Retinal phototherapy system and method having fixed parameters - Patents.com

Abstract

A retinal phototherapy or photostimulation system includes a laser console that generates at least one treatment beam having parameters for photostimulating or treating a target tissue of the retina without permanent damage. The parameters of the at least one treatment beam are fixed so that they cannot be altered by a healthcare provider. A projector or camera projects the at least one treatment beam onto at least a portion of the retina, and a scanning mechanism controllably directs the at least one treatment beam to a treatment area of ​​the retina.

Description

[Technical field]

[0001] The present invention is directed generally to retinal phototherapy or photostimulation. More particularly, the present invention is directed to retinal phototherapy or photostimulation systems and methods in which parameters of one or more treatment beams are selected and fixed so as not to be altered by a healthcare provider to ensure that targeted tissue in the retina is photostimulated or treated while not causing permanent damage to the retinal tissue. [Background technology]

[0002] Until the advent of thermal retinal photocoagulation, there was generally no effective treatment for many retinal diseases, such as diabetic retinopathy. In visible endpoint photocoagulation, absorption of laser light heats the pigmented tissue at the laser's sight. Thermal conduction spreads this temperature increase from the retinal pigment epithelium (RPE) and choroid to the overlying non-pigmented and adjacent unexposed tissue. The laser lesion is visualized so that the treatment site can be tracked. In fact, it has been thought that actual tissue damage and scar formation is necessary to achieve the effects of conventional photocoagulation. Photocoagulation has proven to be an effective means of creating retinal scars and has been the technical standard for decades in the treatment of various retinal diseases, including macular photocoagulation for diabetic macular edema and other retinal diseases.

[0003] It has been universally accepted for almost 50 years and remains the prevailing view that ectopic retinal damage is necessary for effective laser treatment of retinal vascular disease. Current retinal photocoagulation therapy has clear advantages over no treatment, but it does have drawbacks and drawbacks, as it produces visible gray-white to white retinal burns and scars. Conventional photocoagulation is often painful. Local anesthesia may be required, which carries its own risks. Alternatively, treatment may be separated into stages over an extended period of time to minimize treatment pain and postoperative inflammation. Transient vision loss is common following conventional photocoagulation. In fact, thermal tissue damage may be the sole cause of many potential complications of conventional photocoagulation that may lead to immediate and late vision loss.

[0004] The fovea / macular region is the part of the eyeball used for color vision and fine detail vision. The fovea is in the center of the macula and is where the concentration of cells needed for central vision is highest. This is where diseases such as age-related macular degeneration cause the most damage, but it is also where traditional photocoagulation light therapy cannot be used because damaging cells in the fovea would severely impair the patient's vision. Therefore, the fovea is avoided in current traditional photocoagulation therapy.

[0005] Recently, the inventors have discovered that subthreshold photocoagulation, without any visible damage or laser lesions or even detectable by any known means, including fundusoscopy, infrared, color, non-infrared, or autofluorescent fundus photography, and standard retro-mode, intravenous fundus fluorescein or indocyanine green angiography, or spectral-domain optical coherence tomography, at the time of treatment or thereafter, can produce beneficial therapeutic results similar to those of traditional tissue-damaging photocoagulation therapy, but without many of its drawbacks and complications.

[0006] The possible mechanism triggered by such subthreshold photocoagulation is to increase the temperature of tissue in a controlled manner to selectively stimulate the activation of heat shock proteins and / or the generation and promotion of protein repair, which may serve as a mechanism to therapeutically treat tissue. This micropulse train is thought to thermally activate heat shock proteins (HSPs) in the targeted tissue. In the case of retinal tissue, this process thermally activates HSPs in the RPE layer, located immediately behind the retinal layer containing the visually sensitive rods and cones, and these activated HSPs reset the diseased retina to a healthy state by removing damaged proteins that are repaired. The result is improved RPE function, enhanced retinal function and autoregulation, recovery of acute inflammation, reduction of chronic inflammation, and organizational immune regulation. The laser-triggered effects may slow, halt, or reverse the progression of retinal diseases, improve visual function, and reduce the risk of vision loss.

[0007] It was determined that with proper operating parameters, subthreshold photocoagulation treatments can be applied and ideally applied to the entire retina, including sensitive areas such as the fovea, without visible tissue damage or the resulting drawbacks or complications of conventional visible retinal photocoagulation treatments. It is believed that increasing tissue temperature in such a controlled manner to selectively stimulate heat shock protein activation may have benefits in other tissues as well. Selection of proper operating parameters, including wavelength, power, pulse train duration and duty cycle of the laser treatment beam, is important to provide therapeutic or stimulating benefits while not damaging the retinal tissue to which the laser light is applied. Selecting and using laser beam generation and application parameters outside of the acceptable ranges may result in damage to the patient's retina. Furthermore, even within the safely determined parameters, improper selection of combinations of laser treatment light parameters may result in less than ideal therapeutic or photostimulation. Summary of the Invention [Problem to be solved by the invention]

[0008] Thus, there is a continuing need for a retinal phototherapy or photostimulation system that generates one or more laser treatment beams having operating and application parameters selected to treat or photostimulate retinal tissue while not causing permanent damage to the retinal tissue. There is also a need for such a system in which optimal parameters are selected and, once set, are fixed and cannot be altered by a healthcare provider or the like to ensure that the treatment beam is effective and safe. The present invention fulfills these needs and provides other related advantages. [Means for solving the problem]

[0009] The present invention resides in a retinal phototherapy or photostimulation system and associated methods. The system generally comprises a laser console that generates at least one pulsed treatment beam. The at least one treatment beam has parameters of wavelength, power, pulse train duration, and duty cycle to photostimulate or treat retinal target tissue without causing permanent damage to the retinal target tissue. The parameters of the at least one treatment beam are fixed such that they cannot be altered by a healthcare provider.

[0010] At least one treatment beam has a wavelength between 530 nm and 1300 nm, a duty cycle of less than 10%, and a pulse train duration of 0.1 to 0.6 seconds. At least one treatment beam may have a wavelength between 750 nm and 850 nm, a duty cycle of 2% to 5%, and a pulse train duration of 0.15 to 0.5 seconds. The laser console has a power output of between 1.0 to 3.0 watts.

[0011] The projector or camera projects the at least one treatment beam onto at least a portion of the retina, the at least one treatment beam capable of forming a treatment spot on the retina having a size between 100 and 1000 micrometers.

[0012] A scanning mechanism controllably directs the at least one treatment beam to a treatment region of the retina. The at least one treatment beam can be applied to multiple target tissue regions, with adjacent target tissue regions separated to avoid thermal tissue damage.

[0013] The at least one therapeutic beam stimulates activation of heat shock proteins in the target tissue. The at least one therapeutic beam raises the target tissue temperature to a desired level while maintaining the average temperature rise of the target tissue below a predetermined level over a period of time so as not to cause permanent damage to the target tissue. The at least one therapeutic beam may raise the target tissue temperature to no more than 11° C. to achieve a therapeutic or prophylactic effect. The at least one therapeutic beam may raise the target tissue temperature to between 6° C. and 11° C., at least during application of the pulsed energy source to the target tissue. The average temperature rise of the target tissue over several minutes is maintained below a predetermined level so as not to cause permanent damage to the target tissue. The average temperature rise of the target tissue over several minutes, such as 6 minutes or less, may be maintained below 1° C.

[0014] Other features and advantages of the present invention will become apparent from the following more detailed description, taken in conjunction with the accompanying drawings, which illustrate, by way of example, the principles of the invention.

[0015] The accompanying drawings illustrate the invention, in which: [Brief description of the drawings]

