoil momentum [55]. The characteristics

of the PW (peak pressure, rise time and duration)

depend on the laser parameters (wavelength, pulse

duration and fluence) and the optical and mechanical

properties of the target material. The efficiency of the

PW generation, conversion of light energy to the

mechanical energy of the PW, is given by the coupling

coefficient. The coupling coefficient is defined as the

total momentum transfer to the target during ablation

divided by the laser pulse energy. From the coupling

coefficient the peak pressure can be calculated. Eq. (1)

gives the peak pressure generated during ablation of

polymers and metals as a function of irradiance,

wavelength and pulse duration [52]:

P0 ¼ bI0:7

ðkffiffiffis

pÞ0:3

ð1Þ

where P0 is the peak pressure of the wave, b the

proportionality constant which depends on the mate-

rial properties, I the laser irradiance, k the laser

wavelength and s the laser pulse duration. Eq. (1)

has been shown to hold over a wide range of laser

irradiance (3 MW cm� 2–70 TW cm� 2), wavelength

(248 nm–10.6 Am), pulse duration (500 ps–1.5 ms)

and pulse energy (100 mJ–10 kJ). During ablation,

most of the energy of the incident light is absorbed by

the plasma and only a small fraction of the irradiance

is converted into the mechanical energy of the PW.

However, by designing stratified targets and overlays

(confined ablation) the momentum transfer to the

target can be increased and thus create PW more

efficiently or of higher peak pressure [56]. The

presence of the overlay also changes the other char-


acteristics of the PW, such as the rise time, duration

and decay.

The PW rise time depends on the optical penetra-

tion depth of the target material, the dynamics of the

plasma, and the interactions of the plasma with the

surface of the target. There is no simple mathematical

formula, which can predict the rise time from the laser

parameters and properties of the target. Nevertheless,

PW with different rise times have been generated by

appropriate selection of laser wavelength and target

material. Fig. 3 shows five different waveforms pro-

duced by different combinations of laser wavelength,

target material, and confinement. The waveforms were

measured using a polyvinylidene fluoride (PVDF)

transducer. The rise time of the PW can be varied from

Fig. 3. Examples of PW waveforms generated by combination of

laser wavelengths, target material and confinement. Mode-locked

Nd:YAG (1064 nm) on aluminum foil (Al), ArF (193 nm) on black

polystyrene (PS), KrF (248 nm) on polyimide (PI), Q-switched ruby

(694 nm) on PS under direct ablation and under confined ablation

(Conf). The peak pressure of the waveforms has been normalized to

the same value for clarity.

5 ns (ablation of an aluminum foil by a mode-locked

Nd:YAG laser) to f 200 ns (ablation of a polystyrene

target by a Q-switched ruby laser under confinement).

In addition, the rise time can also be modified by taking

advantage of the steepening of the leading edge of the

PW that occurs naturally during non-linear propagation

of the PW in a medium.

From the point of view of hydrodynamics, pressure

waves, the type we have used for drug delivery, is

finite-amplitude waves. The difference from small-

amplitude or infinitesimal-amplitude waves is that the

sound velocity (c) is not constant (independent of

pressure). The sound velocity is given by Eq. (2):

c ¼ffiffiffiffiffiffidP

dq

sð2Þ

where P is the pressure and q is the density of the

material. For normal fluids, the curve of pressure

versus density is a concave upward (adiabatic com-

pression) and the slope increases with increasing

compression. This can affect significantly the propa-

gation of the PW. The rise time of a PW propagating

through a medium is altered by the linear and non-

linear propagation [47]. The linear attenuation, which

increases as a function of frequency, attenuates pre-

dominantly the high frequency components and

causes the rise time to increase. On the other hand,

in non-linear propagation the leading edge of the PW

becomes steeper as the PW propagates through the

medium. This is the result of the dependence of both

the sound and particle velocity (the bulk displacement

velocity) on pressure. The local velocity, which is the

sum of the sound and particle velocity, increases along

the leading edge of the PW which causes the leading

edge of the waveform to steepen that is the rise time to

decrease. The relative strength of the linear attenua-

tion and non-linear coefficient of the medium as well

as the initial peak pressure, the initial rise time, and

the distance traveled in the medium will determine the

final value of the rise time. Thus, we can take

advantage of the non-linear propagation to produce

PW with continuously adjustable rise times. Fig. 4

shows the shortening of the rise time, which occurs

after propagation of a distance of 800 Am in water.

The rise time decreased from the original value of 100

ns, behind the target, to 60 ns at 800 Am. The non-

linear propagation of the PW has been recently

Fig. 4. The rise time of a PW emerging from the target and after

propagation of a distance of 800 Am in water. The rise time

decreased from the original value of 100 ns, behind the target, to

60 ns.


applied to generate shock waves in order to induce

permeabilization of the nuclear envelope [28]. The

permeabilization of the nuclear envelope required a

higher pressure gradient (higher peak pressure, shorter

rise time or both) than the permeabilization of the cell

plasma membrane.

