JP7125117B2 - Method for thermally treating biological tissue using a pulsed energy source - Google Patents

Description

The present invention is generally directed to methods for thermally treating biological tissue. In particular, the present invention is directed to methods of applying a pulsed energy source to living tissue to stimulate activation of heat shock proteins and promote protein repair without tissue damage.

The inventors have determined the average temperature rise of the living tissue over a period of minutes to controllably raise the temperature of the living tissue up to a predetermined temperature range while not permanently damaging the target tissue. It has been found that maintaining below the level has a therapeutic effect on living tissue, especially damaged or diseased living tissue. Raising tissue temperature in such a controlled manner is believed to selectively stimulate the activation of heat shock proteins and/or the production and promotion of protein repair, which may be useful for therapeutic treatment of tissue. serves as a mechanism for

Heat shock proteins (HSPs) are a family of proteins produced by cells in response to exposure to stressful conditions. High levels of heat shock protein production are triggered by exposure to various types of environmental stress conditions such as infection, combustion, exercise, exposure of cells to toxins, starvation, hypoxia or water deprivation. obtain.

Heat shock proteins are known to play a role in responding to many abnormal conditions in the body's tissues, including viral infection, combustion, malignant transformation, exposure to oxidants, cytotoxins, and anoxia. It is The function of some heat shock proteins as intracellular chaperones for other proteins and members of the HSP family is essential for their maintenance and monitoring of cellular proteins under non-stressful conditions. It is expressed or activated at low to moderate levels because of its important role. These activities are part of the cell's own repair system, called the cell's stress or heat shock response.

Heat shock proteins are typically named according to their molecular weight. For example, Hsp60, Hsp70, and Hsp80 refer to a family of heat shock proteins with sizes of about 60, about 70, and about 80 kilodaltons, respectively. These work in many different ways. For example, Hsp70 has peptide binding and ATPase domains that stabilize the protein structure in its unfolded, ready-to-assemble state. Mitochondrial Hsp60s form ring-like structures that encourage protein assembly to their native state. Hsp90 plays a regulatory role in repressor genes by binding to cellular tyrosine kinases, transcription factors, and glucocorticoid receptors. Hsp27 suppresses protein aggregation.

Hsp70 heat shock proteins are members of the extracellular and membrane-bound heat shock proteins involved in antigen binding and presentation of antigens to the immune system. Hsp70 has been shown to inhibit the activity of the influenza A virus ribonucleoprotein and block viral replication. Tumor-derived heat shock proteins induce specific protective immunity. Experimental and clinical observations indicate that heat shock proteins are involved in the regulation of autoimmune arthritis, type 1 diabetes mellitus, arteriosclerosis, multiple sclerosis, and other autoimmune responses. rice field.

In response, the target tissue temperature is selectively and controllably raised to a predetermined temperature range for a short period of time while maintaining the average temperature elevation of the target tissue at the predetermined temperature for a longer period of time. It is believed to be advantageous to be able to This is believed to trigger the heat shock response to increase the number or activity of heat shock proteins in body tissues in response to infection or other abnormalities. However, this must be done in a controlled manner so as not to damage or destroy the tissue or body area being treated. The present invention satisfies these needs and provides other related advantages.

The present invention is directed to methods for thermally treating living tissue by applying a pulsed energy source to the target tissue to therapeutically treat the target tissue. Pulsed energy sources have energy parameters including wavelength or frequency, duty cycle, and pulse train duration. The energy parameters are selected to raise the target tissue temperature up to 11° C. to achieve a therapeutic effect, where the average temperature increase of the tissue over several minutes does not permanently damage the target tissue. maintained below a predetermined level.

The energy source parameters may be selected such that the target tissue temperature is raised to between approximately 6° C. and 11° C. at least during application of the pulsed energy source to the target tissue. The increase in average temperature of the target tissue over several minutes is maintained at 6°C or less, such as 1°C or less over several minutes.

Pulsed energy source energy parameters are selected such that approximately 20-40 Joules of energy are absorbed into each cubic centimeter of target tissue. Application of a pulsed energy source to the target tissue induces a heat shock response and stimulates activation of heat shock proteins in the target tissue without damaging the target tissue.

The device may be inserted into a body cavity to apply pulsed energy to tissue. Pulsed energy may be applied to an outer region of the body that is adjacent to the target tissue or that has a blood supply near its surface.

Pulsed energy sources may also include radio frequencies. The radio frequency may be between approximately 3-6 megahertz (MHz). It may have a duty cycle of between approximately 2.5% and 5%. It may have a pulse train duration of approximately between 0.2-0.4 seconds. Radio frequencies may be generated with a device having a coil radius of approximately 2-6 mm and an ampere-turn of approximately 13-57.

Pulsed energy sources may include microwave frequencies of 10-20 gigahertz (GHz). Microwaves may have a pulse train duration of approximately 0.2-0.6 seconds. Microwaves may have a duty cycle of approximately 2% to 5%. Microwaves may have an average power of approximately 8-52 Watts.

A pulsed energy source may also include a pulsed beam of light, such as laser light. The pulsed light beam may have a wavelength of approximately 530 nm to 1300 nm, more preferably 800 nm to 1000 nm. A pulsed beam may have a power of approximately 0.5-74 Watts. The pulsed light beam has a duty cycle of less than 10%, preferably between 2.5% and 5%. A pulsed beam may have a pulse train duration of approximately 0.1-0.6 seconds.

Pulsed energy sources may also include pulsed ultrasound. Ultrasound has a frequency of approximately 1-5 MHz. Ultrasound has a train duration of approximately 0.1-0.5 seconds. Ultrasound may have a duty cycle of approximately 2% to 10%. Ultrasound has a power of approximately 0.46-28.6 Watts.

Other features and advantages of the invention will become apparent from the following detailed description, taken in conjunction with the accompanying drawings, which illustrate, by way of example, the principles of the invention.

The following accompanying drawings illustrate the invention.
FIG. 4 is a graph illustrating the average power of a laser source compared to the radius of the laser source and the pulse train duration of the laser; FIG. FIG. 4 is a graph illustrating the average power of a laser source compared to the radius of the laser source and the pulse train duration of the laser; FIG. 4 is a graph illustrating time versus temperature decay as a function of laser source radius and wavelength; 4 is a graph illustrating time versus temperature decay as a function of laser source radius and wavelength; 4 is a graph illustrating peak ampere turns for various radio frequencies, duty cycles, and coil shape radii; 4 is a graph illustrating peak ampere turns for various radio frequencies, duty cycles, and coil shape radii; 4 is a graph illustrating peak ampere turns for various radio frequencies, duty cycles, and coil shape radii; 4 is a graph illustrating peak ampere turns for various radio frequencies, duty cycles, and coil shape radii; FIG. 4 is a graph depicting time for temperature rise to decay as compared to the radio frequency coil radius. Fig. 3 is a graph depicting average microwave power compared to microwave frequency and pulse train duration; Fig. 3 is a graph depicting average microwave power compared to microwave frequency and pulse train duration; Fig. 3 is a graph depicting time versus temperature decay for various microwave frequencies; FIG. 4 is a graph depicting average ultrasound source power compared to frequency and pulse train duration; FIG. Fig. 3 is a graph depicting time of temperature decay for various ultrasonic frequencies; Fig. 3 is a graph depicting time of temperature decay for various ultrasonic frequencies; FIG. 4 is a graph depicting the volume of the focal heating area compared to ultrasound frequency; FIG. FIG. 4 is a graph comparing the equation of temperature divided by pulse duration for an ultrasonic energy source; FIG. FIG. 10 is a graph illustrating the logarithmic amplitude of injury and HSP activation Arrhenius integrals as a function of temperature and pulse duration. FIG. FIG. 10 is a graph illustrating the logarithmic amplitude of injury and HSP activation Arrhenius integrals as a function of temperature and pulse duration. FIG. Fig. 2 is a schematic diagram of a light generating unit for generating a time sequence of pulses according to the present invention, comprising an optical waveguide extending therefrom; 1 is a cross-sectional view of a photostimulation delivery device for delivering electromagnetic energy to target tissue in accordance with the present invention; FIG. 1 is a schematic diagram illustrating a system used to generate a laser beam in accordance with the present invention; FIG. 1 is a schematic diagram of an optical system used to generate a geometric pattern of laser light according to the present invention; FIG. FIG. 10 is a schematic diagram illustrating an alternative embodiment of a system used to generate a laser beam for treating tissue in accordance with the present invention; FIG. 4 is a schematic diagram illustrating yet another embodiment of a system used to generate laser light to treat tissue in accordance with the present invention; 1 is a cross-sectional schematic view of the end of an endoscope inserted into a nasal cavity to treat tissue in accordance with the present invention; FIG. 1 is a schematic partial cross-sectional view of a bronchoscope extending through the trachea into the bronchi of the lungs and performing a procedure there in accordance with the present invention; FIG. 1 is a schematic illustration of a colonoscope providing light stimulation to an intestinal or colonic region of the body in accordance with the present invention; FIG. 1 is a schematic illustration of an endoscope being inserted into and performing a procedure on the stomach in accordance with the present invention; FIG. 1 is a partially sectioned perspective view of a capsule endoscope used in accordance with the present invention; FIG. 1 is a schematic diagram of pulsed high intensity focused ultrasound high intensity for treating tissue inside the body in accordance with the present invention; FIG. 1 is a schematic diagram for providing therapy to a patient's bloodstream through an earlobe in accordance with the present invention; FIG. 1 is a cross-sectional view of a stimulation therapy device of the present invention used in delivering light stimulation to blood through an earlobe in accordance with the present invention; FIG.