[0016] [Figure 1A] 1 is a graph showing the average power of a laser source compared to the source radius of the laser and the duration of a pulse train. [Figure 1B] 1 is a graph showing the average power of a laser source compared to the source radius of the laser and the duration of a pulse train. [Figure 2A] 1 is a graph showing temperature decay time as a function of laser source radius and wavelength. [Figure 2B] 1 is a graph showing temperature decay time as a function of laser source radius and wavelength. [Diagram 3] FIG. 2 is a block diagram of a light generating unit having a light pipe extending therefrom for generating a timed train of pulses in accordance with the present invention. [Figure 4] 1 is a diagram illustrating a system used to generate a therapeutic laser beam in accordance with the present invention. [Diagram 5] FIG. 2 is a block diagram of an optical system used to generate a laser light geometric pattern in accordance with the present invention. [Figure 6] FIG. 2 is a top view of an optical scanning mechanism used in accordance with the present invention. [Figure 7] FIG. 7 is a partially exploded view of the optical scanning mechanism of FIG. 6, showing its various components. [Figure 8] 1 illustrates controlled offset exposure of an exemplary geometric pattern grid of laser spots for treating target tissue in accordance with an embodiment of the present invention. [Figure 9] 1 is a diagram illustrating the use of a geometric object in the form of a controllably scanned line to treat a region of target tissue. [Figure 10] 10 is a diagram similar to FIG. 9, but showing rotated geometric lines or bars for treating target tissue. [Figure 11] FIG. 2 is a block diagram illustrating an alternative embodiment of a system used to generate a therapeutic laser beam for treating tissue in accordance with the present invention. [Figure 12] FIG. 2 is a block diagram illustrating yet another embodiment of a system used to generate a therapeutic laser beam to treat tissue in accordance with the present invention. [Figure 13A] 1 is a graph depicting the behavior of HSP cell line components over time after a rapid increase in temperature. [Figure 13B] 1 is a graph depicting the behavior of HSP cell line components over time after a rapid increase in temperature. [Figure 14A] 1 is a graph depicting the behavior of HSP cell line components during the first minute after a rapid temperature increase. [Figure 14B] 1 is a graph depicting the behavior of HSP cell line components during the first minute after a rapid temperature increase. [Figure 14C] 1 is a graph depicting the behavior of HSP cell line components during the first minute after a rapid temperature increase. [Figure 14D] 1 is a graph depicting the behavior of HSP cell line components during the first minute after a rapid temperature increase. [Figure 14E] 1 is a graph depicting the behavior of HSP cell line components during the first minute after a rapid temperature increase. [Figure 14F] 1 is a graph depicting the behavior of HSP cell line components during the first minute after a rapid temperature increase. [Figure 14G] 1 is a graph depicting the behavior of HSP cell line components during the first minute after a rapid temperature increase. [Figure 14H] 1 is a graph depicting the behavior of HSP cell line components during the first minute after a rapid temperature increase. [Figure 15A] 1 is a graph showing the change in activated and unactivated HSP concentrations in a cytoplasmic reservoir over a one minute interval in accordance with the present invention. [Figure 15B] 1 is a graph showing the change in activated and unactivated HSP concentrations in a cytoplasmic reservoir over a one minute interval in accordance with the present invention. [Figure 16] 1 is a graph depicting percent improvement versus interval between treatments in accordance with the present invention. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0017] As shown in the accompanying drawings and more fully described herein, the present invention is directed to systems and methods for delivering pulsed energy, such as one or more light beams, having energy parameters selected to create a thermal time course in tissue that elevates tissue temperature over a short period of time to a level sufficient to achieve a therapeutic effect, while maintaining average tissue temperature over an extended period of time below a predetermined level so as to avoid permanent tissue damage. The creation of a thermal time course is believed to stimulate the activation or production of heat shock proteins, promoting protein repair without damage.

[0018] The inventors have discovered that electromagnetic radiation can be applied to retinal tissue in a manner that does not destroy or damage the retinal tissue while achieving a beneficial effect against ocular diseases. More specifically, a laser beam can be produced that is therapeutic yet sublethal to retinal tissue cells, thus avoiding damaging photocoagulation in the retinal tissue providing preventative and protective treatment of the retinal tissue of the eye. It is believed that this may be due, at least in part, to the stimulation and activation of heat shock proteins and promotion of protein repair in the retinal tissue.

[0019] Various parameters of the light beam must be considered and the combination of parameters selected to achieve a therapeutic effect while not causing permanent damage to the tissue. These parameters include the laser wavelength, the average laser power, the total pulse duration, and the duty cycle of the pulse train.

[0020] The selection of these parameters may be determined by requiring that the Arrhenius integral for HSP activation be greater than 1. The Arrhenius integral is used to analyze the effect of the action on biological tissue. At the same time, the selected parameters must not cause permanent damage to the tissue. Therefore, the Arrhenius integral for damage may also be used, where the solved Arrhenius integral is less than 1.

[0021] Alternatively, satisfy FDA / FCC constraints on energy deposition per gram of tissue and temperature rise measured over a period of minutes so as to avoid permanent damage to tissue. Generally speaking, tissue temperature rises between 6°C and 11°C can produce a therapeutic effect, such as by activating heat shock proteins, but maintaining the average tissue temperature for an extended period of time, such as over several minutes, e.g., 6 minutes, below a given temperature, e.g., 6°C or even 1°C in certain circumstances, will not cause permanent damage to tissue.

[0022] The inventors have discovered that generating subthreshold sublethal micropulse laser light with a wavelength greater than 530 nm and a duty cycle of less than 10% at a given intensity or power and a given pulse length or exposure time produces the desired retinal photostimulation without any visible burn area or tissue destruction. More specifically, a laser light with a wavelength between 530 nm and 1300 nm, and in a particularly preferred embodiment between 750 nm and 850 nm, with a duty cycle of about 2% to 5%, a pulse train duration of 0.15 to 0.5 seconds, at a given intensity or power (e.g., between 100 and 590 watts per square centimeter at the retina, or about 1 watt per laser spot for each treatment spot at the retina) and a given pulse length or exposure time (e.g., between 100 and 600 milliseconds) creates a sublethal "true subthreshold" retinal photostimulation in which all areas of the retinal pigment epithelium exposed to the laser irradiation are preserved and available to contribute therapeutically. In other words, the inventors have discovered that a method of elevating retinal tissue to at least therapeutic levels, but below cell or tissue lethal levels, produces the benefits of prior art methods without destroying, burning, or otherwise damaging the retinal tissue, which is referred to herein as Subthreshold Diode Micropulse Laser Treatment (SDM).

[0023] SDM does not cause laser-induced retinal damage (photocoagulation), has no known side effects, and has been reported to be an effective treatment for many retinal diseases, including diabetic macular edema (DME), proliferative diabetic retinopathy (PDR), macular edema due to branch retinal vein occlusion (BRVO), central serous chorioretinopathy (CSR), and for the restoration of drug resistance and preventive treatment of progressive degenerative retinopathies, such as dry age-related macular degeneration, Stargardt's disease, cone dystrophy, and retinitis pigmentosa. The safety of SDM is such that its transocular use in eyes with 20 / 20 visual acuity can reduce the risk of vision loss due to early stage foveolar-involved DME.

[0024] A possible mechanism by which SDM functions is the generation or activation of heat shock proteins (HSPs). Despite the nearly infinite variety of possible cellular abnormalities, all types of cells share a common and highly conserved repair mechanism: heat shock proteins (HSPs). HSPs are almost instantly induced, within seconds to minutes, by almost any type of cellular stress or injury. In the absence of lethal cellular injury, HSPs are highly effective in repairing surviving cells and returning them to a more normal functional state. HSPs are transient, generally peaking within hours and lasting for days, but their effects can last for longer periods. HSPs suppress inflammation, a common factor in many diseases.

[0025] Laser treatment can induce HSP generation or activation and alter cytokine expression. The more sudden and severe the non-lethal cellular stress (e.g., laser irradiation), the more rapid and robust the HSP activation. Thus, the repetitive bursts of cold thermal spikes at very steep rates of change (-7°C increase per 100 μs micropulse, or 70,000°C / s) produced by each SDM exposure are particularly effective in stimulating HSP activation, especially compared to non-lethal exposures for subthreshold treatments with continuous wave lasers, which can reproduce only low mean tissue temperature increases.

[0026] Laser wavelengths below 530 nm result in increasingly cytotoxic photochemical effects. Between 530 nm and 1310 nm, and especially around 810 nm, SDM produces photothermal, rather than photochemical, cellular stress. Thus, SDM can affect tissue without damaging it. Thus, the clinical benefit of SDM is primarily produced by subpathologic photothermal cellular HSP activation. In dysfunctional cells, HSP stimulation by SDM normalizes cytokine expression, resulting in improved structure and function. The therapeutic effect of this “low-intensity” laser-tissue interaction is then amplified by the application of a “high-intensity” laser to densely / fusionally treat a large tissue area, including the entire area of ​​pathology, thus recruiting all dysfunctional cells in the targeted tissue area and maximizing the therapeutic effect. These principles define the therapeutic strategy of SDM described herein.

[0027] HSP stimulation in normal cells tends not to have a significant clinical effect, since normally functioning cells have no need for repair. The “pathoselectivity” of near-infrared laser effects, such as SDM, on various cell types, affecting diseased cells but not normal cells, is consistent with clinical observations of SDM. SDM has been reported to have a clinically broad therapeutic window, consistent with the American National Standards Institute “maximum tolerable exposure” predictions, and unique among retinal laser therapies. Although SDM can cause direct photothermal effects, such as entropy protein unfolding and disaggregation, SDM appears to be optimized for the stimulation of clinically safe and effective HSP-mediated repair.

[0028] As mentioned above, SDM stimulation of HSPs is non-specific with respect to disease processes, whereas the outcome of HSP-mediated repair is, by its nature, specific to the dysfunctional state. HSPs tend to solve the problem, whatever it may be. Thus, the observed efficacy of SDM in retinal diseases, such as BRVO, DME, PDR, CSR, age-related and hereditary retinopathies, and drug-resistant NAMD, varies widely. Conceptually, this function can be considered as a kind of “reset to defaults” mode of SDM action. For a wide range of diseases where cellular function is critical, SDM normalizes cellular function by inducing a “reset” (to “factory default settings”) via HSP-mediated cellular repair.

[0029] We have found that SDM treatment of patients with age-related macular degeneration (AMD) can slow or even halt the progression of AMD. Most patients see significant improvements in dynamic functional log-MAR mesoptic visual acuity and mesoptic contrast visual acuity after SDM treatment. SDM appears to work by targeting, preserving, and "normalizing" (toward normal) retinal pigment epithelium (RPE) function.