Eventually, a PW becomes a shock wave. The

salient feature of a shock wave is a very fast rise

time, which for all practical purposes amounts to a

discontinuity in pressure, density, particle velocity

and internal energy [57,58]. In water, the rise time of

a shock wave is of the order of a few picoseconds,

corresponding to a shock front thickness of a few

nanometers [59]. The large pressure gradients and

the high translational velocity of the molecules

within the shock front account for the unique inter-

actions of the shock waves with matter. The propa-

gation distance (L) required for the formation of a

shock wave as well as the rise time can be deter-

mined from the theory of non-linear acoustics [47],

Eqs. (3) and (4):

L ¼ lqc2

ePð3Þ

s ¼ qcmeP

ð4Þ

where l is the spatial width of the pressure transient

(temporal duration multiplied by the sound velocity),

q the density of the medium, c the sound velocity, ethe non-linear coefficient, P the peak pressure, s the

rise time and m the dissipative coefficient.

There are a number of different ways to produce

PW, such as extracorporeal shock wave lithotripters,

shock wave tubes as well as the use of energetic

materials. They all produce high amplitude PW. The

specific characteristics can vary over a large range of

values and depend on the design of the particular

device. Extracorporeal shock wave lithotripters and

shock tubes, in particular, have been used for drug

delivery. The former to permeabilize the cell mem-

brane and facilitate the delivery of macromolecules

into the cytoplasm [60,61], the latter for both trans-

dermal and cytoplasmic drug delivery [44,62,63].

Finally, the use of energetic materials could potential-

ly be useful in disposable devices for transdermal drug

delivery.

3. Biological effects of pressure waves

The research on drug delivery with PW is grounded

on the extensive investigations of the biological effects

of PW carried out in the last four decades. How do

pressure waves interact with cells and tissue? Surpris-

ingly, the mechanisms of these interactions are not well

known. The study of the effects of forces, whether

compression tension or shear, on cells has been the

subject of many investigations [64]. In the case of

pressure waves, however, the additional complication

is that the applied forces act on a very short time scale,

tens or hundreds of nanoseconds. Therefore, the

mechanisms of interactions of pressure waves with

cells and tissue are not expected to be the same. In

the past, most of the studies on the photomechanical

effects were phenomenological and focused on the

damage induced by pressure waves. Nevertheless,

they provide some background information of the

damage thresholds of cells and tissue. The use of

high power lasers in medicine over the last 20 years

has yielded a wealth of information on side effects,

collateral damage, potential hazards and safety

issues.

The first observations of lesions from exposure to

PW produced by high power lasers were reported in

Fig. 5. Schematic of the washer-target contraption. The laser

radiation was totally absorbed by the target and produced a single

PW. The PW propagated through drug solution, which also acted as

the acoustic coupling medium, impinged on the skin and permeabi-

lized the SC. The molecules diffused into the viable epidermis under

the concentration gradient until the barrier function of the SC

recovered.


the 1960s [65]. Of particular interest were observa-

tions that showed tissue injury in the brain and the

abdomen of mice located some distance from the

irradiated site. The injury was attributed to the pres-

sure waves generated during the irradiation of tissue.

Many investigations were undertaken in the ensuing

years. The interpretation of the observations, however,

was difficult because in many cases tissue injury was

mediated by other mechanical effects, such as plasma,

cavitation and streaming produced during laser irra-

diation. Over the last decade, we have developed a

methodology for the study of the effects induced by

PW on cells and tissue. This methodology eliminates

problems associated with direct laser irradiation. The

basic idea is to generate the PW by irradiation of

appropriate targets and launch them into the cell

cultures or tissue. In this approach cells or tissue are

not exposed directly to laser radiation [26,29,31].

The most interesting finding of the studies on the

interactions of PW with biological systems is that

these interactions depend on the characteristics of the

PW. Traditionally, it has been thought that the inter-

actions of PW with tissue were non-specific. We have

found that those interactions are specific and depend

on the PW parameters, such as peak pressure, rise

time or duration [30,44,66–68]. We have not fully

explored the extent of the specificity of the interac-

tions of the PW with different biological systems. This

is clearly a very important issue since, perhaps for the

first time, PW can be used to target tissues, cells or

even subcellular sites. The main target of the PW

(though not necessarily the only one) appears to be the

cell plasma membrane. These studies have elucidated

some of the modes of interactions of PW with cells

and subcellular structures and have led to the devel-

opment of methods for drug delivery. Interestingly,

both membrane permeabilization [66] and cell injury

depend on the rise time [30]. However, the pressure

threshold for cell injury appears to be higher than the

threshold for cell permeabilization, at least for the cell

lines that have been investigated [31].

Our conjecture is that the PW interact the strongest

with those cell structures that are of the same dimen-

sions as the spatial length of the PW. For example,

PW of a short rise time are effective in permeabilizing

the cell membrane [66]. In this case, there are indica-

tions that the membrane permeabilization is mediated

by membrane proteins (aquaporins) which form a

water-filled channel and are present in most mamma-

lian cells [69]. On the other hand, PWof long duration

generated can also permeabilize the cell membrane

[44]. The mechanisms of permeabilization in this case

are not known but since the membrane permeabiliza-

tion correlates with the impulse of the PW it is

possible that cell deformation might be responsible

for the permeabilization of the cell membrane.

4. Experimental arrangement

The experimental device for transdermal delivery

is shown in Fig. 5. A flexible washer, f 1.5 mm thick

and f 7 mm inner diameter, was used as a reservoir

for the solution to be delivered through the SC. A

target made out of black polystyrene f 1 mm thick

was used to generate the PW by ablation. The target

was placed in contact with the solution, which also

acted as the acoustic coupling medium. The PW

propagated from the target to the solution and im-

pinged on the surface of skin. A Q-switched ruby laser

(694 nm wavelength and f 28 ns pulse duration) was

used in the experiments. Pressure waves were gener-

ated by either direct or confined ablation. For con-

fined ablation, a thin overlay made of transparent

plastic (f 1 mm thick) was bonded on top of the

polystyrene target. In confined ablation, the expansion

of the generated plasma is delayed by the overlay. The


plasma maintains the applied pressure even after the

laser is switched off. This results in generation of PW

with higher peak pressure and longer duration, for a

given radiant exposure, than those generated by direct

ablation.

