NO334437B1 - Microporation of human skin for drug delivery and monitoring applications - Google Patents

Abstract

A method is described for improving the permeability of the skin (120, 274) to an analyte for diagnostic purposes or to a medicament for therapeutic purposes, using microporation and possibly sound energy and a chemical enhancer. If sound energy is selected, it can be modulated by frequency modulation, amplitude modulation, phase modulation and / or combinations thereof. Microporation is achieved by (a) melting the stratum corneum (274) with localized rapid heating of water so that the water evaporates, thereby eroding cells; (b) puncturing the stratum corneum (274) with a microlance calibrated to form a micropore with a diameter of up to 1000 (am); (274) (d) hydraulic puncture of the stratum corneum (274) with a high pressure fluid jet to form a micropore up to about 1000 nm in diameter; or (e) puncture of the stratum corneum (274) with short electrical pulses for to form a micropore of up to about 1000 μm in diameter.

Description

BACKGROUND OF THE INVENTION

This invention relates to a minimally invasive to noninvasive method of increasing skin permeability through microporation of the stratum corneum, which can be combined with sound energy, chemical permeation enhancers and pressure, to selectively enhance the outward flow of analytes from the body for monitoring such or for drug delivery into the body.

The stratum corneum is mainly responsible for the skin's well-known barrier properties. It is thus this layer that forms the biggest barrier against transdermal flow of drugs or other molecules into the body and of analytes out of the body. The stratum corneum, the outer horny layer of the skin, is a complex structure of compact keratinized cell debris separated by lipid regions. Compared to the oral or gastric mucosa, the stratum corneum is significantly less permeable to molecules externally as well as internally within the body. The stratum corneum is formed by keratinocytes that comprise the majority of epidermal cells that lose their nucleus and become corneocytes. These dead cells comprise the stratum corneum, which only has a thickness of approx. 10-30^m and which, as mentioned above, is a very resistant waterproof membrane that protects the body against the penetration of substances from the outside and against the outward migration of liquids and dissolved molecules. The stratum corneum is continuously renewed through the shedding of corneum cells during desquamation and the formation of new corneum cells through the keratinization process.

The flow of a drug or analyte through the skin can be increased by changing either the resistance (diffusion coefficient) or the driving force (diffusion gradient). Flow can be enhanced by using so-called penetration enhancers or chemical enhancers. Chemical enhancers are well known in the art, and a more detailed description will be given later.

Another method of increasing the skin's permeability to drugs is iontophoresis. iontophoresis involves the application of an external electric field and local delivery of an ionized form of drug or a non-ionized drug carried with the water flow associated with ion transport (electro-osmosis). Although permeation enhancement through iontophoresis has been effective, the technique is associated with problems with control of drug delivery and irreversible skin damage.

Sound energy has also been used to improve the permeability of the skin and synthetic membranes to drugs and other molecules. Ultrasound has been defined as mechanical pressure waves with frequencies higher than 20 kHz, H. Lutz et al., Manual of Ultrasound 3-12 (1984). Ultrasound energy is generated by vibrating a piezoelectric crystal or other electromechanical element, by passing an alternating current through the material, R. Brucks et al., 6 Pharm. Res. 697 (1989). The use of ultrasound energy to increase the skin's permeability to drug molecules is called sonophoresis or phonophoresis.

Although it has been recognized that improved permeability of the skin should theoretically allow molecules to be transported from the inside of the body through the skin to the outside of the body for collection or monitoring, no practicable methods have been described. US Patent 5,139,023 to Stanley et al., discloses an apparatus and method for noninvasive blood glucose monitoring. In this invention, chemical permeation enhancers are used to increase the permeability of mucosal tissue or skin to glucose. Glucose then passively diffuses through the mucosal tissue or skin and is captured in a recipient medium. The amount of glucose in the recipient medium is measured and correlated to determine the blood glucose level. However, as described by Stanley et al., this method is more effective when used on mucosal tissue, such as buccal tissue, which results in detectable amounts of glucose being collected in the recipient medium after a time delay of approx. 10-20 minutes. However, the method described by Stanley et al., results in an extremely long time delay ranging from 2 to 24 hours, depending on the chemical enhancer mixture used, before detectable amounts of glucose can be detected diffusing through human skin (heat-separated epidermis) in vitro. These long time delays can be attributed to the long time required for the chemical permeation enhancers to passively diffuse through the skin and enhance the permeability of the stratum corneum barrier, as well as the long time required for the glucose to passively diffuse out through the skin. Stanley et al., thus clearly do not describe a method for transporting blood glucose or other analytes non in vasi vt through the skin in a way that allows for rapid monitoring, as is required for blood glucose monitoring in diabetics, and for many others analytes in the body, such as blood electrolytes.

Although the use of sound energy for drug delivery is known, the results have been largely disappointing in that the permeability enhancement has been relatively small. There is no general agreement on the effectiveness of sound energy to increase drug flux through the skin. While some studies report successful sonophoresis, J. Davick et al., 68 Phys. Ther. 1672 (1988); J. Griffin et al., 47 Phys. Ther. 594 (1967); J. Griffin & J. Touchstone, 42 Am. J. Phys. With. 77 (1963); J. Griffin et al., 44 Am. J. Phys. With. 20 (1965); D. Levy et al., 83 J. Clin. Invest. 2074); D. Bommannan et al., 9 Pharm. Res. 559 (1992); others have obtained negative results, H. Benson et al., 69 Phys. Ther. 113 (1988); J. McElnay et al., 20 Br. J. Clin. Pharmacol. 4221 (1985); H. Pratzel et al., 13 J. Rheumatol. 1122

(1986). Systems using rodent skin have shown the most promising results, while systems using human skin have generally produced disappointing results. It is well known to the person skilled in the art that the skin of rodents is much more permeable than human skin, and the above results consequently do not tell the person skilled in the art how the sonophoresis can be effectively utilized for transdermal delivery and/or monitoring through human skin.

A significant improvement in the use of ultrasound energy for monitoring analytes and also for delivering drugs to the body is described and claimed in co-pending application serial number 08/152,442 of November 15, 1993, now US Patent 5,458,140 and Serial No. 08/152,174 of December 8, 1993, now US Patent 5,445,611.1 these inventions achieve the transdermal sampling of an analyte or the transdermal delivery of drugs using sound energy that is modulated in intensity, phase or frequency, or a combination of these parameters coupled with the use of chemical permeation enhancers. Also the application of sound energy, possibly with modulation of frequency, intensity and/or phase, to be able to press and/or pump molecules through the stratum corneum via perforations introduced through needle puncture, hydraulic jet, laser, electroporation or other methods, is described.

The formation of micropores (ie, microporation) in the stratum corneum to promote drug delivery has been the subject of various studies and has resulted in the issuance of patents for such techniques.

Jacques et al., J. Invest. Dermatol. 88-93 (1987) describes a method for administering a drug by ablating the stratum corneum of an area of skin by using pulsed laser light of sufficient wavelength, pulse length, pulse energy, pulse number and pulse repetition rate to melt away the stratum corneum without significantly damaging the underlying epidermis , and then apply the drug to the ablation area. This work has resulted in the issuance of US patent 4,775,361 to Jacques et al. Ablation of skin through the use of ultraviolet laser irradiation has previously been reported by Lane et al., 121 Arch. Dermatol. 609-617 (1985). Jacques et al., is limited to the use of a few light wavelengths and expensive lasers.

Tankovich, US Patent 5,165,418 (hereinafter "Tankovich '418") describes a method of obtaining a blood sample by irradiating human or animal skin with one or more laser pulses of sufficient energy to cause vaporization of the skin tissue so that a hole is formed in the skin which extends through the epidermis, and to cut off at least one blood vessel, which means that a quantity of blood can be led out through the hole so that it can be collected. Tankovich '418 is thus insufficient for noninvasive or minimally invasive permeabilization of the stratum corneum, so that a drug can be delivered to the body, or an analyte from the body can be analyzed.

Tankovich et al., US Patent 5,423,803 (hereinafter "Tankovich '803") describes a method for laser removal of superficial epidermal skin cells in human skin for cosmetic applications. The procedure includes applying a light-absorbing "contaminant" to the outer layers of the epidermis and introducing some of this contaminant into the intercellular spaces in the stratum corneum, and irradiating the infiltrated skin with pulses of laser light of sufficient intensity so that the amount of energy absorbed by the contaminant will cause the contaminant to explode with sufficient energy to tear off some of the epidermal skin cells. Tankovich '803 further describes that the contaminant must have a high energy absorption at the wavelength of the laser beam, that the laser beam must be a pulsed beam of less than 1 us duration, that the contaminant must be forced into the upper layers of the epidermis and that the contaminant must explode with sufficient energy to to tear off epidermal cells after absorption of the laser energy. This invention also lacks description or suggestion of a method for drug delivery or analyte collection.

Raven et al., WO 92/00106, describes a method for selectively removing diseased tissue from the body, by administering to the selected tissue a compound which is a strong absorber of infrared radiation of 750-860 nm wavelength, and irradiating the area with equivalent infrared radiation of sufficient strength to cause heat vaporization of the tissue to which the compound was administered, but insufficient to cause vaporization of tissue to which the compound was not administered. The absorbent compound should be soluble in water or serum, such as indocyanine green, chlorophyll, porphyrins, heme-containing compounds or compounds containing a polyene structure, and the power levels are in the range of 50-1000W/cm<2> or even higher.

Konig et al., DD 259351, describes a method for heat treatment of tumor tissue, which comprises placing a medium in the tumor tissue that absorbs radiation in the red and/or near-red infrared spectral range and irradiating the infiltrated tissue with laser light of the correct wavelength . The absorbing medium can be methylene blue, reduced porphyrin or its aggregates, as well as phthalocyanine blue. Methylene blue which absorbs strongly at 600-700 nm and a krypton laser which emits at 647 and 676 nm are exemplified. The power level should be at least 200 mW/cm<2>.

US 5019034 deals with a device for increasing the transport of substances through the skin. Electric pulses with high voltage and short pulse duration are used to pore the skin.

US 4449528 relates to a device that can remove parts of the skin. The device consists of an electrically powered thermal probe that can be heated to melt tissue and then cooled rapidly. The probe is heated in short energy pulses.

US 3207159 includes a device for stimulating the skin by cooling the skin and heating it. Cooling can be carried out using a Peltier device.

US 5342355 A deals with an apparatus for melting away body tissue within a limited area using a heat source in the form of a solid element.

US 4767402 A deals with the use of ultrasound with a frequency of 20kHz to 100 MHz and varying the frequency and intensity over time (modulation) to more effectively transfer molecules through the skin.

It has been shown that by removing the stratum corneum from a small area of the skin by repeatedly applying and removing cellophane tape in the same place, one can easily collect arbitrary amounts of interstitial fluid which can then be examined for a number of interesting analytes. Similarly, "tape-stripped" skin has also been shown to be permeable for transdermal delivery of compounds into the body. Unfortunately, "tape-stripping" leaves an open wound that requires weeks to heal, and for this, as well as other reasons, is not considered acceptable practice for improved transcutaneous transport for wider applications.

As discussed above, it has been shown that pulsed lasers, such as the excimer laser operating at 193 nm, the erbium laser operating near 2.9 µm, or the CO2 laser operating at 10.2 µm, can be used to efficiently melt small holes in the human stratum corneum. These laser ablation techniques allow for a selective and potentially non-traumatic method of opening a delivery and/or sampling hole through the stratum corneum. However, due to the prohibitively high costs associated with these light sources, commercial products based on this concept have not been developed. By specifying a method for directly directing heat energy into the stratum corneum with a very narrowly defined spatial and temporal resolution, the invention described here makes it possible to produce the desired micro-ablation of the stratum corneum by using very reasonable energy sources.

Considering the above problems and/or weaknesses, the development of a method to safely enhance skin permeability for minimally invasive or noninvasive monitoring of analytes in the body within a faster time frame would be a significant advance in the field. Another significant advance in the field would be to provide a method of minimally invasively or noninvasively enhancing the transdermal flow rate of a drug into selected areas of a subject's body.

BRIEF SUMMARY OF THE INVENTION

One purpose of the invention is to reduce the stratum corneum's barrier properties by using perforation, in order to controllably collect analytes from the body through perforations in the stratum corneum to enable monitoring of these analytes.

Poring provides the opportunity to monitor selected analytes in the body through micropores in the stratum corneum in combination with sound energy, permeation amplifiers, pressure gradients and the like.

Poring also provides the opportunity to control transdermal flow rates of drugs or other molecules into the body and possibly into the blood stream, through tiny perforations in the stratum corneum.

Porereing also provides the opportunity to provide a method for delivering drugs into the body through micropores in the stratum corneum in combination with sound energy, permeation amplifiers, pressure gradients and the like.

The invention relates to a system for melting away the stratum corneum layer in the skin of an organism, comprising means (140) for melting away the skin (160) of the organism at a selected area to form at least one micropore, each micropore being 1-1000 µm in diameter and projecting to a desired depth,

characterized in that the means (140) for melting away is a means for heating a selected area of the stratum corneum with a heat source (144) to melt away the stratum corneum, wherein the heat source (144) comprises a fixed element (148) in operative communication with a means (156) for modulating the temperature of said solid element (148) from a reduced temperature state to an elevated temperature state, wherein, in the elevated temperature state, said solid element (148) transfers thermal energy to the skin (160) such that the temperature of the skin surface at the point of contact is greater than 123°C, to melt away a portion of the skin (160), and wherein, in the reduced temperature state, said solid element (148) transfers less thermal energy to the skin (160) than is required to melt away the skin part (160).

BRIEF DESCRIPTION OF THE DIFFERENT SECTIONS OF THE DRAWINGS

FIG. 1 shows a schematic representation of a system for emitting laser diode light and monitoring the poration development. FIG. 2 shows a schematic representation of a closed circulation system for monitoring porosity. FIG. 3A shows a schematic representation of an optical poration system comprising a cooling device. FIG. 3B shows an elevation of a schematic representation of an illustrative cooling device according to FIG. 3A. FIG. 4 shows a schematic representation of an ohmic heating device with a mechanical actuator. FIG. 5 shows a schematic representation of a heating device with a high-resistance current loop. FIG. 6 shows a schematic representation of a device for heating modulation by using inductive heating. FIG. 7 shows a schematic representation of a closed impedance monitor loop that uses impedance changes to determine the degree of porosity. FIG. 8A-D show cross-sections of human skin treated with copper phthalocyanine and then subjected, respectively, to 0, 1, 5 and 50 pulses of 810 nm light with an energy density of 4000 J/cm<2> for a pulse period of 20 ms. FIG. 9-11 show diagrams of the temperature distribution during simulated thermal porations using optical poration. FIG. 12 and 13 show graphical representations of the temperature as a function of time in the stratum corneum and living epidermis, respectively, during simulated heat poration using optical poration. FIG. 14-16 show graphical representations of temperature distribution, temperature as a function of time in the stratum corneum and temperature as a function of time in the living epidermis during simulated heat poration using optical poration when the tissue was cooled before poration. FIG. 17-19 show graphical representations of temperature distribution, temperature as a function of time in the stratum corneum and temperature as a function of time in the living epidermis during simulated heat poration when the tissue was heated with a heating wire. FIG. 20-22 show graphical representations of temperature distribution, temperature as a function of time in the stratum corneum and temperature as a function of time in the living epidermis during simulated heat poration when the tissue was heated with a heating wire and the tissue was cooled before poration. FIG. 23 and 24 show graphical representations of temperature distribution, respectively, temperature as a function of time in the stratum corneum during simulated heat perforations when the tissue is heated optically according to the setting parameters according to Tankovich '803. FIG. 25 shows a graphical representation of interstitial fluid (ISF; O) and blood (<*>) glucose levels as a function of time. FIG. 26 shows a scatter plot of the difference term between the ISF glucose and the blood glucose data of FIG. 25. FIG. 27 shows a histogram of the relative deviation of the ISF in relation to the blood glucose levels from FIG. 25. FIG. 28 shows a cross-section of an illustrative delivery device for delivering a drug to a selected area of a subject's skin. FIG. 29A-C show graphical representations of skin areas affected through the delivery of lidocaine to selected areas, where the stratum corneum is porous (FIG. 29A-B) or non-porous (FIG. 29C). FIG. 30 shows a diagram comparing the amount of interstitial fluid taken from micropores with only suction (O) and with a combination of suction and ultrasound (<*>). FIG. 31, 32 and 33 respectively show a section in perspective of an ultrasound transducer/vacuum apparatus for extracting interstitial fluid, a cross-section of the same apparatus and a schematic cross-section of the same apparatus. FIG. 34A-B show an elevational view of a handheld ultrasonic transducer and an elevational view of the spatula-shaped end thereof, respectively.

