Method and apparatus for optimizing selective photothermolysis

Abstract

To provide a radiation system for selective photothermolysis surgery.SOLUTION: A radiation system for selective photothermolysis surgery comprises: a tunable radiation source configured to generate the surgery radiation pulse and the option-synchronized auxiliary radiation pulse (or modulation) with the adjustable time delay immediately after the surgery radiation pulse; a patient interface 1200 comprising a radiation delivery unit configured to deliver the radiation beam to a tissue 1000, and an ultrasonic detector for acquiring photoacoustic signals excited by the radiation pulse; and a control system 1110 in the tunable radiation source configured to acquire information from the ultrasonic detector, analyze the information and control generation of the radiation pulse having the optimum center wavelength for optimizing the surgical outcome, or the optimum radiation pulse width, or the optimum radiation pulse energy, or the combination of the same.SELECTED DRAWING: Figure 1

Description

[Technical field]

[0001] The present invention relates to surgical techniques, devices and methods, including surgical techniques, devices and methods for selective photothermolysis (SP) surgery. Several techniques can be applied to common surgical systems for heating lesions in tissue, including high intensity focused ultrasound therapy. [Background technology]

[0002] SP, described by Anderson and Parrish in a 1983 SCIENCE paper, utilizes short laser pulses to precisely control collateral thermal or mechanical damage around the light-absorbing lesion without the need to direct a laser microbeam at the surgical target. SP is distinguished from other laser photothermal therapies not only by its innovative scientific concept but also by its unique temperature effect on the surgical target. SP systems deposit short surgical laser pulse energy at the surgical target, generating a spatially and temporally restricted temperature pulse. The pulse duration is equal to the laser pulse duration plus the thermal relaxation time of the surgical target. The temperature of the surgical target in human tissue in vivo returns to normal body temperature immediately after the surgical laser pulse. The short temperature pulse of SP (as short as 10 ns for laser tattoo removal) distinguishes SP from photothermal therapies with slowly changing temperatures that can be monitored by photoacoustic methods in the prior art. However, such photoacoustic methods in the prior art cannot measure the temperature pulse of SP. Also, the photoacoustic method (ultrasound detector) core device has never been applied to an SP system in the prior art. The photoacoustic method disclosed in the present invention is capable of measuring both temperature pulses and slowly changing temperatures, in addition to its unique in vivo calibration capabilities, without suffering from the inaccuracies or invasiveness of existing photoacoustic methods in the prior art. Summary of the Invention [Problem to be solved by the invention]

[0003] The technical challenges associated with SP laser surgery include maximizing the laser energy deposition ratio of the surgical target to the natural chromophore, limiting the laser energy deposition to the surgical target, and optimizing the laser pulse energy. Non-optimized SP laser parameters are associated with poor laser surgery outcomes. Compromised SP laser surgery outcomes have been well known for decades, but there is no good solution in the prior art of SP. [Means for solving the problem]

[0004] The present invention relates to surgical techniques, devices and methods for optimizing selective photothermolysis (SP) surgery. Some of the techniques can be applied to common surgical systems for heating lesions in tissue, including high intensity focused ultrasound therapy. Note that a tunable light source in this specification broadly refers to a light source that has tuning capabilities in its central wavelength, or light pulse width, or light pulse energy, or a combination thereof.

[0005] In one aspect, a pulsed light SP surgery system comprises a tunable pulsed light source for generating a pulsed surgical light beam under the control of a control system, and a patient interface operable to contact a target tissue. The patient interface comprises a light delivery unit, an acoustic detector, and an interface medium. The light delivery unit shapes a light beam profile, delivers the light beam with an articulated arm, adjusts the diameter of the light beam, and transmits the light beam to the tissue surface through the interface medium. The pulsed light beam excites photoacoustic waves that propagate through the interface medium and are detected by the acoustic detector. The detected photoacoustic signals are digitized and analyzed by the control system to generate a surgical light pulse with an optimal central wavelength or optimal light pulse energy or a combination thereof for optimal SP surgery results.

[0006] In another aspect, a pulsed light SP surgery system that generates auxiliary light pulses synchronized with surgical light pulses and surgical radiation pulses with an adjustable time delay therebetween comprises a tunable pulsed light source and a patient interface operable to contact a target tissue. The patient interface comprises a light delivery unit, an acoustic detector, and an interface medium. The light delivery unit shapes a beam profile of the light beam, delivers the light beam with an articulated arm, adjusts the diameter of the light beam, and transmits the light beam to the tissue surface through the interface medium. The light beam excites photoacoustic waves that propagate through the interface medium and are detected by the acoustic detector. The auxiliary light pulses enable a series of temperature-related measurements. The detected photoacoustic signals are digitized and analyzed by a control system to generate surgical light pulses with an optimal center wavelength, or optimal pulse width, or optimal pulse energy, or a combination thereof, for optimal SP surgery results.

[0007] In another aspect, a SP surgical planning system includes a tunable pulsed light source capable of generating sub-therapeutic surgical light pulses and auxiliary light pulses synchronized with surgical radiation pulses with an adjustable time delay therebetween, and a patient interface operable to contact a target tissue. The patient interface includes a light delivery unit, an acoustic detector, and an interface medium. The light delivery unit shapes a beam profile of the light beam, delivers the light beam with an articulated arm, adjusts the diameter of the light beam, and transmits the light beam to the tissue surface through the interface medium. The light beam excites photoacoustic waves that propagate through the interface medium and are detected by the acoustic detector. The auxiliary light pulses enable a series of temperature-related measurements. The detected photoacoustic signals are digitized and analyzed by a control system to determine optimal surgical light pulse parameters to be used in an otherwise conventional pulsed light SP surgical system.

[0008] In another aspect, a pulsed radiation SP surgery system includes two tunable radiation sources, a patient interface operable to contact a target tissue, and a control system. One tunable radiation source generates pulsed surgical radiation under the control of the control system to generate a temperature pulse in a lesion in the tissue or an exogenous contrast agent attached to the lesion in the tissue. The other tunable radiation source generates a pulsed or modulated temperature-sensitive radiation beam that can be absorbed by the lesion in the tissue or an exogenous contrast agent attached to the lesion in the tissue under the control of the control system. The patient interface includes a radiation delivery unit for delivering the radiation beam to the tissue, and an ultrasound detector for acquiring a photoacoustic signal excited by the radiation beam. The control system acquires information from the ultrasound detector in the patient interface, analyzes the information, and controls the generation of a surgical radiation beam with an optimal central wavelength or frequency, or an optimal pulse width, or an optimal pulse energy, or a combination thereof, for an optimal SP surgery result.

[0009] In another aspect, a method for tuning a surgical laser wavelength and optimizing SP laser treatment of an unknown surgical target includes the steps of determining a set of surgical wavelength points for spectroscopic scanning and configuring both the patient interface and the acoustic detector, delivering multiple sub-therapeutic surgical laser pulses for each wavelength, acquiring and averaging photoacoustic signals, and repeating for all wavelength points, reconstructing two-dimensional depth-resolved relative extinction coefficient information for all wavelength points, calculating relative extinction coefficient curves for all absorbers and identifying the unknown surgical target, calculating relative energy deposition ratio curves for the unknown surgical target versus natural chromophores, and determining optimal surgical wavelengths for different types of surgical targets.

[0010] In another embodiment, the tissue Gruneisen The method for calibrating the temperature-dependent relative logarithmic function of JPEG7170 parameters includes the steps of measuring the equilibrium temperature of the tissue, measuring the photoacoustic signal of the surgical target excited by a temperature-sensing laser pulse having a constant laser pulse energy, calculating a baseline signal as the logarithm of the amplitude of the photoacoustic signal at the equilibrium temperature, sending only the surgical laser pulse and measuring the excited photoacoustic signal, sending both the surgical laser pulse and the temperature-sensing laser pulse and measuring the excited photoacoustic signal by the dual pulse, calculating the logarithm of the amplitude of the photoacoustic signal excited by the temperature-sensing laser pulse after subtraction and logarithm operations, and calculating the temperature-sensing laser pulse as the logarithm of the equilibrium temperature. The method includes the steps of: acquiring data points after subtracting a baseline signal from the logarithm of the amplitude of the photoacoustic signal excited by the surgical laser pulse; determining whether there is laser-induced cavitation; if there is no laser-induced cavitation, waiting until the temperature returns to equilibrium; increasing the surgical laser pulse by a factor k i ; repeating the above procedure to acquire more data points of the relative logarithm function of the Gruneisen parameters of the tissue until laser-induced cavitation is observed by the acoustic detector from the nonlinearity of the signal; and calculating the absolute temperature rise due to the surgical laser pulse and fitting the curve of the relative logarithm function of the Gruneisen parameters of the tissue with sufficient data points.

[0011] In another aspect, a method for photoacoustic sensing of a dynamic temperature (temperature pulse) of a surgical target in tissue at the end of a short surgical laser pulse includes the steps of measuring body temperature before a laser surgical intervention, sending a temperature sensing laser pulse with a constant laser pulse energy, measuring the amplitude of the excited photoacoustic signal and calculating its logarithm as a baseline signal, sending only the surgical laser pulse and measuring the excited photoacoustic signal, sending both the surgical laser pulse and the temperature sensing laser pulse at the end of the surgical laser pulse and measuring the excited photoacoustic signal with a dual pulse, calculating the photoacoustic signal amplitude excited by the temperature sensing laser pulse at an unknown temperature at the end of the surgical laser pulse, separating the temperature dependent part from other temperature independent parts by a logarithmic operation, calculating the relative logarithmic function value by subtracting the baseline signal, determining the unknown temperature at the end of the surgical laser pulse from the calibrated relative logarithmic function of the Gruneisen parameters of the tissue, and terminating the temperature sensing operation.