Peak pressures up to 1000 bar have been generated

by ablation of plastic targets in our lab. The PW were

measured using a piezoelectric transducer, with a 9-

Am thick PVDF film as the active element, and

recorded by a digital oscilloscope using a 1 MV

termination. The transducer was calibrated by mea-

suring the signal generated by a known momentum

transfer. A small steel ball bearing was dropped on the

transducer and allowed to bounce. The impact force

was calculated from the conservation of momentum,

the mass of the ball bearing and the time between

bounces [53]. The temporal resolution depends on the

thickness of the active element. The 9-Am thickness of

the PVDF film corresponds to a temporal resolution of

f 5 ns. The resolution was validated by measuring

the rise time of a PW with a known rise time [53]. The

peak pressure on the skin (PS) was calculated from

the measured peak pressure in the water (PW) and the

known acoustic impedance of water (ZW= 1.48� 106

kg m� 2 s� 1) and skin (ZS = 1.54� 106 kg m� 2 s� 1)

using Eq. (5):

PS

PW

¼ 2ZW

ðZS þ ZWÞð5Þ

Fig. 6 shows the sequence of steps for the appli-

cation of a PW for transdermal drug delivery on the

forearm of a volunteer. (A) A rubber washer was

attached to the skin with grease. (B) The washer was

filled with the drug solution to be delivered into the

skin. (C) The target material, black polystyrene, was

placed on top of the washer in contact with the

solution. (D) The articulating arm of the laser was

positioned over the target and the laser fired. The laser

radiation was totally absorbed by the target and

produced a single PW. The PW propagated through

Fig. 6. (A) A rubber washer was attached to the skin with grease.

(B) The washer was filled with the drug solution to be delivered into

the skin. (C) The target material was placed on top of the washer in

contact with the solution. (D) The articulating arm of the laser was

positioned over the target and the laser fired once to generate a

single PW. (E) The target and the washer were removed and the skin

was wiped clean.


the solution, which also acted as the acoustic coupling

medium, impinged on the skin and permeabilized the

SC. The permeabilization of the SC was strictly due to

the PW. Molecules diffused into the viable skin under

the concentration gradient until the barrier function of

the SC was recovered. (E) The target and the washer

were removed and the skin was wiped clean. There

was no change in the appearance of the skin after the

application of a PW. The same procedure was also

followed in the animal studies.

5. Transdermal drug delivery

There are a number of advantages of transdermal

drug delivery with PW:

1. The PW is only applied for a very short time (100

ns–1 As). The PW does not transport the drug

through the SC. It only transiently permeabilizes

the SC. The delivery of the drug takes place by

diffusion under the concentration gradient.

2. The recovery of the barrier function can be easily

modulated by changing the characteristics of the

PW or using chemical enhancers, such as sodium

lauryl sulfate (SLS). This allows to control the

amount of the drug delivered into the viable skin.

3. The size of the probe that has been delivered into

the epidermis, to this date, is 100 nm latex

particles. This is, to the best of our knowledge,

the largest size probe that has ever been delivered

by any transdermal delivery method.

4. Pressure waves have been shown to transiently

permeabilize the plasma membrane as well as the

nuclear envelope of cells. Furthermore, different

PW parameters are required for each of these

applications. Therefore, it is conceivable to apply

one type of PW to permeabilize the SC for

transdermal delivery and follow with a second

PW of different characteristics for delivery into the

cytoplasm or even into the cell nucleus.

5. Pressure waves can be thought of as a generic

technology platform for drug delivery into many

different biological systems (skin, cells, microbial

biofilms).

A very useful probe for human transdermal meas-

urements is y-aminolevulinic acid (ALA). ALA has

been approved by FDA and is currently used in

photodynamic therapy. ALA can be delivered by

iontophoresis and its pharmacokinetics for humans

as well as for a number of experimental animals are

known [70]. ALA is a small charged molecule and

therefore the rate of penetration is limited in normal

SC. It is the precursor of protoporphyrin IX (PpIX),

which is an intermediate in heme biosynthesis [71].

The synthesis of PpIX is the rate-limiting step in the

synthesis of heme and is regulated by the inhibition of

ALA synthase by the heme. However, application of

exogenous ALA enables the cells to bypass this rate-

limiting step and produce excess amounts of PpIX

because the rate of PpIX production is faster than the

rate of conversion of PpIX to heme. Furthermore,

since PpIX is produced in the viable skin, the presence

of ALA in the SC does not interfere with the measure-

ments of PpIX. The peak of PpIX fluorescence is at

634 nm (excitation 405 nm) while ALA does not

absorb or fluoresce at this wavelength. Therefore, the

transport of ALA through the SC can be followed by

monitoring the PpIX fluorescence. A fiber-based

spectrofluorometer can be used to measure the PpIX

fluorescence over time. Fig. 7 shows the fluorescence

emission spectrum of PpIX of a site on the forearm of

a volunteer exposed to a single PW in the presence of

ALA solution (5% w/v in water). The fluorescence

emission of a control site (treated in an identical

manner but without exposure to a PW) is also shown

for comparison.