[0030] SDM has also been shown to halt or reverse the development of diabetic retinopathy pathology without causing treatment-related damage or side effects, despite the persistence of systemic diabetes. Based on this, it has been hypothesized that SDM may function by inducing a return to more normal cellular function and cytokine expression in diabetes-affected RPE cells, similar to pressing the "reset" button on an electronic device to return it to factory default settings. Based on the above information and studies, SDM treatment may directly affect cytokine expression via activation of heat shock proteins (HSPs) in target tissues.

[0031] As shown above, subthreshold diode micropulse laser (SDM) photostimulation is effective in stimulating the direct repair of slightly misfolded proteins in ocular tissues. Besides HSP activation, another possible reason for this is that the temperature spike caused by the micropulse in the form of a thermal time course allows the diffusion of water inside the protein, which in turn allows 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 bound hydrogen bonds by about 1,000-fold. Therefore, this process could be advantageously applied to other tissues and diseases as well.

[0032] 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 causing permanent damage to the tissue. Using a beam energy source such as a laser beam as an example, the full parameters of the laser wavelength, duty cycle and duration of the pulse train must be considered. 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 can affect at least one other parameter.

[0033] Thus, in accordance with the present invention, to maintain safety and prevent user error, a retinal phototherapy or photostimulation system embodying the present invention has at least one treatment parameter, including wavelength, power, pulse train duration and duty cycle, preselected and fixed so that it cannot be altered by a healthcare provider or other end user. In this manner, a pulsed energy generator, such as a laser console, generates a pulsed energy source, such as a pulsed therapeutic laser beam, having a combination of parameters selected to photostimulate and / or treat retinal tissue without destroying, permanently damaging or even damaging the retinal tissue in any way. Since adjustment or selection of one parameter may affect at least one of the other parameters, the appropriate or most beneficial preselected parameters of wavelength, power, pulse train duration and duty cycle of at least one therapeutic beam may be selected and fixed prior to delivery to a healthcare provider or other end user. The preselected combination of parameters is safe and can be optimized. Because the healthcare provider or other end user cannot change the fixed operating parameters, there is no possibility of user error or inadvertent creation of a combination of operating parameters of the light therapy beam that could potentially damage retinal tissue.

[0034] For example, the at least one treatment beam may have a wavelength selected between 530 nm and 1300 nm, a duty cycle of less than 10%, and a pulse train duration of 0.1 seconds to 0.6 seconds. Within the aforementioned range of wavelengths, the duty cycle may be 2% to 5%. The peak laser power may be between 0.5 watts and 3.0 watts. The treatment spot size on the retinal tissue formed by the one or more treatment beams may be between 100 and 1000 micrometers. More specifically, the at least one treatment beam may have a wavelength between 750 nm and 850 nm, a duty cycle between 2% and 5%, and a pulse train duration between 0.15 seconds and 0.5 seconds. As an example, the system may have fixed parameters of 810 nm, a duty cycle of 5%, a pulse train duration of 0.3 seconds, a power of 1.8 watts, and a resulting spot size of 500 micrometers. An acceptable alternative for a wavelength of approximately 810 nm would be a duty cycle of 2%, a pulse train duration of 0.5 seconds, a peak laser power of 2.19 watts, and a resulting retinal treatment spot of 400 micrometers. Other specific combinations within the above ranges are also possible. However, preselecting and fixing these operating parameters prevents end users, such as healthcare providers, from adjusting any of the parameters outside of the acceptable ranges that may cause damage or not provide effective treatment.

[0035] Figures 1A and 1B show graphs showing the average power in watts compared to the radius of the laser source (between 0.1 cm and 0.4 cm) and the duration of the pulse train (between 0.1 and 0.6 seconds). Figure 1A shows a wavelength of 880 nm, and Figure 1B a wavelength of 1000 nm. From these figures, it can be seen that the required power decreases monotonically as the radius of the source decreases, as the total train duration increases, and as the wavelength decreases. The preferred parameters of the radius of the laser source are 1 mm to 4 mm. For a wavelength of 880 nm, the minimum value of the power is 0.55 watts, the radius of the laser source is 1 mm, and the total time of the duration of the pulse train is 600 ms. The maximum value of the power for a wavelength of 880 nm is 52.6 watts for a radius of the laser source of 4 mm and a total pulse train duration of 100 ms. However, if one chooses a laser with a wavelength of 1000 nm, the minimum power value is 0.77 Watts when the radius of the laser source is 1 mm and the total pulse train duration is 600 ms, and the maximum power value is 73.6 Watts when the radius of the laser source is 4 mm and the total pulse duration is 100 ms. The corresponding peak power during an individual pulse is obtained by dividing the average power by the duty cycle.

[0036] The volume of the tissue area heated is determined by the wavelength, the absorption length in the relevant tissue, and by the beam width. The total pulse duration and average laser power determine the total energy delivered to heat the tissue, and the duty cycle of the pulse train gives the associated spike, or peak, power, related to the average laser power. Preferably, the energy parameters of the pulsed energy source are selected such that approximately 20 to 40 Joules of energy are absorbed by each cubic centimeter of target tissue. In the thin melanin layer of the retinal pigment epithelium, the absorption length is very small.

[0037] It has been determined that the target tissue can be heated to approximately 11°C for a short period of time, such as less than 1 second, and can be maintained in a lower temperature range, such as less than 6°C, or even 1°C or less, for a longer period of time, such as several minutes, while still producing the therapeutic effects of the present invention. The selection of the duty cycle and total pulse train duration provides an interval during which the heat can dissipate. Duty cycles of less than 10%, preferably between 2.5% and 5%, and total pulse durations between 100 milliseconds and 600 milliseconds have been found to be effective. Figures 2A and 2B show the decay times from 10°C to 1°C for laser sources with radii between 0.1 cm and 0.4 cm and wavelengths of 880 nm in Figure 2A and 1000 nm in Figure 2B. It can be seen that the decay times are shorter when the 880 nm wavelength is used, but both wavelengths fall within the acceptable requirements and operating parameters to achieve the benefits of the present invention without causing permanent tissue damage.

[0038] It has been found that HSP activation occurs when the average temperature increase in the desired target area during the entire irradiation period is at least 6° C., up to 11° C., and preferably about 10° C. Control of the target tissue temperature is determined by selecting source and target parameters such that the Arrhenius integral of HSP activation is greater than 1, while ensuring compliance with conservative FDA / FCC requirements to avoid damage or the Arrhenius integral of damage being less than 1.

[0039] To meet conservative FDA / FCC constraints to avoid permanent tissue damage, for light and other electromagnetic radiation sources, the average temperature rise of the target tissue over any 6 minute period is no more than 1°C. Figures 2A and 2B above show the typical decay time required for the temperature of a heated target area to drop from a temperature rise of about 10°C to 1°C by thermal diffusion, as seen in Figure 2A, for a wavelength of 880 nm and a source diameter of 1 mm, with a temperature decay time of 16 seconds. For a source diameter of 4 mm, the temperature decay time is 107 seconds. As shown in Figure 2B, for a wavelength of 1000 nm, the temperature decay time is 18 seconds for a source diameter of 1 mm and 136 seconds for a source diameter of 4 mm. This is well within the time that the average temperature rise can be maintained over a period of several minutes, e.g., no more than 6 minutes. During application of the energy source to tissue, the temperature of the target tissue increases very quickly, e.g., in a fraction of a second, such as to approximately 10 °C; however, the relatively low duty cycle provides a relatively long time 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 time period consisting of several minutes, such as 6 minutes or less, without permanent tissue damage.

[0040] The pulse train mode of energy delivery has a clear advantage over single pulse or stepwise modes of energy delivery as far as activating repair HSPs and promoting protein repair are concerned. Two considerations contribute to this advantage:

[0041] First, a major advantage for HSP activation and protein repair in the SDM energy delivery mode is the creation of a spike temperature of around 10° C. This large increase in temperature has a significant impact on the number of HSPs activated and the Arrhenius integral, which quantitatively describes the rate of water diffusion into proteins that promotes protein repair, because temperature enters an exponential function with a large amplification effect.

[0042] It is important to not allow the temperature rise to remain at a high value (above 10°C) for an extended period of time, as this would violate FDA and FCC requirements that the average temperature rise be less than 1°C over several minutes.

[0043] The SDM mode of energy delivery uniquely satisfies both of these aforementioned considerations through judicious selection of power, pulse time, pulse interval, and volume of the treated target area. The volume of the treatment area comes into play because the temperature must decay fairly quickly from as high as 10°C in order for the long-term average temperature rise to not exceed the long-term FDA / FCC limit of 1°C or less for electromagnetic radiant energy sources.

[0044] Referring now to Figure 3, a light generating unit 10, such as a laser having a desired wavelength and / or frequency and other operating parameters, is used to generate electromagnetic radiation, such as laser light, in a controlled, pulsed manner to be delivered to a projector via a light tube or pipe 12. The light generating unit 10, consisting of, for example, a laser console, generates one or more pulsed energy sources, such as pulsed laser beams, having the above-mentioned operating parameters preselected and fixed so as not to be altered by a healthcare provider or other end user. It will be understood by those skilled in the art that the light tube 12 may be connected to a camera, projector, or the light transmitted by the light generating unit 10 may be transmitted directly to such a projector, camera, etc., without the need for a light tube or pipe, etc.