[0012] In another aspect, a method for measuring a thermal relaxation time of a surgical target in tissue includes measuring a photoacoustic signal of the surgical target at body temperature and body temperature with a temperature sensitive laser pulse and calculating a logarithm of the photoacoustic signal amplitude as a baseline signal; sending a sub-therapeutic surgical laser pulse and measuring the excited photoacoustic signal; sending both a sub-therapeutic surgical laser pulse and a temperature sensitive laser pulse at a precise delay time, measuring the excited photoacoustic signal of the dual pulse and calculating the amplitude of the photoacoustic signal excited by the temperature sensitive laser pulse; calculating a relative logarithmic function value by separating the temperature dependent portion from the temperature independent portion by a logarithmic operation and subtracting the baseline signal; determining the temperature of the surgical target at the precise delay time; waiting until the surgical target temperature returns to body temperature; determining whether enough delay time points have been obtained for curve fitting; if not, repeating the above procedure to measure the temperature at another delay time until enough delay time points have been obtained; fitting a temperature vs. delay time curve and determining the thermal relaxation time of the surgical target; and terminating the thermal relaxation time measurement operation.

[0013] In another aspect, a method for optimizing surgical laser pulse energy during laser photocoagulation surgery includes the steps of selecting a surgical target, measuring a thermal relaxation time of the surgical target, optimizing a surgical laser pulse width according to the thermal relaxation time of the surgical target, sending only a surgical laser pulse and measuring an excited photoacoustic signal with the surgical laser pulse, sending both a surgical laser pulse and a temperature sensing laser pulse and measuring an excited photoacoustic signal by a dual pulse, calculating an amplitude of the photoacoustic signal excited by the temperature sensing laser pulse, isolating a temperature-dependent part by a logarithmic operation, calculating a relative logarithmic function value by subtracting a baseline signal obtained in the thermal relaxation time measurement step, determining a temperature of the surgical target, determining whether the temperature has reached a predetermined surgical temperature of photocoagulation, and if not, waiting until the temperature recovers to body temperature and increasing the surgical laser pulse energy, repeating the above temperature measurement procedure until the temperature reaches the predetermined surgical temperature of photocoagulation, and terminating the optimization of the surgical laser pulse energy during the laser photocoagulation surgery.

[0014] In another aspect, a method for optimizing surgical laser pulse energy during laser photodisruption surgery includes the steps of selecting a surgical target, delivering a surgical laser pulse, measuring the excited photoacoustic signal, plotting a point on a curve of photoacoustic signal amplitude versus laser pulse energy, determining whether there is laser induced cavitation, and if there is no laser induced cavitation, waiting until the temperature returns to body temperature, increasing the surgical laser pulse energy, and repeating the above procedure from step 2 until laser induced cavitation is observed and the surgical laser pulse energy is optimized.

[0015] In another aspect, a method for determining a surgical laser pulse energy to reach a predetermined temperature without laser surgery includes the steps of selecting a surgical target, measuring a thermal relaxation time of the surgical target, optimizing a surgical laser pulse width according to the thermal relaxation time of the surgical target, sending a sub-therapeutic surgical laser pulse and measuring an excited photoacoustic signal, sending both a sub-therapeutic surgical laser pulse and a temperature sensing laser pulse, measuring the excited photoacoustic signal in a dual pulse, calculating the amplitude of the photoacoustic signal excited by the temperature sensing laser pulse, calculating a relative logarithmic function value by isolating the temperature-dependent part by logarithmic operation and subtracting the baseline signal obtained in the previous thermal relaxation time measurement step, calculating the dynamic temperature rise (temperature pulse) due to the sub-therapeutic surgical laser pulse and the surgical laser pulse energy required to rapidly heat the surgical target to the predetermined temperature, and completing the optimization of the surgical laser pulse energy to reach the predetermined temperature without laser surgery.

[0016] In yet another aspect, a method of optimized SP laser surgery with skin cooling includes the steps of tuning a surgical laser wavelength to maximize the SP surgery effect, optimizing the surgical laser pulse width, optimizing the surgical laser pulse energy, measuring body temperature, applying a temperature sensing laser pulse, measuring an excited photoacoustic signal of an epidermal target, adjusting skin cooling parameters, applying skin cooling and a delayed temperature sensing laser pulse, measuring the temperature of the epidermal target, determining whether the epidermal target has been cooled to a predetermined temperature, and if not, returning to the above procedure of adjusting skin cooling parameters and measuring the epidermal target temperature until it is cooled to the predetermined temperature, and performing the optimized SP laser surgery. Effect of the Invention

[0017] SP surgery systems have been operating in clinics for decades with unsatisfactory surgical results. Such SP surgery includes laser treatment of vascular malformations and pigmented lesions in the human body. The innovative SP surgery system and method disclosed in this invention overcomes the limitations of existing SP surgery systems and allows for the first time personalized and optimal treatment. [Brief description of the drawings]

[0018] [Figure 1] We present an example of a pulsed light selective photothermolysis (SP) surgical system, where the inclusion of an acoustic detector is key to optimizing SP. [Diagram 2] 13A-13C show examples of patient interfaces having different configurations to facilitate optimized selective photothermolysis by a laser. [Diagram 3] 13A-13C show examples of patient interfaces having different configurations to facilitate optimized selective photothermolysis by a laser. [Figure 4] 13A-13C show examples of patient interfaces having different configurations to facilitate optimized selective photothermolysis by a laser. [Diagram 5] 13A-13C show examples of patient interfaces having different configurations to facilitate optimized selective photothermolysis by a laser. [Figure 6] 13A-13C show examples of patient interfaces having different configurations to facilitate optimized selective photothermolysis by a laser. [Figure 7] 13A-13C show examples of patient interfaces having different configurations to facilitate optimized selective photothermolysis by a laser. [Figure 8] 13A-13C show examples of patient interfaces having different configurations to facilitate optimized selective photothermolysis by a laser. [Figure 9] 13A-13C show examples of patient interfaces having different configurations to facilitate optimized selective photothermolysis by a laser. [Figure 10]13A-13C show examples of patient interfaces having different configurations to facilitate optimized selective photothermolysis by a laser. [Figure 11] An example of wavelength tuning manipulation to optimize the selective photothermolysis of an unknown dye is shown. [Figure 12] An example of constructing a relative logarithmic function of the temperature dependence of the Gruneisen parameters for tissue is given. [Figure 13] An example of dynamic temperature (temperature pulse) sensing operation of a surgical target in tissue after it has been rapidly heated by a short surgical laser pulse is shown. [Figure 14] An example is presented that illustrates optimizing surgical laser pulse duration by measuring in vivo the thermal relaxation time of a surgical target in tissue. [Figure 15] 1 shows an example of in vivo surgical laser pulse energy optimization for laser photocoagulation during laser surgery. [Figure 16] 1 shows an example of in vivo surgical laser pulse energy optimization for laser photodisruption during laser surgery. [Figure 17] An example is provided for determining optimal surgical laser pulse energy to reach a given temperature without performing laser surgery. [Figure 18] An example of the operation of optimized selective photothermolysis laser surgery with skin cooling is shown. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0019] Before addressing its clinical problems, it is important to understand the laser and tissue interaction mechanism of selective photothermolysis (SP) laser surgery. Both thermal and mechanical damage can be utilized in SP. Initially, no measurable effects occur when the tissue temperature is raised to 37o-42oC by the SP laser pulse. If the temperature continues to rise to 42o-50oC, the tissue becomes hyperthermic. If hyperthermia continues for several minutes, most of the tissue may become necrotic. Enzyme activity reduction and cell immobility start at 50oC. Protein and collagen denaturation occurs at 60oC, causing tissue coagulation and cell necrosis. Cell membrane permeability is significantly increased at 80oC. Water molecules are vaporized at 100oC. Acoustic shock waves associated with laser-induced cavitation can cause cavitation and mechanical rupture of tissue. Another type of mechanical damage can be caused by intense photoacoustic waves generated by optical absorbers upon short laser pulse excitation. The main commercial applications of SP include laser tattoo removal, laser treatment of vascular malformations, and laser retinal photocoagulation.

[0020] Laser tattoo removal is usually performed with very short laser pulses in the nanosecond or picosecond range. However, the proposed mechanisms behind laser tattoo removal have their physical, chemical and biological origins. The pigmented particles of the tattoo undergo a rapid temperature increase and volume expansion upon energy deposition or excitation by a short laser pulse. However, most of the temperature increase and volume expansion is lost after a short time determined by the thermal relaxation time of these particles (the time it takes for 50% of the thermal energy to be dissipated). Photoacoustic waves are generated along with the volume change of these particles. The laser energy is converted into both thermal and mechanical energy of the photoacoustic waves. In many cases, the large laser pulse energy absorbed by the pigmented particles can cause optical breakdown, plasma generation, chemical reactions between the plasma and the pigmented particles, cavitation and the generation of acoustic shock waves. These pigmented particles can be pyrolytically transformed or shattered into smaller particles by the photoacoustic and acoustic shock waves. Necrosis of the hosting cells and damage to the surrounding tissue can be thermally and mechanically induced during this process. Eventually, the wound healing process may remove partial pigmented particles by rephagocytosis and alter the skin scattering coefficient of the affected tissue, which may obscure deeper pigmented particles.

[0021] For laser tattoo removal applications, the color of a tattoo depends on many factors, including its light absorption spectrum, the light scattering and absorption coefficients of the tissues above and below the pigment, the depth of the pigment, and the anatomical location of the pigment. It has been reported that tattoos with significantly different light absorption spectra can appear the same color to the naked eye. Clearly, the current practice of selecting a surgical laser wavelength based on the color of the tattoo is not justified. Meanwhile, only a few laser wavelengths (694 nm ruby ​​laser, 755 nm alexandrite, 1064 nm Nd:YAG, and 532 nm second harmonic Nd:YAG) are available on the market for laser tattoo removal. Even if the absorption spectrum of the tattoo pigment happens to be known, it is highly likely that it will not match with existing lasers on the market. Anderson and Parrish predicted a tunable laser for SP in 1983. However, such a laser is not yet available, since no one knows what wavelength needs to be tuned to. Also, the choice of therapeutic laser pulse energy is determined by the clinician's experience. Both clinical problems of laser tattoo removal are addressed by the present invention.

