A system adapted to stimulate the activation of heat shock proteins using ultrasonic energy.

Abstract

A method is provided for applying a pulsed energy source to living tissue to stimulate the activation of heat shock proteins and promote protein repair without damaging the tissue. A method for thermally treating biological tissue includes providing a pulsed energy source having energy parameters selected to raise the target temperature to a level that achieves a therapeutic effect, while maintaining the increase in average temperature of the tissue over time below a predetermined level so as not to permanently damage the target tissue. Application of the pulsed energy source to the target tissue induces a heat shock response and stimulates the activation of heat shock proteins in the target tissue to therapeutically treat the target tissue.

Description

Technical Field

[0001] The present invention generally relates to methods for heat treating biological tissue. In particular, the present invention is directed to methods of applying a pulsed energy source to biological tissue to stimulate the activation of heat shock proteins and promote protein repair without damaging the tissue.

[0002] The inventors have discovered that by controllably raising the temperature of biological tissue to a maximum predetermined temperature range while maintaining the increase in the average temperature of the biological tissue below a predetermined level over several minutes so as not to permanently damage the target tissue, there is a therapeutic effect on biological tissue, especially damaged or diseased biological tissue. Raising the tissue temperature in such a controlled manner is believed to selectively stimulate the production and promotion of the activation of heat shock proteins and / or protein repair, which serves as a mechanism for therapeutically treating tissue.

[0003] Heat shock proteins (HSPs) are a family of proteins produced by cells in response to exposure to stressful conditions. The production of high levels of heat shock proteins can be 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.

[0004] Heat shock proteins are known to play a role in responding to many abnormal conditions in body tissues, including viral infection, combustion, malignant transformation, exposure to oxidants, cytotoxins, and oxygen deprivation. The functions of some heat shock proteins as intracellular chaperones for other proteins and members of the HSP family are expressed or activated at low to moderate levels due to their essential roles in protein maintenance and monitoring of cellular proteins under non-stressful conditions. Such activation is part of the cell's own repair system called the cellular stress response or heat shock response.

[0005] Heat shock proteins are typically named according to their molecular weight. For example, Hsp60, Hsp70, and Hsp80 refer to families of heat shock proteins with sizes of approximately 60, 70, and 80 kilodaltons, respectively. They act in many different ways. For instance, Hsp70 has peptide bonds and an ATPase domain that stabilizes unfolded, assemblable protein structures. Mitochondrial Hsp60s form ring-like structures that facilitate protein assembly to its 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.

[0006] Hsp70 heat shock proteins are members of the extracellular and membrane-bound heat shock protein group involved in antigen binding and its presentation to the immune system. Hsp70 has been shown to inhibit the activity of influenza A virus ribonucleoprotein, thereby blocking viral replication. Tumor-derived heat shock proteins induce specific protective immune responses. Experimental and clinical observations have shown that heat shock proteins are involved in the regulation of autoimmune arthritis, type 1 diabetes mellitus, arteriosclerosis, multiple sclerosis, and other autoimmune responses.

[0007] Accordingly, it is believed to be advantageous to selectively and controllly raise the target tissue temperature to a predetermined temperature range in a short period of time, while simultaneously maintaining an average temperature increase of the target tissue at a predetermined temperature over a longer period. This is believed to trigger a 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 area of ​​body being treated. The present invention satisfies these needs and provides other relevant advantages. [Overview of the project]

[0008] This invention relates to a method for thermally treating biological tissue by applying a pulsed energy source to target tissue in order to therapeutically treat the target tissue. The pulsed energy source has 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 is maintained below a predetermined level so as not to permanently damage the target tissue.

[0009] Energy source parameters may be selected so that the target tissue temperature is raised to at least between approximately 6°C and 11°C during the application of the pulsed energy source to the target tissue. The average temperature increase of the target tissue over several minutes is maintained at or below 6°C, such as less than 1°C, for several minutes.

[0010] The pulsed energy source energy parameters are selected so that approximately 20-40 joules of energy are absorbed per cubic centimeter of target tissue. Applying the pulsed energy source to the target tissue induces a thermal shock response, stimulating the activation of thermal shock proteins in the target tissue without damaging it.

[0011] The device may be inserted into a cavity in the body to apply pulsed energy to tissue. The pulsed energy may be applied to an external region of the body adjacent to the target tissue, or to an external region of the body near its surface where there is a blood supply.

[0012] The pulse energy source may include radio frequencies. These radio frequencies may be between approximately 3 and 6 megahertz (MHz). They may have a duty cycle between approximately 2.5% and 5%. They may have a pulse train duration between approximately 0.2 and 0.4 seconds. The radio frequencies may be generated by a device having a coil radius of approximately 2 to 6 mm and an amperage of approximately 13 to 57.

[0013] The pulsed energy source may include microwave frequencies of 10–20 gigahertz (GHz). The microwave may have a pulse train duration of approximately 0.2–0.6 seconds. The microwave may have a duty cycle of approximately 2–5%. The microwave may have an average power of approximately 8–52 watts.

[0014] The pulsed energy source may include pulsed light, such as laser light. The pulsed light may have a wavelength of approximately 530 nm to 1300 nm, more preferably 800 nm to 1000 nm. The pulsed light may have a power of approximately 0.5 to 74 watts. The pulsed light has a duty cycle of less than 10%, preferably 2.5% to 5%. The pulsed light may have a pulse train duration of approximately 0.1 to 0.6 seconds.

[0015] The pulsed energy source may include pulsed ultrasound. The ultrasound has a frequency of approximately 1–5 MHz. The ultrasound has a pulse duration of approximately 0.1–0.5 seconds. The ultrasound may also have a duty cycle of approximately 2%–10%. The ultrasound has a power of approximately 0.46–28.6 watts.

[0016] Other features and advantages of the present invention will become apparent from the following detailed description, along with the accompanying drawings illustrating the principles of the invention, for example. [Brief explanation of the drawing]

[0017] The following attached drawings illustrate the present invention. [Figure 1A] This graph illustrates the average power of a laser source compared to the radius of the laser source and the duration of the laser pulse train. [Figure 1B] This graph illustrates the average power of a laser source compared to the radius of the laser source and the duration of the laser pulse train. [Figure 2A] This graph illustrates the relationship between temperature attenuation and time, depending on the radius and wavelength of the laser source. [Figure 2B] This graph illustrates the relationship between temperature attenuation and time, depending on the radius and wavelength of the laser source. [Figure 3] This graph illustrates peak amperage cycles for various radio frequencies, duty cycles, and coil radii. [Figure 4] This graph illustrates peak amperage cycles for various radio frequencies, duty cycles, and coil radii. [Figure 5] This graph illustrates peak amperage cycles for various radio frequencies, duty cycles, and coil radii. [Figure 6] This graph illustrates peak amperage cycles for various radio frequencies, duty cycles, and coil radii. [Figure 7] This graph shows the time it takes for the temperature rise to decay compared to the radio frequency coil radius. [Figure 8] This graph shows the average microwave power compared to the microwave frequency and the duration of the pulse train. [Figure 9] This graph shows the average microwave power compared to the microwave frequency and the duration of the pulse train. [Figure 10] This graph shows the time relationship between temperature and attenuation for various microwave frequencies. [Figure 11] This graph shows the average power of an ultrasonic source compared to its frequency and pulse train duration. [Figure 12] This graph shows the temperature decay time for various ultrasonic frequencies. [Figure 13] This graph shows the temperature decay time for various ultrasonic frequencies. [Figure 14] This graph shows the volume of the heating region at the focal point compared to the ultrasonic frequency. [Figure 15] This graph compares the equations for temperature divided by pulse duration for ultrasonic energy sources. [Figure 16]This graph illustrates the logarithmic amplitude of the Arrhenius integral of damage and HSP activation as a function of temperature and pulse duration. [Figure 17] This graph illustrates the logarithmic amplitude of the Arrhenius integral of damage and HSP activation as a function of temperature and pulse duration. [Figure 18] This is a schematic diagram of a photogenerating unit according to the present invention, which generates time-series pulses and includes an optical waveguide extending therefrom. [Figure 19] This is a cross-sectional view of a photostimulation delivery device for delivering electromagnetic energy to a target tissue according to the present invention. [Figure 20] This is a schematic diagram illustrating a system used to generate a laser beam according to the present invention. [Figure 21] This is a schematic diagram of an optical system used to generate geometric patterns of laser light according to the present invention. [Figure 22] This is a schematic diagram illustrating an alternative embodiment of a system used to generate a laser beam for processing tissue according to the present invention. [Figure 23] This is a schematic diagram illustrating yet another embodiment of a system used to generate a laser beam for processing tissue according to the present invention. [Figure 24] This is a schematic cross-sectional view of the end of an endoscope inserted into the nasal cavity for tissue treatment, according to the present invention. [Figure 25] This is a schematic partial cross-sectional view of a bronchoscope that extends through the trachea to the bronchi of the lungs, where treatment is performed, according to the present invention. [Figure 26] This is a schematic diagram of a colonoscope that applies light stimulation to the intestinal or colon region of the body according to the present invention. [Figure 27] This is a schematic diagram of an endoscope inserted into the stomach to perform a procedure, according to the present invention. [Figure 28] This is a partial cross-sectional perspective view of a capsule endoscope used in accordance with the present invention. [Figure 29] This is a schematic diagram of pulsed high-density focused ultrasound high-intensity for treating internal body tissues according to the present invention. [Figure 30] This is a schematic diagram for providing treatment to a patient's blood flow through the earlobe, according to the present invention. [Figure 31] This is a cross-sectional view of the stimulation therapy device of the present invention, used to deliver light stimulation to the blood through the earlobe, according to the present invention. [Modes for carrying out the invention]

