Patent Description:
SP as described by Anderson and Parrish in a paper published by SCIENCE in <NUM>, utilizes short laser pulses to precisely control collateral thermal or mechanical damages around light-absorptive lesions without the need of aiming a laser micro-beam at surgical targets. SP could target not only nature chromophores within vasculature, skin, retina and other human tissues but also labeled single cells and their ultra-structures if both a tunable laser and cell-specific dye delivery system are available. A SP laser surgery has two distinct features, a large surgical area and a short surgical laser pulse that deposits most of the laser pulse energy into surgical targets. Thus non-surgical targets within a large surgical area remain healthy after SP while all surgical targets are damaged. In contrast, laser surgeries of non-SP category include laser surgeries that use continuous wave (CW) laser to photocoagulate surgical targets without limiting damaging area and laser surgeries that use a small, high energy laser beam to non-selectively evaporate or sublime all illuminated tissue or tissue at the laser beam focus. Typical non-SP laser surgery examples include photothermal cancer therapy with CW lasers, laser-assisted in-situ keratomileusis eye surgery, and femto-second laser-assisted cataract surgery. Typical SP laser surgery examples include laser treatment of vascular malformation, some laser retinal photocoagulation surgeries, and some aesthetical laser surgeries such as laser tattoo removal.

An example of non-SP device: The US patent publication No. <CIT> discloses a photoacoustic probe for port wine stain (PWS), burn and melanin depth measurements is comprised of optical fibers for laser light delivery and a piezoelectric element for acoustic detection. The probe induced and measured photoacoustic waves in acryl amide tissue phantoms and PWS skin in vivo. Acoustic waves were denoised using spline wavelet transforms, then deconvolved with the impulse response of the probe to yield initial subsurface pressure distributions in phantoms and skin. The waves were then analyzed for epidermal melanin concentration, using a photoacoustic melanin index (PAMI) related to the amount of laser energy absorbed by melanin. Propagation time of the photoacoustic wave was used to determine the depth of blood perfusion underlying necrotic, burned tissue. Thus, the photoacoustic probe can be used for determining PWS, burn and melanin depth for most patients receiving laser therapy.

Another example of non-SP device: The European patent application No. <CIT> teaches a photoablator (<NUM>) for perforating a nail, having a laser source (<NUM>), optics (<NUM>) for focusing a laser beam emitted by the laser source (<NUM>) in a target plane, and beam positioning means (<NUM>) for positioning the laser beam in relation to the nail. The photoablator (<NUM>) further comprises a treatment area definition unit (<NUM>) for automatically defining an area of the nail to perforate by means of the laser beam. By means of the treatment area definition unit (<NUM>) the nail of a target finger or toe of a patient can be perforated respecting the specific conditions of the nail. Taking into account these conditions, the treatment area including a suitable treatment microhole pattern can be quickly defined by the treatment area definition unit (<NUM>), wherein according computing means can be used for this purpose.

<CIT> discloses a laser treatment device with an ultrasonic detector.

Technical challenges associated with a SP laser surgery include maximizing laser energy deposition ratios of surgical targets to nature chromophores, confining laser energy deposition into surgical targets, and optimizing laser pulse energy. Non-optimized SP laser parameters are associated with unsatisfactory laser surgical outcomes. Although the compromised SP laser surgical outcomes are well-known for decades, no good solution exists in the prior art of SP.

This disclosure relates to surgery techniques, apparatus and methods for optimizing selective photothermolysis (SP) surgeries. Some techniques can be applied to general surgical systems that heat up lesions in tissue, including high-intensity-focused-ultrasound therapies. It is noted that a tunable light source in this document broadly means a light source with tuning capabilities in its central wavelength, or light pulse width, or light pulse energy, or a combination of them.

In one aspect, a selective photothermolysis device is provided and includes a tunable radiation source configured to emit radiation; a patient interface; and a control system. The patient interface includes: a radiation delivery unit configured to deliver said radiation to a tissue; and an ultrasonic detector for detecting photoacoustic waves excited by the radiation from one or more surgical targets in the tissue. The control system is configured to: acquire characteristics of the tissue and the one or more surgical targets based on measurements on detected photoacoustic waves; determine optimal characteristics of the tunable radiation source for optimal surgical outcome on the one or more surgical targets in the tissue; prescribe the characteristics and the optimal characteristics, and adjust the tunable radiation source for optimal surgical outcome based on the optimal characteristics.