[0022] Laser treatment of vascular malformations begins with argon laser (488 nm and 514 nm) treatment of port wine stain (PWS) in the 1970s. The blue-green light of the argon laser is preferentially absorbed by hemoglobin in PWS vessels. The laser pulse energy deposited in the vessels is primarily converted to heat, causing thrombosis and destruction of PWS vessels. The first generation of argon lasers had relatively long pulse durations (~0.01 s) and caused nonspecific tissue thermal damage of epidermal tissue. Thus, scarring was a frequent complication of first generation argon laser treatment of PWS. Selective photothermolysis of PWS vessels was achieved by the first generation of pulsed dye lasers (PDL) (577 nm or 585 nm, 0.45 ms), which selectively photocoagulated PWS vessels and reduced the overlying epidermal tissue, reducing the incidence of side effects. Due to the selective absorption of hemoglobin, when PDL laser energy is deposited in intraluminal blood, heat diffuses into the vessel wall, causing necrosis of the vessel wall and subsequent extravasation of red blood cells into the adjacent dermis. Dermal collagen fills the space of the photocoagulated PWS vessel through the wound healing process. Removal of the photocoagulated PWS lesion results in the blanching of the PWS. Second-generation PDL technology employs larger spot size, higher energy density, variable pulse duration, and dynamic cooling to treat PWS more effectively. Currently, second-generation PDL with dynamic pulse duration and dynamic epidermal cooling by liquid cryogen spray is the treatment of choice for PWS. However, the laser needs to be operated by an experienced clinician who adjusts the laser pulse width and laser pulse energy based on experience. In fact, the average success rate of complete clearance is less than 20%. The selection of pulsed laser parameters is the most challenging clinical problem in the laser treatment of vascular malformations.

[0023] The above clinical problems in laser tattoo removal and laser treatment of vascular malformations are obviously related to the distribution of optical absorbers in tissue. Some experienced clinicians utilize the sound generated during laser-tissue interaction to assist in laser tattoo removal surgery. However, humans can only hear sound waves from 20Hz to 20,000Hz. In laser tattoo removal, laser-tissue interaction generates high-frequency ultrasound. Most of those frequency components are far above 20,000Hz. In other words, most of the useful information is completely ignored. By adding an ultrasound detector to "listen" to the response from the tissue under SP laser surgery, the present invention can address the above clinical problems. The science behind photoacoustic waves during laser-tissue interaction is photoacoustics, which is the key technology of the present invention.

[0024] Photoacoustic technology is named after Alexander Graham Bell, who discovered the photoacoustic effect in 1880. The generation of photoacoustic waves consists of the following steps: the conversion of absorbed pulsed or modulated radiation into thermal energy; the time change in temperature, which rises when the laser pulse energy is absorbed and falls when the laser pulse ends and heat dissipates; and the subsequent volume expansion and contraction that cause pressure changes (i.e., photoacoustic waves). Hordvik et al. reported a photoacoustic technique to determine the optical absorption coefficient of solids in 1977. Photoacoustic spectroscopy has been applied to a variety of conventional spectroscopic measurements, as reviewed by West et al. in 1983. More recent developments in photoacoustic technology have been motivated by biomedical imaging applications. Major photoacoustic technology developments in the biomedical imaging field include the invention of the acoustic and optical resolution photoacoustic microscope by Maslov et al., and the Fabry-Perot photoacoustic sensor based on photoacoustic tomography by Zhang et al. The significantly improved imaging performance (sensitivity, resolution, depth, speed) of the above mentioned photoacoustic imaging systems and the improvements of industry-specific acoustic transducer arrays for various photoacoustic tomography configurations will have a profound impact on biology and medicine.

[0025] However, the penetration of photoacoustic technology into SP laser surgery is very limited. No one has tried to build a wavelength-tunable laser and apply such a laser to tattoo removal with an acoustic detector. The choice of laser wavelength in the laser treatment of vascular malformations may be less critical than in laser tattoo removal. However, optimizing other parameters of the laser surgery system for the laser treatment of vascular malformations is still a very challenging task. In fact, SP laser surgery does not require removing the surgical target or changing the spatial location of the surgical target immediately after SP laser surgery. A long wound healing process is required to remove the damaged tissue via rephagocytosis. Photoacoustic imaging of the lesion before and immediately after SP laser surgery shows little change in the lesion image. Therefore, simple photoacoustic imaging of the lesion is of no value for optimizing SP laser surgery. Viator et al. demonstrated the feasibility of imaging deep port wine stain lesions with photoacoustic tomography without further application of the acquired lesion depth information to optimize the laser treatment of port wine stains. To optimize the laser treatment of port-wine stains in children, Rao et al. imaged the size and depth of port-wine stain vessels in child patients with an optical resolution photoacoustic microscope, used the lesion information to build a physical model of the port-wine stain lesion, and used extensive computer simulations to derive optimal laser treatment parameters (pulse width and pulse energy). Other imaging modalities, such as optical Doppler tomography and optical coherence angiography, rely on blood flow or optical speckles caused by blood flow to obtain information of port-wine stain lesions. However, the lack of blood flow immediately after laser surgery could not confirm complete photocoagulation of the diseased vessels. It was hypothesized that partially coagulated diseased vessels may remain refractory even after laser treatment. Another limitation of these optical imaging modalities is the shallow imaging depth of 1-2 mm. In contrast, the present invention takes a simple experimental approach to address the clinical problem of SP.

[0026] The disclosed techniques, methods and devices of the present invention are based on the physical principles of the photoacoustic effect and its temperature dependence. In early literature on photoacoustic technology, the temperature-dependent photoacoustic effect was utilized for a range of temperature-related measurements, such as measuring flame temperature and measuring solid thermal diffusivity. Esenaliev et al. reported real-time photoacoustic monitoring of temperature in ex vivo canine tissue in 1999. Larin et al. reported photoacoustic laser monitoring of ex vivo canine liver cooling and freezing in 2002. Shah et al. reported photoacoustic temperature monitoring in ex vivo porcine tissue in 2008. Oraevsky et al. described photoacoustic imaging methods and real-time photoacoustic monitoring of changes in tissue properties for medical diagnosis and improved temperature calibration methods in US5840023A, US6309352B1, and US2015 / 0216420A1. In the continuous wave laser thermotherapy described by Oraevsky et al., the continuous wave laser does not excite the photoacoustic waves and therefore does not interfere with the asynchronous photoacoustic temperature sensing process. However, the short surgical light pulse of the SP surgical system heats the surgical target rapidly within its short pulse duration, and the surgical target cools quickly. Accurately measuring such dynamic temperature rise (temperature pulse) due to the energy deposition of the short laser pulse of the short surgical laser in living tissue is a different scenario from the temperature sensing application of slowly changing temperature described in the prior art. Technically, it is necessary to synchronize the temperature sensing light pulse with a precise short time delay to the surgical light pulse. Also, the strong surgical light pulse generates a strong photoacoustic signal upon absorption by the optical absorber in the tissue. The photoacoustic signal excited by the SP surgical light pulse interferes with the photoacoustic signal excited by the temperature sensing light pulse. The photoacoustic temperature measurement method described in Oraevsky et al. and other prior art works for quasi-continuous wave surgical pulses and slowly changing temperatures, but it does not work for the dynamic temperature rise (temperature pulse) generated in SP surgery using short surgical light pulses.

[0027] In summary, the prior art is deficient in addressing clinical problems in SP surgery, such as laser tattoo removal and laser treatment of vascular malformations. In order to bridge the gap between the scientific theory and clinical practice of SP, surgical techniques, devices and methods are disclosed below.

[0028] <Example> By way of example, Figure 1 illustrates an example of an innovative pulsed light SP surgery system. This system distinguishes itself from conventional SP photosurgical systems in the prior art by the inclusion of an acoustic detector. The SP pulsed light surgery system comprises a tunable pulsed light source 1100 for generating a light beam 1010 under the control of its control system 1110, and a patient interface 1200 operable to contact tissue 1000. In one implementation of Figure 1, the light beam 1010 includes only a surgical pulsed light beam. To efficiently excite the photoacoustic signal, the pulse width of the pulsed light beam is 10 ms. -7 Less than a second or 10 -8 Less than a second or 10 -9The patient interface 1200 includes an optical delivery unit 1210, an acoustic detector 1220, and an interface medium 1230. The optical delivery unit 1210 shapes the optical beam profile, delivers the optical beam with an articulated arm, adjusts the diameter of the optical beam, and transmits the optical beam to the tissue 1000 surface through the interface medium 1230. The interface medium 1230 is preferably transparent to the optical beam and conductive to ultrasound. The pulsed optical beam 1010 excites a photoacoustic wave 1020 that propagates through the interface medium 1230 and is detected by the acoustic detector 1220. The detected photoacoustic signal 1030 is digitized and analyzed by the control system 1110 for control of the tunable pulsed optical beam 1100. Note that the tunable light source in this specification broadly means a light source that has tuning capability in its center wavelength, or optical pulse width, or optical pulse energy, or a combination thereof. It is not difficult to make a tunable light source itself. However, the missing piece is how to tune the central wavelength and other surgical light pulse parameters depending on the surgical target in the tissue. The inclusion of an acoustic detector is exactly the missing piece in the conventional techniques for SP surgery. The inclusion of an acoustic detector makes it possible for the first time to utilize a tunable light source in SP surgery. Both the tunable light source and the acoustic detector enable optimized SP surgery, a goal that has been desired for decades.