As shown in the accompanying drawings and as described more fully herein, the present invention raises tissue temperature to a sufficient level to achieve a therapeutic effect in a short period of time while permanently increasing tissue temperature. Laser, ultra-high energy parameters selected to induce a thermal time course in tissue to maintain an average tissue temperature below a predetermined level for an extended period of time to avoid excessive tissue damage. Systems and methods for delivering pulsed energy sources such as sound waves, ultraviolet radio frequencies, and microwave radio frequencies are of interest. Forming a thermal time course is believed to stimulate the activation or production of heat shock proteins and promote protein repair without causing any damage.

The inventors of the present invention have found that electromagnetic radiation in the form of laser light of various wavelengths can be applied to retinal tissue in a manner that does not destroy or damage retinal tissue, while having beneficial effects on eye diseases. discovered that it is possible. It is believed that this may be due, at least in part, to stimulation and activation of heat shock proteins and promotion of protein repair in retinal tissue. No. 14/607,959 filed Jan. 28, 2015; U.S. Patent Application No. 13/798,523 filed Mar. 13, 2013; No. 13/481,124, the contents of which are incorporated herein by reference as if fully referenced.

The inventors have discovered a laser beam that is therapeutic but still sublethal to the cells of the retinal tissue and thus does not disrupt photocoagulation in the retinal tissue resulting in prophylactic and protective treatment of the retinal tissue of the eye. discovered that it is possible to generate The various parameters of the light beam must be taken into account and selected so that the combination of selected parameters achieves a therapeutic effect without permanently damaging tissue. These parameters include laser wavelength, laser source radius, average laser power, total pulse duration, and pulse train duty cycle.

The choice of these parameters may be determined by requiring greater than Arrhenius 1 for HSP activation. The Arrhenius integral is used to analyze the effects of actions on living tissue. See, for example, The CRC Handbook of Thermal Engineering, ed. See Frank Kreith, Springer Science and Business Media (2000). At the same time, the parameters selected should not permanently damage the tissue. Therefore, we may use the Arrhenius integral for damage, and the solved Arrhenius integral is less than one. Alternatively, FDA/FCC constraints on energy deposition per gram of tissue and temperature rise as measured over minutes must be met to avoid permanent tissue damage. For example, FDA/FCC requirements for energy deposition and temperature rise are commonly used, eg, for electromagnetic sources, see www. fda. gov/medicaldevices/deviceregulationandguidance/guidancedocuments/ucm073817. See Anastosio and P.A. for ultrasound sources at http://attacha. La Rivero, ed. , Emerging Imaging Technologies. See CRC Press (2012). Generally speaking, tissue temperature elevations between 6°C and 11°C can produce therapeutic effects, such as by activating heat shock proteins, and under certain circumstances, for extended periods of time, e.g. By maintaining the average tissue temperature below a predetermined temperature, such as below 6° C. and 1° C., for several minutes, such as 6 minutes, permanent tissue damage is not achieved.

We have found a subthreshold and sublethal micropulse laser beam having a wavelength greater than 532 nm and a duty cycle of less than 10% at a predetermined intensity or power and a predetermined pulse width or duration of exposure. creates the desired retinal light stimulation without visible burn areas or tissue disruption. Specifically, it has a wavelength of 550 nm-1300 nm, especially in a preferred embodiment, 810 nm-1000 nm, a duty cycle of approximately 2.5%-5%, and a predetermined intensity or power (retina 100-590 watts per square centimeter (or approximately 1 watt per laser spot for each treatment spot on the retina) and a predetermined pulse width or exposure duration (such as 100-600 milliseconds or less). , forms a sub-lethal “precise subthreshold” retinal photostimulation, where all areas of the retinal pigment epithelium exposed to laser irradiation are protected and available to contribute therapeutically. In other words, the inventors elevate retinal tissue to at least therapeutic levels, but below cell or tissue lethal levels, without destroying, burning, or otherwise damaging the retinal tissue. It has been found that by allowing , the advantages of the halo effect of conventional methods are recreated. This is referred to herein as subthreshold diode micropulse laser therapy (SDM).

SDM does not produce laser-induced retinopathy (photocoagulation), has no known adverse treatment effects, and is associated with many retinopathies (diabetic macular edema (DME), proliferative diabetic retinopathy ( PDR), macular edema due to branch retinal vein occlusion (BRVO), central serous chorioretinopathy (CSR), reversal of drug resistance, and dry age-related macular degeneration, Stargardt disease, cone dystrophy, and retinal pigment It has been reported to be an effective treatment for prophylactic therapy of progressive degenerative retinopathies such as degeneration). The safety of SDM is such that it can be used transfoveally in eyes with 20/20 vision to reduce the risk of vision loss due to DME involving the early fovea. is.

A mechanism by which SDM may act is the generation or activation of heat shock proteins (HSPs). Despite the almost endless variety of possible cellular abnormalities, all cell types share a common, highly conserved repair mechanism: heat shock proteins (HSPs). HSPs are induced instantly, within seconds to minutes, by almost any type of cellular stress or injury. In the absence of lethal cell damage, HSPs are extremely effective in repairing and restoring living cells to a more normal functional state. HSPs are temporary, generally peaking in hours and lasting for days, but their effects can be long lasting. HSPs reduce inflammation, a common factor in many disorders.

Laser therapy can induce the generation or activation of HSPs and alter cytokine expression. The more sudden and severe the non-lethal cell stress (such as laser irradiation), the more rapid and robust the HSP activation. Therefore, sudden repetitive cold thermal spikes with very rapid rate of change (up to 7°C with each 100 μs micropulse, or 70,000°C/s) produced by each SDM exposure are , especially in stimulating HSP activation compared to non-lethal exposure to subthreshold treatment with continuous wave lasers, which may duplicate only low average tissue temperature rises.

Laser wavelengths below 550 nm produce increasingly cytotoxic photochemical effects. At 810 nm, SDM produces photothermal, rather than photochemical, cell stress. Therefore, the SDM can affect tissue without damaging it. Therefore, the clinical benefits of SDM are generated primarily by subpathological photothermal cellular HSP activation. In dysfunctional cells, HSP stimulation by SDM leads to normalized cytokine expression and consequently improved structure and function. The therapeutic effect of this "low-intensity" laser/tissue interaction is then amplified by "high-density" laser application, by densely/intensively treating large tissue areas, including all areas of pathology. Replenish all dysfunctional cells in the targeted tissue area, thereby maximizing treatment efficacy. These principles define the treatment strategies for SDM described herein.

Since normally functioning cells do not require repair, HSP stimulation in normal cells tends to have no significant clinical effect. The "patho-selectivity" of near-infrared laser effects, such as SDM, on a variety of cell types, affecting diseased cells but not normal cells, has been demonstrated in clinical practice of SDM. consistent with observations. SDM is reported to have a unique clinically broad therapeutic spectrum among retinal laser modalities, consistent with the predictions of the American National Standards Institute "Maximum Permissible Exposure." Although SDM can also cause direct photothermal effects such as unfolding and disaggregation of entropic proteins, SDM has been optimized for clinically safe and effective stimulation of HSP-mediated repair. It seems that

As noted above, while SDM stimulation of HSPs is non-specific with respect to disease processes, the outcome of HSP-mediated repair is due to its own properties specific to dysfunctional states. HSPs tend to fix whatever is wrong. Thus, the observed effects of SDM in retinal conditions as widely different as BRVO, DME, PDR, CSR, age-related and genetic retinopathies, and drug-resistant NAMD. Conceptually, this capability can be viewed as a kind of "Reset to Default" mode of SDM operation. For a wide range of disorders where cellular function is critical, the SDM triggers a "reset" (to "factory default settings") through HSP-mediated cellular repair. normalize cell function by subtraction,

The inventors have discovered that SDM treatment of patients suffering from age-related macular degeneration (AMD) can slow progression and even stop progression of AMD. Most of the patients showed significant improvement in dynamic functional logMAR mesopic visual acuity and mesopic contrastive visual acuity after SDM treatment. SDM is believed to act by targeting, protecting, and "normalizing" (approaching normality) the function of the retinal pigment epithelium (RPE).