[0045] Referring now to FIG. 4, there is shown a schematic diagram of a system for generating electromagnetic energy radiation, such as laser light, including an SDM. 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 a preferred embodiment. The laser generates a laser beam that may be passed through optics, such as an optical lens or mask, or multiple optical lenses and / or masks 24, as needed. The laser projector optics 24 passes the shaped beam to a delivery device 26 for projecting the laser beam light onto the target tissue of the patient. It will be understood that the box labeled 26 may represent both the laser beam projector or delivery device as well as the viewing system / camera, or may consist of two different components in use. The viewing system / camera 26 provides feedback to a display monitor 28, which may also include the computerized hardware, data input and controls, etc. necessary to operate the laser 22, the optics 24, and / or the projection / viewing component 26.

[0046] Referring now to FIG. 5, in one embodiment, multiple beams are generated, each having selected parameters such that the target tissue temperature can be controllably increased to therapeutically treat the target tissue without destruction or permanent damage to the target tissue. This can be done, for example, by passing a laser beam 30 through an optical system that diffracts or otherwise generates multiple laser beams from a single laser beam 30 having selected parameters. For example, the laser beam 30 can be passed through a collimator lens 32 and then through a mask 34. In a particularly preferred embodiment, the mask 34 comprises a diffraction grating. The mask / diffraction grating 34 generates a geometric object, or more typically a geometric pattern of multiple laser spots or other geometric objects generated simultaneously. This is represented by multiple laser beams labeled with reference number 36. Alternatively, the multiple laser spots may be generated by multiple fiber optic waveguides. Either method of generating the laser spots allows for a very large number of laser spots to be generated simultaneously over a very large treatment area. Indeed, a very large number of laser spots, perhaps hundreds, even thousands or even more, may be generated simultaneously to cover a given area of ​​target tissue, or even the entirety of the target tissue in some cases. A wide array of small, separated laser spots applied simultaneously may be desirable, as it thus avoids certain drawbacks and treatment risks known to be associated with large laser spot applications.

[0047] By using optical features with feature sizes comparable to the wavelength of the laser employed, for example using diffraction gratings, it is possible to exploit quantum mechanical effects that allow for the simultaneous application of a very large number of laser spots over a very large target area. The individual spots produced by such diffraction gratings are all of similar optical geometry relative to the input beam with minimal power variation for each spot. As a result, multiple laser spots with sufficient irradiance to provide a harmless yet effective therapeutic application are produced simultaneously over a large target area. The present invention also contemplates the use of patterns produced by other geometric objects and other diffractive optical elements.

[0048] The laser light passing through the mask 34 diffracts and produces a periodic pattern away from the mask 34, shown by the laser beams labeled 36 in FIG. 21. In this manner, a single laser beam 30 is formed into hundreds or even thousands of individual laser beams 36 to create a desired pattern of spots or other geometric objects. These laser beams 36 can pass through additional lenses, collimators, etc. 38 and 40 to transmit the laser beams to form the desired pattern. Such additional lenses, collimators, etc. 38 and 40 can further transform and redirect the laser beam 36 as needed.

[0049] By controlling the shape, spacing and pattern of the optical mask 34, any pattern can be constructed. The patterns and exposure spots can be created and modified as desired by experts in the field of optical engineering according to the application requirements. Photolithography techniques, particularly those developed in the field of semiconductor manufacturing, can be used to simultaneously create geometric patterns of spots or other objects.

[0050] Because the parameters and methodology of the present invention produce a therapeutically effective, yet non-destructive, non-permanently damaging treatment, the present invention can use a large number of simultaneously generated therapeutic beams or spots (e.g., tens or even hundreds). Hundreds or even thousands of simultaneous laser spots can be generated and created to form patterns that are simultaneously applied to tissue, but the requirement of not overheating the tissue limits the number of treatment spots or beams that can be used simultaneously in accordance with the present invention. For each individual laser beam or spot to be effective, a minimum average power over the train duration is required. At the same time, however, the tissue cannot exceed a certain temperature rise without being damaged. For example, using an 810 nm wavelength laser, the number of spots generated and used simultaneously can be as low as 1 to as high as about 100 when used with a duty cycle of 0.04 (4%) and a total train duration of 0.3 seconds (300 milliseconds). Water absorption increases with increasing wavelength. For shorter wavelengths, e.g., 577 nm, laser power can be lower. For example, at 577 nm, the power can be reduced by a factor of four for the present invention to be effective. Thus, when using a 577 nm wavelength laser light, there can be as little as one laser spot, or up to about 400 laser spots, without scratching or damaging the tissue.

[0051] Typically, the system of the present invention incorporates a guidance system to ensure complete and holistic retinal treatment by retinal photostimulation of the entire retina, or selected portions of the retina. Fixation targets, tracking mechanisms, and fixation / tracking / registration systems configured in conjunction with system operation can be incorporated into the present invention. The geometric pattern of simultaneous laser spots can be offset sequentially to achieve a confluent and complete treatment of the surface.

[0052] This can be done in a controlled manner using an optical scanning mechanism 50. Figures 6 and 7 show an optical scanning mechanism 50 in the form of a MEMS mirror, having a base 52 with electronically actuated controllers 54 and 56 that serve to tilt and pan the mirror 58 when electricity is applied or removed thereto. Application of electricity to the controllers 54 and 56 causes the mirror 58 to move and thus the simultaneous pattern of laser spots or other geometric objects reflected thereon to move accordingly on the patient's retina. This can be done in an automated manner using an electronic software program to adjust the optical scanning mechanism 50, for example, until complete coverage of the retina, or at least the portion of the retina desired to be treated, is exposed to the phototherapy. The optical scanning mechanism can also be a small beam diameter scanning galvo mirror system, for example sold by Thorlabs, or a similar system. Such a system can scan the laser in a desired offset pattern.

[0053] The pattern of spots is offset with each exposure to create a space between the previous exposure to allow for heat dissipation and prevent possible thermal damage or tissue destruction. Thus, as shown in FIG. 8, a pattern shown for illustrative purposes as a grid of 16 spots is offset with each exposure so that the laser spots occupy a different space than the previous exposure. It will be understood that the schematic use of circles or empty and filled dots is for schematic purposes only to show previous and subsequent exposures to areas of the pattern of spots in accordance with the present invention. The spacing of the laser spots prevents overheating and damage to the tissue. It will be understood that this occurs until the entire target tissue being treated has been exposed to phototherapy or until the desired effect is achieved. This can be done, for example, by applying an electrostatic torque to a micromachined mirror as shown in FIG. 6 and FIG. 7. By combining the use of small laser spots separated by exposure free areas, the prevention of heat accumulation, and a grid with a large number of spots per side, it is possible to treat large target areas intact and invisibly with short exposure durations much more quickly than is possible with current technology.

[0054] By rapidly and sequentially redirecting or offsetting the entire simultaneously applied grid array of spots or geometric objects, complete coverage of the target can be achieved quickly and without thermal tissue damage. This offset can be algorithmically determined to ensure the fastest treatment time and the least risk of thermal tissue damage, depending on the laser parameters and the desired application.

[0055] Additionally, quantum mechanical behavior may be observed due to the small apertures employed in the diffraction grating or mask, which allows for arbitrary distribution of the laser input energy. This allows for the generation of any geometric shape or pattern, such as multiple spots, lines, or any other desired pattern in a grating pattern. Other methods of generating geometric shapes or patterns, such as using multiple optical fibers or micro-lenses, may also be used with the present invention. The time savings from the use of simultaneous projection of geometric shapes or patterns may allow for a 1.2 cm projection to achieve full retinal treatment in a single clinical setting or treatment session. 2 This allows for novel sized treatment fields, such as areas of the retina. Although the generation and application of multiple laser beams to simultaneously create corresponding treatment spots of multiple target tissues has its advantages, it will be understood that the present invention may be utilized with only a single treatment laser beam that is controllably moved over the desired treatment area of ​​the retina.

[0056] 9, instead of a geometric pattern of small laser spots, the present invention contemplates the use of other geometric objects or patterns. For example, a single line 60 of laser light can be created, formed by a series of contiguous or closely spaced spots. An offset optical scanning mechanism can be used to sequentially scan the line over the area indicated by the downward arrow in FIG. 9.

[0057] 10, the same geometric object of line 60 can be rotated as shown by the arrows to create a circular field of phototherapy. However, a potential downside to this approach is that the central region may be repeatedly exposed and reach unacceptable temperatures. However, this issue can be resolved by increasing the time between exposures or by creating gaps in the line so that the central region is not exposed.

[0058] In the field of photobiology, it has become apparent that various biological effects may be achieved by exposing target tissue to lasers of different wavelengths. The same may also be achieved by sequential application of multiple lasers of different or same wavelengths in sequence with variable separation periods and / or different irradiation energies. The present invention contemplates the simultaneous or sequential application of multiple laser, light, radiation wavelengths (or modes) to maximize or customize the desired therapeutic effect. This method also minimizes potential deleterious effects. The optical methods and systems shown and described above provide for simultaneous or sequential application of multiple wavelengths.

[0059] FIG. 11 illustrates a schematic of a system for coupling multiple therapeutic light sources to the pattern generating optical subassembly described above. Specifically, this system 20' is similar to the system 20 described in FIG. 4 above. The primary difference between the alternative system 20' and the previously described system 20 is the inclusion of multiple laser consoles, each of whose outputs is fed into a fiber coupler 42. Each laser console can provide a laser beam having different parameters, such as different wavelengths. The fiber coupler produces a single output that is passed to the laser projector optics 24, as described in the previously described system. The coupling of the multiple laser consoles 22 into a single optical fiber is accomplished using a fiber coupler 42, as known in the art. Other known mechanisms for combining multiple light sources are available and can be used in place of the fiber coupler described herein.