[0029] To fully utilize the following disclosed method for optimized SP surgical results, it is desirable to utilize a more advanced tunable light source 1100 capable of generating surgical light pulses and synchronized auxiliary temperature-sensing light pulses under the control of its control system 1110. In most implementations of FIG. 1, the temperature-sensing light beam includes a temperature-sensing light pulse with a tunable time delay relative to the surgical light pulse. However, the temperature-sensing light beam can be an intensity-modulated light beam to excite photoacoustic waves when depth-resolved tissue information is not required for the application, or a chirped intensity-modulated light beam to enable very low resolution depth discrimination. To efficiently excite a photoacoustic signal, the pulse width of the temperature-sensing light pulse is within 10 -7 Less than a second or 10 -8 Less than a second or 10 -9 The time required for the temperature-sensing light beam to reach the target temperature is less than a second. In some implementations, the central wavelength and pulse width of the temperature-sensing light beam can be fixed. The temperature-sensing light pulse energy is at a sub-therapeutic level. For the tunable light source 1100, its tuning capability needs to be matched to the needs of the specific SP surgical application. The most practical implementation of such a more advanced tunable light source is to integrate the tunable surgical light source unit and the tunable temperature-sensing light source unit into a single package with a shared power subsystem, a shared cooling subsystem, and a shared control system 1110. The light beams of the light source units should be combined and sent out from the same optical output port. In some implementations, it is preferable to use a low-cost fixed-wavelength pulsed solid-state laser to generate the temperature-sensing light beam.

[0030] Yet another implementation of FIG. 1 can be a surgical planning system that provides optimized surgical laser parameters to an otherwise conventional pulsed light SP surgical system. Such a surgical planning system includes a tunable pulsed light source that can generate a sub-therapeutic pulsed light beam, or a pulsed or modulated temperature-sensing light beam, or both, under the control of its control system, and a patient interface that is operable to contact a target tissue. The patient interface includes a light delivery unit, an acoustic detector, and an interface medium. The light delivery unit shapes the beam profile of the light beam, delivers the light beam with an articulated arm, adjusts the diameter of the light beam, and transmits the light beam to the tissue surface through the interface medium. The light beam propagates through the interface medium and excites photoacoustic waves that are detected by the acoustic detector. The detected photoacoustic signals are digitized and analyzed by the control system to determine optimal surgical light pulse parameters to be used in another conventional pulsed light SP surgical system. The advantage of such a surgical planning system is that its laser pulse repetition rate can be much higher than a surgical system, and the time to obtain optimized surgical laser parameters is much shorter.

[0031] The key of the present invention is to incorporate the ultrasound detector into the SP surgery system. The pulsed light SP surgery system in FIG. 1 can be generalized as a pulsed radiation SP surgery system by replacing the surgical light pulse with any form of radiation (radio waves, microwaves, infrared light, visible light, ultraviolet light, X-rays, gamma rays) pulse that generates a temperature pulse in the lesion in the tissue or in the exogenous contrast agent attached to the lesion, replacing the temperature-sensing light pulse with any form of pulsed or modulated radiation (radio waves, microwaves, infrared light, visible light, ultraviolet light, X-rays, gamma rays) that can be absorbed by the lesion in the tissue or in the exogenous contrast agent attached to the lesion in the tissue and effectively excite photoacoustic waves, and replacing the light delivery unit of the patient interface with a radiation beam delivery unit for delivering the radiation beam to the tissue. All the techniques and methods in FIG. 1 are applied to the generalized radiation SP surgery system. Since laser tattoo removal and laser treatment of vascular malformations are the main interests of the present invention, the examples shown below are mainly based on a pulsed laser surgery system.

[0032] 2-10 show examples of patient interfaces with different configurations to facilitate optimized SP laser surgery. FIG. 2 shows a schematic example of a patient interface comprising a single-element ultrasound transducer 2200 and an interface medium 2300 in acoustic contact with tissue 2400. Since the optical delivery unit to be shown in the patient interface is no different from that of a conventional laser SP system, for simplicity, the optical delivery unit is omitted in FIG. 2. The laser beam 2100 (the surgical laser beam, or the temperature-sensing laser beam, or both) can be selectively delivered to the tissue 2400 surface according to the requirements of the method. In some implementations, the laser beam 2100 may include only the surgical laser beam. The single-element ultrasound transducer 2200 is positioned to detect the photoacoustic waves without blocking the laser beam 2100. The transducer 2200 can be based on conventional ultrasound transducers or optical detection technology. The interface medium 2300 allows the transmission of the laser beam and the photoacoustic waves. The interface medium 2300 can be stored in a more simplified configuration in which a single element ultrasound transducer is in direct acoustic contact with the tissue at a tissue surface area immediately adjacent to the tissue surface area illuminated by the laser beam. One advantage of this simplified configuration is that it may simultaneously allow delivery of skin cooling agent through free space. Spectral photoacoustic signals are acquired by the single element ultrasound transducer 2200 from one-dimensional depth-resolved space within the tissue and are digitized and analyzed by the control system 1110 of FIG. 1.

[0033] FIG. 3 shows another schematic example of a patient interface comprising a linear array ultrasonic transducer 3200 and an interface medium 3300 in acoustic contact with tissue 3400. The optical delivery unit to be shown in the patient interface is not different from that of a conventional laser SP device, so for simplicity, the optical delivery unit is omitted in FIG. 3. A laser beam 3100 (a surgical laser beam, or a temperature-sensing laser beam, or both) can be selectively delivered to the tissue 3400 surface according to the requirements of the method. The linear array ultrasonic transducer 3200 is arranged to detect the photoacoustic waves without blocking the laser beam 3100. The linear array transducer 3200 can be based on conventional ultrasonic transducers or optical detection technology. The interface medium 3300 allows the transmission of the laser beam and the photoacoustic waves with minimal losses. The interface medium 3300 can also be stored in a more simplified configuration. In this configuration, the linear array ultrasound transducer is in direct acoustic contact with the tissue at a tissue surface area immediately adjacent to the tissue surface area illuminated by the laser beam delivered through free space. One advantage of this simplified configuration is that it may simultaneously enable delivery of skin cooling agent through free space. Spectral photoacoustic signals are acquired by the linear array ultrasound transducer 3200 from said two-dimensional depth-resolved space within the tissue and are digitized and analyzed by the control system 1110 shown in FIG. 1.

[0034] FIG. 4 shows another schematic example of a patient interface comprising a linear array ultrasonic transducer 4200, an interface medium 4300 in acoustic contact with tissue 4400, and a rotation stage 4500. The rotation stage 4500 mounts the linear array ultrasonic transducer 4200 and the interface medium 4300 and rotates around the central axis of the illuminated elliptical area of ​​the tissue surface to obtain three-dimensional tissue information. The arrow 4600 indicates the rotation direction of the rotation stage 4500. Since the optical delivery unit to be shown on the patient interface is no different from that of a conventional laser SP system, the optical delivery unit is omitted in FIG. 4 for simplicity. The laser beam 4100 (a surgical laser beam, or a temperature-sensing laser beam, or both) can be selectively delivered to the tissue 4400 surface according to the requirements of the method. The laser beam 4100 is delivered with a flexible multimode fiber or a flexible fiber bundle and accessory optics (not shown in FIG. 4) to enable rotation. The linear array ultrasound transducer 4200 is positioned to detect the photoacoustic waves without blocking the laser beam 4100. The linear array transducer 4200 can be based on conventional ultrasound transducers or optical detection technology. The interface medium 4300 allows the transmission of the laser beam and the photoacoustic waves with minimal losses. The interface medium 4300 can also be stored in a more simplified configuration. In this configuration, the linear array ultrasound transducer is in direct acoustic contact with the tissue at the tissue surface immediately adjacent to the tissue surface area illuminated by the laser beam delivered through free space. One advantage of this simplified configuration is that it can simultaneously enable the delivery of skin cooling agent through free space. Spectral photoacoustic signals are acquired by the linear array ultrasound transducer 4200 from three-dimensional depth-resolved space within the tissue and are digitized and analyzed by the control system 1110 shown in FIG. 1.

[0035] FIG. 5 shows another schematic example of a patient interface comprising a single-element ultrasound transducer 5200, an acoustic wave reflector 5500, and an interface medium 5300 in acoustic contact with tissue 5400. The optical delivery unit to be shown in the patient interface is not different from that of a conventional laser SP system, so for simplicity, the optical delivery unit is omitted in FIG. 5. A laser beam 5100 (a surgical laser beam, or a temperature-sensing laser beam, or both) can be selectively delivered to the tissue 5400 surface according to the requirements of the method. The single-element ultrasound transducer 5200 is positioned to detect the photoacoustic waves reflected by the acoustic reflector 5500 without blocking the laser beam 5100. The use of the acoustic reflector 5500 allows the laser beam and the transducer to be placed on the same side of the tissue without blocking each other. The single-element ultrasound transducer 5200 can be based on conventional ultrasound transducers or optical detection technology. The interface medium 5300 allows the transmission of the laser beam and the photoacoustic waves with minimal losses. Spectral optoacoustic signals are acquired from one-dimensional depth-resolved space within the tissue by the single element ultrasound transducer 5200 and are digitized and analyzed by the control system 1110 shown in FIG.