SDM has also been shown to stop or reverse the symptoms of the diabetic retinopathy disease state, despite persistent systemic diabetes mellitus, without treatment-related damage or side effects. On this basis, SDM has been shown to induce a return to normal cellular function and cytokine expression in diabetic RPE cells (pressing the 'reset' button on electronic devices to restore factory settings). similar to). Based on the above information and studies, SDM treatment may directly affect cytokine expression through heat shock protein (HSP) activation in targeted tissues.

Because heat shock proteins play a role in responding to many abnormal conditions in body tissues other than those of the eye, similar systems and methods are advantageously used in treating such abnormal conditions, infections, and the like. believed to be possible. Therefore, the present invention is directed to the use of ultrasound or electromagnetic radiation to treat abnormal conditions, including inflammation, autoimmune diseases, and cancer, which are reachable by fiber optics in endoscopes or surface probes and by focused electromagnetic/sound waves. Intended for controlled application. For example, cancers on the surface of the prostate, which are most likely to metastasize, are accessible by the fiber optics of the rectoscope. Colon tumors are accessible by fiber optic systems such as those used in colonoscopy.

As shown above, subthreshold diode micropulse laser (SDM) photostimulation was effective in stimulating direct repair of slightly misfolded proteins in eye tissue. In addition to HSP activation, this can also occur in other ways. This is because the temperature spikes induced by micropulses in the form of thermal time courses allow diffusion of water inside the protein, which prevents the formation of peptide-peptide hydrogen bonds that prevent the protein from returning to its native state. This is because it enables cutting. Diffusion of water into the protein increases the number of inhibitory hydrogen bonds approximately 1000-fold. Therefore, it is believed that this process can be advantageously applied to other diseases as well.

As explained above, the energy source applied to the target tissue has energy and operating parameters that must be determined and selected to achieve a therapeutic effect while not permanently damaging the tissue. As an example, using a beam energy source such as a laser beam, the laser wavelength, duty cycle, and total pulse train duration parameters must be taken into account. Other parameters that can be considered include the radius of the laser source as well as the average laser power. Adjusting or selecting one of these parameters may have an effect on at least one other parameter.

FIGS. 1A and 1B show graphs of average power in watts comparing laser source radius (between 0.1 cm and 0.4 cm) and pulse train duration (0.1 to 0.6 seconds). exemplified. FIG. 1A shows a wavelength of 880 nm and FIG. 1B has a wavelength of 1000 nm. It can be seen from these figures that the power required decreases monotonically as the radius of the laser source decreases, the total train duration increases, and the wavelength decreases. A preferred parameter for the radius of the laser source is between 1 mm and 4 mm. For a wavelength of 880 nm, the power minimum is 0.55 Watts, the laser source radius is 1 mm, and the total pulse train duration is 600 ms. With a laser source radius of 4 mm and a total pulse train duration of 100 ms, the maximum power for a wavelength of 880 nm is 52.6 Watts. However, if we choose a laser with a wavelength of 1000 nm, the minimum power value is 0.77 Watts, the radius of the laser source is 1 mm, the total pulse train duration is 600 ms, and the laser With a source radius of 4 mm and a total pulse duration of 100 ms, the maximum power value is 73.6 Watts. The corresponding peak power during each pulse is obtained from the average power by dividing by the duty cycle.

The volume of the tissue portion heated is determined by the wavelength, the absorption length in the relevant tissue, and the beam width. The total pulse duration and average laser power determine the total energy delivered to warm the tissue, and the duty cycle of the pulse train gives the associated spike or peak, the power associated with the average laser power. Preferably, the pulsed energy source energy parameters are selected such that approximately 20-40 Joules of energy are absorbed into each cubic centimeter of target tissue.

The absorption length is very small in the thin melanin layer in the pigment epithelium of the retina. In other parts of the body the absorption length is generally not very small. For wavelengths ranging from 400 nm to 2000 nm, the penetration depth and skin range from 0.5 mm to 3.5 mm. The penetration depth into human mucous tissue ranges from 0.5 mm to 6.8 mm. Accordingly, the heated volume is confined to the outer or inner surface on which the radiation source is placed, where the depth equals the penetration depth and the lateral dimension equals the lateral dimension of the radiation source. . Since the light energy source is used to treat diseased tissue near an outer surface or near an internally accessible surface, a source having a radius of 1 mm to 4 mm and operating at a wavelength of 880 nm A penetration depth of approximately 2.5 mm is provided, and a source operating at a wavelength of 1000 nm provides a penetration depth of approximately 3.5 mm.

For a short period of time, such as less than 1 second, to maintain the average temperature of the target tissue in a low temperature range, such as less than 6° C. or less than 1° C., for an extended period of time, such as minutes, while producing the therapeutic effects of the present invention. It has been found that target tissue can be heated up to about 11° C. over a period of time. The selection of duty cycle and total pulse train duration provides a time interval during which heat can dissipate. A duty cycle of less than 10%, preferably between 2.5% and 5%, has been found to be effective, with a total pulse duration of 100 ms to 600 ms. 2A and 2B show the decay time from 10° C. to 1° C. for a laser source with a radius of 0.1 cm to 0.4 cm whose wavelength is 880 nm in FIG. 2A and 1000 nm in FIG. 2B. Illustrate. Although the decay time is shorter when using the 880 nm wavelength, either wavelength is within acceptable requirements and operating parameters to achieve the benefits of the present invention without causing permanent tissue damage. come in.

It has been found that an average temperature rise of the desired target area of at least 6° C. to a maximum of 11° C., and preferably increasing to approximately 10° C., during the total irradiation period results in HSP activation. Control of the target tissue temperature is determined by choosing source and target parameters whereby the Arrhenius integral for HSP activation is greater than 1 and at the same time damage or conservative to avoid damage Arrhenius integrals less than 1. Ensure compliance with FDA/FCC requirements.

To meet conservative FDA/FCC constraints that avoid permanent tissue damage, the average temperature rise of the target tissue over a period of 6 minutes is no more than 1°C for light beams and other sources of electromagnetic radiation. Figures 2A and 2B above illustrate the typical decay times required for the temperature in the heated target region to decrease by thermal diffusion from a temperature rise of approximately 10°C to 1°C. As seen in , the temperature decay time is 16 seconds when the wavelength is 880 nm and the source diameter is 1 millimeter. The temperature decay time is 107 seconds when the source diameter is 4 mm. As shown in FIG. 2B, when the wavelength is 1000 nm, the temperature decay time is 18 seconds when the source diameter is 4 mm, and the temperature decay time is 136 seconds when the source diameter is 1 mm. This is well within the time of the average temperature rise being maintained over several minutes, such as 6 minutes or less. Although the temperature of the target tissue is raised very quickly, such as a fraction of a second, to approximately 10° C., during application of the energy source to the tissue, a relatively low duty cycle is applied to the tissue. A relatively short pulse train duration provides a relatively long period between pulses of applied energy, and a relatively short pulse train duration provides sufficient temperature within a relatively short period of time, including minutes, such as six minutes or less, without permanent tissue damage. Ensure diffusion and attenuation.

Parameters vary with individual energy sources, including microwaves, infrared lasers, radio frequencies and ultrasound, because tissue absorption properties vary with these different types of energy sources. Although tissue water content can vary with tissue type, the uniformity of tissue properties observed in normal or near-normal conditions has allowed the publication of tissue parameters that are widely used by clinicians when designing treatments. be done. Below are tables illustrating the properties of electromagnetic waves in biological media, Table 1 for muscle, skin and tissue with high water content and Table 2 for fat, bone and tissue with low water content.