[0060] In this system 20', the multiple light sources 22 follow a similar path as described in the previous system 20, i.e., collimated, diffracted, recollimated, and directed to the projector apparatus and / or tissue. In this alternative system 20', the diffractive elements must function differently than described above depending on the wavelength of light passing through, resulting in slightly varying patterns. The variation is linear with respect to the wavelength of the light source being diffracted. In general, the difference in diffraction angles is small enough that different overlapping patterns can be directed along the same optical path through the projector apparatus 26 to the tissue for treatment.

[0061] The resulting pattern changes slightly for each wavelength, so the sequential offset to achieve complete coverage is different for each wavelength. This sequential offset can be performed in two modes. In the first mode, all wavelengths of light are applied simultaneously without identical coverage. An offset steering pattern is used to achieve complete coverage for one of the multiple wavelengths. Thus, the light of the selected wavelength achieves complete coverage of the tissue, while the application of the other wavelengths achieves incomplete or overlapping coverage of the tissue. In the second mode, each light source of the various wavelengths is applied in sequence with an appropriate steering pattern to achieve complete coverage of the tissue at that particular wavelength. This mode eliminates the possibility of simultaneous treatment with multiple wavelengths, but allows identical coverage at each wavelength to be achieved by optical methods. This avoids incomplete or overlapping coverage for any of the light wavelengths.

[0062] These modes can also be mixed and matched: for example, two wavelengths can be applied simultaneously to achieve complete coverage at one wavelength and incomplete or overlapping coverage at another wavelength, followed by sequential application of a third wavelength to achieve complete coverage.

[0063] FIG. 12 diagrammatically illustrates yet another alternative embodiment of a system 20″ of the present invention. This system 20″ is configured substantially similarly to the system 20 illustrated in FIG. 4. The primary difference is the inclusion of multiple pattern generating subassembly channels tuned to specific wavelengths of light sources. Multiple laser consoles 22 are arranged in parallel, each directly connected to its own laser projector optics 24. As above, the operating parameters of the laser consoles 22 and the parameters of the resulting laser therapy beam are preselected and fixed so that they cannot be altered by a healthcare provider or other end user in accordance with the present invention. The parameters of one or more, or even each, of the laser therapy beams may be different from one another, but are still preselected and fixed so that they cannot be altered.

[0064] The laser projector optics for each channel 44a, 44b, 44c is an entire set of optics tuned for the particular wavelength produced by the corresponding laser console 22, consisting of a collimator 32, a mask or grating 34, and a recollimator 38, 40, as described in connection with FIG. 11 above. The output from each set of optics 24 is then directed to a beam splitter 46 for combination with other wavelengths. Those skilled in the art will know that beam splitters used in reverse can be used to combine multiple beams of light into a single output. The combined channel output from the final beam splitter 46c is then directed through the projector device 26.

[0065] In this system 20'', the optical elements in each channel are adjusted to produce a precisely specified pattern for that channel's wavelength. As a result, when all channels are combined and properly aligned, a single steering pattern can be used to achieve complete coverage of tissue for all wavelengths. System 20'' can use as many channels 44a, 44b, 44c, etc. and beam splitters 46a, 46b, 46c, etc. as there are wavelengths of light used for treatment.

[0066] Implementations of system 20'' can exploit different symmetries to reduce the number of alignment constraints. For example, the proposed grid pattern is periodic in two dimensions and is steered in two dimensions to achieve perfect coverage. As a result, if the pattern for each channel is identical as specified, the actual pattern for each channel does not need to be aligned to the same steering pattern to achieve perfect coverage for all wavelengths. Each channel is only optically aligned to achieve efficient combining.

[0067] In system 20'', each channel begins with a light source 22, which can be from an optical fiber as in other embodiments of the pattern generating subassembly. This light source 22 is directed to an optical assembly 24 for collimation, diffraction, recollimation, and to a beam splitter that combines the channel with a main output.

[0068] It will be understood that the laser light generating systems shown in Figures 4-12 are exemplary, and other devices and systems can be utilized to generate a source of SDM laser light that can be operatively passed to the projector apparatus.

[0069] The proposed treatment with trains of electromagnetic pulses has two major advantages compared to previous treatments incorporating single short or sustained (long) pulses. First, the short (preferably sub-second) individual pulses in the train activate cellular resetting mechanisms such as HSP activation with a larger reaction rate constant than those operating on longer (minute or hour) time scales. Second, the repeated pulses in the treatment provide large thermal spikes (on the order of 10,000) that allow the cellular repair systems to more quickly overcome the activation energy barrier that separates the dysfunctional cellular state from the desired functional state. The net result is a "lowering of the treatment threshold" in the sense that lower applied average power and total applied energy can be used to achieve the desired therapeutic goal.

[0070] The power limitations of current micropulse diode lasers require rather long exposure durations. The longer the exposure, the more critical the ability of the central spot to dissipate heat relative to the unexposed tissue at the edge of the laser spot becomes. Therefore, the micropulse laser beam of an 810 nm diode laser should have an exposure envelope duration of no more than 500 ms, preferably around 300 ms. Of course, if the micropulse diode laser becomes more powerful, the exposure duration should be shortened accordingly.

[0071] Apart from the power limit, another parameter of the present invention is the duty cycle, i.e. the frequency of the micropulse train, or the length of the thermal relaxation time between successive pulses. The use of a 10% duty cycle or higher, adjusted to deliver a micropulse laser at similar irradiance at similar MPE levels, has been found to significantly increase the risk of lethal cell injury. However, duty cycles below 10%, and preferably below 5%, show sufficient thermal rise and treatment at the level of the MPE cells to stimulate a biological response, but remain below the level expected to produce lethal cell injury. However, the lower the duty cycle, the longer the exposure envelope duration, which can exceed 500 ms in some cases.

[0072] Each micropulse lasts for a fraction of a millisecond, typically 50 to 100 microseconds in duration. Thus, with exposure envelope durations of 300-500 milliseconds, and duty cycles of less than 5%, there is a significant amount of time between micropulses to allow for thermal relaxation time between successive pulses. Typically, a thermal relaxation time delay of 1-3 milliseconds, preferably about 2 milliseconds, is required between successive pulses. For adequate treatment, cells are typically exposed or hit at each location between 50-200 times, preferably between 75-150 times, and with relaxation or interval times of 1-3 milliseconds, the total time according to the above-described embodiments to treat a given area exposed to the laser spot is typically less than 1 second, for example between 100 milliseconds and 600 milliseconds on average. The thermal relaxation time is required to avoid overheating the cells in that location or spot and to prevent the cells from being damaged or destroyed.

[0073] Adjacent exposed regions must be spaced at least a predetermined minimum distance apart to avoid thermal tissue damage. Such distance is at least 0.5 diameters away from the immediately preceding treated location or region, and more preferably 1-2 diameters away. Such spacing is relative to the actual treated location in the previous exposed region. It is contemplated by the present invention that a relatively large region may actually include multiple offset exposed regions therein.

[0074] Although the invention is described for use in conjunction with a micropulse laser, in theory, a continuous wave laser could potentially be used in place of a micropulse laser. However, with a continuous wave laser, there are concerns about overheating as the laser moves from location to location in that the laser does not stop, and there can be heat leakage and overheating between treatment areas. Thus, while it is theoretically possible to use a continuous wave laser, in practice it is not ideal and a micropulse laser is preferred.

[0075] As described above, the controlled method of adding energy to a target tissue is intended to increase the temperature of the target tissue to therapeutically treat the target tissue without destroying or permanently damaging the target tissue. Such heating activates HSPs, and the thermally activated HSPs are believed to act to reset the diseased tissue to a healthy state, such as by removing and / or repairing damaged proteins. The inventors believe that maximizing such HSP activation will improve the therapeutic effect on the target tissue. As such, understanding the behavior and activation of HSPs and HSP-based species, their generation and activation, the temperature ranges for activating HSPs, and the time frames for HSP activation or generation and inactivation can be utilized to optimize thermal treatment of biological target tissues.

[0076] As discussed above, the target tissue is heated by the pulse of energy for a short period of time, such as 10 seconds or less, typically less than 1 second, for example between 100 ms and 600 ms. The time that the energy is actually applied to the target tissue is typically much shorter than this to allow for an interval of time for thermal relaxation so that the target tissue is not overheated and damaged or destroyed. For example, as discussed above, the laser light pulse can last on the order of microseconds with an interval of relaxation time of a few milliseconds.

[0077] Thus, understanding the subsecond behavior of HSPs may be important for the present invention. Thermal activation of HSPs in SDM typically involves the associated Arrhenius integral, Ω=∫dt A exp[-E / k B T(t)] [1] where the integral is over the treatment time and A is the Arrhenius rate constant for HSP activation, E is the activation energy, T(t) is the temperature of the thin RPE layer, including the laser-induced temperature rise.