[0018] As shown in the accompanying drawings and as fully described herein, the present invention relates to systems and methods for delivering pulsed energy sources such as lasers, ultrasound, ultraviolet radio frequencies, and microwave radio frequencies, having energy parameters selected to induce a thermal time course in tissue in order to raise tissue temperature to a level sufficient to achieve a therapeutic effect in a short period of time, while maintaining an average tissue temperature below a predetermined level for a long period of time to avoid permanent tissue damage. It is believed that the formation of a thermal time course stimulates the activation or production of heat shock proteins and promotes protein repair without causing any damage.

[0019] The inventors of the present invention have discovered 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 the retinal tissue, while simultaneously exerting beneficial effects against eye diseases. This is believed to be at least in part due to the stimulation and activation of heat shock proteins and the promotion of protein repair in retinal tissue. This is disclosed in U.S. Patent Application No. 14 / 607,959 filed on 28 January 2015, U.S. Patent Application No. 13 / 798,523 filed on 13 March 2013, and U.S. Patent Application No. 13 / 481,124 filed on 25 May 2012, the contents of which are incorporated herein by reference as if sufficiently referenced.

[0020] The inventors discovered that it is possible to generate a laser beam that is therapeutic but still below the lethal dose to retinal tissue cells, and therefore does not damage photocoagulation in retinal tissue, resulting in a preventative and protective treatment of the retinal tissue of the eye. The combination of selected parameters must be carefully chosen to achieve a therapeutic effect without permanently damaging the tissue. These parameters include the laser wavelength, the radius of the laser source, the average laser power, the total pulse duration, and the duty cycle of the pulse train.

[0021] The selection of these parameters may also be determined by requiring them to be greater than Arrhenius 1 for HSP activation. The Arrhenius integral is used to analyze the effects of the action on biological tissue. See, for example, The CRC Handbook of Thermal Engineering, ed. Frank Kreith, Springer Science and Business Media (2000). At the same time, the selected parameters must not permanently damage the tissue. Therefore, the Arrhenius integral for damage may be used, and the solved Arrhenius integral is less than 1. Alternatively, FDA / FCC constraints on energy deposition per unit gram of tissue and temperature rise, such as those measured over several minutes, must be met to avoid permanent tissue damage. For example, FDA / FCC requirements regarding energy deposition and temperature rise are commonly used and can be referenced, for instance, for electromagnetic sources at www.fda.gov / medicaldevices / deviceregulationandguidance / guidancedocuments / ucm073817.htm#attacha, and for ultrasonic sources at Anastosio and P. LaRivero, ed., Emerging Imaging Technologies. CRC Press (2012). Generally speaking, tissue temperature increases between 6°C and 11°C can produce therapeutic effects by activating heat shock proteins, and under specific environmental conditions, maintaining an average tissue temperature below a predetermined temperature, such as 6°C and 1°C, for an extended period, for example, several minutes such as 6 minutes, can prevent permanent tissue damage.

[0022] The inventors discovered that desirable retinal photostimulation can be created without visible burns or tissue damage by generating a subthreshold, sublethal micropulse laser beam with a wavelength greater than 532 nm and a duty cycle of less than 10%, with a predetermined intensity or power and a predetermined pulse width or exposure duration. Specifically, a laser beam having a wavelength of 550 nm–1300 nm, and in particular, in a preferred embodiment, 810 nm–1000 nm, with a duty cycle of approximately 2.5%–5%, and a predetermined intensity or power (100–590 watts per square centimeter of retina, or approximately 1 watt per laser spot for each treatment spot on the retina), and a predetermined pulse width or exposure duration (e.g., 100–600 milliseconds or less) forms a sublethal, “precisely subthreshold” retinal photostimulation, where all areas of retinal pigment epithelium exposed to the laser irradiation are protected and available for therapeutic contribution. In other words, the inventors discovered that the advantages of the halo effect of conventional methods are re-established by raising retinal tissue to at least a therapeutic level, but below a lethal level for cells or tissue, without destroying, burning, or otherwise damaging the retinal tissue. This is referred to herein as subthreshold diode micropulse laser therapy (SDM).

[0023] SDM does not produce laser-induced retinal damage (photocoagulation), has no known adverse therapeutic effects, and has been reported to be an effective treatment for many retinal disorders (diabetic macular edema (DME), proliferative diabetic retinopathy (PDR), macular edema due to branch retinal vein occlusion (BRVO), central serous chorioretinopathy (CSR), drug resistance reversal, and prophylactic treatment for progressive degenerative retinopathy such as dry age-related macular degeneration, Stargardt disease, cone dystrophy, and retinitis pigmentosa). The safety of SDM is such that it can be used transfoveally in eyes with 20 / 20 visual acuity to reduce the risk of vision loss due to DME, including early fovea.

[0024] The mechanism by which SDM can act is the production or activation of heat shock proteins (HSPs). Despite the almost limitless 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 stress or damage to almost any type of cell. In the absence of lethal cellular damage, HSPs are extremely effective in repairing and restoring living cells to a more normal functional state. HSPs are transient, generally peaking within hours and lasting for several days, although their effects can be long-lasting. HSPs reduce inflammation, a common factor in many disorders.

[0025] Laser therapy can induce the generation or activation of HSPs and alter cytokine expression. The more sudden and severe the non-lethal cellular stress (such as laser irradiation), the more rapid and robust the HSP activation becomes. Therefore, the sudden, repetitive low-temperature thermal spikes brought about by each SDM exposure (up to 7°C with each 100 μs micropulse, or 70,000°C / sec) are particularly effective in stimulating HSP activation compared to non-lethal exposure to subthreshold treatment with continuous-wave lasers, which may only duplicate low mean tissue temperature increases.

[0026] Laser wavelengths below 550 nm increasingly produce cytotoxic photochemical effects. At 810 nm, SDM generates photothermal rather than photochemical cellular stress. Therefore, SDM can affect tissue without damaging it. Thus, the clinical benefits of SDM are primarily generated by the subpathogenic photothermal activation of HSPs in cells. In dysfunctional cells, HSP stimulation by SDM results in standardized cytokine expression and, consequently, improved structure and function. Subsequently, the therapeutic effect of this "low-intensity" laser / tissue interaction is amplified by "high-density" laser application, maximizing the treatment effect by densely treating large tissue areas containing all diseased areas, thereby replenishing all dysfunctional cells in the targeted tissue area. These principles define the treatment strategies for SDM described herein.