The following is a summary of the results:

1. A single PW was sufficient to increase the

permeability of the SC and produce a strong

fluorescence signal from PpIX.

2. The amount of ALA transported across the SC in-

creased with the ALA concentration (5–20% w/v).

3. The permeabilization of the SC was transient. The

barrier function of the SC always recovered.

4. The permeabilization of the SC depended on the

peak pressure. The onset of the permeabilization of

the SC was observed at f 350 bar and increased

with increasing pressure.

The observation of PpIX fluorescence from the

skin indicates that the epidermis was still viable after

the application of a PW. It has been shown previously

that the keratinocytes are a sensitive indicator of tissue

Fig. 7. The fluorescence emission spectrum obtained from human

skin (the inner volar forearm) exposed to a single PW in the

presence of ALA. The fluorescence emission of PpIX peaks at 634

nm (excitation 405 nm). The fluorescence emission spectra were

recorded before the application of PW baseline (BL) as well at 3 and

5 h post-treatment (3PW and 5PW). The fluorescence emission

spectra of a control site, a site treated in an identical manner except

that no PW was applied, at 3 and 5 h (3C and 5C) are also shown for

comparison. The 5 h control overlaps with the baseline.

Fig. 8. The fluorescence intensity of PpIX at 634 nm as a function

of time following the application of a single PW. The three sites on

the forearm were exposed to a PW of different peak pressure

(diamond) 390 bar, (squares) 440 bar, and (circles) 500 bar.


injury caused by PW [37]. However, transmission

electron microscopy of biopsies taken either immedi-

ately [72] or 24 h after the application of a PW [67]

showed that there was no observable injury to the

keratinocytes. Therefore, it appears again that the

pressure threshold for the permeabilization of the SC

is lower than that for injury of the viable skin.

Interestingly, when three PW were applied on the

skin the level of PpIX fluorescence was lower than

that of the single PW (unpublished observations).

Since the permeabilization of the SC could not be

lower with three PW than with one, the conversion of

ALA to PpIX had to be lower. This implies that

multiple PW might have caused some cell injury,

though this injury could not be seen in biopsies with

transmission electron microscopy. The conversion of

ALA to PpIX may prove to be a simple assay of cell

injury.

Fig. 8 shows the fluorescence intensity of PpIX at

634 nm from a volunteer as a function of time after the

application of a single PW, for three different peak

pressures. It should be kept in mind that the fluores-

cence intensity of PpIX reflects the kinetics of the

conversion of ALA to heme. The fluorescence intensity

of PpIX does not yield the instantaneous production of

PpIX but the accumulation of PpIX over timeminus the

conversion of PpIX to heme. It is also possible that

some of the ALA that entered the viable skin was taken

up by the vasculature and removed from the epidermis.

The peak of the fluorescence intensity versus time

shifts toward the longer times with increased accumu-

lation of PpIX. This is not surprising because as the

accumulation of PpIX increases it may take a longer

time to clear it from the skin.

The permeabilization of the SC and subsequent

delivery of ALA depended on the peak pressure. The

pressure threshold for the SC permeabilization was

observed at approximately 350 bar and increased

dramatically at the highest peak pressure. It should

be pointed out that this pressure threshold value is for

a particular site (inner volar forearm) and volunteer. In

fact, the threshold pressure varied among sites, indi-


viduals and skin conditions. In addition, the pressure

threshold for SC permeabilization depends on the

other characteristics of the PW. For example, it

appears that the pressure threshold for the permeabi-

lization of the SC is lower for long duration PW [62].

The application of PW did not cause any pain or

discomfort [67]. Pressure waves of 300 ns duration

did not produce any sensation whatsoever. On the

other hand, PW of 1 As duration generated a sensation

but not pain. With respect to skin changes after the

application of a PW, the 300 ns PW did not produce

any changes of the skin, while the 1 As PW produced

a minor erethyma, which disappeared within 10–15

min. One interesting observation was that the amount

of ALA delivered through the SC correlated with the

transepidermal water loss (TEWL). The higher the

TEWL the more ALA was delivered into the viable

skin (unpublished observations). Higher rates of

TEWL may indicate higher levels of hydration, which

would be consistent with our current understanding of

the mechanism of permeabilization of the SC.

Fig. 9 shows the results of transdermal delivery of

40-kDa rhodamine-B dextran with a single PW in a rat

animal model. The reservoir on the rat skin was filled

with an aqueous solution of rhodamine-B dextran

(500 AM) and a single PW (f 730 bar peak pressure,

f 490 ns pulse duration FWHM, and f 190 ns rise

time) was applied. The solution remained in contact

with the skin for 5 min and then was removed. Full

thickness skin biopsies were obtained approximately 1

h post-treatment, embedded in OCT and frozen in dry

ice. Control sites were treated in the same way except

that no PW was applied. The fluorescence image from

a frozen biopsy and the corresponding transmission

image of a site exposed to a single PW are shown in

Fig. 9A, B, respectively [68]. There is a broad band of

fluorescence from the rhodamine-B dextran that

extends from the SC to a depth of f 300 Am into

the dermis. The thickness of the SC, epidermis and

dermis in the rat are 9–15, 8–18 Am and f 1 mm,

respectively. For comparison, the thickness for the

human skin is 10–15 Am for the SC, 50–100 Am for

the epidermis and 2–3 mm for the dermis. Fig. 9C, D

show the fluorescence and transmission images from

the frozen biopsy of a control site, respectively. The

fluorescence of the control site originates in the SC.