[0078] The laser-induced temperature increase, and therefore the activation Arrhenius integral, depends on both the treatment parameters (e.g., laser power, duty cycle, total train duration) and the RPE characteristics (e.g., absorption coefficient, density of HSPs). It has been found clinically that effective SDM treatments are obtained when the Arrhenius integral is around 1.

[0079] In the Arrhenius integral formalism, only the forward reaction, i.e. the reaction of HSP activation, is considered. The reverse reaction, whereby activated HSPs return to an inactivated state, is not considered at all. For the subsecond durations typical of SDM treatment, this seems quite adequate. However, for longer times (e.g., 1 min or more), this formalism is not a good approximation. For such long times, a whole series of reactions occurs, resulting in a much smaller effective rate of HSP activation. This is the case during the proposed intervals between SDM applications of about 1 min in this disclosure.

[0080] In the published literature, the generation and destruction of heat shock proteins (HSPs) in cells is usually described over longer time periods by a set of 9–13 coupled mass-balanced differential equations that describe the behavior of the various molecular species involved in the life cycle of HSP molecules. These coupled equations are usually solved by computer to show the temporal behavior of HSPs and other species after a sudden increase in temperature.

[0081] All these equations are conservation equations based on the reactions of various molecular species involved in the activity of HSPs. To describe the behavior of HSPs in intervals of about 1 min during repeated applications of SDM, we use the equations described in Journal of the Royal Society Interface 10, No. 88, 20130527 (Rybinski et al. (2013)), M. Rybinski, Z. Szymanska, S. Lasota, A. Gambin (2013), Modeling the effects of hyperthermia treatment. The species considered in Rybinski et al. (2013) are listed in Table 1.

[0082] Table 1. Description of HSP species from Rybinski et al. (2013) HSP: Ubiquitous heat shock protein with a molecular weight of 70 Da (free and activated state) HSF Heat shock (transcription) factor with no DNA binding activity HSF3 (trimer) Heat shock factor formed from HSF and capable of binding to DNA HSE Heat shock element, a DNA site that initiates transcription of HSPs when bound by HSF3 mRNA Messenger RNA molecules that generate HSPs S HSP binding substrate, damaged protein P Properly folded protein HSP.HSF Complex of HSP bound to HSF (inactive HSP) HSF3.HSE A complex of HSF3 bound to HSE that induces transcription and the production of new HSP mRNA molecules. HSP.S A complex of HSPs attached to damaged proteins (HSPs that actively repair proteins)

[0083] These ten coupled simultaneous mass conservation equations are summarized below as equations [2]-

[11] . d[HSP] / dt=(l1+k 10 )[HSPS]+l2[HSPHSF]+k4[mRNA]-k1[S][HSP]-k2[HSP][HSF]-l3[HSP][HSF3]-k9[HSP] [2] d{HSF] / dt=l2[HSPHSF]+2l3[HSP][HSF3]+k6[HSPHSF][S]-k2[HSP][HSF]-3k3[HSF] 3 -l6[HSPS][HSF] [3] d[S] / dt=k 11 {[P]+l1[HSPS]+l6[SPS][HSF]-k1[S][HSP]-k6[HSPHSF][S] [4] d[HSPHSF] / dt=k2[HSP][HSF]+l6[HSPS][HSF]+l3[HSP][HSF3]-l2[HSPHSF]-k6[HSPHSF][S] [5] d[HSPS] / dt=k1[S][HSP]+k6[HSPHSF][S]-(l1+k 10 [HSPS]-l6[HSPS][HSF] [6] d[HSF3] / dt=k3[HSF] 3 +l7[HSF3][HSE]-l3[HSP][HSF3]-k7[HSF3][HSE] [7] d[HSE] / dt=l7[HSF3][HSE]-k7[HSF3][HSE] [8] d[HSF3HSE] / dt=k7[HSF3][HSE]-l7[HSF3][HSE] [9] d[mRNA] / dt=k8[HSF3HSE]-k5[mRNA]

[10] d[P] / dt=k 10 [HSPS]-k 11 [P]

[11]

[0084] In these equations, [ ] indicates the cell concentration of the amount in brackets. For Rybinski et al. (2013), the initial concentrations at the equilibrium temperature of 310 K are shown in Table 2.

[0085] Table 2. Initial species values ​​at 310 K for a typical cell in arbitrary units [Rybinski et al. (2013)]. The arbitrary units were chosen by Rybinski et al. for computational convenience, i.e., to bring the quantities of interest into the range of 0.01–10. [HSP(0)] 0.308649 [HSF(0)] 0.150836 [S(0)] 0.113457 [HSPHSF(0)] 2.58799 [HSPS(0)] 1.12631 [HSF3(0)] 0.0444747 [HSE(0)] 0.957419 [HSF3HSE(0)] 0.0425809 [mRNA(0)] 0.114641 [P(0)] 8.76023

[0086] The rate constants from Rybinski et al. (2013) are shown in Table 3.

[0087] Table 3. Rate constants from Rybinski et al. (2013) are the rate constants for any concentration unit in the previous table, expressed as min -1 As shown in Fig. l1=0.0175 k1=1.47 l2=0.0175 k2=1.47 l3=0.020125 k3=0.0805 k4=0.1225 k5=0.0455 k6=0.0805 l6=0.00126 k7=0.1225 l7=0.1225 k8=0.1225 k9=0.0455 k 10 =0.049 k 11 =0.00563271

[0088] The initial concentration values ​​in Table 2 and the rate constants in Table 3 were determined by Rybinski et al. (2013) to correspond to experimental data on the behavior of the entire HSP system when the temperature increases by about 5 °C over a period of several (e.g., 350) minutes.

[0089] Note that the initial concentration of HSPs is {100 x 0.308649 / (8.76023 + 0.113457 + 1.12631)} = 3.09% of the total number of proteins present in the cell.

[0090] The rate constants in Table 3 are used by Rybinski et al. at T = 310 + 5 + 315 K, but very similar rate constants may exist at other temperatures. In this regard, the qualitative behavior of the simulations is similar for a wide range of parameters. For convenience, we assume that the rate constant values ​​in Table 3 are a good approximation of the values ​​at the equilibrium temperature of T = 310 K.

[0091] The behavior of the different components in Rybinski et al.'s cell is displayed in Figure 13 for 350 min when the temperature is suddenly increased by 5 K at t = 0 from the ambient 310 K.

[0092] With continued reference to FIG. 13, the behavior of HSP cell line components for 350 minutes following a rapid temperature increase from 37° C. to 42° C. is shown.

[0093] Here, concentrations of components are presented in arbitrary units convenient for calculations. S denotes denatured or damaged protein not yet affected by HSP, HSP denotes free (activated) heat shock protein, HSP:S denotes activated HSP attached to damaged protein and carrying out repair, HSP:HSF denotes (inactive) HSP attached to heat shock factor monomer, HSF denotes heat shock factor monomer, HSF3 denotes a heat shock factor trimer that can penetrate the nuclear membrane and interact with heat shock elements on DNA molecules, HSE:HSF3 denotes a heat shock factor trimer that attaches to heat shock elements on DNA molecules and initiates the transcription of new mRNA molecules, and mRNA denotes a messenger RNA molecule arising from HSE:HSF3, leading to the production of new (activated) HSP molecules in the cytoplasm of the cell.

[0094] Figure 13 shows that the initial concentration of activated HSPs is the result of the release of HSPs sequestered in the molecule HSPHSF in the cytoplasm, and that new HSPs are not made from the cell nucleus via mRNA until 60 minutes after the temperature increase occurs. Figure 13 also shows that activated HSPs attach very quickly to damaged proteins and begin the repair process. In the depicted cell, the sudden temperature increase also temporarily increases the concentration of damaged proteins, with the peak concentration of damaged proteins occurring approximately 30 minutes after the temperature increase.

[0095] Figure 13 shows what the formula of Rybinski et al. predicts for the variation of 10 different species over a period of 350 minutes. However, the present invention is concerned with SDM application with respect to the variation of species over the much shorter O (minute) interval between two applications of SDM at any single retinal locus. Although a preferred embodiment of SDM in the form of laser light therapy has been analyzed and described, it will be understood that it is also applicable to other energy sources.

[0096] 14A-H, the behavior of HSP cell line components during the first minute after a rapid temperature increase from 37° C. to 42° C. is shown using the equations of Rybinski et al. (2013) with the initial values ​​and rate constants in Tables 2 and 3. The horizontal axis shows time in minutes and the vertical axis shows concentration in the same arbitrary units as in FIG.

[0097] Figure 15 shows that the nuclear source of HSPs plays virtually no role during the 1 min period, and that the major source of new HSPs in the cytoplasm arises from the release of sequestered HSPs from the reservoir of HSPHSF molecules. It is also shown that a good proportion of the newly activated HSPs attach to damaged proteins and initiate the repair process.

[0098] The initial concentrations in Table 2 are not the equilibrium values ​​of the species, i.e. they do not give d[...] / dt=0, as is evident from the curves in Figures 13 and 14. It can be seen that the equilibrium values ​​giving d[...] / dt=0 corresponding to the rate constants in Table 3 are those given in Table 4.