[0027] Because normally functioning cells do not require repair, HSP stimulation in normal cells tends not to have a significant clinical effect. The "patho-selectivity" of near-infrared laser effects such as SDM, which affect diseased cells but not normal cells across various cell types, is consistent with clinical observations of SDM. SDM has been reported to have a unique clinically broad therapeutic area among retinal laser modalities, consistent with the American National Standards Institute's "Maximum Permissible Exposure" predictions. While SDM can also induce direct photothermal effects such as entropic protein unfolding and deaggregation, SDM appears to be optimized for clinically safe and effective stimulation of HSP-mediated repair.

[0028] As described above, while SDM stimulation by HSPs is nonspecific with respect to disease processes, the results of HSP-mediated repair are due to their inherent nature to be specific to the dysfunctional state. HSPs tend to repair whatever the mistake may be. Hence the observed effects of SDM in a wide range of retinal conditions, such as BRVO, DME, PDR, CSR, age-related and genetic retinopathy, as well as drug-resistant NAMD. Conceptually, this ability can be considered a kind of "reset to default" mode of SDM action. For widespread disorders in which cellular function is critical, SDM normalizes cellular function by triggering a "reset" (to "factory default settings") through HSP-mediated cellular repair.

[0029] The inventors discovered that SDM treatment in patients suffering from age-related macular degeneration (AMD) can slow its progression and even halt it. Most patients showed significant improvement in dynamic functional logMAR cleavage visual acuity and cleavage contrast visual acuity after SDM treatment. SDM is thought to work by targeting, protecting, and "normalizing" (bringing closer to normal) the function of the retinal pigment epithelium (RPE).

[0030] SDM has also been shown to stop or reverse the signs of diabetic retinopathy, even in the presence of persistent diabetes mellitus throughout the body, without treatment-related damage or side effects. Based on this, it is hypothesized that SDM may act in diabetic RPE cells by inducing a return to normal cellular function and cytokine expression (similar to pressing the "reset" button on an electronic device to restore factory settings). Based on the above information and research, SDM treatment may also directly affect cytokine expression by activating heat shock proteins (HSPs) in the targeted tissue.

[0031] Because heat shock proteins play a role in responding to many abnormal conditions in body tissues other than eye tissue, it is believed that similar systems and methods can be advantageously used to treat such abnormal conditions, infections, etc. Therefore, this invention relates to the controlled application of ultrasound or electromagnetic radiation to treat abnormal conditions, including inflammation, autoimmune diseases, and cancer, reachable by fiber optics of endoscopes or surface probes, or by focused electromagnetic / sound waves. For example, cancers on the surface of the prostate, which are most likely to metastasize, can be reached by fiber optics of a rectoscope. Colon tumors can be reached by fiber optic systems, such as those used in colonoscopy.

[0032] As shown above, subthreshold diode micropulse laser (SDM) light stimulation was effective in stimulating the direct repair of slightly misfolded proteins in eye tissue. In addition to HSP activation, this can also occur in another way, because the temperature spikes caused by micropulses in the form of a thermal time course allow for the diffusion of water within the protein, which in turn allows for the cleavage of peptide-peptide hydrogen bonds that prevent the protein from returning to its native state. The diffusion of water into the protein increases the number of inhibitory hydrogen bonds by approximately 1000 times. Therefore, it is believed that this process could be advantageously applied to other diseases as well.

[0033] As described 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 without permanently damaging the tissue. As an example, when using a light energy source such as a laser beam, parameters such as laser wavelength, duty cycle, and total pulse train duration must be taken into consideration. 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 of the others.

[0034] Figures 1A and 1B illustrate graphs showing average power in watts, comparing the radius of the laser source (between 0.1 cm and 0.4 cm) and the pulse train duration (0.1 to 0.6 seconds). Figure 1A shows a wavelength of 880 nm, and Figure 1B shows a wavelength of 1000 nm. These figures show that the required power decreases monotonically as the radius of the laser source decreases, the total train duration increases, and the wavelength decreases. The preferred parameter for the radius of the laser source is 1 mm to 4 mm. For a wavelength of 880 nm, the minimum power is 0.55 watts, the laser source radius is 1 mm, and the total pulse train duration is 600 milliseconds. When the laser source radius is 4 mm and the total pulse train duration is 100 milliseconds, the maximum power for a wavelength of 880 nm is 52.6 watts. However, when selecting a laser with a wavelength of 1000 nm, the minimum power value is 0.77 watts, the laser source radius is 1 mm, and the total pulse train duration is 600 milliseconds. When the laser source radius is 4 mm and the total pulse duration is 100 milliseconds, the maximum power value is 73.6 watts. The corresponding peak power between individual pulses is obtained from the average power by dividing by the duty cycle.

[0035] The volume of tissue to be 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 heat the tissue, the duty cycle of the pulse train gives the associated spike or peak, and the average laser power gives the power. Preferably, the pulse energy source energy parameters are selected so that approximately 20 to 40 joules of energy are absorbed per cubic centimeter of target tissue.

[0036] The absorption length is very small in the thin melanin layer of the retinal pigment epithelium. In other parts of the body, the absorption length is generally not so small. At wavelengths ranging from 400 nm to 2000 nm, the penetration depth in skin ranges 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 volume heated is limited to the outer or inner surface where the radiation source is placed, where the depth is equal to the penetration depth and the lateral dimension is equal to the lateral dimension of the radiation source. Since the light energy source is used to treat affected tissue near the outer surface or near an internally accessible surface, a source with a radius of 1 mm to 4 mm and operating at a wavelength of 880 nm results in a penetration depth of approximately 2.5 mm, and a source operating at a wavelength of 1000 nm results in a penetration depth of approximately 3.5 mm.

[0037] It has been found that the target tissue can be heated to a maximum of approximately 11°C for a short period of time, such as less than 1 second, in order 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 a long period of time, such as several minutes, while producing the therapeutic effect of the present invention. The selection of the duty cycle and total pulse train duration provides a time interval in which heat can dissipate. Duty cycles of less than 10%, preferably between 2.5% and 5%, have been found to be effective, along with total pulse durations of 100 milliseconds to 600 milliseconds. Figures 2A and 2B illustrate the time it takes to decay from 10°C to 1°C for a laser source with a radius of 0.1 cm to 0.4 cm, with a wavelength of 880 nm in Figure 2A and 1000 nm in Figure 2B. When using the 880 nm wavelength, the decay time is shorter, but either wavelength falls within the range of acceptable requirements and operating parameters for achieving the advantages of the present invention without causing permanent tissue damage.

[0038] It was found that an average temperature increase in the desired target region, increasing from at least 6°C to a maximum of 11°C, and preferably to approximately 10°C, during the total irradiation period, results in HSP activation. Control of the target tissue temperature is determined by selecting source and target parameters, thereby ensuring compliance with conservative FDA / FCC requirements to avoid damage or damage Arrhenius integrals that are greater than 1 and simultaneously avoid damage or damage Arrhenius integrals that are less than 1.

[0039] To meet conservative FDA / FCC constraints to avoid permanent tissue damage, the average temperature rise of the target tissue over a 6-minute period is less than 1°C for light and other electromagnetic radiation sources. Figures 2A and 2B above illustrate typical decay times required for the temperature in a heated target area to decrease by thermal diffusion from a temperature rise of approximately 10°C to 1°C. As seen in Figure 2A, the temperature decay time is 16 seconds when the wavelength is 880 nm and the source diameter is 1 mm. When the source diameter is 4 mm, the temperature decay time is 107 seconds. As shown in Figure 2B, the temperature decay time is 18 seconds when the wavelength is 1000 nm and 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 required for the average temperature rise to be maintained over several minutes, such as less than 6 minutes. The temperature of the target tissue rises very quickly, to about 10°C, such as a fraction of a second, during the application of the energy source to the tissue. However, the relatively low duty cycle provides a relatively long period between pulses of energy applied to the tissue, and the relatively short pulse train duration ensures sufficient temperature diffusion and decay within a relatively short period, including a few minutes, such as less than 6 minutes, without causing permanent tissue damage.