DETAILED DESCRIPTION

The scope of the invention appears from the claims.

It should be noted that the singular forms "one", "one" and "it" in this description and the accompanying claims include the plural unless the context clearly indicates otherwise. For example, reference to a method of delivering "a drug" includes delivery of a mixture of two or more drugs, reference to "an analyte" includes one or more such analytes and reference to "a permeation enhancer" includes reference to a mixture of two or more permeation enhancers.

In the present description and requirements, the following terminology will be used if it does not otherwise appear from the requirements according to the definitions given below.

In this context, "poration", "microporation" or any similar term means the formation of a small hole or pore in the stratum corneum in

a selected skin area on an individual, to reduce the barrier properties of this skin layer against the passage of analytes from below the skin surface for analytical purposes or the passage of active permeant or drug into the body for therapeutic purposes. Preferably, the hole or pore will not be larger than approx. 1 mm in diameter, and more preferably, no larger than approx. 100 um in diameter, and protrude sufficiently into the stratum corneum to break the barrier properties of this layer without adversely affecting the underlying tissues.

In this context, "ablation" means the controlled cell removal caused by kinetic energy released when the volatile components of the cell have been heated to the point where vaporization occurs, and the resulting rapid volume expansion as a result of this phase change causes cells and possibly some neighboring cells " blown away” from the ablation site.

In this context, "puncture" or "micropuncture" means the application of mechanical, hydraulic or electrical devices to perforate the stratum corneum.

To the extent that "ablation" and "puncture" achieve the same poration purpose, i.e. creating a hole or a pore in the stratum corneum without significantly damaging the underlying tissues, these terms can be used interchangeably.

In this context, "penetration enhancement" or "permeation enhancement" means an increase in the skin's permeability to a drug, analyte, dye or other molecule (also termed "permeant"), i.e. to increase the rate at which a drug, analyte or molecule permeates the stratum corneum with and facilitates the poration of the stratum corneum, extraction of analyte through the stratum corneum or release of drug through the stratum corneum and into the underlying tissues. The enhanced permeation effected through the use of such enhancers can be observed, for example, by observing the diffusion of a dye, such as a permeant, through animal or human skin, using a diffusion apparatus.

In this context, "chemical enhancer", "penetration enhancer", "permeation enhancer", and the like, include all enhancers that increase the flow of permeant, analyte or other molecule through the skin and are limited only by functionality. In other words, all cell envelope disordering compounds and solvents and any other chemical enhancer are meant to be included.

In this context, "dye", "staining" and the like will be used interchangeably and refer to a biologically appropriate chromophore that exhibits strong absorption in the emission range of a pulsed light source used to melt away tissue in the stratum corneum to form micropores therein .

In this context, "transdermal" or "percutaneous" means the passage of a permeant into and through the skin to achieve effective therapeutic blood or deep tissue levels of a drug, or the passage of a molecule occurring in the body ("analyte") through the skin as that the analyte molecule can be collected on the outside of the body.

In this context, the term "permeant", "drug" or "pharmacologically active agent" or any other similar term means any chemical or biological material or compound suitable for transdermal administration by previously known methods and/or according to the methods that appear in the present invention, which induce a desired biological or pharmacological effect, which may include, but is not limited to (1) having a prophylactic effect on the organism and preventing an unwanted biological effect, e.g. prevention of infection, and (2) alleviation of a condition caused by a disease, such as alleviation of pain or inflammation as a result of disease, and/or (3) either alleviation, reduction or complete elimination of the disease from the organism. The effect can be local as in producing a local anesthetic effect, or it can be systemic. This invention is not directed to new permeants or new classes of active substances. On the other hand, it is limited to the method of delivery of agents or permeants that exist in the area or that can be established as active substances later and which are suitable for delivery according to the present invention. Such substances include large classes of compounds that are normally delivered to the body, including through body surfaces and membranes, including skin. In general, this includes: anti-infectives such as antibiotics and antivirals; analgesics and analgesic combinations; anorexics; anthelmintics; anti-arthritis agents; antiasthmatics; anticonvulsants: antidepressants; antidiabetic agents; antidiarrheal; antihistamines; anti-inflammatory agents; anti-migraine remedies; anti-nausea agents; antineoplastic agents; antiparkinsonian agents; anti-itching agents; antipsychotics; antipyretics; antispasmodics; anticholinergics; sympathomimetics; xanthine derivatives; cardiovascular preparations, including potassium and calcium channel blockers, beta blockers, alpha blockers and antiarrhythmics; antihypertensive agents; diuretics and antidiuretics; vasodilators including general coronary, peripheral and cerebral vasodilators; central nervous system stimulants; vasoconstrictors; cough and cold preparations, including mucosal decongestants; hormones such as estradiol and other steroids, including corticosteroids; hypnotics; immunosuppressants; muscle relaxants; parasympatholytics; psychostimulants; sedatives and tranquillizers. According to the method according to the present invention, both ionized and non-ionized drugs as well as drugs of both high and low molecular weight can be delivered.

In this context, an "effective" amount of a pharmaceutically active agent means a sufficient amount of a compound to produce the desired local or systemic effect and behavior with a reasonable benefit/risk ratio accompanying any medical treatment. An "effective" amount of a permeation or chemical enhancer means in this context an amount that is chosen so that it gives the desired increase in skin permeability and the desired penetration depth, administration rate and amount of drug delivered.

In this context, "carriers" refer to carrier materials without significant pharmacological action in the amounts used, which are suitable for administration with other pharmaceutical agents, and include any such materials known in the art, e.g. any liquid, gel, solvent, liquid diluent or solubilizer which is non-toxic in the amounts used and does not produce any adverse interaction with the drug to be administered. Examples of suitable carriers for use in this context are water, mineral oil, silicone, inorganic gels, aqueous emulsions, liquid sugars, waxes, petroleum jelly and a number of other oils and polymeric materials.

In this context, a "biological membrane" is meant to stand for a membrane material occurring inside a living organism, which separates one area of the organism from another and, in many cases, which separates the organism from its external environment. Skin and mucous membranes are thus included.

In this context, "individual" refers to both a human and an animal for which the present invention is applied.

In this context, "analyte" means any chemical or biological material or compound, which can pass through a biological membrane by the technology that appears in the present invention or by previously known technology, of which one wishes to know the concentration or activity of the body. Glucose is a specific example of an analyte in that it is a sugar suitable for passage through the skin, and individuals such as diabetics may wish to know their blood glucose levels. Other examples of analytes include compounds such as sodium, potassium, bilirubin, urea, ammonia, calcium, lead, iron, lithium, and salicylates.

In this context, "transdermal flow rate" means the rate of passage of any analyte through the skin of an individual, human or animal, or the rate of passage of any drug, pharmacological agent, dye or pigment into and through the skin of an individual, human or animals.

In this context, the terms "intensity amplitude", "intensity" and "amplitude" are used synonymously and refer to the amount of energy produced by the sound energy system.

In this context, "frequency modulation" or "sweep" means a continuous, graded or stepwise variation in the amplitude or frequency of ultrasound in a given period of time. A frequency modulation is a graded or stepwise varied frequency in a given period of time, for example 5.4-5.76 MHz for 1 sec. or 5-10 MHz for 0.1 sec. or 10-5 MHz for 0.1 sec., or any other frequency range or time span relevant to a particular application. A complex modulation can include simultaneous variation of both frequency and intensity. For example, FIG. 4A and 4B of US Patent 5,458,140 respectively represent amplitude and frequency modulation imposed simultaneously by a single sound energy transducer.

In this context, "phase modulation" means that the timing of the signal has been changed relative to its original state shown in FIG. 4C of US Patent 5,458,140. The frequency and amplitude of the signal can be the same. A phase modulation can be realized with a variable delay so that it selectively retards or accelerates the signal temporally in relation to its previous state or to another signal.

The sound energy in its various applications, such as frequency, intensity or phase modulation, or combinations thereof, and the use of chemical amplifiers, combined with modulated sound energy, as described here, can vary over a frequency range of between approx. 5 kHz to 100 MHz, of which the range between approx. 20 kHz and 30 MHz are preferred.

In this context, "non-invasive" means that it does not require the introduction of a needle, a catheter or other invasive instruments into a body part.

In this context, "minimally invasive" refers to the use of mechanical, hydraulic or electrical devices that penetrate the stratum corneum to form a small hole or micropore without causing significant damage to the underlying tissues.

Devices for porating the stratum corneum

The formation of a micropore in the stratum corneum can be achieved by means of various known devices, as well as certain devices described herein which constitute improvements thereof.

The use of laser ablation as described by Jacques et al. in US Patent 4,775,361 and by Lane et al, supra, clearly provides a device for ablation of the stratum corneum using an excimer laser. At a wavelength of 193 nm and 14 ns pulse width, it turned out that approx. 0.24 to 2.8 µm of the stratum corneum could be removed with each laser pulse at a radiation exposure of between approx. 70 and 480 mJ/cm<2>. As the pulse energy increases, more tissue is removed from the stratum corneum, and fewer pulses are required to completely perforate this layer. The lower threshold of radiation exposure that must be absorbed by the stratum corneum within the limitation of heat relaxation time to cause appropriate microbursts resulting in tissue ablation is approx. 70 mJ/cm<2>in 50 milliseconds (ms). In other words, a total of 70 mJ/cm<2> must be emitted within a 50 ms window. This can be done with a single pulse of 70 mJ/cm<2> or in 10 pulses of 7 mJ/cm<2> or by continuous irradiation with 1.4 W/cm<2> during 50 ms. The upper limit for radiation exposure is that which will cause ablation of the stratum corneum without damage to underlying tissue, and it is determined empirically based on light source, light wavelength and other variables of which the skilled person will have experience and knowledge.

By "release" is meant that the specified amount of energy is absorbed by the tissue to be removed. At the excimer laser wavelength of 193 nm, approximately 100% absorption occurs within the first 1 or 2 µm of stratum corneum tissue. Assuming that the stratum corneum is approx. 20 um thick, it is absorbed at longer wavelengths, such as e.g. 670 nm, only approx. 5% of incident light in the 20 u,m layer. This means that approx. 95% of the high-energy beam enters the tissue below the stratum corneum, where it is likely to cause significant damage.

The ideal is to use only as much energy as is needed to perforate the stratum corneum without causing bleeding, thermal or other damage to underlying tissue, from which analytes are to be extracted, or drugs or other permeants are delivered to.

It would be beneficial to use more economical energy sources than energy from excimer lasers. Excimer lasers that emit light at wavelengths in the far UV range are significantly more expensive to operate and maintain than, for example, diode lasers that emit light at wavelengths in the visible range and in the IR range (600 to 1800 nm). At longer wavelengths, however, the stratum corneum becomes increasingly transparent, and absorption occurs primarily in the underlying tissues.

The present invention constitutes a quick and painless method for eliminating the barrier function of the stratum corneum, so that it facilitates the transcutaneous transport of therapeutic substances into the body by local application, or gives access to the analysis of analytes in the body. The method makes use of a procedure that begins by applying a small area of heat contact to the intended area of the stratum corneum.

The heat source must have several important properties which will now be described. Firstly, the heat source must be of such a size that contact with the skin is limited to a small area, generally with a diameter of approx. 1 to 1000 µm. Second, it must be able to modulate the temperature of the stratum corneum at the point of contact from ambient skin surface temperature (33°C) to more than 123°C and then return to approx. ambient skin temperature with cycling times that minimize damage to surrounding living tissue and sensation in the individual concerned. This modulation can be produced electronically, mechanically or chemically.

In addition, an inherent depth-limiting feature of the microporation process can be simplified if the heat source has both sufficiently small thermal mass and limited energy source that its temperature is raised so that when it is placed in contact with tissue with more than 30% water content, the heat dispersion in these is sufficient to limit the maximum temperature of the heat source to less than 100°C. This feature effectively stops the heat evaporation process as soon as the heating probe has penetrated the stratum corneum and into the lower layers of the epidermis.

With the heat source placed in contact with the skin, it is cycled through a series of one or more temperature modulations from a starting point at ambient skin temperature to a maximum temperature in excess of 123°C to approximately ambient skin temperature. In order to reduce or eliminate the individual's sensation of the microporation process, the duration of the pulses is limited, and the distance between the pulses is made long enough so that the living tissues in the skin, and especially the weakened skin tissues, can be cooled to an average temperature of less than approx. 45°C. These parameters are based on the heating time constants for the living epidermal tissue (approx. 30-80 ms) which is located between the heating probe and the weakened tissue in the underlying dermis. The result of this application of pulsed heat energy is that sufficient energy is directed into the stratum corneum within the small target spot, so that the local temperature of this tissue volume is raised sufficiently to the super-evaporation point for the tissue-bound water content in the stratum corneum. As the temperature rises above 100°C, the water content of the stratum corneum (typically 5% to 15%) in this local area is induced to vaporize and expand very rapidly to cause a steam-triggered removal of the stratum corneum corneocytes located near this evaporation. US patent 4,775,361 states that a stratum corneum temperature of 123°C represents a threshold at which this type of flash evaporation occurs. As successive pulses of heat energy are applied, additional layers of the stratum corneum are removed until a micropore is formed through the stratum corneum down into the next layer of the epidermis, the stratum lucidum. By limiting the duration of the heat pulse to less than 1 heating time constant of the epidermis and allowing any heat energy directed into the epidermis to escape for a sufficiently long time, the temperature increase in the living epidermal layers is minimal. Through this, the entire microporation process can take place without the person feeling anything and without damage to the underlying and surrounding tissue.

The present invention comprises a method for the painless formation of microscopic holes, i.e. micropores, with a diameter from approx. 1 to 1000 um in the stratum corneum of human skin. The key to successful completion of the procedure is the creation of a suitable heat energy source or heat probe, which is kept in contact with the stratum corneum. The most significant technical challenge in manufacturing a suitable heat probe is the development of a device that has the desired contact with the skin and that can be thermally modulated with a sufficiently high frequency.

It is possible to make a suitable heating probe by locally applying to the stratum corneum a suitable light-absorbing compound, such as a dye, chosen for its ability to absorb light at the wavelength emitted by a chosen light source. In this case, the selected light source can be a laser diode that emits at a wavelength that is not normally absorbed by the skin tissue. By focusing the light source to a small spot on the surface of the local dye layer, the temperature of the colored area can be modulated by varying the intensity of the luminous flux that is focused on it. It is possible to utilize the energy from laser sources that emit at longer wavelengths than an excimer laser, by first locally applying a suitable light-absorbing compound to the stratum corneum, such as e.g. a dye, chosen for its ability to absorb light at the wavelength emitted by the laser source. The same concept can apply at any wavelength and you just have to choose a suitable dye and the right optical wavelength. One need only look in any suitable reference work to find suitable dyes and wavelengths for the dye's maximum absorbance. One such reference work is Greene, The Sigma-Aldrich Handbook of Stains, Dyes and Indicators, Aldrich Chemical Company, Inc., Milwaukee, Wisconsin (1991). For example, copper phthalocyanine (Pigment Blue 15; CPC) absorbs at approx. 800 nm; copper phthalocyanine tetrasulfonic acid (Acid Blue 249) absorbs at approx. 610nm; indocyanine green absorbs at approx. 775 nm and cryptocyanine absorbs at approx. 703 nm. CPC is particularly suitable for this embodiment for the following reasons: it is a very stable and inert compound, already approved by the FDA for use as a dye in implantable sutures; it absorbs very strongly at wavelengths from 750 nm to 950 nm which goes well with a number of cheap semiconductor emitters, such as laser diodes and LEDs, while this area of the optical bandwidth is not absorbed directly by the skin tissue to a significant extent. CPC has a very high evaporation point (>550°C in vacuum) and passes directly from solid phase to vapor phase without any liquid phase. The CPC has a relatively low heat diffusivity constant which allows the light energy focused on it to selectively heat only the area that is directly in focus, with very little lateral spread of the "heat spot" into the surrounding CPC, thereby contributing to the spatial delimitation of the contact between heat and probe.