The absorption length of radio frequencies in body tissue is long compared to body length. As a result, the heating area is determined by the dimensions of the coil that is the radio frequency energy source rather than by the absorption length. At a long distance r from the coil, the magnetic (near) field from the coil drops off as ¹ /r3. At shorter distances, the electric and magnetic fields can be expressed in terms of magnetic vector potentials, which in turn can be expressed in closed form in terms of elliptic integrals of the first and second kind. Heating occurs only in areas comparable in size to the dimensions of the coil source itself. Therefore, if it is desired to preferentially heat areas characterized by radii, the source coils are chosen to have similar radii. Due to the ¹ /r3 decay of the magnetic field, the heating decays very rapidly outside the hemispherical region of the radius. It is reasonable to consider a coil radius of approximately 2-6 mm since it has been proposed to use radio frequencies in diseased tissue that are only available externally or intraluminally.

In addition to the radius of the source coil, the ampere-turns (NI) in the source coil give the magnitude and spatial extent of the magnetic field, and the radio frequency is the factor that relates the magnitude of the electric field to the magnitude of the magnetic field. Heating is proportional to the product of the conductivity and the square of the electric field. For target tissues of interest near the external or internal surface, conductivity is for skin and mucus tissue. In addition to the pulse train duty cycle, the total train duration of the pulse train is a factor that affects how much total energy is delivered to the tissue.

Preferred parameters for the radio frequency energy source are a coil radius of 2-6 mm, a radio frequency in the range of 3-6 MHz, a total pulse train duration of 0.2-0.4 seconds and between 2.5% and 5% was determined to be a duty cycle of Figures 3-6 show how the ampere-turns change as these parameters are varied to give a temperature increase that produces an Arrhenius integral of approximately 1 for HSP activation. With reference to FIG. 3, peak ampere-turns ( NI) is 13 at a coil radius of 0.2 cm and 20 at a coil radius of 0.6 cm. For a frequency of 3 MHz, the peak ampere-turn is 26 when the pulse train duration is 0.4 seconds, the coil radius is 0.6 cm, and the duty cycle is 5%, as illustrated in FIG. However, at the same 5% duty cycle, the peak ampere-turn is 40 when the coil radius is 0.2 cm and the pulse train duration is 0.2 seconds. A 2.5% duty cycle is used in FIGS. This yields 18 ampere-turns for 6 MHz radio frequency with a coil radius of 0.6 cm and a pulse train duration of 0.4 s, as illustrated in FIG. 5, with a coil radius of only 0.2 cm. , yielding 29 ampere-turns when the pulse train duration is 0.2 seconds. With reference to FIG. 6, at a duty cycle of 2.5% and a radio frequency of 3 MHz, the peak ampere-turn is 36 when the pulse train duration is 0.4 seconds and the coil radius is 0.6 cm. , and 57 when the pulse train duration is 0.2 seconds and the coil radius is 0.2 cm.

The time in seconds for the temperature rise to decay from approximately 10° C. to approximately 1° C. for coil radii between 0.2 cm and 0.6 cm is illustrated for a radio frequency energy source in FIG. The temperature decay time is approximately 37 seconds when the radio frequency coil radius is 0.2 cm and approximately 233 seconds when the radio frequency coil radius is 0.5 cm. When the radio frequency coil radius is 0.6 cm, the decay time is approximately 336 seconds, which is still within the decay time tolerance, but at its upper limit.

Microwaves are another electromagnetic energy source that can be utilized in accordance with the present invention. The microwave frequency determines the tissue penetration distance. The gain of the conical microwave horn is large compared to microwave wavelengths, which indicates that under these circumstances the energy is mostly radiated with a narrow forward load. Typically, microwave sources used in accordance with the present invention have a length dimension on the order of one centimeter or less, and thus are smaller than the wavelength, in which case the microwave source is approximated as a dipole antenna. )be able to. Such small microwave sources are easier to insert into internal body cavities and can also be used to radiate external surfaces. In that case, the heating area can be approximated by a hemisphere with a radius equal to the microwave absorption length in the body tissue being treated. Since microwaves are used to treat tissue near the surface available from external surfaces or internal cavities, frequencies in the range of 10-20 GHz are used, where corresponding penetration distances range from approximately 2 mm to Only between 4 mm.

The temperature rise of tissue using a microwave energy source is determined by the average power of the microwaves and the total pulse train duration. The duty cycle of a pulse train determines the peak power in a single pulse in the train of pulses. Resulting pulse train durations of 0.2 seconds and 0.6 seconds are preferred because the radius of the energy source is less than approximately one centimeter and frequencies between 10 and 20 GHz are typically used.

The required power decreases monotonically as the train duration increases and the microwave frequency increases. For a frequency of 10 GHz, the average power is 18 Watts when the pulse train duration is 0.6 seconds and 52 Watts when the pulse train duration is 0.2 seconds. For a microwave frequency of 20 GHz, an average power of 8 Watts is used when the pulse train is 0.6 seconds, and the average power can be 26 Watts when the pulse train duration is only 0.2 seconds. . The corresponding peak power is obtained by simply dividing the average power by the duty cycle.

Referring now to FIG. 8, a graph shows average microwave power in watts of microwaves with a frequency of 10 GHz and pulse train durations between 0.2 and 0.6 seconds. FIG. 9 is a similar graph, but showing the average microwave power for microwaves having a frequency of 20 GHz. Thus, it can be seen that the average microwave source power changes as the total train duration and microwave frequency change. However, the rule condition is that the Arrhenius integral for HSP activation in the heating region is approximately unity.

Referring to FIG. 10, a graph illustrates time in seconds for temperature decay from approximately 10° C. to approximately 1° C. compared to microwave frequencies between 58 MHz and 20000 MHz. The minimum and maximum temperature decay for the preferred range of microwave frequencies is 8 seconds when the microwave frequency is 20 GHz and 16 seconds when the microwave frequency is 10 GHz.

The use of ultrasound as an energy source allows for heating of tissues of varying depths within the body, including superficial and even deeper tissues. The absorption length of ultrasound within the body is fairly long, as evidenced by its widespread use for imaging. Accordingly, the ultrasound waves can be focused to a target region deep within the body, and the heating of the focused ultrasound beam is primarily concentrated in the generally cylindrical focal region of the beam. The heating region has a volume determined by the focal waist of the Airy disk and the length of the focal waist region, which is a confocal parameter. Multiple beams from sources at different angles can also be used, with heating occurring in overlapping focal regions.

For ultrasound, the relevant parameters for determining tissue temperature are the ultrasound frequency, total train duration, and transducer power when the focal length and diameter of the ultrasound transducer are considered. The frequency, focal length, and diameter determine the volume of the focal region in which the ultrasonic energy is concentrated. It is the focal volume that contains the target volume of tissue for treatment. Transducers with diameters of approximately 5 cm and focal lengths of approximately 10 cm are readily available. Advantageous focal size is achieved when the ultrasound frequency is 1-5 MHz and the total train duration is 0.1-0.5 seconds. For example, for a 10 cm focal length and a 5 cm transducer diameter, the focal volume is 0.02 cc at 5 MHz and 2.36 cc at 1 MHz.

Referring now to FIG. 11, a graph illustrates average source power in Watts compared to frequency (between 1 MHz and 5 MHz) and pulse train duration (0.1-0.5 seconds). A transducer focal length of 10 cm and a source diameter of 5 cm were assumed. The power required to give an Arrhenius integral for an HSP activation of approximately 1 decreases monotonically with increasing frequency and increasing total train duration. Considering the preferred parameters, the minimum power for a frequency of 1 GHz and a pulse train duration of 0.5 seconds is 5.72 Watts, while the maximum power for a frequency of 1 GHz and a pulse train duration of 0.1 seconds is 28.6 watts. For a frequency of 5 GHz, 0.046 Watts are required for a pulse train duration of 0.5 seconds and 0.23 Watts are required for a pulse train duration of 0.1 seconds. The corresponding peak power during each pulse is obtained simply by dividing by the duty cycle.

FIG. 12 illustrates the time in seconds for temperature diffusion or decay from approximately 10° C. to approximately 6° C. when the ultrasonic frequency is 1-5 MHz. FIG. 13 illustrates the time in seconds to decay from approximately 10° C. to approximately 1° C. for ultrasonic frequencies of 1-5 MHz. For a preferred focal length of 10 cm and a transducer diameter of 5 cm, the maximum time for temperature decay is 366 seconds when the ultrasonic frequency is 1 MHz, and the minimum temperature decay is when the microwave frequency is 5 MHz. 15 seconds. Since the FDA only requires that the temperature rise be less than 6°C for a test duration of minutes, a decay time of 366 seconds at 1 MHz is possible resulting in a 1°C rise over several minutes. As can be seen in Figures 12 and 13, the decay time to a 6°C rise is approximately 70 times much shorter than a 1°C rise.