[0040] The parameters vary depending on the individual energy source, including microwaves, infrared lasers, radio frequencies, and ultrasound, because the absorption characteristics of tissues differ with these different types of energy sources. While tissue water content may vary depending on the type of tissue, uniformity of tissue properties is observed under normal or near-normal conditions, which has made it possible to publish tissue parameters widely used by clinicians when designing procedures. The following are tables illustrating the properties of electromagnetic waves in biomedium, with Table 1 relating to muscle, skin, and tissues with high water content, and Table 2 relating to fat, bone, and tissues with low water content.

[0041] [Table 1]

[0042] [Table 2]

[0043] The absorption length of radio frequencies in body tissue is long compared to body length. Consequently, the heating region is determined not by the absorption length, but rather by the dimensions of the coil, which is the source of radio frequency energy. At long distances r from the coil, the magnetic (near) field from the coil is 1 / r 3 It attenuates (drops off). Over shorter distances, the electric and magnetic fields may arise from the point of the magnetic vector potential, which may, in turn, arise in closed form from the points of the first and second kind of elliptic integrals. Heating occurs only in a region that is size-comparable to the dimensions of the coil source itself. Therefore, if it is desirable to preferentially heat a region characterized by radius, the source coil is chosen to have a similar radius. 1 / r of the magnetic field 3 Due to attenuation, heating decays very rapidly outside the hemispherical region of the radius. Since it has been proposed to use radio frequencies in affected tissue accessible only from the outside or inside, it is reasonable to consider a coil radius of approximately 2–6 mm.

[0044] In addition to the radius of the source coil, the number of amperes (NI) in the source coil gives the magnitude and spatial range of the magnetic field, and the radio frequency is a 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 an object near an external or internal surface, conductivity is for skin and mucous tissues. In addition to the duty cycle of the pulse train, the total train duration of the pulse train is a factor that affects how much total energy is delivered to the tissue.

[0045] Preferred parameters for the radio frequency energy source were determined to be 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 a duty cycle between 2.5% and 5%. Figure 3-6 shows how much the ampere-count changes when these parameters are modified to give a temperature rise that results in an Arrhenius integral of approximately 1 for HSP activation. In relation to Figure 3, for an RF frequency of 6 MHz, a pulse train duration of 0.2–0.4 seconds, a coil radius between 0.2–0.6 cm, and a duty cycle of 5%, the peak ampere-count (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, as illustrated in Figure 4, the peak ampere-count is 26 when the pulse train duration is 0.4 seconds, the coil radius is 0.6 cm, and the duty cycle is 5%. However, with the same 5% duty cycle, the peak amperage 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 Figures 5 and 6. This yields 18 amperages for a 6 MHz radio frequency with a coil radius of 0.6 cm and a pulse train duration of 0.4 seconds, as illustrated in Figure 5, and 29 amperages when the coil radius is only 0.2 cm and the pulse train duration is 0.2 seconds. In relation to Figure 6, with a 2.5% duty cycle and a 3 MHz radio frequency, the peak amperage 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.

[0046] Figure 7 illustrates 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, with respect to a radio frequency energy source. 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 acceptable range for decay time, but is at its upper limit.

[0047] Microwaves are another electromagnetic energy source that can be utilized according to the present invention. The frequency of the microwaves determines the tissue penetration distance. The gain of a conical microwave horn is large compared to the microwave wavelength, which indicates that under these circumstances, the energy is radiated, usually with a narrow forward load. Typically, microwave sources used according to the present invention have a length dimension of about 1 centimeter or less, and therefore the microwave source is smaller than the wavelength, in which case the microwave source can be approximated as a dipole antenna. 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 region can be approximated as 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 an external surface or internal cavity, frequencies in the range of 10-20 GHz are used, where the corresponding penetration distance is only between about 2 mm and 4 mm.

[0048] 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 the pulse train determines the peak power in a single pulse within the pulse train. Since the radius of the energy source is approximately less than 1 centimeter and frequencies between 10 and 20 GHz are typically used, the resulting pulse train durations of 0.2 seconds and 0.6 seconds are preferred.

[0049] As the train duration increases and the microwave frequency increases, the required power decreases monotonically. 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 duration 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.

[0050] Referring to Figure 8, the graph shows the average microwave power in watts for a microwave with a frequency of 10 GHz and a pulse train duration between 0.2 and 0.6 seconds. Figure 9 is a similar graph, but shows the average microwave power for a microwave with 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 lawful condition is that the Arrhenius integral for HSP activation in the heating region is approximately 1.

[0051] Referring to Figure 10, the graph illustrates the 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 are 8 seconds when the microwave frequency is 20 GHz and 16 seconds when the microwave frequency is 10 GHz.

[0052] By utilizing ultrasound as an energy source, it becomes possible to heat tissues of varying depths within the body, including surface tissues and considerably deeper tissues. The absorption length of ultrasound within the body is quite long, as evidenced by its widespread use for imaging. Therefore, ultrasound can be focused on deep target regions within the body, and the heating of a focused ultrasound beam is concentrated primarily in the beam's approximately cylindrical focal region. The heated 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, and heating occurs in overlapping focal regions.

[0053] Regarding ultrasound, the relevant parameters for determining tissue temperature are the ultrasound frequency, total column duration, and transducer power, when the focal length and diameter of the ultrasound transducer are taken into consideration. The frequency, focal length, and diameter determine the volume of the focal region where the ultrasound energy is concentrated. The focal volume is the volume of the target tissue for treatment. Transducers with a diameter of approximately 5 cm and a focal length of approximately 10 cm are readily available. Favorable focal dimensions are achieved when the ultrasound frequency is 1–5 MHz and the total column duration is 0.1–0.5 seconds. For example, with a focal length of 10 cm and a transducer diameter of 5 cm, the focal volume is 0.02 cc at 5 MHz and 2.36 cc at 1 MHz.

[0054] Referring here to Figure 11, the graph illustrates the average source power in watts compared to frequency (between 1 MHz and 5 MHz) and pulse train duration (0.1 to 0.5 seconds). A transducer focal length of 10 cm and a source diameter of 5 cm were assumed. The power required to give the Arrhenius integral for approximately 1 HSP activation decreases monotonically as the frequency and total train duration increase. Considering favorable 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 between individual pulses is obtained simply by dividing by the duty cycle.

[0055] Figure 12 illustrates the time in seconds for temperature diffusion or attenuation from approximately 10°C to approximately 6°C when the ultrasonic frequency is 1–5 MHz. Figure 13 illustrates the time in seconds for attenuation 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 attenuation is 366 seconds when the ultrasonic frequency is 1 MHz, and the minimum temperature attenuation is 15 seconds when the microwave frequency is 5 MHz. Because the FDA only requires that the temperature rise be less than 6°C over a test time of minutes, a 366-second attenuation time at 1 MHz, which results in a 1°C rise over several minutes, is possible. As can be seen in Figures 12 and 13, the attenuation time to a 6°C rise is approximately 70 times shorter than the attenuation time to a 1°C rise.

[0056] Figure 14 illustrates the volume of the focal heating area in cubic centimeters compared to ultrasonic frequencies between 1 and 5 MHz. Considering ultrasonic frequencies in the range of 1 to 5 MHz, the corresponding focal size for these frequencies ranges from 3.7 mm to 0.6 mm, and the focal area length ranges from 5.6 cm to 1.2 cm. The corresponding therapeutic volume ranges from approximately 2.4 cc to 0.02 cc.

[0057] Examples of parameters considering the Arrhenius integral of desirable HSP activation greater than 1 and the Arrhenius integral of damage less than 1 are a total ultrasonic power of 5.8–17 watts, a pulse duration of 0.5 seconds, a pulse interval of 5 seconds, and a total of 10 pulses within a total pulse train time of 50 seconds. The treatment volume of the target is approximately 1 mm on one side. By applying ultrasound in multiple simultaneously applied adjacent but separately spaced rows, larger treatment volumes may be treatable with an ultrasonic system similar to a laser-diffracted optical system. Multiple focused ultrasonic beams converge on a very small treatment target within the body, thereby allowing minimal heating in the target except for overlapping beams. This area is heated, and the transient high-temperature spike stimulates HSP activation and promotes protein repair. However, considering the relatively small area being treated in a given time, in addition to the pulsed aspect of the present invention, the treatment complies with FDA / FCC requirements for a long-term (minutes) average temperature rise <1K. A key difference between this invention and existing thermotherapy treatments for pain and muscle tension is that existing techniques lack high T-spikes, and it is necessary to efficiently activate HSPs and promote protein repair to provide healing at the cellular level.