The purpose of this review is not to provide a complete list of suitable dyes as such can be easily obtained by the person skilled in the art from readily available data.

The same applies to any desired pulsed light source. For example, the method can be implemented with a mechanically closed, focused incandescent lamp as the pulsating light source. Various catalogs and brochures list a wide range of lasers operating in the near-UV, visible and near-IR range. Representative lasers are Hammamatsu Photonic Systems Model PLP-02 operating at an output power of 2 x 10"<8>J, Hammamatsu Photonic Systems Model PLP-05 operating at an output power of 15 J at a wavelength of 685 nm; SDL, Inc. SDL-3250 series pulsed laser operating at an output power of 2 x 10<6>J at a wavelength of approximately 800-810 nm; SDL, Inc. Model SDL-8630 operating at an output power of 500 mW at a wavelength of approximately 670 nm; Uniphase Laser Model AR-081-15000 operating with an output power of 15,000 mW at a wavelength of 790-830 nm; Toshiba America Electronic Model TOLD9150 operating with an output power of 30 mW at a wavelength of 690 nm; and LiCONIX, Model Diolite 800-50 which operates with a power of 50 mW at a wavelength of 780 nm.

For the present invention, a pulsed laser light source can emit radiation over a wide wavelength range varying from between approx. 100 nm to 12,000 nm. Excimer lasers will generally emit over a range of between 100 and 400 nm. Commercial excimer lasers with wavelengths in the range from approx. 193 nm to 350 nm are commercially available. Preferably, a laser diode will have an emission range of between approx. 380 and 1550 nm. A frequency-doubled laser diode will have an emission range of between approx. 190 and 775 nm. Longer wavelengths varying from between approx. 1300 and 3000 nm can be used by making use of a laser diode-pumped optical parametric oscillator. It must be expected that with the research taking place in the field of laser technology that these areas will eventually expand.

Energy emitted or absorbed need not be obtained from a laser, as any light source, whether from a laser, a short arc lamp such as a xenon flash lamp, an incandescent lamp, a light emitting diode (LED), the sun or a other source, can be used. The particular instrument used to emit electromagnetic radiation is less important than the wavelength and energy associated with it. Any suitable instrument that can deliver the necessary energy at suitable wavelengths, i.e. in the range from approx. 100 nm to approx. 12,000 nm, can be considered to lie within the scope of the invention. The essential feature is that the energy must be absorbed by the light-absorbing compound in order to cause local heating thereof with subsequent conduction of sufficient heat to the tissue to be melted away within the permitted time frame.

In an illustrative embodiment, the heat probe itself is formed from a thin layer, preferably approx. 5 to 1000 µm thick, of a solid non-biologically active compound that is applied topically to a selected area of a subject's skin, which is large enough to cover the site where a micropore is to be created. The particular formulation of the chemical compound is chosen so that it exhibits high absorption over the spectral range of the light source selected to energize the light absorbing compound. The probe can be, for example, a sheet of a solid compound, a film treated with a high-melting absorbent compound or a direct application of the light-absorbing compound to the skin as a deposit or as a suspension in a carrier. Regardless of the configuration of the light-absorbing heat probe, it must exhibit a sufficiently low lateral heat diffusion coefficient so that any local rise in temperature will remain spatially confined, and the dominant form of heat loss will be via direct conduction into the stratum corneum through the point of contact between the skin and the probe.

The necessary temperature modulation of the probe can be achieved by focusing a light source on the light-absorbing compound and modulating the intensity of this light source. If the absorbed energy within the illuminated area is sufficiently high, it will cause the light-absorbing compound to quickly heat up. The amount of emitted energy and subsequently both the heating rate and the maximum temperature at the focal point of the light-absorbing compound can be easily modulated by varying the pulse width and the maximum power of the light source. In this embodiment, it is only the small volume of light-absorbing compound heated by the focused, incident optical energy that forms the heating probe, any additional light-absorbing compound that may have been applied over a larger area than the relevant poration site being random. By using a light-absorbing solid compound with a relatively high melting point, such as copper phthalocyanine (CPC), which remains in the solid state up to a temperature of more than 550°C, the heat probe can be quickly brought up to a temperature of several hundred degrees C and still remain in contact with the skin so that the heat energy can be directed into the stratum corneum. This embodiment also includes choosing a light source with an emission spectrum where very little energy would normally be absorbed in the skin tissue.

After the light-absorbing compound has been locally applied to the intended area, the heat probe is formed when the light source is activated so that the focal narrowing of the beam is placed so that it coincides with the surface of the treated area. The energy density of the light at the focal constriction and the degree of absorption taking place in the light-absorbing compound are adjusted to be sufficient to bring the temperature of the light-absorbing compound within the area of the small spot defined by the focal point of the light source to more than 123°C within a few milliseconds. As the temperature in the heating probe increases, the wire into the stratum corneum emits energy to these tissues, whereby the local temperature of the stratum corneum is raised. When sufficient energy has been delivered to this small area of the stratum corneum to bring the local temperature above the boiling point of the water that these tissues contain, a lightning-quick evaporation of this water occurs, whereby the stratum corneum at this point is removed.

By switching the light source on and off, the temperature of the heat probe can be rapidly modulated and the selective ablation of these tissues achieved so that a very precisely sized hole can be created that selectively only penetrates through the first 10 to 30 µm of the skin.

A further feature of this embodiment is that by choosing a light source which would normally give the skin or underlying tissue very little absorbed energy, and by designing focusing and emission optics with a sufficiently high numerical aperture, the small amount of emitted light which is not accidentally absorbed in the heat probe itself, quickly spread as it penetrates deeper into the body. Since there is very little absorption at the emitted wavelengths, the light source emits practically no energy directly to the skin. This three-dimensional dilution of coupled energy in the tissue due to radiation scattering and the low level of absorption in the untreated tissue results in a completely benign interaction between the light beam and the tissue without causing damage.

In a preferred embodiment of the invention, a laser diode with an emission wavelength of 800+30 nm is used as light source. A heat probe can be formed by local application of a transparent adhesive tape, which is treated on the adhesive side with a 0.5 cm spot formed by a deposit of finely ground copper phthalocyanine (CPC). CPC exhibits extremely high absorption coefficients in the 800 nm spectral range, and generally absorbs more than 95% of the beam energy from a laser diode.

FIG. 1 shows a system 10 for emitting light from such a laser diode to a selected skin area of an individual and for controlling the development of the poring process. The system comprises a laser diode 14, connected to a controller 18, which controls the intensity, duration and distance between the light pulses. The laser diode emits a beam 22, which is directed towards one or more collecting lenses 26, which focus the beam onto a mirror 30. The beam is then reflected by the mirror towards one or more objective lenses 34, which focus the beam towards a pre-selected point 38. This in advance selected point corresponds to the plane of an xyz podium 42, and the objective hole 46, therein, so that a preselected area of the skin of an individual can be irradiated. This xyz podium is connected to the controller so that the position of the xzy podium can be controlled. The system also comprises a control system comprising a CCD camera 50 connected to a monitor 54. The CCD camera is confocal in line with the objective lens so that the development of the poration process can be controlled visually on the monitor.

In another illustrative embodiment of the invention, a system for sensing photodiodes and collecting optics is provided which is confocal in line with the ablation light source. FIG. 2 shows a sensor system 60 for use in this embodiment. The system comprises a light source 64, for emitting a light beam 68, which is controlled by an optical delivery system 72, which focuses the beam towards a preselected point 76, such as the skin surface 80, of an individual. Part of the light that hits the skin is reflected and other light is emitted from the irradiated area. A portion of the reflected and emitted light passes through a filter 84, and then through a collecting optics system 88, which focuses the light towards a photodiode 92. A controller 96 is connected to both the laser diode and the photodiode to respectively control the laser diode's output and detect the light that reaches the photodiode. Only selected parts of the spectrum emitted from the skin pass through the filter. By analyzing the displacements in the reflected and emitted light from the intended area, the system is able to detect when the stratum corneum has been breached, and this feedback is then used to control the light source, which deactivates the light pulses when the microporation of the stratum corneum is achieved. By using this type of active system for closed feedback, a self-regulating universally applicable device is achieved that produces uniformly sized micropores in the stratum corneum, with minimal power requirements despite variations from one individual to another.

In another illustrative embodiment, a cooling device is incorporated into the system's contact with the skin. FIG. 3A shows an illustrative schematic representation of this. In this system 100, a light source 104 (connected to a controller 106) emits a light beam 108, which passes through, and is focused by, an optical delivery system 112. The beam is focused by the optical delivery system to a preselected point 116, such as a selected skin area 120, on an individual. A cooling device 124, such as a peltier device or other cooling devices, is in contact with the skin to cool the surface thereof. In a preferred embodiment of the cooling device 124

(FIG. 3B) there is a central hole 128 through which the focused light beam passes to contact the skin. Of FIG. 3A it appears that a cooling element 132 is also preferably placed in contact with the cooling device. By using a cooling device where a small hole in its center coincides with the focal point of the light, the skin tissue in the general area where the pore is to be produced can be pre-cooled to 5°C to 10°C. This prior cooling provides a greater margin of safety for the use of the system in that what the user may feel, and the possibility of possible lateral damage to the epidermis immediately below the poration site, is significantly reduced compared to embodiments without cooling. For monitoring applications, prior cooling also reduces evaporation of interstitial fluid and may also result in advantageous physical properties, such as reduced surface tension of such interstitial fluid. In addition, it is known that cooling the tissue causes a localized increase in blood flow in the cooled tissue which thereby promotes diffusion of analytes from the blood into the interstitial fluid.

The method can also be used for other microsurgical techniques, where the light-absorbing compound/heat probe is applied to the area to be removed, after which the light source is used to selectively modulate the temperature of the probe at the selected target site under the influence of the tissue via the produced evaporation-ablation process.

A further feature of the invention is to use the light source to help seal the micropore after its usefulness is over. When it comes to monitoring an internal analyte, a micropore is created whereupon a certain amount of interstitial fluid is extracted through this opening. After a sufficient amount of interstitial fluid has been collected, the light source is reactivated with a reduced power to facilitate rapid deposition or coagulation of the interstitial fluid in the micropore. By forcing coagulation or sedimentation of the fluid in the pore, this opening is effectively sealed so that the risk of infection is reduced. The use of the light source both for the formation of the micropore and the sealing of this is also in itself a sterile process without any physical penetration of any device or apparatus into the body. The heat shock induced by the light energy also kills any microbes that may occur at the ablation site.

This optical sterilization can be extended to include a further step in the process, where the light source is first applied without focusing, thereby covering the target area with an illuminated area projecting 100 µm or more beyond the actual site of micropore generation. By selecting the area over which the unfocused beam is applied, the flux density can be correspondingly reduced to a level well below the ablation threshold, but high enough to effectively sterilize the skin surface. After sufficiently long exposure of the larger area, either in a continuous step or during a series of pulses, to the sterilization beam, the system is then configured to the sharply focused ablation mode, whereby the optical microporation process begins.

Another illustrative embodiment of the invention consists in producing the necessary heat probe from a metallic solid, such as e.g. a wire of small diameter. As in the above-described embodiment, the contacting surface of the heating probe must be able to have the temperature modulated from ambient skin temperature (33°C) to temperatures higher than 123°C within the necessary allowable time of preferably between approx. 1 and 50 ms at the high temperature (on time) and at least approx. 10 to 50 ms at the low temperature (off time). In particular by being able to modulate the temperature up to more than 150°C in an "on" time of approx. 5 ms and an off time of 50 ms, a very effective heat ablation is achieved with little or no sensation of pain for the person concerned.

Several methods for temperature modulation of the wire-shaped heating probe's contact area can be successfully used. For example, a short length of wire can be brought up to the desired high temperature through an external heating element, such as an ohmic heating element used in the tip of a soldering iron. FIG. 4 shows an ohmic heating device 140, with a mechanical actuator. The ohmic heating device consists of an ohmic heat source 144, connected to a wire-shaped heating probe 148. The ohmic heat source is also connected through an insulating plate 152, to a mechanical modulation device 156, which e.g. a coil. In this configuration, an equilibrium state can be reached where the tip of the wire probe will stabilize at a certain equilibrium temperature determined by the physical parameters of the structure, i.e. the temperature of the ohmic heat source, the length and diameter of the wire, the temperature of the air surrounding the wire and the material that the wire consists of. When the desired temperature is achieved, the modulation of the temperature of the selected area of the skin 160, of an individual, is effected directly via the mechanical modulating device to alternatively place the hot wire tip in contact with the skin for, preferably 5 ms on-time and then pull it back into the air for preferably a 50ms off time.

Another illustrative example (FIG. 5) shows a device 170, which includes a current source 174, connected to a controller 178. The current source is connected to a current loop 182, which includes a wire 186, formed into a structure so that it produces a point with high resistance. Preferably, the wire is held in a stand 190, and an insulator 194 separates different parts of the current loop. The desired temperature modulation is then achieved by only modulating the current through the wire. If the thermal mass of the wire element is of the right size and the heat sink produced by the electrodes connecting it to the current source is sufficient, the heating and cooling times for the wire element can be achieved in a few milliseconds. By bringing the wire into contact with the selected skin area 198, the stratum corneum is heated to achieve the intended ablation.

In FIG. 6, another illustrative example of poration of the stratum corneum with a heating wire is shown. In this system 200, the wire 204 can be placed in a modulated varying magnetic field formed by a coil of wire 208, the field coil. By energizing the alternating current in the field coil by means of a controller 212, connected to this, eddy currents of sufficient intensity can be induced in the wire-shaped heating probe so that it is heated directly via the internal ohmic losses. This is essentially a miniature version of an inductive heating system which is commonly used for heat treatment of tool ends or inducing outgassing from the electrodes in vacuum tubes or electron flash. The advantage of the inductive heating method is that the energy emitted to the wire-shaped heating probe can be precisely controlled and easily modulated via the field coil's electronic control. If the thermal mass of the wire probe as such, and the thermal mass of the stratum corneum in contact with the tip of the probe, is known, control of the emitted inductive energy can produce very precise control of the temperature at the point of contact 216, with the skin 220. Since the skin tissue is practically non-magnetic at the lower frequencies at which inductive heating can be achieved, if correctly selected frequencies are used in the field coil, this varying electromagnetic field will have no effect on the skin tissue.

If a mechanically controlled contact modulation is used, a further feature can be realized by incorporating a control system with a single closed loop where the electrical impedance between the probe tip and the individual's skin is controlled. In this way, the probe position can be brought into contact with the individual's skin, which manifests itself in the step-by-step resistance reduction as soon as the contact is closed, and is held there for the desired "on time", after which it can be withdrawn. Several types of linear actuators are suitable for this form of closed-loop control, such as a voice coil mechanism, a simple coil, a rotary system with a cam or angle arm, and the like. The advantage is that as the heat ablation develops, the position of the heat probe tip can also be advanced into the skin, which always ensures good contact to facilitate efficient transfer of the necessary heat energy. Furthermore, the change in the conduction properties of the stratum corneum and epidermis can be used to provide an elegant closed-loop confirmation that the poration process is complete, i.e. when the resistance indicates that the epidermis has been reached, it is time to stop the poration process.