In another aspect, a method is provided an includes: maneuvering a patient interface to be configured against a tissue, and allowing an ultrasonic detector within the patient interface to be configured in acoustic contact with the tissue; acquiring characteristics of the tissue and one or more surgical targets and determining optimal characteristics of a tunable radiation source for optimal surgical outcome on the one or more surgical targets in the tissue based on measurements on photoacoustic waves detected by the ultrasonic detector; and prescribing the characteristics and the optimal characteristics or adjusting the tunable radiation source based on the optimal characteristics for optimal surgical outcome.

SP surgery systems have been operated in clinics with unsatisfactory surgical outcome for decades. Such SP surgeries include laser treatment of vascular malformation and pigmented lesions in human bodies. The innovative SP surgical systems and methods disclosed overcome the limitations of existing SP surgical systems and allow personalized, optimal treatment for the first time.

It is important to understand laser-tissue interaction mechanisms of selective photothermolysis (SP) laser surgery before addressing its clinical problems. Both thermal and mechanical damages could be utilized in SP. Initially, no measurable effects could be caused when tissue temperature is elevated to <NUM>°-<NUM> by SP laser pulses. Tissue is in hyperthermia status when temperature keeps rising to <NUM>°-<NUM>. A large portion of tissue might undergo necrosis if the hyperthermia lasts for several minutes. Enzyme activity reduction and cell immobility start from <NUM>. Denaturation of proteins and collagen occurs at <NUM> and leads to coagulation of tissue and necrosis of cells. Cell membrane permeability will significantly increase at <NUM>. Water molecules will be vaporized at <NUM>. It may lead to cavitation and tissue mechanical rupture by acoustic shock-waves associated with the laser-induced cavitation. Another type of mechanical damage could be caused by the strong photoacoustic waves generated by light absorbers upon the short laser pulse excitations. Major SP commercial applications include laser tattoo removal, laser treatment of vascular malformation and laser retinal photocoagulation.

Laser tattoo removal is usually performed with very short laser pulses in nanosecond or even picosecond regime. However, the proposed mechanisms behind laser tattoo removal have their physical, chemical and biological origins. Pigmented particles of tattoo will experience rapid temperature rise and volume expansion upon the energy deposition or excitation by a short laser pulse. However, most of the temperature rise and the volume expansion will be lost after a short period of time, determined by thermal relaxation time (time taken for <NUM>% of heat energy to be dissipated away) of these particles. Photoacoustic waves are generated along with the volume changes of these particles. Laser energy is transformed into both thermal energy and mechanical energy carried by the photoacoustic waves. In many cases, large laser pulse energy absorbed by pigmented particles may cause optical breakdown, plasma generation, chemical reactions between plasma and pigmented particles, cavitation and generation of acoustic shock-waves. These pigmented particles might be pyrolytically altered or shattered into smaller particles by the photoacoustic waves and acoustic shock-waves. Hosting cell necrosis and surround tissue damage might be induced thermally and mechanically during this process. In the end, the wound healing process might remove partial pigmented particles through rephagocytosis and alter the dermal scattering coefficients of the affected tissue, which might make the deeper pigmented particles less visible.

For laser tattoo removal application, the color of a tattoo depends on many factors including its optical absorption spectrum, optical scattering and absorption coefficients of the tissue above and below the pigments, the depth of the pigment and anatomical location of the pigments. It was reported that tattoos with significantly different optical absorption spectra could present themselves with the same color to naked eyes. Obviously, current practice of selecting surgical laser wavelength based on the color of a tattoo is not justified. On the other hand, there are only a handful of laser wavelengths (<NUM> ruby laser, <NUM> Alexandrite, <NUM> Nd: YAG and <NUM> second harmonic Nd: YAG) available in the market for laser tattoo removal. Even if the absorption spectra of the pigments of the tattoo is happened to be known, there are significant chances that it is not matching with any existing laser in the market. Anderson and Parrish envisioned a tunable laser for SP in <NUM>. However, such a laser is not available yet because nobody knows what wavelength should be adjusted to. Additionally, the selection of treatment laser pulse energy is also determined by the clinician's experience. Both clinical problems of laser tattoo removal are addressed by this invention.