FIG. 14 illustrates the volume of the focal heating area in cubic centimeters compared to ultrasound frequencies between 1 and 5 MHz. Considering ultrasound frequencies in the range of 1-5 MHz, the corresponding focal spot sizes for these frequencies range from 3.7 mm to 0.6 mm, with focal region lengths ranging from 5.6 cm to 1.2 cm. Range. Corresponding treatment volumes range between approximately 2.4 cc and 0.02 cc.

An example of parameters that take into account a desirable Arrhenius integral of HSP activation greater than 1 and an Arrhenius integral of damage less than 1 is 5.8-17 Watts total ultrasound power, 0.5 sec pulse duration, 5 sec , with a total of 10 pulses within a total pulse train time of 50 seconds. The target treatment volume will be approximately 1 mm on one side. By applying ultrasound in multiple, simultaneously applied, adjacent but separately spaced rows, larger treatment volumes can be treated with ultrasound systems analogous to laser-diffracted optical systems. could be. Multiple focused ultrasound beams converge on a very small treatment target within the body, allowing minimal heating other than overlapping beams at the target. This area is heated and a transient high temperature spike stimulates HSP activation and promotes protein repair. However, given the pulsatile aspect of the present invention, as well as the relatively small area being treated in a given period of time, the treatment is consistent with FDA/FCC requirements for long term (minutes) average temperature rise <1K. there is An important difference of the present invention from existing hyperthermia treatments of pain and muscle tension is that existing techniques lack high T spikes and efficiently activate HSPs to provide healing at the cellular level. and to promote protein repair.

As far as therapeutic HSP activation and protein repair promotion are concerned, the pulse train mode of energy delivery has distinct advantages over single pulse or stepwise modes of energy delivery. There are two considerations about this advantage:
First, a significant advantage for HSP activation and protein repair in the SDM energy delivery mode results from the generation of spike temperatures of approximately 10°C. This large temperature increase has a large impact on the Arrhenius integral, which quantitatively describes the number of HSPs that are activated, and the rate of water diffusion into proteins, which promotes protein repair. This is because temperature enters an exponential function which has a large amplifying effect.

It is important that the temperature rise does not stay at a high value (10°C or higher) for a long time, because this means that the average temperature rise must be less than 1°C (6°C for ultrasound) over a period of minutes. because it violates FDA and FCC requirements.

The SDM mode of energy delivery uniquely meets both of these previous considerations through judicious selection of power, pulse time, pulse interval, and volume of the target area to be treated. The temperature rises fairly rapidly by approximately 10°C so that the long-term average temperature rise does not exceed the long-term FDA/FCC limits of 6°C for ultrasonic frequencies and 1°C or less for electromagnetic radiation energy sources. The volume of the treatment area enters because it must decay from its high value of .

For a region of length dimension L, the time it takes for the peak temperature to e-fold in tissue is approximately L ² /16D, where D=0.00143 cm ² /sec is: It is a typical thermal diffusion coefficient. For example, for L=1 mm, the decay time is approximately 0.4 seconds. Thus, for a 1 mm area on the side, a train of 10 pulses, each 0.5 s in duration, with 5 s pulse spacing, exceeds an average long-term temperature rise of 1°C. The desired instantaneous high temperature rise can be achieved without This is further demonstrated below.

The reason RF electromagnetic radiation is not a good choice for deep body SDM type treatments as ultrasound is the limited amount of heating. Long skin depth (penetration distance) and ohmic heating along the entire skin depth results in a large amount of heating, whose thermal inertia activates HSPs and achieves high spike temperatures that promote protein repair. , and rapid temperature decay meeting long-term FDA and FCC limits for average temperature rise.

Ultrasound is already used to therapeutically heat areas of the body to relieve pain and muscle tension. However, the heating does not follow an SDM-type protocol and does not have temperature spikes responsible for HSP excitation.

Now consider a group of focused ultrasound beams directed at a target region deep within the body. To simplify the calculations, we assume that the beam is replaced by a single source with a spherical shape that is focused at the center of the sphere. The absorption length of ultrasound can be quite long. Table 3 below shows typical absorption coefficients for ultrasound at 1 MHz. The absorption coefficient is approximately proportional to frequency.

Assuming that the geometric deformation of the incident beam due to focusing dominates the deformation due to attenuation, the intensity of the incident ultrasound at a distance r from the focal point is calculated approximately as follows:
I(r)=P/(4πr ² ) [1]
where P denotes the total ultrasonic power. The temperature rise at the end of a short pulse of duration t _p at r is
dT(t _p )=Pαt _p /(4πC _v r ² ) [2]
where α is the absorption coefficient and _Cv is the specific heat capacity. This is true until r reaches a point where the thermal diffusion distance at _tp becomes comparable to r, or until the diffraction limit of the focused beam is reached. For shorter r, the temperature rise is essentially independent of r. As an example, assume that the diffraction limit is reached at a radial distance less than the distance determined by thermal diffusion. here,
r _dif =(4Dt _p ) ^1/2 [3]
where D is the thermal diffusion coefficient and for r<r _dif the temperature rise at _tp is:
If r<r _dif then dT(r _dif ,t _p )=3Pα/(8πC _v D)[4]
Therefore, at the end of the pulse, the temperature rise can be recorded:
dt _p (r)={Pαt _p /(4πC _v }[(6/r _dif ² )U{r _dif −r)+(1/r ² )U(r−r _dif )] [5]
Green's function for the heat diffusion equation:
G(r, t)=(4ΩDt) ^−3/2 exp[−r ² /(4Dt)][6]
was applied to this initial temperature distribution, the temperature dT(t) at the focus r=0 at time t was found to be:
dT(t)=[dT _o /{(1/2)+(π ^1/2 /6)}][(1/2)(t _p /t) ^3/2 +(π ^1/2 /6) (t _p /t)] [7]
with dT _o =3Pα/(8πC _v D) [8]

A good approximation to equation [7] is provided by, as can be seen in FIG.

これは、標的処置ゾーンでのｄＴ（ｔ）／ｄＴ_ｏに対する方程式［７］と［９］の比較である。下の曲線は方程式［９］の近似式である。Ｎパルスの列に対するアレニウス積分は、ここで、方程式［９］によって与えられた温度上昇を用いて評価することができる。この式では、以下であり：

This is a comparison of equations [7] and [9] for dT(t)/dT _o at the target treatment zone. The lower curve is an approximation of equation [9]. The Arrhenius integral for a train of N pulses can now be evaluated using the temperature rise given by equation [9]. This formula is:

式中、ｄＴ（ｔ－ｎｔ_Ｉ）は、ｔがｔ－ｎｔ_Ｉに取って代わり、ｔ_Ｉがパルスの間隔を示している、方程式［９］の式である。

where dT(t−nt _I ) is the expression of equation [9] where t replaces t−nt _I and t _I denotes the interval between pulses.

The Arrhenius integral can be approximately evaluated by dividing the integration interval into parts where the temperature spike occurs and parts without the temperature spike. Summing over the temperature spike contribution can be simplified by applying Laplace's end point formula to the integration over the temperature spike. Furthermore, the integral over the part when there are no spikes shows that the spike-free temperature rise reaches an asymptotic value very quickly, so that a good approximation can be obtained by exchanging the varying time rise for its asymptotic value. It can be simplified by noting that When these approximations are obtained, equation [10] becomes:
Ω=AN [{t _p (2k _B T _o ² /(3EdT _o )} exp [−(E/k _B )1/(T _o +dT _o +dT _N (Nt _I ))]
+exp [−(E/k _B )1/(T _o +dT _N (Nt _I ))]] [12]
where:

（方程式［１３］中の２．５は、（Ｎ－ｎ）^－３／２のｎに対する合計から生じ、対象の典型的なＮに対する調和数（Ｎ，３／２）の大きさである。）

(The 2.5 in equation [13] results from the summation of (N−n) ^−3/2 over n and is the magnitude of the harmonic number (N,3/2) over N typical of interest. )

It is interesting to compare this expression with that for SDM applied to the retina. The first term is very similar to that from the spike contribution in the retinal case, except that the effective spike spacing is reduced by a factor of 3 for this 3D focused beam case. The second condition is that it contains much less dT _N (Nt _I ) than in the retina. Here, the background temperature rise was comparable in magnitude to the spike temperature rise. However, here in the focused beam case, the background temperature rise is much smaller by the ratio (t _p /t _I ) ^3/2 . Since the background temperature rise, which is similar to that in the case of continuous ultrasonic heating, is insignificant compared to the spike contribution, this suggests that the spike contribution to the activation or generation of HSPs and promotion of protein repair is not significant. emphasizes its importance. At the end of the pulse train, even this low background temperature rise is quickly dissipated by thermal diffusion.