The fluorescence images of the sites exposed to a

PW show that the probe was distributed uniformly

into the epidermis and dermis. This implies that the

permeabilization of the SC was uniform rather than

limited to a few sites. There are a number of dark

areas (low concentration of the fluorescent probe) in

the fluorescence images. Comparison with the trans-

mission micrographs show that they coincide with

hair follicles. The reduced delivery in the hair follicles

is probably due to the difference of the acoustic

impedance between the dermis and the hair follicle

cavity. In addition to the 40-kDa dextran molecules,

fluorescent latex particles 20 and 100 nm in diameter

have been transported through the SC with a PW

[26,73]. The transdermal delivery of 100-nm particles

is, to the best of our knowledge, the largest size probe

that has ever been delivered by any transdermal

delivery method. Thus, PW can facilitate the delivery

into the epidermis of very large molecules such as

proteins and DNA plasmids. In addition, novel probes

such as quantum dots could be delivered into the

epidermis and dermis. Another potential application is

the transdermal delivery of encapsulated drugs. Drugs

can be incorporated in time-release microspheres and

delivered over an extended period of time.

As we have mentioned earlier, the biological

effects of PW depend on their characteristics. In

transdermal delivery, the permeabilization of the

SC depends on the duration of the PW. Fig. 10

shows a comparison of the penetration depth of the

40-kDa rhodamin-B dextran in a rat by PW of

different duration [68]. The two types of PW used

in these experiments had different temporal charac-

teristics shown in Fig. 11. The fluorescence photo-

micrograph shows the transdermal delivery with the

long duration (490 ns FWHM) PW. The transdermal

delivery of the short duration (110 ns FWHM) PW is

shown in the insert. The fluorescence photomicro-

graph and the insert are shown under the same

magnification. There is a dramatic difference in the

penetration depth of the 40-kDa dextran produced by

the two types of PW. The penetration of the dextran,

when the long duration PW was used, reached to a

depth of f 300 Am into the dermis. In comparison,

the penetration depth, when the short duration PW

was applied, was only f 50 Am.

It is of interest to compare the permeabilization

pattern produced by the PW with that of ultra-

sound. Mitragotri and coworkers [74] have shown

that as the frequency of the ultrasound increased

Fig. 9. The fluorescence photomicrographs from (A) the biopsy of the rat skin exposed to a PW and (C) a control site. An aqueous solutions of

rhodamine-B dextran (40-kDa) at 500 AM concentration was used for transdermal delivery. The corresponding transmission photomicrographs

from (B) the site exposed to a PW and (D) the control site. Scale bar 200 Am.


from 19.6 to 58.9 kHz, the pattern of permeabiliza-

tion of the SC changed from a single spot to

multiple foci, thus, presenting a more uniform

pattern. Furthermore, the permeabilization threshold

Fig. 10. Transdermal delivery of the 40 kDa rhodamin-B dextran

with PW of different characteristics. The fluorescence photomicro-

graph shows the transdermal delivery with the long duration (490 ns

FWHM) PW. The transdermal delivery of the short duration (110 ns

FWHM) PW is shown in the insert. Scale bar 200 Am.

Fig. 11. (A) A long duration PW generated by confined ablation.

The pulse duration is f 490 ns FWHM and (B) a PW generated by

direct (unconfined) ablation. The pulse duration is f 110 ns

FWHM. The two waveforms have been normalized to the same

peak pressure for clarity. The actual peak pressures are (A) 730 bar;

(B) 600 bar.


increased with frequency. These observations are

consistent with the fact that PW with a frequency

bandwidth, which covers the range from kHz to

tens of MHz, can produce a uniform permeabiliza-

tion of the SC and subsequent uniform delivery

into the epidermis and dermis. It appears, therefore,

that the two technologies of transdermal delivery,

namely ultrasound and PW, may converge at high

frequencies.

Transdermal drug delivery with PW has been

reported by Ogura et al. [75]. The photosensitizer,

Photofrin II, was delivered by a PW (520 bar peak

pressure, 1 As duration) in a rat animal model. What is

worth mentioning is that when the skin was heated

prior to the application of the PW to a surface

temperature of 47 jC the penetration depth increased

from 75 to 100 Am. This was probably caused by the

increased fluidity of the lipids at the elevated temper-

ature. This observation is consistent with the view

that the PW target the weaker lipid domains rather

than the corneocytes which are protected by the

cornified envelope and the covalently bound lipid

envelope.

Fig. 12. The recovery of the SC barrier with ( + SLS) and without

(� SLS) sodium lauryl sulfate. The SC recovery was measured by

monitoring the fluorescence intensity at 596 nm (excitation 568 nm).

The fluorescence intensity was measured after tape-stripping to

reduce the fluorescence of dextran in the SC. The fluorescence of the

skin without dextran (baseline) was also subtracted. The fluorescence

intensity (averageF S.D.) at time points of 2, 30 and 60 min post-

exposure are shown. The measurements of the recovery of the SC

without the use of SLS are shown for comparison. For the time point 0

min, the PW was applied with the reservoir filled with an aqueous

solution of dextran. For the time points 2, 10, 15, and 30 min post-

exposure, the PW was applied with the reservoir filled with water.