[0058] Insofar as therapeutic activation of HSPs and promotion of protein repair are involved, pulsed train modes of energy delivery have advantages over single-pulse or stepwise modes of energy delivery. There are two considerations for these advantages: First, a major advantage of the SDM energy delivery mode for HSP activation and protein repair arises from the generation of a spike temperature of approximately 10°C. This significant temperature increase greatly affects the Arrhenius integral, which quantitatively describes the number of activated HSPs, and the rate of water diffusion to proteins that promote protein repair. This is because the temperature enters an exponential function, resulting in a large amplification effect.

[0059] It is important that the temperature rise does not remain high (above 10°C) for an extended period, because this would violate FDA and FCC requirements that the average temperature rise over a period of minutes must be less than 1°C (6°C in the case of ultrasound).

[0060] The SDM mode of energy delivery uniquely satisfies both of these prior considerations through a wise selection of power, pulse time, pulse interval, and volume of the target area being treated. The volume of the treatment area must decay fairly rapidly from its high of approximately 10°C so that the long-term average temperature rise does not exceed the long-term FDA / FCC limits of 6°C for ultrasound frequency and 1°C for electromagnetic radiation energy source.

[0061] For a region of length L, the time it takes for the peak temperature to become e times (e-fold) in the tissue is approximately L. 2 / 16D, where D = 0.00143 cm 2 The coefficient of thermal diffusion is a typical value per second. For example, when L=1mm, the decay time is approximately 0.4 seconds. Therefore, for a 1mm area on the side surface, a sequence of 10 pulses, each with a duration of 0.5 seconds and a pulse interval of 5 seconds, can achieve the desired instantaneous high temperature rise without exceeding an average long-term temperature rise of 1°C. This is further demonstrated below.

[0062] The reason why RF electromagnetic radiation is not a good choice for SDM-type treatments in deep regions of the body as ultrasound is the limitation of the heating amount. The long skin depth (penetration distance) and ohmic heating along the entire skin depth result in a large amount of heating, and its thermal inertia does not allow for the achievement of high spike temperatures that activate HSP and promote protein repair, nor for the rapid temperature decay that meets the long-term FDA and FCC limits for average temperature rise.

[0063] Ultrasound is already used to therapeutically heat regions of the body to relieve pain and muscle tension. However, the heating does not follow an SDM-type protocol and does not have the temperature spikes that are a factor in the excitation of HSP.

[0064] Now, consider a group of focused ultrasound beams oriented at a deep target region within the body. To simplify the calculations, assume that the beam is replaced by a single source with a spherical shape that focuses at the center of a 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 the frequency.

[0065]

Table 3

[0066] Assuming that the geometric deformation due to focusing dominates over the deformation due to attenuation, the intensity of the incident ultrasound at a distance r from the focus is calculated approximately as follows: I(r)=P / (4πr 2 ) [1] where P represents the total ultrasonic power. The temperature rise at the end of a short pulse of duration t at r is as follows, p and dT(t p )=Pαt p / (4πC v r 2 ) [2] where α is the absorption coefficient and Cv t is the specific heat capacity. p This holds true until r reaches a point where the thermal diffusion distance at r is 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, suppose the diffraction limit is reached at a radial distance shorter than the distance determined by thermal diffusion. Here, r dif =(4Dt p ) 1 / 2 [3] In the equation, D is the thermal diffusion coefficient, and r <r dif Regarding t p The temperature increase is as follows: r <r dif If so, dT(r dif ,t p ) = 3Pα / (8πC v D)[4] Therefore, the temperature rise can be recorded at the end of the pulse: dt p (r) = {Pαt p / (4πC v}[(6 / r dif 2 )U{r dif -r)+(1 / r 2 )U(rr dif ) [5] Green's function for the thermal diffusion equation: G(r,t)=(4ΩDt) -3 / 2 exp[-r 2 / (4Dt)][6] Applying this to the initial temperature distribution, we found that the temperature dT(t) at the focus r=0 at time t is as follows: 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] together dT o =3Pα / (8πC v D) [8]

[0067] A good approximation for equation [7] is provided by the following, as can be seen in Figure 15:

[0068]

number

[0069]

number

[0070] The Arrhenius integral can be approximated by dividing the integration interval into a portion with temperature spikes and a portion without temperature spikes. The sum of the contributions of the temperature spikes can be simplified by applying Laplace's endpoint formula to the integral with respect to the temperature spikes. Furthermore, the integral for the portion without spikes can be simplified by noting that the temperature rise without spikes reaches its asymptotic value very rapidly, and as a result, a good approximation can be obtained by exchanging the changing time rise with its asymptotic value. 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 +dTo +dT N (Nt I )) +exp[-(E / k B )1 / (T o +dT N (Nt I ))]]

[12] In the formula, the following applies:

[0071]

number

[13] is (Nn) -3 / 2 It arises from the sum of n and is the magnitude of the harmonic (N, 3 / 2) for a typical N of the subject.

[0072] It is interesting to compare this expression with the expression for SDM applied to the retina. The first condition (term) is very similar to the condition from the spike contribution in the retina, except that the effective spike interval is reduced by a factor of 3 in the case of this 3D focused beam. The second condition is dT N (Nt I The fact that it includes ) is far less than in the case of the retina. Here, the background temperature rise was on a scale comparable to the spike temperature rise. However, here in the case of a focused beam, the background temperature rise is a ratio (t p / t I ) 3 / 2 This is by far smaller. The background temperature rise, which is similar to the rise in the case of continuous ultrasonic heating, is not significant compared to the spike contribution, highlighting the importance of the spike contribution to HSP activation or generation and the promotion of protein repair. At the end of the pulse train, even this low background temperature rise disappears rapidly due to thermal diffusion.

[0073] Figures 16 and 17 show the pulse duration t. p = 0.5 seconds (sec), pulse interval t I =10 seconds, and dT for the total number of pulses N=10 oThis shows the magnitude of the logarithm of the Arrhenius integral for damage and for HSP activation or generation, corresponding to the pulse duration t. p = 0.5 seconds, pulse interval t I =10 seconds, and single pulse dT for the total number of ultrasonic pulses N=10. o The logarithm of the Arrhenius integral [Equation 12] for damage and HSP activation as the temperature rises in Kelvin degrees from [A]. Figure 16 shows the Arrhenius constant A = 8.71 x 10⁻¹⁰. 33 sec -1 and E=3.55x10 -12 The logarithm of the damage integral in ergs is shown. Figure 17 shows the Arrhenius constant A = 1.24 x 10⁻¹⁰. 27 sec -1 and E=2.66x10 -12 The logarithm of the HSP activation integral in ergs is shown. The graphs in Figures 16 and 17 are dT o Until it exceeds 11.3K Ω damage It does not exceed 1, while Ω hsp However, it shows a value greater than 1 over the entire interval shown, which is a desirable condition for cell repair without damage.

[0074] Equation [8] is given by α = 0.1 cm -1 At that time, dT o However, it has been shown that this can be achieved with a total ultrasonic power of 5.8 watts. This is easily achievable. Even if α is increased by 2 or 3 times, the resulting power is still easily achievable. Volume of the region where the temperature rise is constant (i.e., r=r d =(4Dt p ) 1 / 2 The corresponding volume is 0.00064 cc. This corresponds to a cube with one side measuring 0.86 mm.

[0075] This simple example demonstrates that focused ultrasound should be available to stimulate HSPs for deep tissue repair within the body, with readily achievable equipment: Total ultrasonic power: 5.8 watts - 17 watts Pulse duration: 0.5 seconds Pulse interval: 5 seconds Total column duration (N=10): 50 seconds To facilitate the handling of larger volumes, the SAPRA system can be used.