[0036] FIG. 6 shows another schematic example of a patient interface comprising a linear array ultrasonic transducer 6200, an acoustic reflector 6500, and an interface medium 6300 in acoustic contact with tissue 6400. The optical delivery unit to be shown in the patient interface is not different from that of a conventional laser SP system, so for simplicity, the optical delivery unit is omitted in FIG. 6. A laser beam 6100 (a surgical laser beam, or a temperature-sensing laser beam, or both) can be selectively delivered to the tissue 6400 surface according to the requirements of the method. The linear array ultrasonic transducer 6200 is arranged to detect the photoacoustic waves reflected by the acoustic reflector 6500. The use of the acoustic reflector 6500 allows the laser beam and the transducer to be placed on the same side of the tissue without blocking each other. The linear array ultrasonic transducer 6200 can be based on conventional ultrasonic transducers or optical detection technology. The interface medium 6300 allows the transmission of the laser beam and the photoacoustic waves with minimal losses. Spectral optoacoustic signals are acquired from the two-dimensional depth-resolved space within the tissue by the linear array ultrasound transducer 6200 and are digitized and analyzed by the control system 1110 shown in FIG.

[0037] FIG. 7 shows another schematic example of a patient interface comprising a linear array ultrasonic transducer 7200, an acoustic reflector 7500, an interface medium 7300 in acoustic contact with tissue 7400, and a rotation stage 7600. The rotation stage 7600 mounts the linear array ultrasonic transducer 7200, the acoustic reflector 7500, and the interface medium 7300 and rotates around the axis of the laser beam. The arrow 7700 indicates the rotation direction of the rotation stage 7600. Since the optical delivery unit to be shown on the patient interface is no different from that of a conventional laser SP system, for simplicity, the optical delivery unit is omitted in FIG. 7. The laser beam 7100 (the surgical laser beam, or the temperature-sensing laser beam, or both) can be selectively delivered to the tissue 7400 surface according to the requirements of the method. The linear array ultrasonic transducer 7200 is arranged to detect the photoacoustic waves reflected by the acoustic reflector 7500. The use of an acoustic reflector 7500 allows the laser beam and the transducer to be placed on the same side of the tissue without blocking each other. The linear array ultrasound transducer 7200 can be based on conventional ultrasound transducers or optical detection technology. The interface medium 7300 allows the transmission of the laser beam and the photoacoustic waves with minimal loss. Spectral photoacoustic signals are acquired by the linear array ultrasound transducer 7200 from the three-dimensional depth-resolved space within the tissue and are digitized and analyzed by the control system 1110 shown in FIG. 1.

[0038] FIG. 8 shows another schematic example of a patient interface comprising a Dammann grating 8200, a two-dimensional array ultrasonic transducer 8400, and an interface medium 8500 in acoustic contact with tissue 8600. The optical delivery unit to be shown in the patient interface is no different from that of a conventional laser SP system, so for simplicity, the optical delivery unit is omitted in FIG. 8. A laser beam 8100 (a surgical laser beam, or a temperature-sensing laser beam, or both) can be selectively delivered to the tissue 8600 surface according to the requirements of the method. The laser beam is converted into a two-dimensional array laser beam 8300 by the Dammann grating 8200. The two-dimensional array laser beam passes through the free space not taken up by the two-dimensional array ultrasonic transducer 8400 and the interface medium 8500 before reaching the tissue 8600. The excited photoacoustic wave is detected by the two-dimensional array ultrasonic transducer 8400. The two-dimensional array ultrasound transducer 8400 can be based on conventional ultrasound transducers or optical detection technology. The interface medium 7300 allows the transmission of the laser beam and photoacoustic waves with minimal loss. Spectroscopic photoacoustic signals 8700 are acquired by the two-dimensional array ultrasound transducer 8400 from the three-dimensional depth-resolved space within the tissue, and are digitized and analyzed by the control system 1110 shown in FIG. 1.

[0039] FIG. 9 shows another schematic example of a patient interface comprising a dichroic mirror 9200, a scanning lens 9500, a two-dimensional galvano scanner 9600, and a Fabry-Perot sensor 9300 in acoustic contact with tissue 9400. The light delivery unit to be shown in the patient interface is not different from that of a conventional laser SP system, so for simplicity, the light delivery unit is omitted in FIG. 9. A laser beam 9100 (a surgical laser beam, or a temperature-sensing laser beam, or both) can be selectively delivered to the tissue 9400 surface according to the requirements of the method. Both the dichroic mirror 9200 and the Fabry-Perot sensor 9300 are transparent to the laser beam 9100 (a surgical laser beam, or a temperature-sensing laser beam, or both). The ultrasound detection laser beam 9700 is scanned by the two-dimensional galvanometer scanner 9600 and the scanning lens 9500, and is reflected by the dichroic mirror 9200 to the Fabry-Perot sensor 9300 to detect the photoacoustic waves. The Fabry-Perot sensor 9300 includes two dielectric mirror layers and an acoustic wave sensing layer between the two dielectric mirrors. In other implementations, the ultrasound detection laser beam 9700 can be in the form of a one-dimensional array laser beam or a two-dimensional array laser beam. The optical system for sending out the ultrasound detection laser beam 9700 and detecting the reflected ultrasound detection laser beam 9700 modulated by the photoacoustic waves is omitted in FIG. 9 for simplicity. The spectroscopic photoacoustic signals can be acquired from one-dimensional, two-dimensional, or three-dimensional depth-resolved space in the tissue and digitized and analyzed by the control system 1110 shown in FIG. 1.

[0040] FIG. 10 shows another schematic example of a patient interface with a vacuum structure 10200, a circular array ultrasonic transducer 10400, and an interface medium 10300 in acoustic contact with tissue 10500. The optical delivery unit to be shown in the patient interface is no different from that of a conventional laser SP system, so for simplicity, the optical delivery unit is omitted in FIG. 10. A laser beam 10100 (a surgical laser beam, or a temperature-sensing laser beam, or both) can be selectively delivered to the tissue 10500 surface according to the requirements of the method. The circular array transducer 10400 can be based on a conventional ultrasonic transducer or optical detection technology. The circular array ultrasonic transducer 10400 is in acoustic contact with the wall of the vacuum structure 10200. The vacuum structure 10200 is designed to suck a portion of tissue into it in a similar manner to a body cupping device. An interface medium 10300, such as water, can be injected into the bottom of the vacuum structure to immerse the tissue sucked into the vacuum structure. The laser beam 10100 illuminates the tissue inside the vacuum structure and excites photoacoustic waves. The circular array ultrasound transducer 10400 can acquire spectroscopic photoacoustic signals from a two-dimensional depth-resolved space in the tissue. The circular array ultrasound transducer 10400 can be adjusted in elevation to acquire spectroscopic photoacoustic signals from a three-dimensional depth-resolved space in the tissue. The spectroscopic photoacoustic signals are digitized and analyzed by the control system 1110 shown in FIG. 1.

[0041] FIG. 11 shows an example of wavelength tuning operation for optimizing SP laser treatment of an unknown dye. The optimal surgical laser wavelength should maximize the laser energy deposition ratio of the unknown dye and the natural chromophore. The following procedure is designed for a laser SP surgical system in which the surgical laser pulse width is short enough to effectively excite the photoacoustic signal. However, if the surgical laser pulse is too long to effectively excite the photoacoustic signal, the following procedure requires the use of a short temperature-sensing laser pulse generated by a more advanced dual-pulse (surgical laser pulse followed by a delayed temperature-sensing laser pulse) laser system to excite the photoacoustic signal. First, a set of wavelength points containing the characteristic peaks and valleys of oxygenated and deoxygenated hemoglobin is established. Second, a tissue region is selected, and the patient interface is manipulated to make acoustic contact with the selected tissue region. Third, the acoustic detector is preferably configured in a mode to acquire two-dimensional depth-resolved photoacoustic tissue information with a single surgical laser pulse. It should be noted that the acoustic detector can be configured to acquire one-dimensional depth-resolved tissue information in a simplified implementation, and the following procedure may require slight modifications. Fourth, the control system delivers multiple sub-therapeutic surgical laser pulses for each wavelength point and acquires photoacoustic signals detected by the acoustic detector. Fifth, averaged two-dimensional depth-resolved tissue information is obtained after tomographic reconstruction of each wavelength point. Sixth, unknown dyes are identified along with natural chromophores after their relative extinction coefficients are calculated and their relative extinction coefficient curves are fitted. If the natural chromophores do not exist in the two-dimensional depth-resolved tissue space, known extinction coefficient curves of natural chromophores from literature can be used. Seventh, the relative energy deposition ratio curves of unknown dyes and natural chromophores are calculated. Finally, the optimal surgical laser wavelength is established with an algorithm that weights different priorities to different natural chromophores for different types of unknown dyes. For different types of unknown dyes, multiple laser treatments using optimized laser wavelengths can be performed consecutively to obtain optimal laser treatment results.The same technique of FIG. 11 applies to pulsed laser (coherent light source) SP surgical systems, other non-coherent pulsed light source SP surgical systems where the central wavelength of the non-coherent pulsed light source is tuned, and other general radiation SP systems that use a wavelength-tunable, pulsed or modulated radiation beam to effectively excite photoacoustic waves from lesions in tissue or exogenous contrast agents attached to lesions in tissue.

[0042] For more advanced dual-pulse (surgical laser pulse followed by a delayed temperature sensing laser pulse) laser SP surgery systems, the present invention provides a method for photoacoustic temperature pulse sensing in living tissue. This method overcomes the limitations of methods in the prior art. This method, detailed in Figs. 12 and 13, can non-invasively measure the dynamic temperature (temperature pulse) of a surgical target in tissue that has been rapidly heated by a short surgical laser pulse after calibration. The calibration procedure is based on the hypothesis that the rapid heating of a surgical target by a laser pulse is a linear process and the maximum temperature rise of the surgical target is proportional to the laser energy deposited in the surgical target when the temperature is measured immediately after the surgical laser pulse.