FIG. 7 shows an illustrative example of such a closed loop impedance monitor. In this system 230, an ohmic heat source 234 is connected to a wire-shaped heat probe 238. The heat source is mounted through an insulation holder 242, on a mechanical modulator 246. A controller 250 is connected to the wire and to the skin 254, where the controller detects impedance- changes in the selected area 258, of the skin, and where the controller stops the pore process when a predetermined level is reached.

In line with the hydraulic poring devices, microlancets are intended to precisely penetrate the stratum corneum for the purpose of administering a permeant, such as a drug, through the formed pore or through the pore and extracting an analyte for analysis. Such a device is considered to be "minimally invasive" compared to devices and/or techniques that are non-invasive. The use of microlancets that penetrate below the stratum corneum for the extraction of blood is well known. Such devices are commercially available from manufacturers such as Beckton-Dickinson and Lifescan and can be used for the present invention by controlling the penetration depth. As an example of a microlancet device for collecting body fluids, see Erickson et al. Internationally Published PCT Application WO 95/10223 (published April 20, 1995). This application shows a device for penetrating the dermis of the skin without penetrating subcutaneous tissue, to collect body fluids for monitoring purposes, such as blood glucose levels.

Poration of the stratum corneum can also be achieved by using sound devices. Sound poration is a variant of the optical devices described above, except that instead of using a light source, a very tightly focused beam of sound energy is emitted to the area of the stratum corneum to be melted away. The same energy levels are required, ie a threshold of 70 mJ/cm<2>/50 ms must still be absorbed. The same pulsed focused ultrasound transducers as described in our parent applications series no. 08/152.442 and 08/152.174 can be used to deliver the necessary energy densities for ablation, such as those used when delivering sound energy that is modulated in intensity, phase or frequency or a combination of these parameters for the transdermal sampling of an analyte or the transdermal delivery of drugs. This has the advantage of allowing the use of the same transducer to push a drug through the stratum corneum or draw a body fluid to the surface for analysis, as that used to first form a micropore.

Also, electroporation or short bursts or pulses of electrical current can be delivered to the stratum corneum with sufficient energy to form micropores. Electroporation is known in the art to produce pores in biological membranes, and electroporation instruments are commercially available. The expert can thus, without extensive experimentation, choose an instrument and conditions for its use, based on the guidelines set out here.

The micropores produced in the stratum corneum by the methods of the present invention allow the transdermal delivery of high flow rates of high molecular weight therapeutic compounds. In addition, these non-traumatic microscopic openings in the body allow access to various analytes in the body, which can then be analyzed to determine their concentrations.

Example 1

In this example, skin samples were prepared as follows. Epidermal membranes were separated from whole skin from human cadavers by the heat-separation method of Klingman & Christopher, 88 Arch. Dermatol. 702 (1963) which involves exposure of the skin in full thickness to a temperature of 60°C for 60 seconds, after which the stratum corneum and part of the epidermis (epidermal membrane) were carefully peeled from the dermis.

Example 2

Heat-separated stratum corneum samples prepared according to the procedure in Example 1 were cut into 1 cm sections. These small samples were then attached to a glass cover slip by placing them on the glass and placing a pressure sensitive adhesive coated disc with a 6 mm hole in the center over the skin sample. The skin samples were then ready for experimental examination. In some cases, the skin samples were hydrated by soaking them for several hours in a neutral buffered phosphate solution or in pure water.

As a test on these untreated skin samples, the output from several different infrared laser diodes emitting at approx. 810, 905, 1480 and 1550 nm, applied to the sample. The output optics were designed to produce a 25 µm diameter focal constriction with a final objective with a numerical aperture of 0.4. The total power delivered to the focal point was measured to be between 50 and 200mW for the 810 and 1480 nm laser diodes, which could be operated in CW (continuous wave) mode. The 905 and 1550 nm laser diodes were designed to produce high maximum power pulses with a length of approx. 10 to 200 nanoseconds and repetition rates up to 500 Hz. For the pulsed lasers, the maximum power levels were measured to be 45W at 905 nm and 3.5W at 1550 nm.

Under these operating conditions, there was apparently no effect on the skin samples from either laser. The intended area was continuously illuminated for 60 seconds and then examined microscopically, which revealed no visible effects. In addition, the samples were placed in a modified Franz cell which is commonly used to test transdermal delivery systems based on chemical permeation enhancers, and the conductivity from one side of the membrane to the other was measured both before and after the irradiation with the laser and showed no change. On the basis of these tests, which were carried out on skin samples from four different donors, it was concluded that at these wavelengths the coupling of the optical energy into the skin tissue was so small that no effects could be detected.

Example 3

In order to assess the possible sensation of a living individual when illuminated with optical energy under the conditions of Example 2, six subjects were used and the radiation of each laser source directed at their fingertips, forearms and back of the hand. When using 810, 905 and 1550 nm lasers, subjects could not feel when the laser was switched on or off. For the 1480 nm laser, there was some sensation during illumination with the 1480 nm laser operated at 70 mW CW, and shortly thereafter a tiny blister formed under the skin as a result of the absorption of the 1480 nm irradiation by one of the water absorption bands. The amount of energy absorbed was apparently sufficient to induce the formation of the blister, but not to cause melting of the stratum corneum. Moreover, the absorption of the 1480 nm light occurred mainly in the deeper fully hydrated (85% to 90% water content) tissues of the epidermis and dermis, not in the relatively dry (10% to 15% water content) tissues of the stratum corneum.

Example 4

Having demonstrated the lack of effect on the skin in its natural state (Example 3), a number of chemical compounds were evaluated for effectiveness in absorbing the light energy and then transmitting the absorbed energy through conduction into the target tissue of the stratum corneum. Compounds investigated included marker; "SHARPIE" branded fade-resistant black, blue and red markers; methylene blue; magenta; epolite #67, an absorbent compound developed for embedding in polycarbonate lenses for laser goggles; tincture of iodine; iodine-polyvinylpyrrolidone complex ("BETADINE"); copper phthalocyanine; and printing ink.

Using both of the CW laser diodes described in Example 2, positive ablation results were determined on in vitro samples of heat-separated stratum corneum prepared according to Example 1, using all of these products, although some worked better than others. In particular, copper phthalocyanine (CPC) and Epolite #67 were among the most effective. A possible reason for the better effect of CPC is its high boiling point of more than 500°C and the fact that it retains its solid form up to this temperature.

Example 5

Since copper phthalocyanine is already approved by the FDA for use in implantable sutures and is also listed in the Merck index as a relatively benign and stable molecule in terms of human biocompatibility, the next step was to combine the local application of CPC and the focused light source on the skin of healthy subjects. A suspension of finely ground CPC in isopropyl alcohol was prepared. The application method consisted of shaking the solution and then depositing a small drop on the target site. As the alcohol evaporated, a fine and even coating of the solid CPC was left on the skin surface.

The apparatus shown in FIG. 1 was then placed against the site where the CPC locally covered the skin, by placing the selected area of the subject's skin against a reference plate. The reference plate consisted of a thin glass window of approx. 3 cm x 3 cm, with a 4 mm hole in the centre. The CPC-covered area was then positioned so that it was in the central hole. A confocal video microscope (FIG. 1) was then used to bring the surface of the skin into sharp focus. Positioning the skin to achieve the sharpest possible focus on the video system also positioned it so that the focal point of the laser system coincided with the skin surface. The operator then fired pulses of laser light while the effects on the target site were viewed on the video monitor. The degree of penetration was estimated visually by the operator by measuring the degree of defocusing of the laser spot in the micropore as the depth of the micropore increased, and this can be dynamically corrected by the operator essentially by following the melted surface down into the tissues by moving the position of the camera/laser the source along the "Z" axis into the skin. At the point where the stratum corneum had been removed down to the epidermis, the appearance of the bottom of the hole changed noticeably by becoming more moist and shiny. After registering this change, the operator deactivated the laser. In many cases depending on the individual's state of hydration as well as other physiological conditions, a remarkable outflow of interstitial fluid occurred in response to the removal of the barrier function of the stratum corneum in this small area. The video system was used to record this visual recording of the access to interstitial fluid in the puncture site.

Example 6

The procedure of Example 5 was followed except that the CPC was applied to a transparent adhesive tape which was then adhered to a selected location on the skin of a subject. The results were essentially the same as in Example 5.

Example 7

Histological experiments were performed on cadaver skin according to well-known methods to determine ablation threshold parameters for given dye mixtures and information on collateral damage. The upper skin surface was treated with a solution of copper phthalocyanine (CPC) in alcohol. After the alcohol had evaporated, a layer of solid CPC with an average thickness of 10 to 20 nm was locally distributed over the skin surface. FIG. 8A shows a full-thickness cross-section of skin prior to laser application, showing the CPC layer 270, stratum corneum 274, and underlying epidermal layer 278. FIG. 8B shows the sample after a single pulse of 810 nm light directed at an 80 µm diameter circle was applied at an energy density of 4000 J/cm<2> for a pulse duration of 20 ms. It should be noted that there is still a significant amount of CPC present on the surface of the stratum corneum even in the center of the crater, 282. It should also be noted that laboratory measurements indicated that only 10% of the light energy incident on the CPC was actually absorbed, while the the other 90% was reflected or scattered. The effective energy flow delivered to the dye layer, which could cause the desired heating, was thus only approx. 400 J/cm<2>. 8C shows the sample after five pulses of 810 nm light were applied, where the stratum corneum barrier was removed without damage to underlying tissue. These results give a good representation of the "ideal" optically modulated heat ablation. FIG. 8D shows the sample after fifty pulses were applied. Damaged tissue 286, occurred in the epidermal layers as a result of charring of tissue that had not melted away and heat denaturation of underlying tissue. FIG. 8A-8C show separations between the stratum corneum and the underlying epidermal layers as a result of an artificial pf dehydration, freezing and preparations for imaging.

Example 8

In order to investigate details of the heat ablation mechanism, a mathematical model of the skin tissue was set up on which the different embodiments of the heat ablation method could be tried. This model calculates the temperature distribution in a layered semi-"infinite" medium with a specified heat flow supply locally on the surface and removal of heat from the surface some distance beyond, i.e. convection is formed between the two. The axisymmetric time-dependent diffusion equation is solved in cylindrical coordinates using the ADI (alternating-direction-implicit) method (note: constant temp. B.C. is applied to the lower limit to serve as z->inf; and zero radial heat flux is applied to the maximum radial limit for to serve as r->inf). The layers are parallel to the surface and are defined as: (1) dye; (2) stratum corneum; (3) underlying epidermis and (4) dermis. The depth into the semi-infinite medium and thermal properties, density (rho), specific heat (c) and conductivity (k) must be specified for each layer.

A heat transfer coefficient, h, was first calculated on the skin based on the "steady" "l-D", temperature distribution determined by the temperature of the surrounding air, skin surface temperature and dermis temperature. It is assumed that no dye is present and produces "h" on the skin surface. The program thereby provides the option of using this "h" on the dye surface or inserting another desired "h" for the dye surface. The "stable" temperature distribution through all the layers (including the dye layer) is then calculated by using the specified "h" at the dye surface. This temperature distribution is the initial state for the time-dependent heating problem. This constitutes "m-file" initial.m. The program then solves with a view to time-dependent temperature distribution by «marching in time», calculates and displays the temperature field for each step.

Each embodiment of the method described here for which empirical data is available has been modeled for at least one set of operating parameters to show how stratum corneum ablation can be achieved in an accurate and controllable manner. The results of the simulations are presented graphically in two different formats: (1) a cross-section of the skin showing different layers of tissue with three isotherms plotted at the top of the figure and delimiting three critical temperature thresholds, and (2) two different temperature-vs-time plots , one for the midpoint of the stratum corneum directly below the target site, and the other for the point at the border of the living cell layers of the epidermis and the underside of the stratum corneum. These plots show how the temperature at each point varies with time as heat pulses are applied, as if one could implant a microscopic thermocouple into the tissue. The application of this model also enables examination of the parametric limit values within which the method can be used, in order to determine the limitation for two important aspects of the method's behaviour. First, general cases are presented that define the framework within which the method can be applied without causing pain or unwanted tissue damage.

For any given heat source described in the various embodiments of the invention, there exists a point where the effect on the individual's skin tissue is no longer optimal, such that the individual feels a sensation of pain or the living cells in the underlying epidermis and/or dermis reach temperatures that , if sustained long enough, will damage these tissues. A test simulation was therefore run where the embodiment with optically heated locally placed copper phthalocyanine (CPC) dye constitutes a basic method to determine how the heating time constants for the various skin layers almost define a window within which the method can be used without pain or damage to the layers of adjacent tissue.

FIG. 9 and 10 show schematic cross-sections of the skin and the local dye layer. In each figure, three separate isotherms are shown: (1) 123°C, the point where evaporation of the water in the tissue leads to an ablation of the tissue; (2) 70°C, the point at which living cells will be damaged if the temperature is maintained for several seconds and (3) 45°C, the average point at which the individual will feel a sensation of pain. This pain threshold is described in several physiological books, but experience shows that the threshold is somewhat subjective. In repeated tests on the same individual, different puncture sites a few millimeters apart can actually show a significant difference in sensation, possibly as a result of the short distance between a nerve ending and the puncture site.

The dimensions in the diagram show the different layers of dye and skin, measured in µm, delimited by smooth boundary lines. While the actual skin tissue has significantly more tortuous boundary lines, the model, on average for the dimensions involved, provides a good approximation of the heat gradients that occur in the tissues in question. The dimensions used in this and all subsequent simulations for the thickness of the CPC layer and the different skin layers are as follows: dye, 10 µm; stratum corneum, 30 µm; underlying epidermis, 70 µm and dermis 100 µm.

Additional conditions used for the model for this particular simulation are shown in the following tables:

When these simulations are carried out, the following conservative assumptions are used:

1. Although it may appear that a part of the stratum corneum has a temperature that already exceeds the ablation threshold for thermal evaporation of the water content, this situation is not taken into account in the model, and the subsequent loss of heat energy in the tissue as a result of this evaporation is not included as a factor in the simulation. This will cause a slight rise in the temperatures shown in the underlying tissues from that point on in the simulation. 2. When part of the copper phthalocyanine (CPC) dye layer is correspondingly shown to have reached the evaporation point of 550°C, this situation is not included in the model, but the temperature is only limited to this level. This will also cause a slight rise in later temperatures in the underlying layers as the simulation develops. Even with these simplifications in the model, the correlation between predicted behavior and the empirically observed behavior based on both clinical studies and histological studies on donor tissue samples is remarkable. The key data to be noted in FIG. 9 and 10 are the duration of the heat pulse, and the location of the three different threshold temperatures that the isotherms show.

In FIG. 9, with a pulse length of 21 ms., the 70°C isotherm cuts exactly the boundary that separates the stratum corneum and the living cell layers of the epidermis. In an in vitro study on donor skin samples, 50 pulses of heat energy delivered at 50 millisecond intervals, under these conditions, caused detectable damage to the upper layer of living cells (see FIG. 8D). However, it was also shown in the in vitro studies that 5 pulses of heat energy under the same operating parameters did not cause any significant damage to the tissues. It seems reasonable that although the nominal damage threshold may be exceeded, at least transiently, this temperature must be maintained for some cumulative time to actually cause any damage to the cells. The essential information that emerges from the simulation is nevertheless that if you keep the "on-time" of the heat pulse for less than 20 milliseconds with the flow density of 400 Joules/cm<2>, the living cells of the underlying epidermis will not suffer any damage themselves if the ablation threshold isotherm is shifted well into the stratum corneum. By using a heat energy source with low field intensity, modulated so that the "on-time" is appropriately short, in other words, the ablation of the stratum corneum can be achieved without damage to adjacent cells in the underlying epidermis (see FIG. 8C). This is possible to a large extent as a result of the significantly different thermal conductivity of these two tissue layers. This means that the stratum corneum, which only contains approx. 10 to 20% water, has a significantly lower heat conduction constant, 0.00123 J/(S<*>cm<*>K) than that of 0.00421 J/(S<*>cm<*>K) for epidermis. This allows the temperature to build up in the stratum corneum while maintaining a narrow spatial demarcation to the point where ablation will occur.