6. The synergy of pressure waves and sodium

lauryl sulfate

The permeabilization of the SC depends on the PW

parameters. Both the PW peak pressure as well as the

duration can modulate the efficiency of the SC per-

meabilization. In addition, molecular transport through

the SC increases dramatically when SLS, an anionic

surfactant, is administered together with a PW [76].

Therefore, the use of SLS may have the same effect as

higher peak pressure or longer duration. The use of

SLS and surfactants, in general, allow greater flexibil-

ity in the choice of parameters for the optimization of

drug delivery which can result in simpler PW gener-

ation schemes and reduced collateral damage. The use

of SLS has been shown to enhance drug delivery in

phonophoresis as well [77].

The use of SLS delayed the recovery of the SC

barrier from < 2 min to >30 min [26,76]. This in-

creased substantially the amount of dextran that was

delivered into the epidermis and dermis. The presence

of the dextran in the skin was measured by fluores-

cence emission spectroscopy. Fluorescence emission

spectra of the exposed and control sites were collected

while the animal (rat) was alive under anesthesia.

Fluorescence emission spectra (excitation 488 nm,

emission 498–600 nm) were obtained using a fiber-

based spectrofluorimeter. Fluorescence emission spec-

troscopy allows the investigation of transdermal de-

livery quickly and non-invasively. The fluorescence

intensity depends on the concentration of the mole-

cules in the tissue and the scattering properties of the

tissue. However, because all experiments were carried

out in adjacent sites on the dorsal side of the rat, the

optical properties of the skin sites were expected to be

comparable. The experimental procedure for measur-

ing the recovery of the barrier function of the SC was

to apply the PW with water or SLS solution (2% w/v)

in the reservoir. Subsequently, the contents of the

reservoir were removed within 1 min and the reservoir

was filled, at different time points, with an aqueous

solution of the 40-kDa rhodamine-B dextran. The

presence or the absence of rhodamine-B dextran in

the epidermis at the particular time point indicated


whether the SC was still permeable or its barrier

function had recovered. Fig. 12 shows the fluorescence

intensity at 596 nm (the peak fluorescence intensity of

rhodamine-B dextran) at different time points post-

treatment with and without SLS.

The SC recovery as measured in these experiments

was the recovery time that corresponded to the partic-

ular probe used. In the present case, the 40 kDa dextran

with a diameter of approximately 9 nm. This recovery

time is not the same as the absolute recovery time of

the barrier function of the SC. In fact, TEWL measure-

ments gave a recovery time of f 15 min when water

was used as the coupling medium (unpublished obser-

vations). This contrasts to < 2 min recovery time for

the 40 kDa dextran. This apparent discrepancy is not

surprising. If the PW creates transient channels in the

SC, as we hypothesize, the transport of molecules

Fig. 13. The fluorescence images obtained from frozen biopsies of (A) a sit

sites. The fluorescence images were obtained under identical conditions bu

a PW and (D) a control site. The PW used in both experiments were id

fluorescence intensity increases from violet (lowest intensity) to red (high

through the SC will continue until the size of the

channels becomes smaller than the size of the mole-

cules. Therefore, large molecules will be blocked from

going through the SC earlier in the recovery process

than small molecules.

The fluorescence images from the frozen biopsies

(Fig. 13) are shown in pseudocolor intensity scale.

The fluorescence intensity is a measure of the

concentration of rhodamine-B dextran in the epider-

mis and dermis. The use of the SLS in combination

with the PW increased substantially both the amount

of dextran in the viable skin and the penetration

depth. The use of SLS also increased the concentra-

tion of the dextran in the SC as shown in the control.

Fig. 14 shows the fluorescence intensity profile of

the 40-kDa rhodamine-B dextran as a function of the

penetration depth for transdermal delivery with and

e exposed to a single PWand (B) a control site. SLS was used in both

t without the use of SLS from frozen biopsies of (C) a site exposed to

entical. The images are shown in pseudocolor intensity scale. The

est intensity). Scale bar 200 Am.

Fig. 14. The fluorescence intensity profile of the 40-kDa rhodamine-

B dextran as a function of the penetration depth for transdermal

delivery with or without SLS. The fluorescence intensity was

measured along the lines shown in Fig. 12A, C. The intensity for

each profile was normalized to its maximum value.


without SLS. The fluorescence intensity was mea-

sured along the lines shown in Fig. 12A, C. The

intensity for each profile was normalized to its

maximum value.

Fig. 15. High magnification electron micrograph of the SC (A) from

a control site and (B) exposed to a single PW with water as the

coupling medium. The expanded lacunar domains within the

intercellular lamellae can be seen.

7. The mechanism of the permeabilization of the

stratum corneum

Transdermal delivery can occur whether the mole-

cules are present during the application of the PW or

introduced after the PW [67]. Given the short duration

of the PW, a few microseconds at the most, the effect of

the PW is probably limited to the permeabilization of

the SC. The diffusion of the drug occurs under the

concentration gradient through the channels produced

by the PW. The barrier function of the SC always

recovers.

In a series of in vivo experiments, sites on the inner

volar forearm of volunteers were exposed to a single

PW with water as the acoustic coupling medium [72].