[0043] FIG. 12 shows an example of calibrating the relative logarithmic function of the temperature dependence of the Gruneisen parameters of tissue. First, start the calibration process by measuring the equilibrium temperature T0 of the tissue. Second, measure the photoacoustic signal of the surgical target excited by the temperature-sensing laser pulse with a constant laser pulse energy, perform a logarithmic operation on the photoacoustic signal amplitude to obtain a baseline signal. Third, send a surgical laser pulse to rapidly heat the surgical target, and measure the photoacoustic signal excited by the surgical laser pulse. Fourth, send both the surgical laser pulse and the temperature-sensing laser pulse, measure the photoacoustic signal excited by the dual pulse, subtract the photoacoustic signal excited by the surgical laser pulse from the photoacoustic signal excited by the dual pulse, calculate the amplitude of the photoacoustic signal excited by the temperature-sensing laser pulse, separate the temperature-dependent part from the other temperature-independent part by a logarithmic operation, subtract the baseline signal, and calculate the relative logarithmic function value logΓ(T0+δT)-logΓ(T0). Γ denotes the Gruneisen parameter of the tissue, and δT denotes the temperature rise caused by the laser pulse heating abruptly for the temperature point of T0+δT. Fifth, if the recorded photoacoustic signal amplitude does not show abrupt increase due to laser-induced cavitation at 100 °C, wait until the tissue temperature returns to the original equilibrium temperature. Then, adjust the abrupt heating laser pulse energy by its ki times and go back to the third step to obtain another relative logarithmic function value of logΓ(T0+kiδT)-logΓ(T0) for the temperature point of T0+kiδT. In order to have a more accurate measurement of 100 °C, it is necessary to keep the increase in the surgical laser energy small. If laser-induced cavitation is observed, proceed to the last step. In the last step, there is. Thus, calculate the absolute temperature rise (k0δT, k1δT, …, kmδT) due to each surgical laser pulse and fit the logΓ(T)-logΓ(T0) function between T0 and 100 °C. where T represents temperature. The equilibrium temperature can be the patient's body temperature. It can also be the equilibrium temperature of ex vivo tissue in an environment of known temperature. In practice, the laser pulse energy varies from pulse to pulse. The above procedure requires compensation by simultaneous laser pulse energy monitoring.As long as both the starting temperature and the temperature being measured are between T0 and 100° C., the calibrated relative logarithmic function Gruneisen parameters of the tissue are valid for temperature sensing operations, as detailed below.

[0044] FIG. 13 shows an example of a dynamic temperature (temperature pulse) sensing operation of a surgical target in tissue after it has been rapidly heated by a short surgical laser pulse. However, if the delay time between the surgical laser pulse and the temperature sensing laser pulse is adjusted, the dynamic temperature variation (temperature pulse) profile of the surgical target along time can be accurately measured by repeating the following dynamic temperature (temperature pulse) sensing procedure. It is important to start from a known body temperature and to know the starting point of the relative logarithmic function of the Gruneisen parameters. First, measure the body temperature before the laser surgical intervention. Second, send the temperature sensing laser pulse and measure the amplitude of the excited photoacoustic signal at the body temperature and calculate its logarithm as the baseline signal. Third, send only the surgical laser pulse and measure the photoacoustic signal excited by the surgical laser pulse. Fourth, send both the surgical laser pulse and the temperature sensing laser pulse and measure the excited photoacoustic signal. Fifth, calculate the pulse amplitude of the photoacoustic signal excited by the temperature sensing laser and separate the temperature-dependent part from the other temperature-independent parts by a logarithmic operation. Sixth, calculate the relative logarithmic function value by subtracting the baseline signal. Finally, the dynamic temperature (temperature pulse) at the end of the surgical laser pulse is determined from the calibrated relative logarithmic function of the Gruneisen parameters of the tissue. Temperature sensing of the surgical target heated by a continuous wave laser only requires two measurements of the photoacoustic signal excited by the temperature-sensing laser pulse, at body temperature and at a certain time point during the CW laser surgery. More accurate measurement results are expected with the more accurate and non-invasive calibration method of Figure 12. The same technique of Figures 12-13 is applicable to pulsed laser (coherent light source) surgical systems, other non-coherent pulsed light source surgical systems, high intensity focused ultrasound therapy systems, and other common radiation SP systems that use a surgical pulsed radiation beam to rapidly heat a lesion in tissue or an exogenous contrast agent attached to a lesion in tissue, and use a pulsed or modulated temperature-sensing radiation beam to effectively excite photoacoustic waves from the lesion in tissue or an exogenous contrast agent attached to a lesion in tissue.

[0045] In applications such as laser treatment of vascular malformations, it is desirable to effectively and rapidly heat the surgical target with a laser pulse whose laser pulse width matches the thermal relaxation time of the surgical target. Most of the energy of the laser pulse is confined to the surgical target instead of being diffused to the surrounding healthy tissue. Computer simulations using tissue models and dimensional information of the surgical target are the only methods available in the research field of laser treatment of vascular malformations to estimate the thermal relaxation time of the surgical target in tissue. However, Figure 14 shows an example of optimizing the surgical laser pulse width by measuring the thermal relaxation time of the surgical target in tissue in vivo. In addition to the surgical laser pulse, a temperature-sensing laser pulse is required to perform the task of measuring the thermal relaxation time of the surgical target. First, the photoacoustic signal of the surgical target at body temperature and at body temperature is measured with the temperature-sensing laser pulse, and the logarithm of the photoacoustic signal is calculated as the baseline signal. Second, a sub-therapeutic surgical laser pulse is sent and the excited photoacoustic signal is measured. Third, both the sub-therapeutic surgical laser pulse and the temperature-sensing laser pulse are sent with a precise delay time, the excited photoacoustic signal of the dual pulse is measured, and the amplitude of the photoacoustic signal excited by the temperature-sensing laser pulse is calculated. Fourth, separate the temperature-dependent part from the temperature-independent part by logarithmic calculation and calculate the relative logarithmic function value by subtracting the baseline signal. Fifth, determine the temperature of the surgical target at the correct delay time. Sixth, wait until the surgical target temperature returns to body temperature and determine whether there are enough delay time points to fit the temperature vs. delay time curve of the surgical target. If there are not enough time points, change the delay time value and repeat the procedure between steps three and six until the temperature vs. delay time curve can be fitted and the thermal relaxation time of the surgical target can be derived. The optimal surgical laser pulse width is determined to be equal to the measured thermal relaxation time of the surgical target.The same technique of FIG. 14 applies to pulsed laser (coherent light source) SP surgical systems, other non-coherent pulsed light source SP surgical systems, and other common radiation SP systems that use a surgical pulsed radiation beam to rapidly heat a lesion in tissue or an exogenous contrast agent attached to a lesion in tissue, and a pulsed or modulated temperature-sensing radiation beam to effectively excite photoacoustic waves from the lesion in tissue or an exogenous contrast agent attached to a lesion in tissue.

[0046] In conventional laser SP surgery, the surgical laser pulse energy is selected according to the clinician's past experience. However, the inclusion of an acoustic detector makes it possible for the first time to objectively optimize the surgical laser pulse energy. For laser treatment of vascular malformations, the optimal laser pulse energy rapidly heats the selected surgical target to a predetermined temperature for photocoagulation. For applications based on laser photodisruption, such as laser tattoo removal, the optimal surgical laser pulse energy rapidly heats the selected surgical target to 100°C, causing laser-induced cavitation.

[0047] FIG. 15 shows an example of optimizing the surgical laser pulse energy in a tunable laser during laser photocoagulation SP surgery. Before the laser power optimization, assume that the laser wavelength has already been optimized. First, select a surgical target with two-dimensional depth-resolved tissue information excited by a temperature-sensing laser pulse and acquired by an acoustic detector. Second, optimize the laser pulse width after measuring the thermal relaxation time of the surgical target. Third, send a surgical laser pulse and measure the excited photoacoustic signal. Fourth, send both a surgical laser pulse and a temperature-sensing laser pulse, measure the photoacoustic signal excited by the dual pulse, and calculate the amplitude of the photoacoustic signal excited by the temperature-sensing laser pulse. Fifth, separate the temperature-dependent part with a logarithmic operation, calculate the relative logarithmic function value by subtracting the baseline signal acquired in the previous step, and determine the temperature of the surgical target. Sixth, determine whether the surgical target reaches the predetermined temperature for photocoagulation. If not, wait until the temperature returns to its original body temperature, increase the laser pulse energy, and return to the third step. When the temperature reaches a predetermined temperature for photocoagulation, the operation is terminated with the laser pulse energy optimized for photocoagulation of the surgical target. The same technique of FIG. 15 applies to pulsed laser (coherent light source) SP surgical systems, other non-coherent pulsed light source SP surgical systems, high intensity focused ultrasound therapeutic systems for optimizing ultrasound beam energy, and other common radiation SP systems that use a surgical pulsed radiation beam to rapidly heat the lesion in tissue or an exogenous contrast agent attached to the lesion in tissue, and a pulsed or modulated temperature-sensing radiation beam to effectively excite photoacoustic waves from the lesion in tissue or an exogenous contrast agent attached to the lesion in tissue.