In FIG. 10, the same simulation scenario is started at the critical damage threshold point illustrated in FIG. 9, carried forward in time. By maintaining the pulse for 58 milliseconds at the same field intensity of 400 J/cm<2>within the 60 µm diameter circumference of the dye being heated, the pain sensitivity isotherm at 45°C passes just into the weakened skin layer constituted by the dermis. In addition, the damage threshold isotherm moves significantly further into the epidermis layer than where it was shown to occur in FIG. 9. By viewing this simulation in relation to the many clinical examinations that have been carried out with this method, an excellent verification of the model's accuracy is achieved in that the model shows almost exactly the duration of "on-time" that the heat probe can be applied to the skin before the person concerned individual feels it. In clinical tests, a controllable pulse generator was used to set the "on-time" and "off-time" of a series of light pulses applied to the local layer of copper phthalocyanine (CPC) dye on the skin. While maintaining a constant "off time" of 80 milliseconds, the "on time" was gradually increased until the subject expressed a slight "pain" sensation. Without exception, all individuals involved in these studies expressed the first "pain" at an "on time" of between 45 and 60 milliseconds, very close to what was predicted by the model. In addition, the place-to-place variation previously mentioned with regard to the sensation of

"pain" noted in these clinical trials. What is designated as "pain" is therefore the point where the first unequivocal sensation can be felt. In one place this can be stated as pain, while an adjacent place by the same person can be stated as only "noticeable".

One element of this clinical examination is the recognition that even in the same place, an uneven pulse train of heat pulses can influence the individual's psycho-physiological neuro-perception to cause a genuine reduction in perceived sensation. For example, a series of shorter heat pulses can be used to saturate the neurons in the area and momentarily deprive this synapse gap of available neurotransmitters and thereby limit the ability to send a "pain" message. This means that a longer pulse that follows these short pulses will be felt less than if it was applied at the beginning of the sequence. A number of experiments were therefore carried out with some arbitrarily generated pulse trains, and the results agreed with this hypothesis. An analogy to this situation can be found in the sensation you get when you first step into a very hot bath, which is initially painful, but which quickly becomes acceptable as you acclimatise to the sensation of heat.

Example 9

An object of this invention is to achieve a painless, microporation of the stratum corneum without causing any significant damage to adjacent living tissue. As described in the simulation illustrated in Example 8 and FIG. 9-10, there appears to be a limit to any given field intensity of heat energy in the ablation target spot in which microporation can be achieved in an equally painless and non-traumatic manner. Both in vivo and in vitro studies have shown that this is the case, and this has, through empirical methods, enabled the development of certain operating parameters that seem to work very well. The following set of simulations shows how the method works when these specific parameters are used.

In the first case, a pulse train of 10 pulses, 10 milliseconds "on-time" separated by 10 milliseconds "off-time" was applied to the CPC-covered skin. FIG. 11 shows the final temperature distribution in this skin tissue immediately after the pulse train was over. As can be seen, the isotherms representing the three critical temperature thresholds show that stratum corneum ablation is achieved without any sensation in the nerves in the skin layer and very little transition of the damage threshold to the living cells in the underlying epidermis. As mentioned earlier, it appears that in order to really cause permanent cell damage, the epidermal cells must not only be heated to a certain point, but that they must also be held at this temperature for a certain period of time, which is generally assumed to be approx. 5 seconds. FIG. 12 and 13 show the temperature respectively in the stratum corneum and the living epidermis, as a function of time, with heating during the "on time" and cooling during the "off time", where all 10 cycles are shown. When comparing this simulation with the conducted in vivo investigations, it must be noted that more than 90% of the poration attempts with the system parameters set to adapt to the simulation, effective poration of the stratum corneum was achieved without pain for the individual concerned, and in microscopic examination of the poration site for several days later, no significant tissue damage was visible. The in vitro studies performed on full-thickness donor skin samples were also consistent with the model's behavior prediction.

Example 10

Both when conducting the empirical in vivo studies and these simulations, it appears that prior cooling of the skin helps to optimize the microporation process in terms of reducing the potential for pain or destruction of adjacent tissue. In practice, this can easily be achieved by using a simple cooling plate placed against the skin before the poring process. For example, placing a Peltier-cooled plate against the 1 cm diameter circle surrounding the poration target site, with the plate held at approx. 5°C for a few seconds, clearly reduce the tissue temperature. A schematic illustration of an experimental device used for this purpose in the laboratory is shown in FIG. 3A-B. By applying exactly the same pulse train of 10 cycles as in the simulation illustrated in Example 9, one can, by comparing FIG. 11 with FIG. 14, FIG. 12 with FIG. 15 and FIG. 13 with FIG. 16, see what improvement can be made with regard to the control of the temperature penetration into the skin tissues. Again, the relatively low heat diffusivity and specific heat of the stratum corneum, compared to the epidermis and dermis, is advantageous. After they are first cooled, the highly hydrated epidermis and dermis tissues require a significantly greater supply of heat energy for temperature elevation, while the stratum corneum, with its relatively dry nature, can quickly be heated to the ablation threshold.

Example 11

After the basic heat conduction mechanism for delivering the energy to the skin tissue and underlying the effective painless ablation and microporation of the stratum corneum is recognized, several different specific methods for achieving the necessary rapid temperature modulations of the contact point can be devised, such as hot wire- the embodiments illustrated in FIG. 4-7.

A basic embodiment described here makes use of an ohmic heating element (FIG. 4), which e.g. the tip of a small cordless soldering iron, with an appropriately sized relatively non-reactive wire wound around it, with a short section of the wire left to protrude from the heater. When electricity is supplied from a cash flow source, the heating element will reach a certain temperature and within a few seconds reach an equilibrium state with the convection loss to the surrounding air. Similarly, the wire that is part of this heating system will reach an equilibrium state so that the very tip of the wire can be raised to just about any arbitrary temperature up to approx. 1000°C with these types of components. The tip can be sized to provide exactly the micropore dimension desired.

In a laboratory context, tungsten wires with a diameter of 80fam attached to the non-replaceable tip of a "WAHL" wireless soldering iron with approx.

2 mm thread protruding from the tip. With a thermocouple, the temperature of the tip has been measured in the equilibrium state, and it has been determined that by varying the constant current settings, equilibrium temperatures of more than 700°C can easily be achieved. To achieve the desired modulation, a fast-responding electromechanical actuator with low mass was connected to the tip so that the position of the wire could be displaced linearly more than 2 mm at a rate of up to 200 Hz. By mounting the entire device on a precision stand, this vibrating tip could be brought into contact with the skin surface under good control so that it was only in contact for less than 10 milliseconds at a time, the "on time", while an "off time" for arbitrarily long periods could be achieved by corresponding setting of the pulse generator. These in vivo investigations showed that the poration could actually be achieved before the patient even knew that the wire tip had been brought into contact with the skin.

To compare the behavior of this embodiment with the optically heated CPC dye embodiment, the following simulations were performed according to the procedure of Example 8. By varying only the initial conditions, the heat wire embodiment could be performed with essentially the same simulation code. Since contact with the wire is essentially instantaneous, there is no time-dependent heat build-up in the CPC layer, and when the wire is physically removed from contact with the skin, there is no residual heat left on the surface, as with the heated CPC layer. Furthermore, since the wire itself delimits the area intended for ablation/micro-poration, there should be no lateral diffusion of heat energy prior to its application to the stratum corneum. Comparisons of the performances of the "hot wire" embodiment are shown in FIG. 17-19.

Example 12

In this example, the procedure of Example 11 was followed, except that the skin was pre-cooled according to the procedure of Example 10. Similarly, pre-cooling of the target site produced similar positive results with the "hot wire" embodiment. The results of simulation of the hot wire method after prior cooling are shown in FIG. 20-22.

Example 13

As discussed during the introduction of this specification, the Tankovich '803 patent appears at first glance to correspond to the present invention. In this example, the simulation model was set to the operating parameters specified in Tankovich '803, ie a pulse width of 1^s and a power level of 40,000,000 W/cm<2>. FIG. 23 and 24 show that no part of the stratum corneum under these conditions reached the threshold value for rapid evaporation of water, 123°C, and that therefore no ablation/microporation of the stratum corneum occurred. By using this type of high peak power, the short pulses against the locally applied dye layer only vaporize the dye from the skin surface without any effect on the skin. This example thus demonstrates that the conditions set forth by Tankovich '803 are without effect for the present invention.

Example 14

In this example, interstitial fluid obtained after poring the skin according to the procedure in Example 6, was collected and analyzed to determine the glucose concentration therein. Data were generated from four non-diabetics and six type I diabetics who underwent a glucose stress test. Their ages ranged from 27 to 43 years. The purpose was to investigate the method's suitability for painlessly extracting sufficient interstitial fluid (ISF) from subjects to enable analysis of ISF samples for glucose content and then comparing these concentrations with glucose levels found in their whole blood.

All subjects had both blood and ISF glucose determinations made with the "ELITE" system from Miles-Bayer. All ten subjects were subjected to identical measurement protocols, with adjustments made for glucose load and insulin injection for subjects with insulin-dependent diabetes.

The trial scheme consisted of recruiting a moderate number of volunteers, some with diabetes and some without diabetes, from whom a number of sample pairs of ISF and whole blood were taken every 3 to 5 minutes during the 3 to 4 hours the investigation lasted. Both the blood and ISF samples were analyzed for glucose, and the statistical relationship between blood glucose levels and interstitial fluid was determined. To examine the hypothesis of a transient delay of ISF glucose levels relative to whole blood glucose levels, subjects were induced to produce a significant and dynamic change in their glucose levels. This was accomplished by allowing each subject to fast for 12 hours before the start of the test and then giving the subjects a glucose load after his or her basal glucose level was established, through a set of three fasting blood and ISF glucose levels. After basal levels were established, the subjects were given a glucose load in the form of sweet juice, according to the following guidelines:

i. For control subjects, the glucose load was calculated on the basis of

1.7 grams of glucose per kg of body weight.

ii. For people with insulin-dependent diabetes, the glucose load was 50 grams of glucose. In addition, after taking the glucose load, the diabetics injected themselves with their normal morning dose of rapid-acting insulin. In the event that the diabetic had fasting glucose levels higher than 300 mg/dL, they were asked to first give themselves their insulin injection, after which their glucose load was given after blood glucose levels had fallen below 120 mg/dL.

Each of the participants was given a full description of the study in the "Informed Consent" document, which they were asked to familiarize themselves with and sign before they formally joined the program. After agreeing to this, they filled out a medical questionnaire. The clinical procedure implemented was: (a) The participant fasted from 21:00 the evening before attendance and consumed only water. There was no occasion for caffeine, cigarettes or fruit juice during this time.

(b) The participant showed up for the test at 09:00 the next day.

(c) The participant was placed in a reclining chair to be able to relax during the examination process. (d) Both whole blood and ISF samples were collected at 3- to 5-minute intervals starting on arrival and continuing for the next 3 to 4 hours. The sample collection time was determined by the time when the subject's blood glucose level had returned to the normal range and stabilized after the glucose load. The ISF samples were taken using the optical poration, ISF pump method, described in more detail below. Each ISF sample amounted to approx. 5^L to ensure that the ELITE test strip was properly filled. The blood samples were taken using a conventional finger-puncture lancet. Both the ISF and blood samples were immediately examined for glucose with the ELITE home glucometer system from Miles-Bayer. To improve the estimate of the "true" blood glucose levels, two separate ELITE tests were performed on each finger prick sample. (e) To facilitate the continuous withdrawal of ISF from the same location on the patient during the entire data collection phase, a 5 by 5 matrix of 25 micropores was created on the subject's forearm, each micropore being between 50 and 80 nm at a distance of 300^m. A 30 mm diameter Teflon disc with a 6 mm hole in the center was attached to the subject's forearm with a pressure-sensitive adhesive and positioned so that the 6 mm hole in the center was located over the micropores of the 5 by 5 matrix. This attachment enabled an easy method whereby the porous area via a small suction tube could be connected to a weak vacuum (250-300 mm Hg) to cause ISF to flow out of the body through the micropores. The upper part of the Teflon disc was provided with a clear glass window which gave the operator the opportunity to directly observe the underlying microporous skin. When a 5 µL ISF globule formed on the skin surface, this could be easily determined by visual observation of the site through this window. This vacuum produced a nominal pressure gradient of approx. 0.35 kg/cm<2>. Without the micropores, no ISF could be extracted from the subject's body at all using only the weak vacuum. (f) After the first three pairs of trials were taken, the subject was given a glucose load in the form of a highly sweetened orange juice. The amount of glucose given was 1.7 grams per kg of body weight for non-diabetics and 50 grams for diabetics. The diabetics also gave themselves an injection of fast-acting insulin (regular) of the correctly calculated dose based on this 50 gram glucose level at the same time as the intake of the glucose load. With the normal 1.5 to 2.5 hour delay between the insulin injection and the maximum effect of the injection, the diabetic was expected to show a rise in glucose levels up to 300 mg/dL and then fall back to the normal range as the insulin took effect. The non-diabetics were expected to exhibit the standard glucose tolerance profiles which generally show a peak in blood glucose levels between 150 mg/dL and 220 mg/dL from 45 minutes to 90 minutes after administration of the glucose load, and then a rapid fall back to their normal baseline level in over the next hour or so. (g) After administration of the glucose load or glucose load and insulin injection, subjects had three simultaneous samples of ISF and of finger-prick whole blood drawn at 5-minute intervals for the next 3 to 4 hours. Sampling was terminated when the blood glucose levels in three successive samples indicated that the subject's glucose level had stabilised.

When reviewing the data, several features were visible. In particular, for any given production rate of ELITE test strips, there is a clear shift in the result indicated in mg/dL of glucose on the glucometer compared to the level indicated by the blood. A higher reading would be expected due to the lack of hematocrit in ISF and the normal differences in electrolyte concentrations between ISF and whole blood. Regardless of the underlying reasons for these shifts in the displayed result, when compared to a reference test, it was determined that the true ISF glucose levels are linearly related to the values produced by the ELITE system, with constant scaling coefficients for each particular production batch of ELITE strips. For the comparison of the ISF glucose levels in relation to the whole blood measurements, the following first order linear correction of the ISF data was entered:

ISFglyucosis<=>0.606<*>ISFelitis<+>19.5.

This scaling of the reading of the ELITE glucometer when measuring ISF glucose levels makes it possible, for the entire data set, to investigate sources of error associated with the use of ISF for the determination of blood glucose levels. Of course, even without any linear scaling, the correlations between the ISF glucose values and the whole blood glucose values are the same as for the scaled version.

Based on the bulk of the published literature regarding ISF glucose as well as preliminary data, it was originally expected that a 15 to 20 minute delay would be seen between ISF glucose levels and those drawn from finger prick whole blood. However, this is not what the data showed when analyzed. Specifically, when each subject's data set was analyzed to determine what time lag was needed to achieve maximum correlation between the ISF glucose levels and blood glucose levels, it was determined that the longest time lag for this sample of subjects was only 13 minutes, and that the average time lag was only 6.2 minutes, where several people also showed a temporal adaptation that was almost instantaneous (approx. 1 minute).

Based on the minimal displacement observed for this data set, the curve shown in FIG. 21 all ten glucose challenge tests chained one after the other on an extended time scale. The data are presented without any time lag and show the high degree of agreement between the ISF and blood glucose levels when the entire clinical data set was processed in exactly the same way. If the total data set is shifted as a whole to find the best estimate of temporal agreement, the correlation between the ISF and blood glucose levels has a maximum at a lag of two (2) minutes at an r value of r=0.97. This is only a minor improvement from the unshifted correlation of r=0.964. During the rest of the analysis, the ISF values were therefore processed without any time shift. That is, each set of blood and ISF glucose levels was treated as simultaneously collected data pairs.