Skin biopsies were obtained immediately after the

experiments and fixed in Karnovsky’s fixative for 1

h at room temperature. The exposed sites, post-fixed

with RuO4, showed many highly expanded SC extra-

cellular domains near continuous lacunar domains

(Fig. 15). The lacunae have been defined as electron

lucent areas embedded within the lipid bilayers span-

ning the SC extracellular domains, and considered as

the putative pores [78]. These domains were not

present at every level, i.e. between every corneocyte,

as normal bilayers could be seen in the extracellular

spaces of corneocytes that were immediately below.

However, it should be kept in mind that the lacunae

seen in electron micrographs are in essence cross

sections of the three dimensional trabecular network

which may form a continuous, permeable lacunar

system. No ultrastructural changes were seen in the

morphology of individual corneocytes. When SLS was

used as the coupling medium to enhance transdermal

delivery the expansion of the lacunae system was

significantly larger. Fig. 16 shows the low magnifica-

Fig. 16. Low magnification electron microscopy of the SC stratum

granulosum junction (A) with water as the coupling medium and

(B) with SLS as the coupling medium. The lacunae system of the

site where SLS was used is noticeably larger.


tion electron microscopy of the SC stratum granulosum

(SG) junction with water and SLS as the coupling

medium, respectively. The lacunar system of the site

where SLS was used was noticeably larger.

The expansion of the lacunar domains could pos-

sibly create transient channels, which enable drug

delivery through the SC and into the epidermis and

dermis. We hypothesize that under the action of a PW

the lacunar system forms a continuous pathway that

allows the passive diffusion of the drug under the

concentration gradient. The actual physical mecha-

nism is not known. Our conjecture is that the free

water within the SC [79] is involved in the permeabi-

lization process. Water can be considered incompress-

ible in the time scale of the duration of the PW. The

free water has to go somewhere. It is possible that

under the high pressure gradient generated by the PW,

the free water is forced in the constricted domains of

the lacunar system expanding them and thus, forming

a continuous pathway.

As it evolved, the mammalian SC is designed to

function as a barrier in diverse and specialized

environments (low and high humidity, extremes of

temperature, UV radiation and other physical fac-

tors), while functioning as a sensory transducer as

well. As a composite biopolymer of proteins and

lipids, arranged in a ‘‘brick and mortar’’ organiza-

tion, the corneocytes provide the scaffolding for the

waterproofing lipid bilayers. Physical forces, such as

PW, ultrasound and electric fields, target the weaker

lipid domains rather than the tough corneocytes

protected by the cornified envelope and the cova-

lently bound lipid envelope. The recent observations

by Ogura et al. [75] that the combination of heating

the skin and PW enhances transdermal delivery is

consistent with this view. Again, within the lipid

domains, it is the hydrophilic regions such as the

lacunae that contain water that are susceptible to

most physical and chemical agents used for perme-

ability enhancement.

Surprisingly, we have been unable to permeabilize

the SC ex vivo, although the same PW would per-

meabilize the SC in vivo. It should be pointed out that

low-frequency ultrasound has been applied success-

fully to cadaver skin for transdermal drug delivery

[24]. However, ultrasound may interact with the SC

differently than the PW. Low-frequency ultrasound

works predominantly through cavitation induced by

the tensile component. Confocal micrographs show

significant bleaching of fluorescein-loaded SC after

ultrasound exposure [80]. The bleaching was attribut-

ed to the oxidation of fluorescein by cavitation-gen-

erated free radicals.

Routine electron microscopy (OsO4 post-fixed) did

not reveal any noticeable differences between the

control sites and the PW exposed sites. The nucleated

epidermis as well as the dermis maintained their

typical ultrastructural features with no indication of

damage either in the extracellular matrix or the

cellular components [72]. In addition, transmission

electron microscopy of human biopsies taken 24

h following the exposure of a PW showed no damage

to the subcellular organelles [67].

Fig. 17. The picture of the back of a guinea pig treated with the

allergen dinitrochlorobenzene with (A) the Finn chamber under

occlusion for 21 h, (B) a single PW with water as the coupling

medium followed by the application of the allergen for 5 min and

(C) the Finn chamber under occlusion for 5 min as a control.


8. Examples of transdermal drug delivery

8.1. Delivery of allergens

Pressure waves can be applied for rapid delivery

of allergens and thus, make it possible to differenti-

ate irritant from allergic contact dermatitis [81].

Presently, the suspect allergen is applied at sub-irritant

concentrations to the skin under occlusion (Finn

chamber) for a period of up to 48 h to maximize

penetration [82]. Once the patch is removed, the site is

clinically examined for morphologic evidence of an

eczematous response. If the subject develops such a

response at a concentration below the irritant threshold

concentration, the eczematous lesion is considered to

be an allergic response to the tested substance. Pres-

sure waves allow the rapid transdermal delivery of

allergens and thus, improve the optimal penetration of

the allergen across the SC. This has the potential to

reduce the exposure time for the clinical manifestation

of the challenge and improve the accuracy of the

procedure.

The allergic skin reaction using PW delivery was

compared to 5 min and 21 h occlusion in a

sensitized hairless albino guinea pig model. The

pigs were sensitized by intradermal injection of

(0.01%) dinitrochlorobenzene and topical adminis-

tration (0.1%, 1 week later) of the hapten. One

month later, testing for the allergic response was

performed by the administration of 10 Al of 0.1%

dinitrochlorobenzene with a PW. The picture of the

back of a pig (Fig. 17) shows two skin sites treated

under occlusion for 21 h and 5 min using the Finn

chamber, respectively. In addition, a single PW was

applied to one site with water as the coupling

medium followed by the application of the allergen

for 5 min. The skin site treated with the Finn

chamber under occlusion for 21 h showed an

erythematous and edematous skin reaction, which

in some cases resulted in skin maceration and

necrosis. These reactions always extended beyond

the contact site of the skin with the allergen. On the

other hand, skin sites treated with a PW showed a

pink, well demarcated erythematous area confined

to the beam diameter at 24 and 48 h after delivery.