Figures 16 and 17 show injury and activation of HSPs as a function of dT _o for pulse duration _tp = 0.5 seconds (sec), pulse interval tI ₌ 10 seconds, and total number of pulses N = 10. or the logarithmic magnitude of the Arrhenius integral for generation. _{to damage and} to Logarithm of the Arrhenius integral [equation 12] for HSP activation. FIG. 16 shows the logarithm of the damage integral with the Arrhenius constant A=8.71×10 ³³ sec ⁻¹ and E=3.55×10 ⁻¹² ergs. FIG. 17 shows the logarithm of the HSP activation integral with the Arrhenius constant A=1.24×10 ²⁷ sec ⁻¹ and E=2.66×10 ⁻¹² ergs. The graphs in FIGS. 16 and 17 show that Ω _damage does not exceed 1 until dT 0 exceeds 11.3 K, while _{Ω hsp} exceeds 1 over the entire interval shown, which indicates damage It is a desirable condition for cell repair without.

Equation [8] shows that a dT _o of 11.5 K can be achieved with a total ultrasonic power of 5.8 Watts when α=0.1 cm ⁻¹ . This is easily achievable. Even if α is increased by a factor of two or three, the resulting power is still easily achievable. The volume of the constant temperature rise region (ie, the volume corresponding to r=r _d =(4Dt _p ) ^1/2 ) is 0.00064 cc. This corresponds to a cube with one side 0.86 mm.

This simple example demonstrates that focused ultrasound should be able to be used to stimulate deep reparative HSPs within the body with readily achievable equipment:
Total ultrasound power: 5.8 Watts - 17 Watts Pulse duration 0.5 seconds Pulse interval 5 seconds Total train duration (N=10) 50 seconds can be used.

The pulsed energy source may be directed to an outer region of the body, which has a blood supply adjacent to or near the surface of the target tissue. Alternatively, the device may be inserted into a body cavity to apply the pulsed energy source to the target tissue. Whether the energy source is applied outside or inside the body and what type of device is utilized depends on the energy source selected and used to treat the target tissue.

Optical stimulation can be effectively delivered to internal surface areas or tissues of the body using endoscopes, such as bronchoscopes, rectoscopes, colonoscopes, etc., in accordance with the present invention. Each of these consists essentially of a flexible tube which itself contains one or more internal tubes. Typically, one of the inner tubes is a light pipe or multiple tube that directs light down the endoscope to illuminate the area of interest and allow the physician to see what is at the illuminated end. Including mode optical fiber. Another inner tube may consist of a wire that carries an electric current to allow the physician to ablate the illuminated tissue. Yet another inner tube may consist of a biopsy tool that allows the physician to cut and hold illuminated tissue.

In the present invention, one of these internal tubes is used as an electromagnetic radiation pipe, such as a multimode optical fiber, to transmit SDM or other electromagnetic radiation pulses that feed the endoscope into the end held by the physician. be done. Referring now to Figure 18, a light generating unit (10), such as a laser having a desired wavelength and/or frequency, is inserted into the body distally of an endoscope (14) illustrated in Figure 19. used to generate electromagnetic radiation, such as laser light, in a controlled and pulsed manner delivered through a light tube or pipe (12) to the end, laser light or other radiation ( 16) is delivered to the target tissue (18) to be treated.

Referring now to Figure 20, a schematic diagram of a system for generating electromagnetic energy radiation, such as laser light, including an SDM is shown. The system, generally referred to by reference numeral (20), includes a laser console (22), such as, for example, an 810 nm near-infrared micropulse diode laser in the preferred embodiment. The laser produces a laser beam that is passed through optics, such as an optical lens or mask, or optionally multiple optical lenses or masks (24). Laser projector optics (24) pass the formed beam through a delivery device (26), such as an endoscope, to project the laser beam onto the target tissue of the patient. The box labeled as (26) may represent both a laser beam projector or delivery device as well as a viewing system/camera such as an endoscope, or may contain two different components in use. understood. The viewing system/camera (26) includes the necessary computerized hardware, data input and It provides feedback to a display monitor (28), which may also include controls and the like.

Referring now to Figure 21, in one embodiment, a laser beam (30) may be passed through a collimating lens (32) and then through a mask (34). In a particularly preferred embodiment, mask (34) comprises a diffraction grating. The mask/grating (34) provides a geometric object, or more typically a geometric pattern of multiple simultaneously generated laser spots, or other geometric object. This is represented by the plurality of laser beams referenced (36). Alternatively, multiple laser spots can be generated by multiple fiber optic waveguides. Both methods of generating laser spots allow simultaneous generation of a large number of laser spots over a very large treatment field. In fact, hundreds, thousands or more of a large number of laser spots can be generated simultaneously to cover a given area of the target tissue, or potentially even the entire target tissue. Application of a broad series of small, discrete laser spots applied simultaneously may be desirable, such as to avoid certain drawbacks and procedural risks known to be associated with the application of large laser spots.

Using optical features with feature sizes comparable to the wavelength of the laser utilized, e.g. using diffraction gratings, allows simultaneous application of a large number of laser spots for very large target areas It is possible to take advantage of quantum mechanical effects that All the individual spots produced by such a grating are optically shaped like the incident beam with minimal power variation for each spot. As a result, multiple laser spots with sufficient irradiance result in harmless yet effective treatment application over a large target area simultaneously. The invention also contemplates the use of other geometric objects and patterns produced by other diffractive optical elements.

Laser light passing through the mask (34) is diffracted resulting in a periodic pattern spaced from the mask (34), shown by the laser beam labeled as (36) in FIG. A single laser beam (30) was then shaped into hundreds or thousands of individual laser beams (36) to create a desired pattern of spots or other geometric objects. These laser beams (36) may be passed through additional lenses, collimators, etc. (38) and (40) to transmit the laser beams and form the desired pattern. Such additional lenses, collimators, etc. (38) and (40) can further transform and redirect the laser beam (36) as required.

Arbitrary patterns can be constructed by controlling the shape, spacing, and pattern of the optical mask (34). Patterns and exposure spots can be arbitrarily created and changed as desired according to application requirements by professionals in the field of optical engineering. Photolithographic techniques, particularly those developed in the field of semiconductor manufacturing, can be used to create simultaneous geometric patterns of spots or other objects.

FIG. 22 diagrammatically illustrates a system for coupling multiple light sources to an optical subassembly that produces the pattern described above. Specifically, this system (20') is similar to the system (20) described in Figure 20 above. The main difference between the alternative system (20') and the system described above (20) is the inclusion of multiple laser consoles, the outputs of which are each fed to a fiber coupler (42). A fiber coupler produces a single output that is passed to the laser projector optics (24) as described in the previous system. Coupling of multiple laser consoles (22) into a single optical fiber is accomplished using fiber couplers (42) as is known in the art. Other known mechanisms for combining multiple light sources are also available and may be used to replace the fiber couplers described herein.

In this system (20'), the multiple light sources (22) follow paths similar to those described in the previous system (20): collimated, diffracted, re-collimated, and projector Oriented to the device and/or tissue. In this alternative system (20'), the diffractive element must function differently than previously described, depending on the wavelength of the light passing through, which results in a slightly different pattern. change. The variation is linear with the wavelength of the light source being diffracted. Generally, the difference in diffraction angles is small enough so that different overlapping patterns are oriented along the same optical path through the projector device (26) towards the tissue for treatment.

Since the resulting pattern varies slightly for each wavelength, the successive offsetting to achieve full coverage is different for each wavelength. Continuous offset can be achieved in two modes. In the first mode, all wavelengths of light are applied simultaneously without the same coverage. An offset steering pattern is used to achieve full coverage for one of the multiple wavelengths. Thus, while the selected wavelength of light achieves complete tissue coverage, the application of other wavelengths achieves either incomplete or overlapping tissue coverage. The second mode applies each light source at varying wavelengths in succession, using appropriate steering patterns to achieve complete tissue coverage for that particular wavelength. This mode precludes the possibility of simultaneous treatment using multiple wavelengths, but makes it possible to achieve identical coverage for each wavelength by optical methods. This avoids incomplete coverage and overlapping coverage for any of the optical wavelengths.