[0099] Table 4. Equilibrium values ​​of species in arbitrary units [Rybinski et al. (2013)] corresponding to the rate constants in Table 3. The arbitrary units were chosen by Rybinski et al. for convenience of calculations and to bring the quantities of interest into the range of 0.01–10. [HSP(equil)] 0.315343 [HSF(equil)] 0.255145 [S(equil)] 0.542375 [HSPHSF(equil)] 1.982248 [HSPS(equil)] 5.05777 [HSF3(equil)] 0.210688 [HSE(equil)] 0.206488 [HSF3HSE(equil)] 0.643504 [mRNA(equil)] 0.1171274 [P(equil)] 4.39986

[0100] The equilibrium concentration of HSPs is 100.times.{0.315343 / (4.39986+5.05777+0.542375)}=3.15% of the total number of proteins present in the cell. This is comparable, but less than the expected 5%-10% of the total number of proteins found by other researchers. However, we have not attempted to adjust the percentages upwards in the hope that the general behavior will not change significantly, as others have shown.

[0101] The inventors have discovered that a first treatment can be performed on a target tissue by repeatedly applying pulsed energy (e.g., SDM) to the target tissue over a period of time to controllably increase the temperature of the target tissue and therapeutically treat the target tissue without destroying or permanently damaging and destroying the target tissue. A "treatment" consists of a total number of applications of pulsed energy to the target tissue over a given period of time, such as tens or even hundreds of applications of light or other energy to the target tissue over a short period of time, such as a period of less than 10 seconds, more typically less than 1 second, such as 100 milliseconds to 600 milliseconds. This "treatment" controllably increases the temperature of the target tissue and activates heat shock proteins and associated components.

[0102] However, what has been found is that when the application of pulsed energy to a target tissue is stopped for a time interval that exceeds the first period constituting a "first treatment", e.g., a time interval that may comprise a few seconds to a few minutes, e.g., 3 seconds to 3 minutes, or more preferably 10 seconds to 90 seconds, and then a second treatment is performed on the target tissue after the time interval within a single treatment session or session, the second treatment also involves repeated reapplication of pulsed energy to the target tissue in a manner that controllably increases the temperature of the target tissue to therapeutically treat the target tissue without destroying or permanently damaging the target tissue, and the amount of activated HSPs and related components within the cells of the target tissue is increased, resulting in a more effective overall treatment of the biological tissue. In other words, the first treatment creates a level of heat shock protein activation in the target tissue, and the second treatment creates a level of heat shock protein activation in the target tissue that is higher than the level created by the first treatment. Thus, performing multiple treatments on a patient's target tissue within a single treatment session or session will enhance the overall treatment of the biological tissue, so long as the second or additional treatment is performed after an interval of time not exceeding a few minutes, but long enough to allow for temperature relaxation so as not to damage or destroy the target tissue.

[0103] This technique may be referred to herein as "stair-stepped" in that the level of activated HSP production increases with subsequent treatments or treatments within the same treatment session. This "stair-stepped" technique may be described by a combination of the Arrhenius integration approach to subsecond phenomena and the Rybinski et al. (2013) treatment of intervals between repeated subsecond applications of SDM or other pulse energies.

[0104] For the proposed stepped SDM (iterative SDM application) proposed in this disclosure, there are some important differences from the situation depicted in FIG.

[0105] Although SDM can be applied prophylactically to healthy cells, in many cases SDM is applied to diseased cells, in which case the initial concentration of damaged proteins [S(0)] can be greater than given in Table 4. We do not explain this, assuming that the qualitative behavior remains unchanged.

[0106] The duration of a single SDM application is only subseconds, rather than the minutes shown in Figure 13. The Rybinski et al. rate constant is much smaller than the Arrhenius constant; the latter gives an Arrhenius integral of the order of 1 for subsecond durations, whereas the Rybinski et al. rate constant is too small to do so. This is an example of different effective rate constants that exist when the time scales of interest are different. The Rybinski et al. rate constant applies to phenomena that occur over minutes, whereas the Arrhenius rate constant applies to subsecond phenomena.

[0107] Therefore, to analyze what happens with the stepped SDM technique proposed to improve the effectiveness of SDM, we decide to combine the Arrhenius integral treatment, which is suitable for subsecond phenomena, with the treatment of Rybinski et al. (2013), which is suitable for phenomena occurring on the order of 1 min intervals between repeated applications of SDM.

[0108] SDM subsecond applications described by the Arrhenius integral formalism

[0109] Interval O (min) between SDM applications as described by Rybinski et al. (2013)

[0110] Specifically, we consider two successive applications of SDM, where each SDM micropulse train has a subsecond duration. We assume that on short time scales of less than 1 second, all inactive HSPs, which are the source of activated (free) HSPs, are contained in HSPHSF molecules in the cytoplasm. Therefore, the first application of SDM is thought to reduce the cytoplasmic reservoir of inactive HSPs in the initial HSPHSF molecule population from [HSPHSF(equil)] to [HSPHSF(equil)]exp[-Ω]. ·Then, increase the initial HSP molecule population from [HSP(equil)] to [HSP(equil)] + [HSPHSF(equil)] (1-exp[-Ω]). and increasing the initial HSF molecule population from [HSF(equil)] to [HSF(equil)]+[HSPHSF(equil)](1-exp[-Ω]). · The equilibrium concentrations of all other species are assumed to remain the same after the first SDM application. Next, the formula of Rybinski et al. is used to calculate what happens to [HSP] and [HSPHSF] in the interval λt=O (min) between the first and second SDM applications, and the initial values ​​of HSP, HSF, and HSPHSF after the first SDM application are [HSP(SDM1)]=[HSP(equil)]+[HSPHSF(equil)](1-exp[-Ω]) [HSF(SDM1)]=[HSF(equil)]+[HSPHSF(equil)](1-exp[-Ω]) and [HSPHSF(SDM1)]=[HSPHSF(equil)](exp[-Ω]) It is said. In the second application of SDM after the interval λt, the values ​​of [HSP], [HSF], and {HSPHSF] after SDM are [HSP(SDM2)]=[HSP(λt)]+[HSPHSF(λt)](1-exp[-Ω]) [HSF(SDM2)]=[HSF(λt)]+[HSPHSF(λt)](1-exp[-Ω]) and [HSPHSF(SDM2)]=[HSPHSF(λt)]exp[-Ω] Here, [HSP(λt)], [HSF(λt)], and [HSPHSF(λt)] are the values ​​calculated from the formula of Rybinski et al. (2013) at time λt. Our current interest is to compare [HSP(SDM2)] with [HSP(SDM1)] and see whether repeated application of SDM at intervals λt following the first SDM application results in more activated (free) HSP in the cytoplasm. The ratio β(λt,Ω)=[HSP(SDM2)] / [HSP(SDM1)]={[{[HSP(λt)]+[HSPHSF(λt)](1-exp[-Ω])} / {[HSP(0)]+[HSPHSF(0)](1-exp[-Ω])} provides a direct measure of the improvement in HSP activation when SDM is repeatedly applied after an interval λt from the first SDM application.

[0111] The concentrations of HSP and HSPHSF vary considerably over the interval t between SDM applications.

[0112] Figures 15A and 15B show the variation of the activated concentration [HSP] and the unactivated HSP [HSPHSF] in the cytoplasmic reservoir during the interval λt = 1 min between SDM applications when the Arrhenius integral of the SDM is Ω = 1 and the equilibrium concentrations are as given in Table 4.

[0113] Although we have only dealt with a single repetition (one step) here, it is clear that the procedure could be repeated to provide multiple step-like events as a means of improving the efficacy of SDM or as other treatments that involve activation of tissue HSPs.

[0114] The effect of varying the magnitude of the Arrhenius integral Ω and the spacing λt between two different treatments spaced apart in time is illustrated by the following examples and results.

[0115] Nine examples generated by the above procedure are presented below. All examples are of treatments consisting of two SDM treatments, where the second treatment occurs at a time λt following the first, and they investigate: · Effect of different magnitudes of Arrhenius integral Ω on SDM treatment [Three different Ω are considered: Ω=0.2, 0.5, and 1.0] The effect of varying the interval λt between two SDM treatments [three different λt are considered: λt=15 s, 30 s, and 60 s]

[0116] As shown above, the activation Arrhenius integral Ω depends on both the treatment parameters (e.g., laser power, duty cycle, total train duration) and the RPE properties (e.g., absorption coefficient, density of HSPs).

[0117] Table 5 below shows the effect of different Ω (Ω = 0.2, 0.5, 1) on the cellular HSP content when the interval between two SDM treatments is λt = 1 min, where the cells have the equilibrium concentrations of the 10 species involved (Rybinski et al., 2013) shown in Table 4.

[0118] Table 5 shows the four HSP concentrations (in arbitrary units according to Rybinski et al.) corresponding to the four different times, respectively. Before the first SDM treatment. [HSP(equil)] Immediately after the first SDM application. [HSP(SDM1)] At the end of the interval λt following the first SDM treatment. [HSP(λt)] Immediately after the second SDM treatment at λt. [HSP(SDM2)] Also shown is the improvement coefficient for a single treatment: β=[HSP(SDM2)] / [HSP(SDM1)]

[0119] Table 5. HSP concentrations at the four times just described in the text. Effect of changes in SDMΩ when two SDMs are applied to cells, when treated with an interval of λt = 0.25 min = 15 s. [HSP(equil)] [HSP(SDM1)] [HSP(λt)] [HSP(SDM2)] β Ω=0.2 0.315 0.67 0.54 0.95 1.27 Ω=0.5 0.315 1.10 0.77 1.34 1.22 Ω=1.0 0.315 1.57 0.93 1.71 1.09

[0120] Table 6 is the same as Table 5, except for the interval between SDM treatments of λt = 0.5 min = 30 s.