After the non-shifted ELITE ISF readings were scaled to reflect the proportion of glucose in the ISF, it was possible to examine the error associated with these data. The simplest method for this is to assume that the average of the two ELITE fingerstick blood glucose readings is actually the absolute correct value and then simply compare the scaled ISF values to these average blood glucose values. The following data were obtained: standard deviation blood ISF, 13.4 mg/dL; coefficient of variation for ISF, 9.7%; standard deviation of the two ELITE determinations, 8.3 mg/dL and the coefficient of variation for blood (Miles), 6%.

As these data show, the measurements based on blood already contain an error component. The manufacturer's published performance data indicate that the ELITE system has a nominal coefficient of variation (CV) of between 55 and 7%, depending on the glucose levels and blood hematocrit value.

A further assessment of the difference between ISF glucose and blood glucose is shown in the form of a dot plot in FIG. 26.1 this figure, the upper and lower limits of the 90% confidence interval are also indicated. It is interesting to note that all data, with only two exceptions, in the sample of blood glucose levels below 100 mg/dL, fall within these 90% confidence interval error ranges. This is important since the consequences of missing a tendency towards hypoglycaemia will be of significant importance for the diabetic. That is, it would be much better to underestimate glucose levels in the 40 to 120 mg/dL range than to overestimate them. If one essentially assumes that the measurement error when the ELITE system is used on ISF is comparable to the measurement error associated with the use of ELITE on whole blood, the discrepancy between ISF glucose and blood glucose can be described as:

By applying this equation to the values shown below, it can be solved for the estimated "true" value of the ISF error part: or solve the equation

A histogram of the relative deviation between ISF and blood glucose levels is shown in

FIG. 27.

Drug delivery through pores in the stratum corneum

The present invention also includes a method for delivering drugs, including drugs that are currently delivered transdermally, through micropores in the stratum corneum. In an illustrative embodiment, delivery is achieved by placing the solution in a reservoir above the poration site. In another illustrative embodiment, a pressure gradient is used to further improve delivery. In yet another illustrative embodiment, sound energy is used with or without a pressure gradient to further enhance the delivery. The sound energy can be controlled according to traditional transdermal parameters or by using acoustic streaming effects, which will be described immediately, to push the delivery solution through the porous stratum corneum.

Example 15

This example shows the use of stratum corneum perforation for the delivery of lidocaine, a local analgesic. The lidocaine solution also contains a chemical permeation enhancer formulation designed to enhance its passive diffusion through the stratum corneum. A drawing of an illustrative delivery apparatus, 300, is shown in FIG. 28, where the apparatus includes a housing 304, which encloses a reservoir 308, for holding a drug-containing solution 312. The upper part of the housing includes an ultrasound transducer 316, which provides sound energy that helps to transport the drug-containing solution through micropores 320, in stratum corneum 324. An opening 328 in the ultrasound transducer makes it possible to put it under pressure to further help transport the drug-containing solution through the micropores in the stratum corneum. The delivery device is placed on a selected area of the person's skin so that it is placed over at least one, and preferably several, micropores. An adhesive layer 332, attached to the lower part of the housing, causes the device to stick to the skin so that the drug-containing solution in the reservoir is in fluid contact with the micropores. Delivery of the drug through the micropores results in transport into the underlying epidermis 336, and dermis 340.

Five subjects were tested for the effectiveness of drug delivery using poration in conjunction with ultrasound. In the experiment, two places were used on the person's left forearm at a distance of approx. 8 cm at equal distance from thumb and upper arm. The site near the thumb will be referred to as site 1, the site furthest from the thumb will be referred to as site 2. Site 1 was used as a control where lidocaine and enhancer solution was applied using an identical delivery device 300, but without any microporation of the stratum corneum or sound energy. Site 2 was perforated with 24 holes at 0.8 mm spacing in a grid within a circle of 1 cm diameter. The micropores in site 2 were produced according to the procedure in Example 6. Lidocaine and low-level ultrasound were applied. The ultrasound applications were made with a custom-ordered Zevex ultrasound transducer set in burst mode with a 0.4 volt peak-to-peak input to the ENI #2100L linear amplifier entering 1000 count bursts at 10 Hz with a 65 .4 kHz fundamental frequency, i.e. a pulse modulated signal with the transducer energized for 15 millisecond bursts and then turned off for the next 85 milliseconds. The measured output from the amplifier to the transducer was 0.090 Watts RMS.

After administration of the lidocaine, sensation measurements were made by rubbing a 30 gauge wire over the test site. The experiments were carried out at both sites, site 1 for 10 to 12 minutes duration and site 2 for two 5 minute intervals applied one after the other at the same site. Both sites were assessed for numbness using a scale from 10 to 0, where 10 indicates no numbness and 0 indicates that the subject indicates complete numbness. The following summary of the results applies to all five people.

The control site, site 1, showed little or no numbness (7 to 10 scale) after 10 to 12 minutes. After approx. After 20 minutes, a certain numbness (scale 3) was observed at site 1, as the solution completely permeated the stratum corneum. Site 1 was cleaned after completion of lidocaine administration. Site 2 showed almost complete numbness (scale 0 to 1) within the 1 cm area containing the perforations. Outside the 1 cm circle, the numbness decreased almost linearly to 1 in a 2.5 cm diameter circle with no numbness outside the 2.5 cm diameter circle. Assessment of site 2 after the second application resulted in a slightly larger circle of total numbness with approx. 1.2 cm diameter, where the numbness decreased linearly to 1 in an irregular oval pattern with a diameter of 2 to 2.5 cm perpendicular to the forearm and a diameter of 2 to 6 cm parallel to the forearm. Outside of this worm mode, no numbness was recorded. A graphical representation of a typical illustrative result obtained on one subject is shown in FIG. 29A-C. FIG. 29A and 29B show the results obtained at location 2 (porous) after 5 and 10 minutes respectively.

FIG. 29C shows the result obtained at site 2 (control without poration).

Sound energy and transdermal flow amplification amplifiers

The physics of sound energy fields produced by acoustic transducers can be utilized in a method where the sound frequency can be modulated to improve flow rates achieved by other methods. As shown in FIG. 1 in US patent 5,445,611, the energy distribution of an acoustic transducer can be divided into near and far fields. The near field, characterized by the length N, is the zone from the first energy minimum to the last energy maximum. The zone distal to the last maximum is the far field. The near (N) field pattern is dominated by local pressure peaks and zeros at a small distance from each other. The length of the near-field zone, N, is a function of the frequency, the size and shape of the transducer surface and the speed of sound in the medium through which the ultrasound passes. For a single transducer, intensity variations within its normal operating range do not affect the nature of the sound energy distribution in any other way than linearly. For a system with several transducers that are all modulated in both frequency and amplitude, however, the relative intensities of the individual transducers affect the energy distribution in the sound medium, regardless of whether it is skin or another medium.

By making a moderate change in the frequency of the sound energy, say in the range of 1 to 20%, the pattern of peaks and troughs remains relatively constant, but the length, N, of the near field zone changes in direct proportion to the frequency. Significant changes in frequency with e.g. a factor of 2 or more, will very likely produce a different set of resonances or modes of vibration in the transducer, causing an important and unpredictably different pattern of the near-field energy. With a moderate change in the sound frequency, the complex pattern of peaks and zeros will thereby be compressed or expanded in an accordion-like manner. By choosing the directions of frequency modulation, the direction of displacements of these local pressure peaks can be controlled. By applying sound energy to the surface of the skin, selective modulation of the sound frequency controls the movement of these local pressure peaks through the skin either towards the interior of the body or towards the body surface. A frequency modulation from high to low directs the pressure peaks into the body, while a frequency modulation from low to high draws the pressure peaks from the body towards the surface and through the skin to the outside of the body.

Assuming typical parameters for this application, such as a 1.27 cm diameter acoustic transducer and a nominal operating frequency of 10 MHz and a similar acoustic impedance to water, a frequency modulation of 1 MHz produces a movement of approx. 2.5 mm of the peaks and zeros in the near-field energy pattern near the stratum corneum. With regard to transdermal and/or transmucosal extraction of analytes, this degree of movement gives access to the area well below the stratum corneum and even to the epidermis, dermis and other tissues below it. For a given transducer, there may be an optimal frequency range within which this frequency modulation is most effective.

The flow of a drug or analyte through the skin can also be increased by changing either the resistance (diffusion coefficient) or the driving force (diffusion gradient). The current can be amplified by using so-called penetration or chemical amplifiers.

Chemical enhancers consist of two primary categories of components, i.e., cell envelope-disrupting compounds and solvent or binary systems containing both cell envelope-disrupting compounds and solvents.

Cell envelope-disrupting compounds are known in this area to be suitable for pharmaceutical preparations for local use and also work by analyte uptake through the skin. These compounds are believed to contribute to skin penetration by disrupting the lipid structure of the stratum corneum cell envelope. An extensive list of these compounds is described in European Patent Application 43,738, published June 13, 1982. It is believed that any cell envelope disrupting compound is useful for purposes of the present invention.

Suitable solvents are i.a. water; diols, such as propylene glycol and glycerol; monoalcohols, such as ethanol, propanol and higher alcohols; DMSO; dimethylformamide; N,N-dimethylacetamide; 2-pyrrolidone; N-(2-hydroxyethyl)-pyrrolidone, N-methylpyrrolidone, 1-dodecyl azacycloheptan-2-one and other N-substituted alkyl-azacycloalkyl-2-ones (azones) and the like.

US Patent 4,537,776, Cooper, av27. August, 1985, contains an excellent summary of the state of the art and background information, detailing the use of certain binary systems for permeant enhancement.

Similarly, European patent application 43,738 referred to above describes the use of selected diols as solvents together with a wide variety of cell envelope-disrupting compounds for the delivery of lipophilic pharmacologically active compounds.

A binary system to enhance metoclopramide penetration is disclosed in UK Patent Application GB 2,153,223 A, published August 21, 1985, and consists of a monovalent alcohol ester of a C8-32 aliphatic monocarboxylic acid (unsaturated and/or branched if C18- 32) or a C6-24-aliphatic monoalcohol (unsaturated and/or branched if C14-24) and an N-cyclic compound, such as e.g. 2-pyrrolidone, N-methylpyrrolidone and the like.

Combinations of enhancers consisting of diethylene glycol monoethyl or monomethyl ether with propylene glycol monolaurate and methyl laurate are stated in US patent 4,973,468 to improve the transdermal delivery of steroids such as progestogens and estrogens. A dual enhancer consisting of glycerol monolaurate and ethanol for transdermal drug delivery is shown in US Patent 4,820,720. US patent 5,006,342 specifies a series of enhancers for transdermal drug administration consisting of fatty acid esters or fatty alcohol ethers of C2 to C4 alkanediols, where each fatty acid/alcohol part of the ester/ether is approx. 8 to 22 carbon atoms. US Patent 4,863,970 discloses penetration-enhancing compositions for topical application, comprising an active permeant contained in a penetration-enhancing carrier containing specified amounts of one or more cell envelope-disrupting compounds, such as oleic acid, oleyl alcohol and glycerol esters of oleic acid; a C2 or C3 alkanol and an inert diluent such as water.

Other chemical enhancers not necessarily associated with binary systems include DMSO or aqueous solutions of DMSO as described in Herschler, US Patent 3,551,554; Herschler, US Patent 3,711,602; and Herschler, US Patent 3,711,606, and azones (N-substituted-alkyl-azacycloalkyl-2-ones) as disclosed in Cooper, US Patent 4,557,943.

Certain chemical amplifier systems can have negative side effects such as e.g. toxicity and skin irritation. US Patent 4,855,298 discloses compositions for reducing skin irritation caused by chemical enhancers, which contain compositions having skin irritant properties, together with a sufficient amount of glycerol to provide an anti-irritant effect.

Since the combination of microporation of the stratum corneum and the application of sound energy accompanied by the use of chemical enhancers can result in an improved extraction rate of analyte or permanent release through the stratum corneum, the particular carriers and in particular the applied chemical enhancers can be selected from a long list of previously known carriers. It is not considered necessary to specifically list what is readily available in the area. The invention is not directed to the use of chemical enhancers as such, and it is believed that all chemical enhancers suitable for the delivery of drugs through the skin will work with dyes by optical microporation and also with sound energy to effect measurable extraction of analyte from within and through the skin surface or release of permeants or drugs through the skin surface.

Example 16

Modulated sound energy and chemical amplifiers were tested on cadaver skin samples for their ability to control transdermal permeation. In these tests, the epidermis membrane was separated from whole skin from a corpse using the heat separation method according to Example 1. The epidermis membrane was cut and placed between two halves of the permeation cell with the stratum corneum directed either towards the upper (donor) compartment or the lower (recipient) ) room. Modified Franz cells were used to hold the epidermis, as shown in FIG. 2 of US patent 5,445,611. Each Franz cell consisted of an upper chamber and a lower chamber held together by one or more clamps. The lower chamber had a sampling opening through which materials could be added or removed. A sample of the stratum corneum is held between the upper and lower vessels when they are clamped together. The upper chamber of each Franz cell is modified to allow placement of an ultrasound transducer 1 cm from the stratum corneum membrane. Methylene blue solution was used as an indicator molecule to assess the permeation of the stratum corneum. A visual record of the process and results of each experiment was achieved through a time-recording magnetic tape format with a video camera and a video cassette recorder (not shown). In addition, samples were taken to measure with an absorption spectrometer the amount of dye that had passed through the stratum corneum membrane during an experiment. Chemical enhancers suitable for use may vary, as indicated above, within a wide variety of solvents and/or cell envelope-disrupting compounds. The particular enhancer used was: ethanol/glycerol/water/glycerol monooleate/methyl laurate in a 50/30/15/2.5/2.5 volume ratio. The system for generating and controlling the sound energy included a programmable 0-30 MHz arbitrary waveform generator (Stanford Research Systems Model DS345), a 20 Watt 0-30 MHz amplifier and two unfocused submersible ultrasonic transducers having maximum resonances at 15 and 25 MHz. Six cells were prepared simultaneously for testing stratum corneum samples from the same donor. After the stratum corneum samples were installed, they were allowed to hydrate with distilled water for at least 6 hours before any testing was performed.

Example 17

Effect of sound energy without chemical amplifiers

As indicated above in Example 16, the heat-separated epidermis was placed in the Franz cells with the epidermis side up and the stratum corneum side down, unless otherwise indicated. The lower chambers were filled with distilled water, while the upper chambers were filled with a concentrated methylene blue solution in distilled water.

Heat-separated epidermis: Immediately after the upper chambers were filled with methylene blue solution, one of the cells was sonicated with the transducer completely submerged. This orientation would, for example, correspond to having the transducer on the opposite side of a skin fold or causing the sound energy to be reflected away from a correspondingly placed reflector plate and used to "push" analyte out of the other side of the fold into a collection device. The sound energy setting was first set at the nominal operating frequency of 25 MHz with an intensity corresponding to a 20 volt P-P (peak-to-peak) input waveform. This corresponds to approximately a 1 Watt average input power to the transducer and correspondingly, assuming the manufacturer's nominal conversion efficiency value of 1% for this particular transducer, an acoustic output power of approx. 0.01 Watt over the 0.78 cm surface of the active area or a sound intensity of 0.13 Watt/cm<2>. Three other control cells were not supplied with sound energy. After 5 minutes that sound energy was turned off. No visual indication of dye flow through the stratum corneum was observed in any of the cells in this interval, indicating levels of less than ca.

0.0015 vol% dye solution in 2 ml_ recipient medium.

Testing of these 3 same control cells and 1 experimental cell continued as follows. The intensity of the sound energy was increased to the maximum possible output from the drive equipment of a 70 volt peak-to-peak input 12 Watt average input power or ("0.13 Watt/cm<2>) of acoustic output intensity. The frequency was also set to modulate or sweep from 30 MHz to 10 MHz. This 20 MHz sweep was performed 10 times per second, ie a sweep rate of 10 Hz. At these input power levels it was necessary to monitor the sound energy transducer to avoid overheating. A contact thermocouple was placed directly on the transducer body and the load cycled on and off to keep the maximum transducer temperature below 42°C. After approx. 30 minute cycling of maximum power at an approx. 50% duty cycle of 1 minute on and 1 minute off, there was still no visually detectable permeation of the stratum corneum with the methylene blue dye.