These modes may also be combined and matched. For example, two wavelengths may be applied simultaneously, one wavelength to achieve complete coverage and another wavelength to achieve incomplete or overlapping coverage, followed by a third wavelength in succession. may be applied to achieve full coverage.

Figure 23 diagrammatically illustrates yet another alternative embodiment of the system (20'') of the present invention. This system (20'') is configured in substantially the same manner as the system (20) depicted in FIG. The main difference is the inclusion of subassembly channels that produce multiple patterns tuned to the specific wavelengths of the light source. A plurality of laser consoles (22) are arranged in parallel, each one leading directly to its own laser projector optics (24). The laser projector optics of each channel (44a), (44b), (44c) include a collimator (32), a mask or grating (34), and a recollimator, as described in connection with FIG. 21 above. A whole set of optics, including (recollimators) (38), (40), are tuned for the specific wavelengths produced by the corresponding laser consoles (22). The output from each set of optical elements (24) is then directed to a beam splitter (46) for combination with other wavelengths. It is known to those skilled in the art that beamsplitters used in reverse can be used to combine multiple beams into a single output. The combined channel output from the final beamsplitter (46c) is then directed through the projector device (26).

In this system (20''), the optical elements for each channel are adjusted to produce a precise specified pattern for that channel's wavelength. Consequently, when the channels are all combined and properly aligned, a single steering pattern can be used to achieve complete tissue coverage for all wavelengths.

The system (20'') includes as many channels (44a), (44b), (44c), etc. and beamsplitters (46a), (46b), (46c) as there are wavelengths of light being used for treatment. etc. can be used.

Implementations of the system (20'') may exploit different symmetries to reduce the number of alignment constraints. For example, the proposed grating pattern is periodic in two dimensions and manipulated in two dimensions to achieve full coverage. As a result, if the pattern for each channel is identical to the specified pattern, the actual pattern for each channel need not be aligned to the same steering pattern to achieve full coverage for all wavelengths. do not have. Each channel only needs to be optically aligned to achieve efficient combining.

In the system (20'') each channel starts with a light source (22), which can be from an optical fiber, such as in other embodiments of the pattern generating subassembly. This light source (22) is directed to an optical assembly (24) for collimation, diffraction, reclimation, and orientation to a beam splitter that combines the channel with the primary output.

It is understood that the laser light generation systems illustrated in FIGS. 20-23 are exemplary. Other devices and systems can be utilized to generate a source of SDM laser light that can be operatively passed to a projector device, typically in the form of an endoscope having a light pipe or the like. Other forms of electromagnetic radiation may also be produced and used, including ultraviolet waves, microwaves, other radio frequency waves, and laser light at predetermined wavelengths. In addition, ultrasound is also generated to produce a thermal time course temperature spike in the target tissue sufficient to activate or generate heat shock proteins in the cells of the target tissue without damaging the target tissue itself. and may be used. To that end, typically a pulsed source of ultrasonic or electromagnetic radiant energy is provided and raises the temperature of the target tissue from 6° C. to 11° C., temporarily only 6° C., over an extended period of time, such as several minutes. It is applied to the target tissue in a manner that raises the temperature, such as to 1°C or less.

For deep tissue that is not close to the internal opening, light pipes are not an effective means of delivering pulsed energy. Pulsed low frequency electromagnetic energy or preferably pulsed ultrasound may then be used to induce a series of temperature spikes in the target tissue.

Thus, in accordance with the present invention, a source of pulsed ultrasound or electromagnetic radiation is applied to target tissue to stimulate the production or activation of HSPs and promote protein repair in living animal tissue. Generally, the electromagnetic radiation can be ultraviolet waves, microwaves, other radio frequency waves, laser light, etc. at predetermined wavelengths. On the other hand, when electromagnetic energy is used to target deep tissue away from natural orifices, the absorption length limits the wavelength to microwave or radio frequency wave wavelengths, depending on the depth of the target tissue. However, ultrasound is preferred over long wavelength electromagnetic radiation for deep tissue targets away from natural orifices.

Ultrasonic or electromagnetic radiation produces a thermal time course in tissue to stimulate HSP generation or activation and promote protein repair without causing damage to the cells and tissues being treated. pulsed so as to The area and/or volume of treated tissue is also controlled such that temperature spikes are of the order of magnitude, e.g., approximately 10°C, while maintaining long-term temperature elevations below the FDA-mandated 1°C limit; minimized. It has been found that if the area or volume of tissue to be treated is too large, the elevated tissue temperature cannot be diffused fast enough to meet FDA requirements. However, in addition to limiting the area and/or volume of the treated tissue, by creating a pulsed source of energy, the treated cells and tissue can escape the excess heat generated within acceptable limits. The object of the present invention is achieved to stimulate the activation or production of HSPs by heating or stressing cells and tissues while allowing for

It is believed that stimulation of HSP production in accordance with the present invention can be effectively used in treating a wide range of tissue disorders, diseases, and even infections. For example, viruses that cause the common cold primarily affect a small portion of the respiratory epithelium in the nasal passages and nasopharynx. Like the retina, the respiratory epithelium is a thin, transparent tissue. Referring to Figure 24, a cross-sectional view of a human head (48) is shown with an endoscope (14) inserted into a nasal cavity (50) and energy (16), such as laser light, directed into the nasal cavity (50). is directed to the tissue (18) to be treated with. The tissue to be treated (18) may be within the nasal cavity (50), including the nasal passages and nasopharynx.

To ensure absorption of laser energy, or other energy sources, the wavelength can be adjusted to the infrared (IR) absorption peak of water, or an adjuvant dye to act as a photosensitizer. can be used for In such cases, the treatment is followed by drinking the adjuvant or applying it topically, waiting a few minutes for the adjuvant to penetrate the surface tissue, and then endoscopically as shown in FIG. It consists of administering laser light or other energy source (16) to the target tissue (18) for several seconds, such as through an optical fiber in a mirror (14). To provide patient comfort, the endoscope (14) may be inserted after application of local anesthetic. If necessary, the above procedure can be repeated periodically, such as 1-2 days.

The treatment stimulates the activation or production of heat shock proteins and promotes protein repair without damaging the cells and tissues being treated. As discussed above, certain heat shock proteins have been found to play an important role in the immune response, as well as the health of targeted cells and tissues. The energy source can be monochromatic laser light, such as 810 nm wavelength laser light, and can be administered in a manner similar to that described in the above-cited patent applications, but with the internal It can also be administered via a microscope or the like. Adjuvant dyes are selected to increase laser light absorption. While this includes particularly preferred methods and embodiments of carrying out the invention, it is recognized that other types of energy and delivery means may also be used to achieve the same objectives in accordance with the invention.

Referring now to Figure 25, a similar situation exists for influenza virus, where the primary target is the epithelium of the upper respiratory tree in segments with diameters greater than about 3.3 mm. , the upper six generations of the upper respiratory tree. A viscous thin layer separates the targeted epithelial cells from the airway lumen, where antigen-antibody interaction occurs, resulting in viral inactivation.

With continued reference to Figure 25, the flexible light tube (12) of the bronchoscope (14) is inserted from the individual's mouth (52) through the throat and trachea (54) into the bronchi (56) of the airway tree. It is Laser light or other energy source (16) is now administered to the tissue in this region of the top segment to treat the tissue and region in the same manner as described above with respect to FIG. 24, and delivered. The wavelength of laser or other energy heats tissue, stimulates the activation or production of HSPs, and promotes protein repair, with attendant benefits, along with the IR absorption peaks of water present in mucus. It is contemplated to be selected to match.

Referring now to FIG. 26, as illustrated, the colonoscope (14) delivers selected laser light or other energy source (16) to the area and tissue to be treated. It may have a flexible light tube (12) inserted into the anus and rectum (58) and into either the large intestine (60) or the small intestine (62). It can be used to help with other gastrointestinal problems in addition to treating colon cancer.

Typically, the procedure involves removing all stool from the bowel, the patient lying on their side, and the physician inserting the long, thin light tube portion (12) of the colonoscope (14) into the rectum, to the region of the colon, large intestine (60) or small intestine (64) to be treated. Through the monitor, the physician can see the path of the inserted flexible member (12) and the tip of the colonoscope (14) in the intestine to see the area to be treated. I was able to see the organization in Using one of the other optical fibers or light tubes, the endoscope tip (64) is directed at the tissue to be treated and the laser light or other source of radiation (16) is directed into the tissue (18). To stimulate HSP activation or production, it is delivered through one of the light tubes of the colonoscope (14) to treat the area of tissue to be treated, as described above.