[0121] Table 6. HSP concentrations at the four times described in the text. Effect of varying SDMΩ on cells for the two SDM treatments when treatments were separated by λt = 0.5 min = 30 s. [HSP(equil)] [HSP(SDM1)] HSP(λt)] [HSP(SDM2)] β Ω=0.2 0.315 0.67 0.44 0.77 1.14 Ω=0.5 0.315 1.10 0.58 1.18 1.08 Ω=1.0 0.315 1.57 0.67 1.59 1.01

[0122] Table 7 is the same as Tables 5 and 6, except that the treatments are separated by 1 minute or 60 seconds.

[0123] Table 7. HSP concentrations at the four times just described in the text. The effect of changing the two SDMΩ of SDM treatments on normal (healthy) cells when treatments were separated by λt = 1 min = 60 s. [HSP(equil)] [HSP(SDM1)] [HSP(λt)] [HSP(SDM2)] β Ω=0.2 0.315 0.67 0.30 0.64 0.95 Ω=0.5 0.315 1.10 0.37 1.06 0.96 Ω=1.0 0.315 1.57 0.48 1.51 0.96

[0124] Tables 5 to 7 show the following: The first treatment of SDM significantly increases [HSP] in all three Ω, but the increase is greater for larger Ω. Although not explicitly shown in the table, the increase in [HSP] occurs at the expense of the cytoplasmic reservoir of sequestered (inactivated) HSPs. [HSPHSF(SDM1)] is much smaller than [HSPHSF(equil)]. · [HSP] is significantly decreased in the interval λt between the two SDM treatments, with the decrease being larger the larger the λt. (The decrease in [HSP] is accompanied by an increase in both [HSPHSF], as shown in Figure 15 and [HSPS] during the interval λt, indicating a rapid replenishment of the cytoplasmic reservoir of inactivated HSPs and a rapid attachment of HSPs to damaged proteins.) For λt less than 60 s, two SDM treatments improved the number of activated (free) HSPs in the cytoplasm more than a single treatment. The smaller λt is, the greater the improvement. However, when λt becomes as large as 60 s, the ratio β = [HSP(SDM2)] / [HSP(SDM1)] becomes less than 1, indicating that no improvement is observed with two SDM treatments compared to a single SDM treatment, although this result may vary depending on the energy source parameters and the type of tissue treated. The improvement for λt<60 s is larger for smaller SDM Arrhenius integral Ω.

[0125] The results for the improvement rate β = [HSP(SDM2)] / [HSP(SDM1)] are summarized in Figure 16, where the relationship between the improvement rate β = [HSP(SDM2)] / [HSP(SDM1)] and the interval λt (sec) between SDM treatments for three values ​​of the SDM Arrhenius integral Ω and for three values ​​of the interval λt = 15 s, 30 s, and 60 s is shown. The top curve is for Ω = 0.2. The middle curve is for Ω = 0.5. The bottom curve is for Ω = 1.0. These results are for the rate constants from Rybinski et al. (2013) in Table 3 and the equilibrium species concentrations in Table 4.

[0126] It should be understood that the results in Tables 5-7 and Figure 16 are for the rate constants of Rybinski et al. (2013) in Table 3, and the equilibrium concentrations in Table 4. The actual concentrations and rate constants in cells may differ from these values, and therefore the numerical results in Tables 5-7 and Figure 16 should be taken as representative, not absolute. However, they are not expected to differ significantly. Thus, performing multiple intra-session treatments, such that a second or subsequent treatment follows a first, at any interval between 3 seconds and 3 minutes, preferably 10 seconds and 90 seconds, on a single target tissue location or region, such as a single retinal locus, should increase the activation of HSPs and associated components and thus increase the overall treatment efficacy of the target tissue. The resulting "staircase" effect achieves a gradual increase in the number of heat shock proteins activated, enhancing the therapeutic efficacy of the treatment. However, if the time interval between the first and subsequent treatments is too long, the "staircase" effect will be diminished or not achieved.

[0127] The techniques of the present invention are particularly useful when the treatment parameters or tissue characteristics are such that the associated Arrhenius integral for activation is low, and when the interval between repeated applications is small, such as less than 90 seconds, and preferably less than 1 minute. Thus, such multiple treatments should be performed within the same treatment session, such as within a single practice, and there can be a time gap between different treatments to obtain the benefits of the techniques of the present invention.

[0128] Although certain 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, and the invention is not therefore limited except as by the appended claims.

Claims

1. 1. A retinal phototherapy or photostimulation system comprising: a laser console that generates at least one pulsed treatment beam, the at least one pulsed treatment beam having wavelength, power, pulse train duration, and duty cycle parameters for photostimulating or treating a retinal target tissue without causing permanent damage to the retinal target tissue, the parameters of the at least one pulsed treatment beam being fixed such that they cannot be altered by a healthcare provider; a projector or camera for projecting the at least one pulsed treatment beam onto at least a portion of the retina; a scanning mechanism for controllably directing the at least one pulsed treatment beam to a treatment area of ​​the retina; A system comprising:

2. 10. The system of claim 1, wherein the at least one pulsed therapeutic beam stimulates activation of heat shock proteins in the target tissue.

3. 10. The system of claim 1, wherein the at least one pulsed therapy beam raises a target tissue temperature to a desired level while maintaining an average temperature rise of the target tissue below a predetermined level over a period of time so as not to permanently damage the target tissue.

4. 3. The system of claim 2, wherein to achieve a therapeutic or prophylactic effect, the at least one pulsed therapy beam raises a target tissue temperature to 11° C. or less and an average temperature rise of the target tissue over several minutes is maintained below a predetermined level so as not to cause permanent damage to the target tissue.

5. 5. The system of claim 4, wherein the at least one pulsed therapy beam raises the target tissue temperature to between 6° C. and 11° C., at least during application of a pulsed energy source to the target tissue, while maintaining the average temperature rise of the target tissue at or below 1° C. over several minutes.

6. 6. The system of claim 5, wherein the average temperature of the target tissue is maintained at or below 1° C. for 6 minutes.

7. the at least one pulsed therapy beam is applied to multiple target tissue regions, adjacent target tissue regions being separated to avoid thermal tissue damage; and / or said laser console having a power output of between 0.5 and 3.0 watts; and / or 10. The system of claim 1, wherein the at least one pulsed treatment beam produces a treatment spot on the retina having a size of 100 to 1,000 micrometers.

8. 10. The system of claim 1, wherein the at least one pulsed treatment beam has a wavelength between 530 nm and 1300 nm, a duty cycle of less than 10%, and a pulse train duration between 0.1 and 0.6 seconds.

9. 10. The system of claim 8, wherein the at least one pulsed treatment beam has a wavelength between 750 nm and 850 nm, a duty cycle between 2% and 5%, and a pulse train duration between 0.15 and 0.5 seconds.

10. The at least one pulsed therapeutic beam stimulates activation of heat shock proteins in the target tissue to increase the target tissue temperature by between 6°C and 11°C to achieve a therapeutic or prophylactic effect; 10. The system of claim 1, wherein the average temperature rise of the target tissue over several minutes is maintained at 1° C. or less over 6 minutes to avoid permanent damage to the target tissue.

11. The at least one pulsed therapy beam is applied to multiple target tissue regions, and adjacent target tissue regions are separated to avoid thermal tissue damage, and / or said laser console having a power output of between 0.5 and 3.0 watts; and / or 11. The system of claim 10, wherein the at least one pulsed treatment beam produces a treatment spot on the retina having a size of 100 to 1,000 micrometers.

12. The system described in claim 10, wherein the at least one pulsed treatment beam has a wavelength between 530 nm and 1300 nm, a duty cycle of less than 10%, and a pulse train duration between 0.1 and 0.6 seconds.

13. The system of claim 12, wherein the at least one pulsed treatment beam has a wavelength between 750 nm and 850 nm, a duty cycle between 2% and 5%, and a pulse train duration between 0.15 and 0.5 seconds.

14. The at least one pulsed therapy beam has a wavelength between 530 nm and 1300 nm, a duty cycle of less than 10%, and a pulse train duration between 0.1 and 0.6 seconds; the at least one pulsed treatment beam produces a treatment spot on the retina having a size of 100 to 1,000 micrometers; 10. The system of claim 1, wherein the at least one pulsed therapy beam is applied to multiple target tissue regions, and adjacent target tissue regions are separated to avoid thermal tissue damage.

15. The at least one pulsed therapy beam has a wavelength between 750 nm and 850 nm, a duty cycle between 2% and 5%, and a pulse train duration between 0.15 and 0.5 seconds, and the laser console has a power output between 0.5 and 3.0 Watts; and / or said at least one pulsed therapeutic beam stimulates activation of heat shock proteins in said target tissue to increase the target tissue temperature by between 6° C. and 11° C. to achieve a therapeutic or prophylactic effect; 15. The system of claim 14, wherein the average temperature rise of the target tissue over several minutes is maintained at 1 degree C or less over 6 minutes so as not to cause permanent damage to the target tissue.

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