[0076] The pulsed energy source may be directed towards an external region of the body, where there is 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 externally or internally to the body, and what type of device is used, depends on the energy source selected and used to treat the target tissue.

[0077] Light stimulation can be effectively transmitted to internal surface areas or tissues of the body using an endoscope such as a bronchoscope, rectoscope, or colonoscope, according to the present invention. Each of these substantially consists of a flexible tube containing one or more internal tubes. Typically, one of the internal tubes includes an optical pipe or multimode optical fiber that guides light down the endoscope to illuminate the area of ​​interest and allow the physician to confirm what is at the illuminated end. Another internal tube may consist of a wire that carries an electric current to allow the physician to cauterize the illuminated tissue. Yet another internal tube may consist of a biopsy tool that allows the physician to cut out and hold the illuminated tissue.

[0078] 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 advance the end end held by the physician. Referring here to Figure 18, a photogenerating unit (10), such as a laser having a desired wavelength and / or frequency, is used to generate electromagnetic radiation, such as laser light, in a controlled and pulsed manner, which is delivered through the optical tube or optical pipe (12) to the distal end of the endoscope (14) illustrated in Figure 19, which is inserted into the body, and the laser light or other radiation (16) is delivered to the target tissue (18) to be treated.

[0079] Referring here to Figure 20, a schematic diagram of a system for generating electromagnetic energy radiation, such as laser light, including an SDM. The system, referred to collectively by reference numeral (20), includes a laser console (22), such as an 810 nm near-infrared micropulse diode laser in a preferred embodiment. The laser generates a laser beam that passes through optical elements (optics), such as an optical lens or mask, or optionally multiple optical lenses or masks (24). The laser projector optical element (24) passes 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 (26) may represent both a laser beam projector or delivery device and a viewing system / camera, such as an endoscope, or it may contain two different components when in use. The visibility system / camera (26) provides feedback to a display monitor (28), which may also include computerized hardware, data input units, and control units necessary for operating the laser (22), optical elements (24), and / or projection / visualization components (26).

[0080] Referring here to Figure 21, in one embodiment, a laser beam (30) may be passed through a collimator lens (32) and then through a mask (34). In a particularly preferred embodiment, the mask (34) includes a diffraction grating. The mask / diffraction grating (34) results in a geometric object, or more typically, a geometric pattern of multiple simultaneously generated laser spots, or other geometric object. This is represented by multiple laser beams designated reference numeral (36). Alternatively, multiple laser spots may be generated by multiple fiber optic waveguides. Any method of generating laser spots allows for the simultaneous generation of a large number of laser spots over a very wide treatment field. In fact, hundreds, thousands, or even more laser spots may be generated simultaneously to cover a given area of ​​target tissue, or potentially the entire target tissue. In some cases, the application of a wide series of simultaneously applied, small, separated laser spots may be desirable, for example, to avoid certain drawbacks and procedural risks known to be associated with the application of large laser spots.

[0081] By using optical features with a shape and size equivalent to the wavelength of the laser being used, for example, by using a diffraction grating, it is possible to leverage quantum mechanical effects that enable the simultaneous application of a large number of laser spots for a very large target area. Each individual spot generated by such a diffraction grating has an optical shape similar to the incident beam, with minimal power variation for each spot. As a result, multiple laser spots with sufficient irradiation doses can simultaneously provide a harmless and effective treatment application over a large target area. The present invention also considers the use of other geometric objects and patterns generated by other diffractive optical elements.

[0082] The laser light passing through the mask (34) is diffracted, resulting in a periodic pattern at a distance from the mask (34), as shown by the laser beam labeled as (36) in Figure 21. The single laser beam (30) is therefore shaped into hundreds or thousands of individual laser beams (36) to produce a desired pattern on a spot or other geometric object. These laser beams (36) can be passed through additional lenses, collimators, etc., (38) and (40) to transmit the laser beam and form the desired pattern. Such additional lenses, collimators, etc., (38) and (40) can further transform and change the orientation of the laser beam (36) as needed.

[0083] By controlling the shape, spacing, and pattern of the optical mask (34), arbitrary patterns can be constructed. Patterns and exposure spots can be arbitrarily created and modified as desired, according to the application requirements of experts in the field of optical engineering. Photolithography techniques, particularly those developed in the field of semiconductor manufacturing, can be used to create simultaneous geometric patterns of spots or other objects.

[0084] Figure 22 schematically illustrates a system for connecting multiple light sources to an optical subassembly that generates the above pattern. Specifically, this system (20') is similar to system (20) described in Figure 20 above. The main difference between the alternative system (20') and the aforementioned system (20) is that it includes multiple laser consoles, each of which is supplied to a fiber coupler (42). The fiber coupler produces a single output that is passed to a laser projector optical element (24) as described in the previous system. The connection of the multiple laser consoles (22) to a single optical fiber is achieved using a fiber coupler (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 coupler described herein.

[0085] In this system (20'), the multiple light sources (22) follow the same path as described in the previous system (20), namely, collimated, diffracted, collimated again, and directed toward the projector and / or tissue. In this alternative system (20'), the diffracting element must function differently from that described previously, depending on the wavelength of the light passing through it, which results in a slight variation in the pattern. The variation is linear with the wavelength of the light source being diffracted. Generally, the difference in diffraction angles is small enough that different overlapping patterns are directed toward the tissue through the projector (26) for treatment.

[0086] Because the resulting pattern varies slightly for each wavelength, the consecutive offsets required to achieve full coverage differ for each wavelength. Consecutive offsets can be achieved in two modes. In the first mode, all wavelengths of light are applied simultaneously without identical coverage. A steering pattern of the offset is used to achieve full coverage for one of the multiple wavelengths. Thus, light of the selected wavelength achieves full coverage of the tissue, while application of other wavelengths achieves either incomplete or overlapping coverage. The second mode involves successively applying each light source of varying wavelengths using an appropriate steering pattern to achieve full coverage of the tissue for that particular wavelength. This mode eliminates the possibility of simultaneous treatment using multiple wavelengths but allows for achieving identical coverage for each wavelength by optical means. This avoids both incomplete and overlapping coverage for any of the optical wavelengths.

[0087] These modes may also be combined and matched. For example, two wavelengths may be applied simultaneously, with one wavelength achieving full coverage and the other achieving incomplete or overlapping coverage, after which a third wavelength may be applied consecutively to achieve full coverage.

[0088] Figure 23 schematically illustrates yet another alternative embodiment of the system (20'') of the present invention. This system (20'') is configured substantially the same as system (20) depicted in Figure 20. The main difference is the inclusion of sub-assembly channels that generate multiple patterns tuned to specific wavelengths of the light source. Multiple laser consoles (22) are arranged in parallel, one of each of which is directly connected to its own laser projector optical elements (24). The laser projector optical elements of each channel (44a), (44b), (44c) include a collimator (32), a mask, or diffraction grating (34), and recollimators (38), (40), as described in relation to Figure 21 above, and the entire set of optical elements is tuned for specific wavelengths produced by the corresponding laser console (22). The output from each set of optical elements (24) is then directed to a beam splitter (46) for combinations with other wavelengths. It is known to those skilled in the art that a beam splitter used in the reverse direction may be used to combine multiple beams into a single output. The combined channel output from the final beam splitter (46c) is then directed via a projector device (26).

[0089] In this system (20''), the optical elements for each channel are tuned to produce a precisely specified pattern for the wavelength of that channel. Consequently, when all channels are combined and properly aligned, a single steering pattern can be used to achieve full applicability of the structure across all wavelengths.

[0090] The system (20'') may use as many channels (44a), (44b), (44c), etc. as the wavelength of light used in the procedure, as well as beam splitters (46a), (46b), (46c), etc.

[0091] In implementing the (20'') system, different symmetries can be utilized to reduce the number of alignment constraints. For example, the proposed grid pattern is periodic in two dimensions and is manipulated in two dimensions to achieve full applicability. As a result, if the pattern for each channel is identical to the specified pattern, the actual pattern of each channel does not need to be aligned to the same steering pattern to achieve full applicability for all wavelengths. Each channel only needs to be optically aligned to achieve an efficient combination.