Referring now to Figure 27, another example in which the present invention can be used to advantage is the "leaky gut" syndrome, a disease of the gastrointestinal (GI) tract that is manifested by inflammation and other metabolic dysfunctions, and frequently is called. The GI tract, like the retina, is vulnerable to metabolic dysfunction and is expected to respond well to the treatments of the present invention. This may be done by sub-threshold diode micropulse laser (SDM) treatment, as discussed above, or other energy sources and means as discussed herein and known in the art. obtain.

With continued reference to FIG. 27, a flexible light tube (12), such as an endoscope, is inserted from the patient's mouth (52) through the throat and tracheal region (54) into the stomach (66) where , the tip or end (64) of the flexible light tube (12) is directed toward the tissue (18) to be treated and the laser light or other energy source (16) is directed toward the tissue (18). oriented to It will be appreciated by those skilled in the art that a colonoscope is also used and can be inserted through the rectum (58) into the stomach (66) or any tissue between the stomach and rectum.

If desired, chromophore dyes can be delivered orally to GI tissue to allow absorption of radiation. For example, if unfocused 810 nm radiation from a laser diode or LED is used, the dye will have an absorption peak at or near 810 nm. Alternatively, the wavelength of the energy source can be tuned to a slightly longer wavelength at the absorption peak of water, thus eliminating the need for external application of chromophores.

It is also contemplated by the present invention that a capsule endoscope (68), such as that shown in Figure 28, may be used to provide a source of radiation and energy in accordance with the present invention. Such capsule endoscopes are relatively small in size, such as approximately one inch in length, to be swallowed by the patient. Once the capsule or pill (68) is swallowed, enters the stomach and passes through the GI tract, at an appropriate location the capsule or pill (68) activates an energy source (72) such as a laser diode and associated circuitry. Power and signals can be received via an antenna (70) or the like to activate and a suitable lens (74) directs the generated laser light or radiation through a radiation transparent cover (76). , focused on the tissue to be treated. Illuminated miniature where the position of the capsule endoscope (68) is confirmed by the physician by various means such as external imaging, signal tracking, or images of the GI tract at which the pill or capsule (68) passes. It is understood that the camera may also be determined. The capsule or pill (68) may be supplied with its own power source, such as by a battery, or may be externally powered via an antenna, so as to treat the tissue and area to be treated. To do so, a laser diode (72) or other energy generating source produces the desired wavelength and pulse energy source.

As in the retinal treatment in previous applications, the radiation is pulsed to take advantage of the micropulse temperature spikes and associated safety, and the power is adjusted so that the treatment is completely harmless to the tissue. can be This involves adjusting the peak power, pulse time, and repetition rate to provide a spike temperature rise of approximately 10°C while maintaining the long-term temperature rise below the FDA-mandated 1°C limit. can contain. If a pill form of delivery (68) is used, the device may be powered by a small rechargeable battery or radio inductive excitation, or the like. Heated or stressed tissue stimulates the activation or production of HSPs and promotes protein repair, providing their attendant benefits.

From the previous examples, the technology of the present invention is limited to the treatment of diseases at near or internal body surfaces that are readily available through fiber optics or other means of light delivery. The reason the application of SDM to activate HSP activity is limited to near-surface or optically accessible regions of the body is that the absorption length of IR radiation or visible light in the body is very short. However, it is diseases deep within the tissue or body that benefit from the present invention. Thus, the present invention contemplates the use of ultrasound and/or radio frequency (RF) and shorter wavelength electromagnetic (EM) radiation, such as microwaves, which have relatively long absorption lengths in body tissues. . The use of pulsed ultrasound is preferred over RF electromagnetic radiation to activate therapeutic HSP activity in abnormal tissue not accessible to superficial SDM and the like. A source of pulsed ultrasound may also be used in anomalies at or near the surface.

Referring now to FIG. 29, using ultrasound, specific regions deep within the body can be specifically targeted by using one or more beams each focused on a target site. Pulsed heating is then applied mostly only to the targeted area, where the beams are focused and overlapped.

As illustrated in FIG. 29, an ultrasound transducer (78) or the like produces a plurality of ultrasound beams (80) that are coupled to the skin via an acoustic impedance matching gel and beams (80 ), through the skin (82) and undamaged tissue in front of the focal point of the target organ (84), exemplified by the liver, and specifically the ultrasound beam (80). target tissue (86). As noted above, pulsed heating is then applied only to targeted focal regions (86) where the focused beams (80) overlap. Tissues in front of and behind the focal region (86) are not appreciably heated or affected.

The present invention contemplates the treatment of superficial or near-surface tissue, such as with laser light, and deep tissue treatment, such as with focused ultrasound beams, as well as the treatment of blood disorders such as sepsis. is doing. As indicated above, focused ultrasound treatment can be used on deep body tissues as well as on the surface, and can also be applied when treating this blood disease. However, SDM and similar treatment options, which are typically limited to treatments at or near the surface of epithelial cells, have been associated with blood disorders in areas where blood is accessible through relatively thin layers of tissue, such as the ear lobe. It is also contemplated that it should be used in treating

Referring now to Figures 30 and 31, treatment of blood disorders simply involves delivery of SDM or other electromagnetic radiation or ultrasound pulses to the earlobe (88), where SDM, or energy Other sources of radiation can travel through the earlobe tissue and into the blood flowing through the earlobe. It will be appreciated that this approach may be performed in other areas of the body where blood flow is relatively high and/or near tissue surfaces such as inside the fingertips, mouth or throat.

Referring now to Figures 30 and 31, an earlobe (88) is shown adjacent a clamping device (90) configured to transmit SDM radiation or the like. This may be, for example, by one or more laser diodes (92) delivering desired frequencies in desired pulses and pulse trains to the ear lobe (88). For example, power may be provided by a lamp drive (94). Alternatively, lamp drive (94) may be the actual source of laser light, which is transmitted to earlobe (88) through appropriate optics and electronics. The clamping device (90) is simply used to clamp onto the patient's earlobe (88) and to force the radiation to be confined to the patient's earlobe (88). This can be done by mirrors, reflectors, diffusers, or the like. This may be controlled by a control computer (96) operated by a keyboard (98) or the like. The system may also include a display and speaker (100) if desired, for example when the above procedure is performed by an operator away from the patient.

The proposed treatment with trains of electromagnetic or ultrasound pulses has two major advantages over initial treatments incorporating either single short pulses or sustain (long) pulses. First, the short (preferably sub-second) individual pulses in the train induce cellular reset mechanisms, such as HSP activation, with larger kinetic constants than those operating on longer (minutes or hours) timescales. to activate. Second, the repeated pulses in the treatment allow the cell's repair system to more rapidly overcome the activation energy barrier that separates the dysfunctional cellular state from the desired functional state, a large thermal spike ( approximately 10,000). The end result is a "lower therapeutic threshold" in the sense that lower applied average power and total applied energy can be used to achieve the desired treatment goal.

Although various embodiments have been described in detail for purposes of illustration, various modifications may be made without departing from the scope and spirit of the invention. Accordingly, the invention is not to be restricted except as by the appended claims.

Claims (6)

Thermally treating living tissue characterized by a radio frequency pulse energy source (16) having a frequency of 3-6 MHz and energy parameters including a duty cycle of 10% or less and a total pulse train duration of less than 1 second . wherein said radio frequency pulsed energy source (16) applies at least said pulsed energy source (16) to target tissue (18) for a total pulse train duration of less than 1 second, maintaining a target tissue temperature elevation of 6°C to 11°C;
the average temperature rise of the target tissue for 6 minutes or less is maintained at 1° C. or less so as not to permanently damage the target tissue ;
A system suitable for heat-treating living tissue, characterized by: 2. The system of claim 1, wherein the radio frequency pulse energy source (16) stimulates activation of heat shock proteins in the target tissue. The system of claim 1, wherein energy parameters of the radio frequency pulsed energy source (16) are selected such that approximately 20-40 Joules of energy are absorbed into each cubic centimeter of target tissue. 2. The system of claim 1, comprising a device (14) insertable into a body cavity to apply the radio frequency pulse energy source to target tissue. 2. The system of claim 1, wherein the radio frequency pulse energy source (16) is applied to a blood supply near target tissue (18). The radio frequency pulse energy source (16) is approximately 2 . With a duty cycle of between 5% and 5% and a pulse train duration of between approximately 0.2 and 0.4 seconds, preferably a coil radius of approximately 2-6 mm and an ampere-turn of approximately 13-57. 6. A system according to any one of claims 1 to 5, wherein the system is produced by a device comprising:

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