A water cooling jacket was then connected to the sound energy transducer to allow extended excitation at the maximum energy level. Using the same three controls and a test cell, sound energy was applied to the test cell at maximum energy for 12 hours. During this time, the temperature of the liquid in the upper chamber only rose to 35°C, only a fraction above the approx. 31 °C normal temperature of the stratum corneum in vivo. No visible indication of dye flow through the stratum corneum was recorded in any of the four cells after application of sound energy as described above for 12 hours.

Example 18

Effects of sound energy without chemical amplifiers

Perforated stratum corneum: Six cells were prepared as described above in Example 16. The clamps holding the upper and lower chambers of the Franz cells were tightened more than normally required to seal the upper chamber from the lower chamber and so that the perforations and "nail holes" were artificially introduced into the heat-separated epidermis samples. When dye solution was added to the upper chamber of each cell, there was an immediate visual indication of leakage of dye into the lower chambers through the perforations formed in the statum corneum. After applying sound energy to cells where the stratum corneum was perforated in this way with small "nail holes", a rapid increase in the transport of liquid through a nail hole in the stratum corneum was observed. The rate of transport of the indicator dye molecules was directly related to whether or not sound energy was applied. That is, the application of sound energy caused an immediate (delay time approx. <0.1 second) pulse of the indicator molecules through the nail holes in the stratum corneum. This pulse of indicator molecules ceased immediately after the sound energy was turned off (a turn-off delay of about <0.1 second). The pulse could be repeated as described.

Example 19

Effects of sound energy and chemical amplifiers

Two different chemical enhancer formulations were used. Chemical EnhancerOne or CE 1 was a mixture of ethanol/glycerol/water/glycerol monooleate/methyl laurate in a volume ratio of 50/30/15/2.5/2.5. These are components generally regarded as safe, i.e. GRAS, by the FDA for use as pharmaceutical excipients. Chemical Enhancer Two or CE2 is an experimental formulation that has been shown to be very effective in enhancing transdermal drug delivery, but generally considered too irritating for long-term transdermal delivery. CE2 contained ethanol/glycerol/water/lauradone/methyl laurate in the volume ratio 50/30/15/2.5/2.5. Lauradone is the lauryl (dodecyl) ester of 2-pyrrolidone-5-carboxylic acid ("PCA") and is also referred to as lauryl-PCA.

Six Franz cells were set up as above (Example 16) except that the heat-separated epidermis was installed with the epidermis layer down, ie the stratum corneum side facing up. Hydration was initiated by exposing each sample to distilled water overnight. To begin the experiment, the distilled water in the lower chambers was replaced with methylene blue solution in all six cells. The upper chambers were filled with distilled water and the cells were observed for approx. 30 minutes, confirming that no passage of dye occurred to ensure that no pinhole perforations occurred in any of the cells. When none were found, the distilled water in the upper chamber was removed from four of the cells. The other two cells served as distilled water controls. The upper chambers of the two test cells were then filled with CE1, and the other two test cells were filled with CE2.

Sound energy was immediately supplied to one of the two CE2 cells. A 25 MHz transducer was used with a frequency that every 0.1 second swept from 10 MHz to 30 MHz with a maximum intensity of "0.13 Watt/cm<2>". After 10-15 minutes of applied sonic energy at a 50% duty cycle, dye flow was visually detected. No dye flow was detected in the other five cells.

Sound energy was then applied to one of the two cells containing CE1 at the same settings. Dye began to appear in the upper chamber within 5 minutes. Sound energy together with a chemical amplifier thus gives a significantly increased transdermal flow rate of a marker dye through the stratum corneum, and likewise reduced the delay time.

Example 20

Effects of sound energy and chemical amplifiers

Formulations of the two chemical enhancers CE1 and CE2 were prepared minus glycerol and these new formulations, designated CE1MG and CE2MG, were tested as above. The glycerol was replaced with water so that the ratio between the other components remained unchanged. Three cells were prepared in modified Franz cells with the epidermis side of the heat-separated epidermis sample facing the upper side of the chambers. These samples were then hydrated in distilled water for 8 hours. After the hydration step, the distilled water in the lower chambers was replaced with either CE1 MG or CE2MG, and the upper chamber was filled with the dye solution. Sound energy was applied to each of the three cells in turn.

After application of the pulsed, frequency-modulated sound energy for a total of less than 10 minutes, a significant increase in the permeability of the stratum corneum samples was observed. The permeability of the stratum corneum was changed relatively uniformly over the exposed area both for the chemical amplifier and for sound energy. No "nail hole" perforations were observed through which the dye could pass the stratum corneum. The transdermal flow rate could be controlled instantaneously by turning the sound energy on or off. When the sound energy was turned off, this appeared to instantly reduce the transdermal flow rate so that no dye was visibly actively transported through the skin sample. Presumably the rate was reduced to that of passive diffusion. When the sonic energy was turned back on, the high flow rate immediately resumed. The modulated mode appeared to produce a regular pulsatile increase in the transdermal flow rate at the modulated rate. When the sound energy was set at a constant frequency, the maximum increase in transdermal flow rate for this configuration appeared to occur at ca. 27 MHz.

After obtaining the same results with all three samples, the cells were drained of all fluid and rinsed with distilled water on both sides of the stratum corneum. The lower chambers were then immediately filled with distilled water and the upper chambers again filled with dye solution. The cells were observed for 30 minutes. No holes were observed in stratum corneum samples, and no large amounts of dye were detected in the lower chambers. A small amount of dye was visible in the lower chambers, probably as a result of dye and enhancer trapped in skin samples from their previous exposures. After a further 12 hours, the amount of dye detected was still very small.

Example 21

Effects of sound energy and chemical amplifiers

Perforated stratum corneum: Three cells were prepared with heat-separated epidermis samples with the epidermis side facing the upper side of the chamber, from the same donor as in Example 16. The samples were hydrated for 8 hours, and the distilled water in the lower chamber was then replaced with either CE1MG or CE2MG. The upper chambers were then filled with dye solution. Pinhole perforations in stratum corneum samples allowed dye to leak through the stratum corneum samples into the underlying enhancer-containing chambers. Sound energy was applied. Immediately after application of the sound energy, the dye molecules were quickly pushed through the pores. As shown above, the rapid flow of dye through the pores could be directly and immediately correlated with the application of the sound energy.

Example 22

Effects of sound energy and chemical amplifiers

An inexpensive sound energy transducer, TDK #NB-58S-01 (TDK Corp.) was tested for its ability to amplify transdermal flow rates. The maximum response for this transducer was determined to be approx. 5.4 MHz with the other local peaks occurring at approx. 7 MHz, 9 MHz, 12.4 MHz and 16 MHz.

This TDK transducer was then tested at 5.4 MHz for its ability to

enhance transdermal flow rate together with CE1MG. Three cells were set up with the epidermis side towards the lower chamber, after which the skin samples were hydrated for 8 hours. The dye solution was placed in the lower chamber. The transducer was placed in the upper chamber immersed in the CE1 MG. By using sweep frequencies from 5.3 to 5.6 MHz as sound energy excitation, significant amounts of dye were carried through the stratum corneum and detected in the cell's collection well within 5 minutes. A local heating occurred, where the transducer reached a temperature of 48°C. In a

control using CE1 MG without sonic energy, a 24 hour exposure resulted in less dye in the collection well than the 5 minute exposure with sonic energy.

This example shows that an inexpensive low-frequency sound energy transducer can strikingly affect transdermal flow rate when used in conjunction with an appropriate chemical amplifier. Although more high frequency sound energy will theoretically concentrate more energy in the stratum corneum, when applied with a chemical amplifier, the lower frequency modulated sound energy can accelerate the transdermal flow rate to make the technology applicable and practical.

Example 23

Demonstration of molecule migration through human skin: The tests described above with the TDK transducer and CE1 MG were repeated at approx. 12.4 MHz, one of the transducer's highest local resonance peaks, with a frequency sweep at a rate of 2 Hz from 12.5 to 12.8 MHz and a sound energy density of less than 0.1 W/cm<2>. The epidermis side of the heat-separated epidermis was turned down, the dye solution was in the lower chamber, and the enhancer solution and sound energy were placed in the upper chamber. Within 5 minutes, a substantial part of the dye had moved through the stratum corneum into the collection well. Ohmic heating in the transducer was significantly less than with the same transducer operated at 5.4 MHz, and caused a temperature increase in the chemical amplifier to only approx. 33°C.

Even at these low power levels, the results obtained with the CE1 MG and sound energy from the TDK transducer were remarkable in the monitoring direction. FIG. 3A and 3B of US Patent 5,445,611 show a plot of data obtained from three separate cells with the transdermal flow rate measured in the direction of monitoring. Even at the 5-minute time point, readily measurable amounts of the dye in the chemical enhancer occurred on the outside of the stratum corneum, indicating transport from the epidermis side through the stratum corneum to the "outside" area of the skin sample.

In order to optimize the use of sound energy or of sound energy/chemical amplifier to collect and monitor analytes from the body, devices are required to determine the amount of the analyte in question. A determination system that takes multiple readings while the device is in the process of extracting analyte using sound energy with or without a chemical amplifier makes it unnecessary to standardize over a broad population base and normalize for different skin characteristics and flow rates. By plotting two or more data points in time as the analyte concentration in the collection system increases, a curve fitting algorithm can be used to determine the parameters that describe the curve with respect to analyte withdrawal or flow rate until equilibrium is reached, thereby establishing the target interval concentration. The general shape of this curve does not vary from one individual to another; only the parameters change. Once these parameters are determined, solving this function with respect to the equilibrium state solution (ie time is infinite), ie when full equilibrium is established, will give the concentration of the analyte in the body. This method thus makes it possible to carry out measurements to the desired level of precision, during the same time for all members of a population regardless of individual variations in skin permeability.

Several detection techniques exist today that can be adapted to this application. See D.A. Christensen in 1648 Proceedings of Fiber Optic, Medical and Fluorescent Sensors and Applications 223-26 (1992). One method involves the use of a pair of optical fibers which are placed close together in approximately parallel. One of the fibers is a source fiber through which light energy is conducted. The second fiber is a detection fiber connected to a light-sensitive diode. When light is passed through the source fiber, part of the light energy, the evanescent wave, occurs on the surface of the fiber, and part of this light energy is absorbed by the detection fiber. The detection fiber conducts the captured evanescent wave energy to the photosensitive diode which measures it. The fibers are treated with a binding agent to attract and bind the analyte to be measured. As analyte molecules bind to the surface (as when the analyte glucose binds to immobilized lectins such as concanavalin A or to immobilized anti-glucose antibodies) the amount of transient wave coupling between the two fibers changes, and the amount of energy captured by the detection fiber and measured by the diode, changes likewise. Several measurements of detected fleeting wave energy over short periods of time contribute to a rapid determination of the parameters that describe the equilibrium curve, and thus make calculation of the concentration of analyte in the body possible. Test results showing measurable current within 5 minutes (FIGS. 3A and 3B of US Patent 5,445,611) with this system indicate that enough data for an accurate final reading is collected within 5 minutes.

In its most basic embodiment, a device that can be used for the delivery of sound energy and collection of analyte comprises an absorbent layer, either of natural or synthetic material, which serves as a reservoir for the chemical amplifier, if used, and for receiving the analyte from the skin surface . The layer or reservoir is held in place, either passively or by means of suitable fasteners, such as a strap or adhesive tape, on the selected area of the skin surface.

A sound energy transducer is placed so that the layer or reservoir is between the skin surface and the transducer and held in place by suitable devices. A power unit is connected to the transducer and activated by switch devices or other suitable mechanism. The transducer is activated to emit sound energy modulated with regard to frequency, phase or intensity, as desired, to emit the chemical amplifier, if used, from the reservoir through the skin surface, followed by collection of the analyte from the skin surface into the reservoir. After the desired fixed or variable period of time, the transducer is deactivated. The layer or reservoir now containing the analyte of interest can be removed to quantify the analyte, for example in a laboratory using a variety of conventional chemical assays or with a transportable device. Alternatively, the mechanism for quantifying the analyte can be built into the device used for collecting the analyte, either as an integral part of the device or as a separate accessory. Devices for monitoring an analyte are described in US patent 5,458,140.

Example 24

An alternative method for the detection of an analyte, such as glucose, after sample collection through the porous skin surface as described above, can be achieved using enzymatic techniques. There are several enzymatic methods for measuring glucose in a biological sample. One method involves oxidation of glucose in the sample with glucose oxidase to develop gluconolactone and hydrogen peroxide. In the presence of a colorless chromogen, the hydrogen peroxide is converted by peroxidase to water and a colored product.

The intensity of the colored product will be proportional to the amount of glucose in the liquid. This color can be determined by using conventional absorbance or reflectance methods. When calibrating with known concentrations of glucose, the amount of color can be used to determine the concentration of glucose in the collected analyte. Through tests to determine the ratio, the concentration of glucose in the person's blood can be calculated. This information can then be used in the same way as the information obtained from a blood glucose test from a finger prick. Results can be available within 5 to 10 minutes.

Example 25

Any system that utilizes a visual display or reading of the glucose concentration will inform the diagnostician or patient of the need for administration of insulin or another appropriate drug. In crisis treatment or other situations where constant monitoring is desired and where corrective action must be taken almost simultaneously, the display device can be connected to suitable signaling devices which appropriately trigger the administration of insulin or another drug. For example, there are insulin pumps that are implanted in the abdominal cavity or other body cavities, which can be activated in response to external or internal stimuli. By utilizing the improved transdermal flow rates that are possible by microporation of the stratum corneum and other techniques in this description, it will alternatively be possible to establish an insulin delivery system transdermally under the control of the flow rates that are modulated by the signal from the glucose sensing system. In this way, a complete biomedical control system, which not only monitors and/or diagnoses a medical need, but also provides for corrective measures, can be created.

Biomedical control systems of a similar nature will be able to be produced for other situations, e.g. maintenance of correct electrolyte balance or administration of analgesics in response to a measured analyte parameter, such as prostaglandins.

Example 26

Like audible sound, sound waves can be subject to reflection, bending and absorption when they encounter another medium with different properties [D. Bommannan et al., 9 Pharm. Res. 559 (1992)]. Reflectors or lenses can be used to focus or otherwise control the distribution of sound energy in a relevant tissue. In many places on the human body there can be a fold of flesh that can hold this system. For example, an earlobe is an appropriate location that would allow the use of a reflector or lens to assist in the exercise of directional control (e.g., "puffing" analytes or permeants through the porous stratum corneum) in a manner similar to what is realized by changing sound frequency and intensity.

Example 27

Several sound energy transducers can optionally be used to selectively control the direction of transdermal flow through the porous stratum corneum either into the body or out of the body. A skin fold, such as an earlobe, makes it possible to place transducers on one side or the other of the fold. The transducers can be activated selectively or phase-directed to improve transdermal flow in the desired direction. Arrays of transducers, or an acoustic circuit, can be designed to use phased array concepts similar to those developed for radar and microwave communication systems to direct and focus the sound energy into the area of interest.

Example 28

In this example, the procedure of Example 19 was followed, except that the heat-separated epidermis sample was first treated with an excimer laser (eg, model EMG/200 from Lambda Physik; 193 nm wavelength, 14 ns pulse width) to melt away stratum corneum according to the methods described in US patent 4,775,361.

Example 29

In this example, the procedure in Example 19 was followed, except that the heat-separated epidermis sample was first treated with 1,1'-diethyl-4,4'-carbocyanine iodide (Aldrich, A,max=703 nm), after which a total 70 mJ/cm<2>/50 ms is delivered to the dye-treated sample with a model TOLD9150 diode laser (Toshiba America Electronic, 30 mW at 690 nm) to melt away the stratum corneum.

Example 30

In this example, the procedure of Example 29 was followed, except that the dye is indocyanine green (Sigma Cat. No. I-2633; Xmax=775 nm) and the laser is a model Diolite 800-50 (LiCONiX, 50 mW at 780 nm ).