The control sites, exposed to the allergen under

occlusion for 5 min, showed no clinically percepti-

ble reaction.

8.2. Systemic delivery of insulin

Pressure waves can also facilitate the transdermal

delivery of drugs for systemic treatment [83]. Insulin

was used as the probe for systemic delivery because

transport through the SC can be easily monitored by

measuring the glucose level in the blood. The animal

model was the streptozotocin diabetic rat. A two-step

procedure was used in order to increase the amount of

insulin delivered into the viable skin [83]. For the first

step, the reservoir was filled with 2% w/v aqueous

solution of SLS. The SLS was allowed to remain in

contact with the skin for 2 min. This step was intended

to enhance the permeabilization of the SC. For the

second step, the SLS solution was removed and the

reservoirs was filled with a solution of porcine insulin

400 U ml� 1 adjusted to pH 4. The second target was

driven by the laser pulse into the reservoir like the

plunger of a syringe. Therefore, the first laser pulse in

this procedure produced a PW, which permeabilized

the SC while the second laser pulse drove the target

into the reservoir by exerting a hydrodynamic pulse

on the insulin solution. Fig. 18 shows the glucose

level of three diabetic rats over time after the proce-

Fig. 18. Blood glucose kinetics following PW delivery of insulin in

diabetic rats. Shown for comparison is the blood glucose kinetics

following intramuscular injection of insulin (0.1 and 0.3 U).


dure. The blood glucose of the diabetic rat dropped

from >350 to < 100 mg dl� 1 or f 80% of the initial

glucose level. Overall, the blood glucose remained

within the normal range for f 3 h. These experiments

suggest that the amount of a drug that can be delivered

through the SC can reach therapeutic levels. Compar-

ison of glucose kinetics after the application of PW

with the kinetics of intramuscular injection of insulin

indicated that the total amount of insulin delivered

through the SC was between 0.1 and 0.3 U [83]. From

the insulin concentration, the skin area treated, and the

time the insulin solution remained in contact with the

Fig. 19. The design concept of a disposable transdermal patch based on th

energy for the generation of the PW, (B) once the SC is permeabilized th

skin, we can estimate that between 0.3 and 0.9 Al ofinsulin solution was transported through the SC. This

corresponds to an average value of skin permeability

between 4� 10� 4 and 1.2� 10� 3 cm h� 1.

The application of the PW did not affect the activity

of insulin. In addition, in experiments where DNA

solutions were exposed to PW no strand breaks were

observed (unpublished observations). This is not sur-

prising, a review of the literature shows that the

pressure required to have any effect on molecules or

induce any chemistry is two orders of magnitude higher

than that used in our experiments [84–86]. Further-

more, in those particular cases the waves were truly

shock waves with rise times of the order of 10–100 ps.

That means that the pressure gradients in our experi-

ments were five to six orders of magnitude lower than

the pressure gradients that are required to produce

chemical reactions.

9. Conclusions

A PW can effectively deliver drugs through the SC,

the cell plasma membrane as well as into microbial

biofilms. Furthermore, the PW parameters required for

efficient drug delivery are different for different bio-

logical systems. In the case of transdermal drug deliv-

ery, the permeabilization of the SC is transient and its

barrier function always recovers. A PW can facilitate

the delivery ofmacromolecules, the size of proteins and

DNA plasmids, in the epidermis and deep into the

dermis. Furthermore, the drug delivery can be of

sufficient quantity to produce systemic treatment. The

mechanism of permeabilization is probably caused by

the disruption of the hydrophilic domains of the SC.

Although the PW can be characterized as broad-

band pulsed ultrasound, we have tended to retain the

e use of energetic material. (A) The energetic material provides the

e drug can diffuse into the epidermis and dermis.


term pressure wave because its characteristics are so

different from the ultrasound that has been tradition-

ally used in phonophoresis. Furthermore, it is highly

probable that the interactions of PW with tissue, cells

and subcellular structures are fundamentally different

than those of ultrasound.

The laser has been the preferred method of gener-

ation of PW in our work. The laser has proven a

unique tool for the study and development of methods

of drug delivery. However, the laser itself is not a

necessary component of drug delivery systems. In fact,

our current concept for transdermal drug delivery is

based on the use of energetic material, which can

produce efficiently high pressure transients with small

amounts of energetic material [45]. Fig. 19 shows the

design concept of a disposable transdermal patch

based on the use of energetic material. The patch

roughly the size of a bandage contains the energetic

material and the drug to be delivered through the SC.

The energetic material provides the energy for the

generation of the PW. Once the SC is permeabilized

the drug can diffuse into the epidermis and dermis.

This scheme of delivery avoids the substantial cost of

the laser system. Furthermore, makes possible a deliv-

ery device, which could be appropriate for home use.

Acknowledgements

The work on drug delivery at the Wellman

Laboratories of Photomedicine was supported by the

DoD Medical Free Electron Program under contracts

N00014-94-1-0927 and F4 9620-01-1-0014.

References

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