[0092] In the system (20''), each channel is started with a light source (22), which may be from an optical fiber in other embodiments of the subassembly that generates the pattern. The light source (22) is oriented into the optical assembly (24) for collimation, diffraction, recollimation, and orientation to a beam splitter that combines the channels with the main output.

[0093] The laser light generation system illustrated in Figure 20-23 is understood to be typical. Other devices and systems can be used to generate a source of SDM laser light that can be operably transmitted to a projector device, typically in the form of an endoscope having an optical pipe or the like. Other forms of electromagnetic radiation, including ultraviolet waves, microwaves, other radio frequency waves, and laser light, may also be generated and used at predetermined wavelengths. Furthermore, ultrasound may also be generated and used to produce a thermal temperature spike in the target tissue over a period of time sufficient to activate or generate heat shock proteins in the cells of the target tissue without damaging the target tissue itself. For this purpose, typically, a pulsed source of ultrasonic or electromagnetic radiation energy is provided and applied to the target tissue in a manner that raises the temperature of the target tissue to 6°C to 11°C, or temporarily to just 6°C or 1°C or less, over a period of time such as several minutes.

[0094] For deep tissues not close to the internal opening, optical pipes are not an effective means of delivering pulsed energy. In such cases, pulsed low-frequency electromagnetic energy, or preferably pulsed ultrasound, may be used to induce a series of temperature spikes in the target tissue.

[0095] Therefore, according to the present invention, a source of pulsed ultrasound or electromagnetic radiation is applied to target tissue to stimulate the generation or activation of HSPs and to promote protein repair in living animal tissue. Generally, electromagnetic radiation can be ultraviolet waves, microwaves, other radio frequency waves, or laser light at predetermined wavelengths. On the other hand, when electromagnetic energy is used to target deep tissues away from natural openings, the absorption length limits the wavelength to microwave or radio frequency wavelengths, depending on the depth of the target tissue. However, ultrasound is preferred over long-wavelength electromagnetic radiation for targeting deep tissues away from natural openings.

[0096] The ultrasonic or electromagnetic radiation is pulsed to stimulate the generation or activation of HSPs and promote protein repair without causing damage to the treated cells and tissues, by inducing a thermal time course in the tissue. The area and / or volume of the treated tissue is also controlled and minimized so that temperature spikes are only a few degrees, e.g., about 10 degrees, while maintaining a long-term temperature rise below the FDA-defined 1 degree limit. It has been found that if the area or volume of the treated tissue is too large, the elevated tissue temperature cannot dissipate quickly enough to meet FDA requirements. However, in addition to limiting the area and / or volume of the treated tissue, the objective of the present invention is achieved by creating a pulsed source of energy, thereby stimulating the activation or generation of HSPs by heating or stressing the treated cells and tissues, while allowing the treated cells and tissues to dissipate the excess heat generated within acceptable limits.

[0097] It is believed that the stimulation of HSP production according to the present invention can be effectively used to treat a wide range of tissue abnormalities, diseases, and even infections. For example, viruses that cause the common cold primarily affect small parts 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-section of a human head (48) is shown, in which an endoscope (14) is inserted into the nasal cavity (50), and energy (16), such as laser light, is directed within the nasal cavity (50) to be treated (18). The treated tissue (18) may be within the nasal cavity (50), including the nasal passages and nasopharynx.

[0098] To ensure absorption of laser energy or other energy sources, the wavelength may be adjusted to the infrared (IR) absorption peak of water, or an adjuvant dye may be used to act as a photosensitizer. In such cases, the procedure then consists of taking the adjuvant or applying it topically, waiting a few minutes for the adjuvant to penetrate the surface tissue, and then administering laser light or other energy source (16) to the target tissue (18) for several seconds via an optical fiber in an endoscope (14), as shown in Figure 24. To provide comfort to the patient, the endoscope (14) may be inserted after the application of a local anesthetic. If necessary, the above procedure may be repeated periodically, such as every one to two days.

[0099] The procedure stimulates the activation or production of heat shock proteins and promotes protein repair without damaging the treated cells and tissues. As discussed above, certain heat shock proteins have been found to play an important role not only in the immune response but also in the health of the targeted cells and tissues. The energy source may be monochromatic laser light, such as laser light with a wavelength of 810 nm, and may be administered in a manner similar to that described in the patent application of the cited references above, but may also be administered by means of an endoscope, as shown in Figure 24. The adjuvant dye is selected to increase laser light absorption. This includes particularly preferred methods and embodiments for carrying out the present invention, but it is recognized that other types of energy and delivery means may also be used to achieve the same objective in accordance with the present invention.

[0100] Referring to Figure 25, a similar situation exists for the influenza virus, where the primary target is the epithelium of the upper respiratory tree, specifically the upper 6th generation, in segments with a diameter greater than approximately 3.3 mm. A viscous thin layer separates the targeted epithelial cells from the airway lumen, and in this layer, antigen-antibody interaction occurs, resulting in viral inactivation.

[0101] Referring to Figure 25, the flexible optical tube (12) of the bronchoscope (14) is inserted from the mouth (52) of the individual through the throat and trachea (54) into the bronchi (56) of the airway tree. Here, laser light or other energy source (16) is administered and delivered to the tissue in this region of the uppermost segment in the same manner as described above with respect to Figure 24. It is considered that the wavelength of the laser or other energy should be selected to match the IR absorption peak of water present in the mucus in order to heat the tissue, stimulate the activation or generation of HSPs, and promote protein repair, along with its associated benefits.

[0102] Referring here to Figure 26, as illustrated, a colonoscope (14) may have a flexible optical tube (12) that is inserted into the anus and rectum (58) and into either the large intestine (60) or small intestine (62) to deliver a selected laser light or other energy source (16) to the area and tissue to be treated. This may be used to assist with other gastrointestinal problems in addition to the treatment of colon cancer.

[0103] Typically, the procedure described above can be performed similarly to a colonoscopy, in that all stool is removed from the intestines, the patient lies on their side, and the physician inserts the long, thin optical tube portion (12) of a colonoscope (14) into the rectum and moves it into the colon, large intestine (60), or small intestine (64) region to be treated. The physician can view the path of the inserted flexible member (12) via a monitor and can also view the tissue at the tip of the colonoscope (14) in the intestine to view the region to be treated. Using one of the other optical fibers or optical tubes, the tip of the endoscope (64) is oriented towards the tissue to be treated, and a source of laser light or other radiation (16) is delivered through one of the optical tubes of the colonoscope (14) to treat the region of tissue to be treated, as described above, in order to stimulate the activation or generation of HSPs in the tissue (18).

[0104] Referring here to Figure 27, another example in which the present invention can be conveniently used is the frequently referred to “leaky gut” syndrome, a disease of the gastrointestinal (GI) tract characterized by inflammation and other metabolic dysfunctions. The GI tract, like the retina, is vulnerable to metabolic dysfunction and is therefore expected to respond well to the treatment of the present invention. This can be carried out by subthreshold diode micropulse laser (SDM) treatment, as discussed above, or by other energy sources and means known in the art, as discussed herein.

[0105] Referring again to Figure 27, a flexible optical tube (12), such as an endoscope, is inserted through 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 optical tube (12) is oriented toward the tissue (18) being treated, and the laser light or other energy source (16) is oriented toward the tissue (18). It will be understood by those skilled in the art that a colonoscope may also be used and inserted through the rectum (58) into the stomach (66) or any tissue between the stomach and the rectum.

[0106] If necessary, chromophore dyes can be orally delivered to GI tissue to enable radiation absorption. 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 wavelength slightly longer at the water absorption peak, thus eliminating the need to apply the chromophore externally.