[0048] A distinct feature of laser photodisruption is the generation of acoustic shock waves due to laser-induced cavitation. We assume that once acoustic shock waves due to laser-induced cavitation are observed, there is no need to further increase the laser pulse energy. For laser photodisruption during laser tattoo removal applications, it is desirable to use a tunable surgical laser instead of the more advanced tunable dual-pulse (surgical laser pulse followed by a delayed temperature-sensing laser pulse) surgical laser. Figure 16 shows an example of in vivo surgical laser pulse energy optimization during laser photodisruption SP surgery. The following procedure is designed for a laser SP surgery system in which the surgical laser pulse width is short enough to effectively excite the photoacoustic signal. We assume that the laser wavelength has already been optimized before the laser pulse energy optimization. First, select a surgical target with two-dimensional depth-resolved tissue information to be excited by the surgical laser pulse and acquired by the acoustic detector. Second, measure the thermal relaxation time of the surgical target and optimize the surgical laser pulse width according to the thermal relaxation time of the surgical target. Third, send the surgical laser pulse and measure the excited photoacoustic signal. Fourth, draw one point on the curve of photoacoustic signal versus laser pulse energy to determine whether there is a sudden increase in photoacoustic signal due to laser-induced cavitation. If there is no sudden increase in photoacoustic signal, we can wait for the surgical target to recover to its original body temperature, increase the surgical laser pulse energy, and return to the third step. If a sudden increase in photoacoustic signal due to laser-induced cavitation is observed, the optimal laser pulse energy is obtained. The same technique in FIG. 16 applies to pulsed laser (coherent light source) SP surgical systems, other non-coherent pulsed light source SP surgical systems, and other common radiation SP systems that use a surgical pulsed radiation beam to suddenly heat the lesion in tissue or the exogenous contrast agent attached to the lesion in tissue, and excite photoacoustic waves from the lesion in tissue or the exogenous contrast agent attached to the lesion in tissue.

[0049] The method of Figs. 15 and 16 provides real-time control of the surgical laser pulse energy during laser SP surgery. However, it is possible to accurately determine the surgical laser pulse energy required for the surgical target to reach a predetermined temperature without actually performing laser surgery. Fig. 17 shows an example of determining the optimal surgical laser pulse energy to reach a predetermined temperature without performing laser surgery. Assume that the laser wavelength has already been optimized. First, select a surgical target with two-dimensional depth-resolved tissue information excited by a temperature-sensing laser pulse and acquired by an acoustic detector. Second, optimize the laser pulse width after measuring the thermal relaxation time of the surgical target. Third, send a sub-therapeutic surgical laser pulse and detect its photoacoustic signal. Fourth, send both a sub-therapeutic surgical laser pulse and a temperature-sensing laser pulse immediately after the end of the surgical laser pulse. Fifth, measure the photoacoustic signal excited by both the surgical laser pulse and the temperature-sensing laser pulse, and calculate the amplitude of the photoacoustic signal excited by the temperature-sensing laser pulse. Sixth, separate the temperature-dependent part with a logarithmic operation and calculate the relative logarithmic function value by subtracting the baseline signal acquired in the previous step. Finally, the temperature rise due to the sub-therapeutic surgical laser pulse is calculated to determine the surgical laser pulse energy required to rapidly heat the target to the desired temperature. The same technique of FIG. 17 applies to pulsed laser (coherent light source) SP surgical systems, other non-coherent pulsed light source SP surgical systems, high intensity focused ultrasound therapeutic systems, and other common radiation SP systems that use a surgical pulsed radiation beam to rapidly heat the lesion in tissue or an exogenous contrast agent attached to the lesion in tissue, and a pulsed or modulated temperature-sensing radiation beam to effectively excite photoacoustic waves from the lesion in tissue or an exogenous contrast agent attached to the lesion in tissue.

[0050] Another potential use of the acoustic detector is to provide in vivo temperature calibration for skin cooling devices that protect the epidermal layer of the skin in laser SP surgery. Skin cooling is widely used in the laser treatment of vascular malformations. Skin cooling can also effectively prevent laser-induced cavitation in the epidermis during laser tattoo removal. Figure 18 shows an example of the operation of optimized laser SP surgery with skin cooling. First, tune the surgical laser wavelength to maximize the SP surgery effect. Second, optimize the surgical laser pulse width according to the measured thermal relaxation time of the surgical target. Third, optimize the surgical laser pulse energy according to the expected surgical effect (photocoagulation or photodisruption). Fourth, measure the body temperature, send a temperature-sensing laser pulse, measure the amplitude of the excited photoacoustic signal of the epidermal target, and calculate the logarithmic value as the baseline signal. Fifth, adjust the skin cooling parameters, apply skin cooling, and send a delayed temperature-sensing laser pulse. Sixth, measure the amplitude of the photoacoustic signal excited by the temperature-sensing laser pulse, take the logarithm, and subtract the baseline signal to calculate the relative logarithmic function value. Seventh, determine the temperature of the epidermal target and whether the epidermal target has been cooled to the predetermined temperature. If not, wait for the temperature to recover and return to step 5. If cooled to the predetermined temperature, perform the optimized SP surgery. The same technique of FIG. 18 applies to pulsed laser (coherent light source) SP surgery systems, other non-coherent pulsed light source SP surgery systems, and other common radiation SP systems that use a surgical pulsed radiation beam to rapidly heat the lesion in tissue or the exogenous contrast agent attached to the lesion in tissue, and use a pulsed or modulated temperature-sensing radiation beam to effectively excite photoacoustic waves from the lesion in tissue or the exogenous contrast agent attached to the lesion in tissue. [Industrial Applicability]

[0051] The above examples are the most comprehensive implementations of the technology disclosed in this invention. Its wide range of industrial applications include laser treatment of vascular malformations, laser treatment of port wine stains, laser tattoo removal, and other innovative SP procedures in oncology, ophthalmology, cardiology, neurology, and dentistry. Partial implementations of the technology disclosed in this invention may already result in highly competitive surgical products in the market. For example, the current tattoo removal laser with the addition of an ultrasound detector and associated hardware and software could become the first generation of artificial intelligence (AI) tattoo removal surgical laser system that ensures optimized surgical results without the need for skin typing or laser parameter selection by a physician or laser technician. More advanced AI tattoo removal surgical systems are possible with tunable laser wavelengths and tunable laser pulse widths. Variations and enhancements of the described implementations, as well as other implementations, can be made based on what one skilled in the art has described for the many potential medical applications.

[0052] <Prior art documents> <Patent Documents> 5840023 filed in January 1996 by A Oraevsky et al. 6309352 B1 filed in October 1998 by Oraevsky et al. 2015 / 0216420 A1 filed in February 2015 by Oraevsky et al. 4303343 filed in February 1980 by Patel et al. No. 62 / 617,681 filed by Rao Bin in January 2018. 7322972 B2 Filed in February 2003 by Viator et al. 2017 / 0014317A1 filed in July 2016 by Youngbull et al. 2002 / 0019625A1 filed by Zion Azar in April 2001. 5759200 Filed by Zion Azar in September 1996. Filed in August 2011 by Nikolai et al., WO 2013 / 033145A1.

[0053] <Non-Patent Document> Adams et al., "Thermal diffusivity and thickness measurements for solid samples utilizing the optoacoustic effect," The Analyst 102, 678 (1977). Barua et al., "Laser-tissue interaction in tattoo removal by Q-switched lasers," J Cutan Aesthet Surg 8, 5-8(2015). Bernstein, "Laser Tattoo Removal," SEMINARS IN PLASTICS SURGERY 21, 175-192(2007). C. T. Andrew, "Applications of photoacoustic sensing techniques," Rev. Mod. Phys. 58, 381-431(1986). Cox et al., "Quantitative spectroscopic photoacoustic imaging: a review," Journal of Biomedical Optics 17(6), 061202(2012). Esenaliev et al., "Real-time optoacoustic monitoring of temperature in tissues", Proc. SPIE 3601, Laser-Tissue Interaction X: Photochemical, Photothermal, and Photomechanical(1999). Gary A. West, "Photoacoustic spectroscopy," Review of Scientific Instruments 54, 797 (1983). Hordvik et al., "Photoacoustic technique for determining optical absorption coefficients in solids," Applied Optics 16, 101-107 (1977). Larin et al., "Monitoring of temperature distribution in tissues with optoacoustic technique in real time", Proc. SPIE 3916, Biomedical Optoacoustics(2000). Liu et al., "In vivo, high-resolution, three-dimensional Imaging of port wine stain microvasculature in human skin," Lasers in Surgery and Medicine 45, 626-632(2013). Maslov et al., "Optical-resolution photoacoustic microscopy for in vivo imaging of single capillaries", Optics Letters 33, 929-931(2008). Maslov et al., "In vivo dark-field reflection-mode photoacoustic microscopy," Optics Letters 30, 625-627(2005). Nelson et al., "Imaging blood flow in human port-wine stain in situ and in real time using optical Doppler tomography," Arch Dermatol. 137, 741-744(2001). Ortiz et al., "Port-wine stain laser treatment and novel approaches," Facial Plast Surg 28, 611-620(2012). Peach et al., "Colour shift following tattoo removal with Q-switched Nd-YAG laser (1064 / 532), " British Journal of Plastic Surgery 52,482-487(1999). Rao et al., "Smart laser treatment of port-wine stain in children," NIH grant abstract, published on Aug 1st, 2012. http: / / grantome.com / grant / NIH / K99-AR062530-01. Shah et al., "Photoacoustic imaging and temperature measurement for photothermal cancer therapy," Journal of Biomedical Optics 13(3), 034024(2008). Viator et al., "In vivo port-wine stain depth determination with a photoacoustic probe," Appl. Opt. 42, 3215-3224(2003). Wang, "Prospects of photoacoustic tomography," Medical Physics 35, 5758 (2008). Wenzel, "Current concepts in laser tattoo removal," Skin Therapy Letter 15, 3-5(2010). Zapka et al., "Noncontact optoacoustic monitoring of flame temperature profiles," Optics Letters 7, 477-479 (1982). Zhang et al., "Backward-mode multiwavelength photoacoustic scanner using a planar Fabry-Perot polymer film ultrasound sensor for high-resolution three-dimensional imaging of biological tissues," Applied Optics 47,561-577(2008).