Example 31

In this example, the procedure of Example 29 was followed, except that the dye is methylene blue and the laser is a model SDL-8630 (SDL Inc.; 500 mW at 670 nm).

Example 32

In this example, the procedure in Example 29 was followed, except that the dye is included in a solution comprising a permeation enhancer, e.g. CE1.

Example 33

In this example, the procedure in Example 29 was followed, except that the dye and enhancer-containing solution are delivered to the stratum corneum by exposure to ultrasound.

Example 34

In this example, the procedure of Example 31 was followed, except that the pulsed light source is a short arc lamp emitting over the broad range from 400 to 1100 nm, but having a bandpass filter placed in the system to limit the output wavelength to about 650 to 700 nm.

Example 35

In this example, the procedure of Example 19 was followed, except that the heat-separated epidermis samples were first punctured with a microlancet (Becton Dickinson) calibrated to produce a micropore in the stratum corneum without reaching the underlying tissue.

Example 36

In this example, the procedure of Example 19 was followed, except that the heat-separated epidermis samples were first treated with focused sound energy in the range of 70-480 mJ/cm<2>/50 ms to melt away the stratum corneum.

Example 37

In this example, the procedure in Example 19 was followed, except that the stratum corneum is first hydraulically punctured with a high-pressure fluid jet to form a micropore with a diameter of up to 100 µm.

Example 38

In this example, the procedure in Example 19 was followed, except that the stratum corneum is first punctured with short electrical pulses to form a micropore with a diameter of up to approx. 100^m.

Example 39

Acoustic flow

A new mechanism and application of sound energy for the delivery of therapeutic substances into the body and/or withdrawal of fluids from the inside of the body into an external reservoir through microporations formed in the stratum corneum layer will now be described. A further aspect of the present invention is the use of sound energy to produce an acoustic flow effect on the fluids that flow around and between the intact cells of the epidermis and dermis of the human skin. Acoustic flow is a well-documented way in which sound energy can interact with a liquid medium. Nyborg, Physical Acoustics Principles and Methods, pp. 265-331, Vol. II part B, Academic Press, 1965. The first theoretical analysis of acoustic flow phenomenon was made by Rayleigh (1884, 1945). In a comprehensive treatment of the subject, Longuet-Higgins (1953-1960) has given a result applicable to two-dimensional flow resulting in the immediate vicinity of any vibrating cylindrical surface. A three-dimensional approximation for an arbitrary surface was developed by Nyborg (1958). As described by Fairbanks et al., 1975 Ultrasonics Symposium Proceedings, IEEE Cat. No. 75, CHO 994-4SU, sound energy and the acoustic flow phenomenon can be of great help in accelerating the fluid flow through a porous medium, exhibiting measurable increases in the flow rate up to 50 times what is possible passively or only with the application of pressure gradients .

All previous efforts at transdermal delivery or extraction using ultrasound have focused on methods of interaction between the sound energy and skin tissue to permeabilize the stratum corneum layer. The exact mode of interaction involved has been thought to be solely due to the local temperature rise in the SC layer, and the resulting melting of the lipid regions in the intercellular openings between the corneocytes (Srinivasan et al.). Other researchers have suggested that micro-cavitations and/or shearing effects in the structure of the stratum corneum open channels through which fluids can flow more easily. The design of the acoustic systems for enhancing transdermal flow rates has generally been based on the prior recognition that the use of an existing therapeutic ultrasound device designed to produce a "depth-warming" effect when applied to a subject in conjunction with local application of a gelatinized or liquid preparation containing the drug to be delivered into the body could produce a quantifiable increase in the rate of flow of the drug into the body. In connection with the method described here to produce micropores in this barrier layer, the application of sound energy can now be seen as a completely new and different way to the classically defined sonophoresis concepts.

On the basis of the experimental discovery mentioned in US patents 5,458,140 and 5,445,611 that when there is or was created, a small hole in the stratum corneum (SC) in the Franz cells that were used in the in vitro studies, the application of a properly operated ultrasonic transducer against the liquid reservoir on one side of the porous SC sample could produce an "acoustic flow", where the liquid at high flow rates could be pumped through this porous membrane.

With the method described here to produce the controlled microporations in the stratum corneum layer of the skin of a living individual, the application of the fluid flow mode of sound/fluid interaction to the induction of fluid flow into or out of the body can now be practically explored. For example, clinical investigations have shown that by producing an array of four 80 µm diameter micropores in a 400 µm square and then applying a weak suction (250 to 300 mm Hg) to the area, an average of approx. 1 \ A. interstitial fluid is induced to leave the body for external collection in an external chamber. By adding to the system a small low-power sound transducer, configured to actively generate inwardly converging concentric circular pressure waves in the 2 to 6 mm of tissue surrounding the puncture site, it has been shown that this ISF flow rate can be increased by 50%.

A desire to produce some form of direct absorption of sound energy in the skin tissue (which is required to produce heating), sound energy frequencies can be determined for which the tissue is practically transparent, that is, the very low frequency range from 1 kHz to 500 kHz. Even at some of the lowest frequencies investigated, significant acoustic flow effects could be observed by using a microscope to observe an in vivo test where the subject's skin was microporated and the ISF was caused to leave the body and accumulate on the skin surface. Activation of the sound transducer showed startling indications of the degree of acoustic flow, as small pieces of particulate matter were transported along with the ISF as it swirled away. Typical degree of motion produced can be described as follows: for a 3 mm diameter circular droplet of ISF on the skin surface, a single visible particle could be seen to complete approx. three complete revolutions per second. This corresponds to a linear fluid velocity of more than 2.5 mm/second. This entire effect was demonstrated with tissue sound power levels of less than 100 mW/cm<2>.

While one can easily see the top of the skin surface and the fluid activity on it, an assessment of what dynamically takes place in the skin tissue layers in response to the inclusion of sound energy in these tissues is much more difficult. One can assume that if such high fluid velocities (e.g. >2.5 mm/sec.) can be so easily induced on the surface, a noticeable increase in the fluid flow in the intercellular channels occurring in the living skin tissues would also be realized as response to this sound energy input. A calculation has currently been made of an increase in extracted ISF through a given set of microporations when low-frequency sound energy was applied to the area in a circle surrounding the perforation site. In this experiment, an ISF collection technique based solely on mild suction (250 to 300 mm Hg) was alternately used with the exact same apparatus but with the sound transducer engaged. During a series of 10 two-minute collection periods, five with only suction and five with both suction and active sound energy, it was recorded that upon activation of the sound source, approx. 50% more ISF in the same period of time. These data are shown in FIG. 30. This increase in ISF flow rate was achieved without the subjects reporting any increase in pain sensation as a result of the sound energy. The apparatus used for this experiment is illustrated in FIG. 31-33. The transducer device of FIG. 31-33 consists of a thick-walled cylinder of piezoelectric material with an internal diameter of approx. 8 mm and a wall thickness of 4 mm. The cylinder is polarized so that when an electric field is applied across the metallized surfaces of the outer and inner diameters, the wall thickness of the cylinder expands or contracts in response to the field polarity. In practice, this configuration results in a device that rapidly squeezes the tissue that has been sucked into the central hole, causing an inwardly directed radial acoustic flow effect on the fluids contained in those tissues. This inward acoustic flow is responsible for bringing more ISF to the microporation site in the center of the hole, where it can leave the body for collection outside.

A similar device shown in FIG. 34A-B were produced and tested and initially led to similar results. In FIG. The 34A-B version was an ultrasonic transducer manufactured by Zevex, Inc. Salt Lake City, Utah, modified by having a spatula-like extension added to the horn. A 4 mm hole was placed in the 0.5 mm thick spatula end of this extension. Upon activation, the main movement is longitudinal along the length of the spatula, resulting in a rapid movement essentially back and forth. The physical disturbance of the metal spatula caused by the placement of the 4 mm hole resulted in a very active but chaotic behavior with large displacement at this point. In use, the subject's skin was sucked up into this hole and the sound energy was then directed into the skin in a manner similar to that illustrated in FIG. 33.

The novel aspect of this new application of ultrasound lies in the following basic areas: 1. The function of ultrasound no longer needs to be focused on permeabilization of the SC barrier membrane as described by Langer, Kost, Bommannan and others. 2. A much lower frequency system can be used which has very little absorption in the skin tissue, but which can still create the desired fluid flow phenomenon in the intercellular passages between the epidermal cells that contain the interstitial fluid. 3. The way in which the interaction with the tissues and the fluids in this takes place is the so-called "streaming" (streaming) which is considered within the theory of sound to be a unique and different mode from the classic vibrational interactions which are capable of exerting shear forces on cell membranes and accelerating it passive diffusion process.

By optimizing the geometric configuration, frequency, power and modulations of the sound transducer, it has been shown that significant increases in fluid flow through porous skin sites can be achieved. The optimization of these parameters is done to explore the non-linearities that govern the fluid flow conditions in this microscope-scale environment. By using frequencies below 200 kHz, large liquid effects can be observed without any detectable heating or other negative effects on the tissue. The sound power levels required to produce these measurable effects are very low with average power levels generally below 100 milliwatts/cm<2>.

The examples above are therefore only representative of systems that can be used when utilizing ultrasound or ultrasound and chemical amplifiers when collecting and quantifying analytes for diagnostic purposes and for transdermal delivery of permeants. The invention is directed to the discovery that poration of the stratum corneum with subsequent correct application of ultrasound, especially when this is accompanied by the use of chemical amplifiers, enables noninvasive or minimally invasive transdermal determination of analytes or delivery of permeants. However, the invention is not only limited to the respective illustrations. There are numerous poration techniques and amplifier systems, some of which may work better than others, for the detection and extraction of specific analytes or for the delivery of permeants through the stratum corneum. However, within the guidelines presented here, the person skilled in the art can easily carry out certain experiments to achieve optimal poration, optimal amplifiers or optimal time, intensity and frequency of the applied ultrasound, as well as modulation of the frequency, amplitude and phase of applied ultrasound. The scope of the invention is therefore only limited by the subsequent claims and functional equivalents thereof.

Claims (28)

1. System for melting away the stratum corneum layer of the skin of an organism, comprising means (140) for melting away the skin (160) of the organism at a selected area to form at least one micropore, each micropore being 1-1000 µm in diameter and projecting to a desired depth, characterized in that the means (140) for melting away is a means for heating a selected area of the stratum corneum with a heat source (144) to melt away the stratum corneum, wherein the heat source (144) comprises a fixed element (148) in operative communication with a means (156) for modulating the temperature of said solid element (148) from a reduced temperature state to an elevated temperature state, wherein, in the elevated temperature state, said solid element (148) transfers thermal energy to the skin (160) such that the temperature of the skin surface at the point of contact is greater than 123°C, to melt away a portion of the skin (160), and wherein, in the reduced temperature state, said solid element (148) transfers less thermal energy to the skin (160) than is required to melt away the skin part (160). 2. System according to claim 1, characterized in that it further comprises a reservoir containing a permeant, said reservoir being of a size for placement to cover the large amount of micropores formed, wherein in use, the permeant passes out of the reservoir and into the formed at least one micropore. 3. System according to claim 1 or 2, characterized in that the heat source (144) is of a size to limit the contact of the heat source (144) with the skin (160) to an area approximately 1-1000 µm in diameter. 4. System according to claims 1, 2 or 3, characterized in that in the reduced temperature state, the heat source (144) does not transfer conductive thermal heat to the skin (160) so that the temperature of the skin surface at the contact point is reduced to approximately ambient skin temperature. 5. System according to claim 1, characterized in that the heat source (144), in the elevated temperature state, has a temperature that is less than 100 °C when it is placed in contact with tissue with more than 30% water content. 6. System according to claims 1, 2, 3 or 4, characterized in that the heat source (144) is cycled through at least one or more cycles of modulations of the temperature between the reduced temperature state and the elevated temperature state. 7. System according to claim 6, characterized in that each cycle for modulations of the temperature is less than between approximately 20 to 80 msec. 8. System according to claims 1, 2, 3, 4 or 6, characterized in that the contact surface of the fixed element (148) is modulated from the reduced temperature state to the elevated temperature state, one pulse-on time, for a time period that is greater than approximately 1 msec. 9. System according to claim 8, characterized in that the pulse-on time is in a time period between approximately 1 to 50 msec. 10. System according to claim 9, characterized in that the fixed element (148) is modulated from the elevated temperature state to the reduced temperature state, a pulse-off time, for a time period that is higher than 10 msec. 11. System according to claim 8, characterized in that the pulse-off time is a time period between approximately 10-50 msec. 12. System according to claim 1 or 2, characterized in that the solid element (148) comprises a metallic solid. 13. System according to claims 1, 2, 3, 4, 5, 6, 8 or 10, further comprising a sensing system (250) which makes it possible to find whether the melting away of the skin (160) has formed at least one micropore with the desired depth in the skin (160). 14. System according to claim 13, characterized in that the sensing system (250) comprises a system for measuring changes in electrical impedance in the skin when at least one pore is formed and in which said sensing system (250) is operatively connected to the melting away means (140) to interrupt the melting away of the skin (160) when the micropore reaches the desired depth. 15. System according to claims 1, 2, 3, 4, 6, 8 or 10, characterized in that said fixed element (148) is placed in direct contact with part of the selected area of the skin (160). 16. System according to claim 1, 2, 3, 4, 6, 8 or 15, characterized in that said modulating means (156) comprises a modulating current loop which modulates the current supplied to said fixed element (148) so that said fixed element (148) is maintained substantially at the elevated temperature condition for less than about 20 msec to reduce sensation to the organism and minimize the volume of tissue affected by the transferred thermal energy. 17. System according to claim 16, characterized in that in the reduced temperature state the modulating current loop removes the current from said fixed element (148). 18. System according to claims 1, 2, 3, 4, 6, 8 or 15, characterized in that said modulating means (156) comprises a modulating magnetic field, the magnetic field being selectively activated to produce electric eddy currents within said fixed element (148) so that the fixed element (148) is maintained substantially at the elevated temperature condition for less than about 20 msec to reduce sensation to the organism and to minimize the volume of tissue affected by the transferred thermal energy. 19. System according to claims 1, 2, 3, 4, 6, 8, 15, 16 or 18, characterized in that it further comprises a means (124) for cooling the selected skin area (160) so that the selected area is in a cooled state. 20. System according to claim 1, 2, 3, 4, 6, 8, 15, 16, 18 or 19, characterized in that the means (124) for cooling comprises the transfer of a Peltier device to said selected area. 21. System according to claims 1, 2, 3, 4, 6, 8, 15, 16, 18 or 19, characterized in that it further comprises a means for applying sonic energy to a selected skin area (160) with at least one micropore therein by a frequency in the range of 5kHz to 100 MHz. 22. System according to claim 21, characterized in that said means for applying sonic energy is modulated by frequency modulation, amplitude modulation, phase modulation or combinations thereof. 23. System according to claims 1, 2, 3, 4, 6, 8, 15, 16, 18, 19 or 21, characterized in that it further comprises a means for applying electroporation to a selected skin area (160) with at least one micropore value. 24. System according to claim 1, characterized in that the fixed element (148) applies conductive energy. 25. System according to claims 1, 2, 3, 4, 6, 8, 15, 16, 18, 19 or 23, characterized in that the means (140) for melting causes exudation of interstitial fluid, and further comprises means for collecting a selected amount of said interstitial fluid. 26. System according to claims 1, 2, 3, 4, 6, 8, 15, 16, 18, 19, 21, 23 or 25, characterized in that it further comprises a means for applying suction to the formed at least one micropore. 27. System according to claims 1, 2, 3, 4, 6, 8, 15, 16, 18, 19, 21, 23 or 24, characterized in that it further comprises a means for contacting the selected area with a composition comprising an effective amount of tattoo ink so that the flux of said tattoo ink in the body is increased. 28. System according to claim 14, characterized in that the sensing system (250) detects a predetermined change in electrical impedance in the skin (160), and wherein, upon achieving the predetermined change, the melting away agent (140) stops.