[0107] The present invention also considers 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 a capsule endoscope is relatively small in size, such as about 1 inch in length, so that it can be swallowed by a patient. Once the capsule or pill (68) is swallowed, enters the stomach, and passes through the GI tube, at the appropriate position, the capsule or pill (68) can receive power and signals via an antenna (70) or the like to activate an energy source (72), such as a laser diode and associated circuits, and a suitable lens (74) focuses the generated laser light or radiation through a transparent radiation cover (76) onto the tissue to be treated. It is understood that the position of the capsule endoscope (68) can be determined by various means such as external imaging and signal tracking, or by a small illuminated camera through which a physician can view an image of the GI tube through which the pill or capsule (68) is passing at that time. The capsule or pill (68) may be powered by its own power source, such as a battery, or may be powered externally via an antenna, so that the laser diode (72) or other energy source can produce a desired wavelength and pulsed energy source to treat the tissue and area to be treated.

[0108] As in the retinal treatment in a previous application, the radiation is pulsed to take advantage of micropulse temperature spikes and associated safety, and the power can be adjusted so that the treatment is completely harmless to the tissue. This may involve adjusting the peak power, pulse duration, and repetition rate to produce a spike temperature increase of approximately 10°C while maintaining a long-term temperature increase below the FDA-defined 1°C limit. If a pill form of delivery (68) is used, the device may be powered by a small rechargeable battery or by wireless inductive excitation, etc. Heated or stressed tissue stimulates the activation or production of HSPs, promoting protein repair and providing associated benefits.

[0109] From the previous examples, the techniques of the present invention are limited to the treatment of diseases near or on the internal surface of the body, which are readily available by fiber optics or other light delivery means. The reason why the application of SDM to activate HSP activity is limited to areas near the body surface or optically accessible areas is that the absorption lengths of IR radiation or visible light in the body are very short. However, it is diseases located in tissues or deep within the body that benefit from the present invention. Therefore, the present invention considers the use of ultrasound and / or radio frequency (RF) and even shorter wavelength electromagnetic (EM) radiation, such as microwaves, which have relatively long absorption lengths in body tissues. To activate therapeutic HSP activity in abnormal tissues that cannot be utilized for surface SDM, etc., the use of pulsed ultrasound is preferred over RF electromagnetic radiation. The source of pulsed ultrasound can also be used on or near the surface of the abnormality.

[0110] Referring to Figure 29, by using ultrasound, specific deep regions within the body can be precisely targeted by using one or more beams, each focused on a target site. Subsequently, pulsed heating is applied mostly only to the targeted region, where the beams are focused and overlapped.

[0111] As illustrated in Figure 29, an ultrasonic transducer (78), etc., generates multiple ultrasonic beams (80), which are coupled to the skin via an acoustic impedance matching gel and penetrate through the skin (82) and undamaged tissue in front of the focal point of the beams (80) into a target organ (84), such as the illustrated liver, and specifically into the target tissue (86) being treated where the ultrasonic beams (80) are focused. As mentioned above, pulsed heating is then applied only to the targeted focus region (86) where the focused beams (80) overlap. Tissues in front of and behind the focus region (86) are not heated or affected to a noticeable extent.

[0112] The present invention considers not only the treatment of surface tissue or near-surface tissue using laser light, etc., and the treatment of deep tissue using, for example, a focused ultrasound beam, but also the treatment of hematological disorders such as sepsis. As shown above, focused ultrasound treatment can be used not only on the surface but also on deep body tissue, and can be applied when treating this hematological disorder (blood). However, it has also been considered that similar treatment options, which are typically limited to the treatment of SDM and the surface or near the surface of epithelial cells, etc., should be used when treating hematological disorders in areas where blood is accessible through relatively thin layers of tissue, such as the earlobe.

[0113] Referring here to Figures 30 and 31, treatment of blood disorders simply requires the transmission of SDM or other electromagnetic radiation or ultrasonic pulses to the earlobe (88), where the SDM or other source of energy can travel through the earlobe tissue into the blood flowing through the earlobe. It will be recognized that this approach can be performed in other areas of the body where blood flow is relatively high, and / or near tissue surfaces such as the fingertips, mouth, or throat.

[0114] Referring here to Figures 30 and 31, the earlobe (88) is shown adjacent to a clamping device (90) configured to transmit SDM radiation, etc. This may be, for example, by one or more laser diodes (92) that transmit desired frequencies in desired pulses and pulse trains to the earlobe (88). Power may be supplied, for example, by a lamp drive (94). Alternatively, the lamp drive (94) may be an actual source of laser light transmitted to the earlobe (88) through appropriate optical elements and electronic devices. The clamping device (90) is simply used to clamp onto the patient's earlobe (88) and to suppress radiation to the patient's earlobe (88). This may be done by mirrors, reflectors, diffusers, etc. This may be controlled by a control computer (96) operated by a keyboard (98), etc. The system may also include a display and speaker (100) if necessary, for example, when the above procedure is performed by an operator away from the patient.

[0115] The proposed treatment using a series of electromagnetic or ultrasonic pulses offers two major advantages over initial treatments incorporating either a single short pulse or a maintenance (long) pulse. First, individual short pulses (preferably less than a second) in a series activate cellular reset mechanisms, such as HSP activation, with larger reaction rate constants than those operating on longer (minute or hour) timescales. Second, repeated pulses in the treatment provide a large thermal spike (approximately 10,000) that allows the cellular repair system to more rapidly overcome the activation energy barrier that separates the dysfunctional cellular state from the desired functional state. The end result is a "lower treatment threshold," meaning that lower applied average power and total applied energy can be used to achieve the desired treatment objective.

[0116] While various embodiments have described the purpose of the illustrations in detail, various modifications may be made without departing from the scope and spirit of the invention. Accordingly, the invention is not limited to the appended claims.

Claims

1. A system adapted to stimulate the activation of heat shock proteins in target tissue without damaging the target tissue, wherein the system is The apparatus comprises an ultrasonic transducer (78) that generates at least one pulsed ultrasonic beam (16, 80) having a frequency of 1 to 5 MHz and a power of 0.46 to 28.6 W, wherein the apparatus emits at least one ultrasonic beam having a pulse train duration of 0.1 to 0.5 seconds and a duty cycle of 2% to 10%, and focuses it onto the target tissue. When applied to the target tissue (18), the at least one pulsed ultrasonic beam raises the temperature of the target tissue (18) by 6 to 11 degrees Celsius during application. A system characterized by the interval of pulses due to the duty cycle and the duration of the pulse train such that the rise in temperature of the target tissue is attenuated by thermal diffusion and the average temperature rise over a 6-minute period during the application of the at least one pulsed ultrasonic beam is limited to 6 degrees Celsius or less, so as not to destroy or damage the target tissue.

2. The system according to claim 1, wherein the at least one ultrasonic beam comprises a plurality of ultrasonic beams that are focused on the target tissue.

3. The system according to claim 1, wherein the at least one pulsed ultrasound beam is applied to a region of the body adjacent to the target tissue, i.e., near the surface of the region of the body, and having a blood supply that includes the target tissue.

4. The system according to claim 3, wherein the region of the body includes an earlobe, and the at least one pulsed ultrasound beam is applied to the blood flowing through the earlobe to treat the blood containing the target tissue.

5. The system according to claim 1, wherein the target tissue temperature is raised by 10 to 11 degrees Celsius during the application of the at least one pulsed ultrasonic beam.

6. The at least one pulsed ultrasonic beam (16, 80) has a pulse train duration of 0.5 seconds and a pulse interval of 5 seconds, and is applied to the target tissue for a total pulse train time of 50 seconds, which is the sum of the pulse train duration and the pulse interval. The system according to claim 1, wherein a power of 5.8W to 17W is applied to the target tissue during the total pulse train time.

7. The system according to claim 1, wherein the ultrasonic transducer has a diameter of 5 cm and a focal length of 10 cm.

8. The system according to claim 1, wherein the volume of the target tissue is 0.02 to 2.4 cc.

9. The system according to claim 1, wherein the temperature of the target tissue decreases by 1 degree Celsius or less within 6 minutes after the application of the at least one pulsed ultrasonic beam.

10. The system according to claim 1, wherein the heat shock protein is activated in the target tissue by heating the target tissue with the at least one pulsed ultrasonic beam without damaging the target tissue.

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