Claims

1. a tunable radiation source (1100) configured to emit radiation; a patient interface (1200) including a radiation delivery unit (1210) configured to deliver the radiation to tissue (1000) and an ultrasound detector (1220) for detecting photoacoustic waves (1020) excited by the radiation from one or more targets within the tissue (1000); a control system (1110) configured to acquire characteristics of the tissue (1000) including a relative logarithmic function of the temperature dependence of the Gruneisen parameters, a skin cooling parameter, and a thermal relaxation time and a dynamic temperature of the one or more targets based on measurements of the detected photoacoustic waves, determine optimal characteristics of the tunable radiation source (1100) for an optimal surgical outcome for the one or more targets in the tissue (1000) including a central wavelength, a pulse width or pulse duration, and a pulse energy, define the characteristics and the optimal characteristics, and adjust the tunable radiation source (1100) towards an optimal surgical outcome based on the optimal characteristics; A selective photothermolysis (SP) surgical system comprising:

2. the optimal characteristics of the tunable radiation source for one or more targets include an optimal central wavelength; inputting a series of radiation wavelengths for a spectral scan; For each of said radiation wavelengths, receiving a photoacoustic signal from an ultrasound detector corresponding to a sub-therapeutic radiation pulse or modulation; reconstructing 2D depth-resolved tissue extinction coefficient information for said radiation wavelengths; and calculating a relative extinction coefficient curve for all radiation absorbers in the tissue and identifying the one or more targets; calculating a relative energy deposition ratio curve of the one or more targets in the tissue relative to the natural chromophores in the tissue; determining the optimal central wavelength for the one or more targets within the tissue; The surgical system of claim 1 .

3. The tissue properties include the relative logarithm of the temperature dependence of the Gruneisen parameters, i. inputting the equilibrium temperature of the tissue; ii. inputting a first photoacoustic signal of a target corresponding to a second auxiliary radiation pulse having a constant radiation pulse energy and calculating a baseline signal as the logarithm of an amplitude of the first photoacoustic signal at the equilibrium temperature; iii. inputting a second photoacoustic signal of the target corresponding to the first radiation pulse; iv. inputting a third photoacoustic signal of the target corresponding to both the first radiation pulse and the second auxiliary radiation pulse, subtracting the second photoacoustic signal from the third photoacoustic signal, and performing a logarithmic operation on the result of the subtraction; v. subtracting the baseline signal from the result of the logarithmic calculation; vi. repeating steps iii through vi to process data corresponding to Ki-fold increased energy of said first radiation pulse, where i is a loop index and Ki is the i-th numerical element of a monotonically increasing numerical sequence {Ki}; vii. calculating the absolute temperature rise corresponding to the first sequence of radiation pulses and fitting a relative logarithmic function of the temperature dependence of the Grueneisen parameters of the tissue; The surgical system of claim 1 .

4. the characteristics of the one or more targets include a dynamic temperature of the one or more targets, the dynamic temperature corresponding to a temperature at a dynamic delay time; Enter your body temperature, inputting an amplitude of a first photoacoustic signal corresponding to a second auxiliary radiation pulse at the body temperature and calculating a logarithm thereof as a baseline signal; inputting a second photoacoustic signal corresponding to only the first radiation pulse; inputting a third photoacoustic signal corresponding to both the first radiation pulse and the second auxiliary radiation pulse; calculating an amplitude of a fourth photoacoustic signal excited by the second auxiliary radiation pulse at the dynamic temperature by subtracting the second photoacoustic signal from the third photoacoustic signal, and separating a temperature-dependent portion of the fourth photoacoustic signal from a temperature-independent portion by a logarithmic operation; calculating a relative logarithmic value by subtracting the baseline signal from the logarithmic result; determining the dynamic temperature based on the body temperature, the calculated relative logarithmic function value, and a calibrated relative logarithmic function curve of the Gruneisen parameters of the tissue; The surgical system of claim 1 .

5. the optimal characteristics of the tunable radiation source for one or more targets include an optimal pulse duration of the tunable radiation source for the one or more targets; i. inputting a body temperature of the one or more targets and a first photoacoustic signal corresponding to a second supplemental radiation pulse at body temperature, and calculating a logarithm of an amplitude of the first photoacoustic signal as a baseline signal; ii. inputting a second photoacoustic signal of the one or more targets corresponding to a sub-therapeutic first radiation pulse; iii. inputting a third photoacoustic signal of the one or more targets corresponding to both the first radiation pulse and the second auxiliary radiation pulse at a precise delay time, and calculating a fourth photoacoustic signal excited by the delayed second auxiliary radiation pulse by subtracting the second photoacoustic signal from the third photoacoustic signal; iv. separating a temperature dependent portion of the fourth photoacoustic signal from a temperature independent portion using a logarithmic operation and calculating a relative logarithmic function value by subtracting the baseline signal from the result of the logarithmic operation; v. determining the temperature of the one or more targets at the precise delay time; and vi. repeating steps iii through v for all delay time points to fit a temperature vs. delay time curve; vii. fitting the temperature versus delay time curve to derive a thermal relaxation time for the one or more targets, and determining the optimal pulse duration of the tunable radiation source for the one or more targets based on the thermal relaxation time; The surgical system of claim 1 .

6. the optimal characteristics of the tunable radiation source for one or more targets include an optimal pulse energy of the tunable radiation source for photodisruption of the one or more targets; i. calculating a thermal relaxation time of the target; ii. inputting a photoacoustic signal of the target corresponding to a first radiation pulse along with a pulse energy; iii. plotting a curve of photoacoustic signal versus radiation pulse energy of the first radiation pulse; iv. repeating steps ii through iii for a series of said first radiation pulses with increasing pulse energy but not inducing cavitation by observing the linearity of said curve; v. determining the optimal pulse energy for photodisruption based on a current pulse energy of the first radiation pulse; The surgical system of claim 1 .

7. the optimal characteristics of the tunable radiation source for one or more targets include an optimal pulse energy of the tunable radiation source to achieve a predetermined temperature of the one or more targets; i. calculating an optimal pulse duration of the tunable radiation source for the one or more targets; ii. inputting a first photoacoustic signal of a target corresponding to a sub-therapeutic radiation pulse; iii. inputting a second optoacoustic signal of the target corresponding to both a first radiation pulse and a second auxiliary radiation pulse, and calculating a third optoacoustic signal of the second auxiliary radiation pulse immediately after the end of the first radiation pulse by subtracting the first optoacoustic signal from the second optoacoustic signal; iv. isolating the temperature dependent portion of the third photoacoustic signal using a logarithmic operation and calculating a relative logarithmic function value by subtracting the baseline signal obtained in step i from the result of the logarithmic operation; v. calculating a temperature rise due to the sub-therapeutic pulse based on body temperature, the relative logarithmic value, and a calibrated relative logarithmic curve of the Gruneisen parameters of the tissue; vi. calculating the required energy of the first radiation pulse to heat the target to a predetermined temperature; The surgical system of claim 1 .

8. the optimal characteristics of the tunable radiation source for one or more targets include an optimal pulse energy of the tunable radiation source for photocoagulation of the one or more targets; i. calculating a thermal relaxation time of the target; ii. inputting a first photoacoustic signal of the target corresponding to a first radiation pulse; iii. inputting a second optoacoustic signal of the target corresponding to both the first radiation pulse and a second auxiliary radiation pulse, and calculating a third optoacoustic signal of the second auxiliary radiation pulse immediately after the end of the first radiation pulse by subtracting the first optoacoustic signal from the second optoacoustic signal; iv. isolating the temperature dependent portion of the third photoacoustic signal using a logarithmic operation and calculating a relative logarithmic function value by subtracting the baseline signal obtained in step i from the result of the logarithmic operation; v. calculating a heating temperature due to the first radiation pulse based on body temperature, the relative logarithmic function value, and a calibrated relative logarithmic function curve of the Gruneisen parameters of the tissue; vi. if the heating temperature has not reached the photocoagulation temperature, repeating steps ii to v; vii. determining the optimal pulse energy of the tunable radiation source for photocoagulation of the one or more targets based on a current pulse energy of the first radiation pulse; The surgical system of claim 1 .

9. the tissue properties include relative logarithmic functions of temperature dependence of Gruneisen parameters; determining a series of temperature points corresponding to a series of short heating laser pulses with increasing laser pulse energy based on the recorded photoacoustic waves and the linear relationship between the temperature increase and the excited photoacoustic waves, until the last laser pulse induces cavitation in the tissue and the target reaches 100°C; calibrating the photoacoustic response of the target corresponding to a temperature-sensing laser pulse at a constant pulse energy immediately after reaching the temperature point and correcting the photoacoustic response corresponding to a short heating laser pulse alone; determining a relative logarithm of the temperature dependence of the Grueneisen parameters of a medium surrounding the target; The surgical system of claim 1 .

10. An optimal selective photothermolysis device, comprising: a tunable radiation source (1100) configured to emit pulsed radiation, or pulsed or modulated temperature-sensing radiation, or both, and to excite photoacoustic waves from one or more targets within the tissue (1000); a patient interface (1200) including a radiation delivery unit (1210) configured to deliver the radiation to the tissue (1000) and an ultrasound detector (1220) for detecting the photoacoustic waves excited by the radiation from one or more targets within the tissue (1000); a control system (1110) configured to acquire characteristics of the tissue (1000) including a relative logarithmic function of the temperature dependence of the Gruneisen parameters, a skin cooling parameter, and a thermal relaxation time and a dynamic temperature of the one or more targets based on measurements of the detected photoacoustic waves, determine optimal characteristics of the tunable radiation source (1100) for an optimal surgical outcome for the one or more targets in the tissue (1000) including a central wavelength, a pulse width or pulse duration, and a pulse energy, define the characteristics and the optimal characteristics, and adjust the tunable radiation source (1100) towards an optimal surgical outcome based on the optimal characteristics. An optimal selective photothermolysis device characterized by:

11. The optimal selective photothermolysis device of claim 10, characterized in that the control system of the optimal selective photothermolysis device determines that the temperature of the surgical target has reached 100°C when the recorded photoacoustic signal amplitude shows a sudden increase due to laser-induced cavitation at 100°C.

Images