Laser treatment of vascular malformation starts from the argon laser (<NUM> and <NUM>) treatment of port-wine stain (PWS) in <NUM>. The blue-green light of argon lasers is preferentially absorbed by hemoglobin within the PWS blood vessels. The deposited laser pulse energy into the vessels is largely converted to heat, causing thrombosis and destruction of the PWS blood vessels. The first generation argon laser had relatively long pulse duration (-<NUM>), which caused non specific tissue thermal damage of epidermis tissue. Thus, scarring was a frequent complication of the first generation argon laser treatment of PWS. Selective photothermolysis of PWS blood vessels was achieved by the first generation pulsed dye laser (PDL) (<NUM> or <NUM>, <NUM> milliseconds) that selectively photocoagulated PWS blood vessels and spared overlying epidermal tissue with a low incidence of side effects. As PDL laser energy is deposited in the intraluminal blood due to selective absorption of hemoglobin, the heat diffuses to the vessel wall and causes vascular wall necrosis and subsequent extravasation of red blood cells into the adjacent dermis. Dermal collagen fills the space of photocoagulated PWS vessels via wound healing process. The removal of photocoagulated PWS lesions leads to the blanching of PWS. The second generation PDL technique adopts larger spot sizes, higher energy densities, variable pulse durations, and dynamic cooling for more effective treatment of PWS. Currently, the second generation PDL with dynamic pulse duration and dynamic epidermal cooling by liquid cryogen sprays is the treatment of choice for PWSs. However, the laser has to be operated by experienced clinicians who adjust laser pulse width and laser pulse energy based on their experiences. In fact, the average success rate for full clearance is below <NUM>%. The selection of pulsed laser parameters is the most challenging clinical problem in laser treatment of vascular malformation.

The above mentioned clinical problems in laser tattoo removal and laser treatment of vascular malformation are obviously related to the distributions of light absorbers inside of tissue. Some experienced clinician takes advantage of the sounds generated during laser and tissue interaction to help laser tattoo removal surgery. However, human only hears sound wave between <NUM> - <NUM>,<NUM>. For laser tattoo removal, laser tissue interaction does generate highfrequency ultrasonic waves. Most of their frequency components are far beyond <NUM>,<NUM>. In other words, most of useful information are completely ignored. By adding an ultrasonic detector to "hear" the responses from tissue under a SP laser surgery, this invention is able to address the above- mentioned clinical problems. The science behind the photoacoustic waves during laser-tissue interactions is photo acoustics, the key technique for this invention.

Photoacoustic techniques originate from Alexander Graham Bell who discovered photoacoustic effect in <NUM>. The generation of photoacoustic wave consists of the following stages including conversion of the absorbed pulsed or modulated radiation into heat energy, temporal change of temperature that rises as laser pulse energy is absorbed and falls when laser pulse ends and the heat dissipates, and volume expansion and contraction following these temperature changes, which generate pressure changes (i.e. photoacoustic wave). Hordvik et al. reported photoacoustic technique for determining optical absorption coefficients in solids in <NUM>. Photoacoustic spectroscopy was applied to a wide variety of conventional spectroscopic measurements as reviewed by West et al. More recent developments of photoacoustic techniques were motivated by biomedical imaging applications. Major photoacoustic technique developments in the biomedical imaging field include the inventions of acoustic-resolution & optical-resolution photoacoustic microscopies by Maslov et al. and the Fabry-Perot photoacoustic sensor based photoacoustic tomography by Zhang et al. The significantly improved image performance (sensitivity, resolution, depth and speed) of the above photoacoustic imaging systems and improvements of acoustic transducer arrays by industry for various photoacoustic tomography configurations generate high impacts in biology and medicine.

However, the penetration of photoacoustic techniques into SP laser surgery is very limited. Nobody tried to build a wavelength tunable laser and apply such a laser for tattoo removal with photoacoustics. Selecting laser wavelength in laser treatment of vascular malformation might be less critical than in laser tattoo removal. But it is still a very challenging task to optimize other parameters of laser surgical systems for laser treatment of vascular malformation. In fact, a SP laser surgery does not necessary get rid of surgical targets or change the spatial location of surgical targets right after a SP laser surgery. It requires a long wound healing process to remove damaged tissues through rephagocytosis. Photoacoustic imaging of lesions before and right after SP laser surgery almost presents no changes in lesion images. Thus, a simple photoacoustic imaging of lesions has 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 for optimizing laser treatment of port-wine stain. In order to optimize laser treatment of port-wine stain in children, Rao et al. proposed to image the port- wine stain vessel size and depth in child patients with optical-resolution photoacoustic microscopy, construct physical model of port-wine stain lesions with lesion information, and derive optimal laser treatment parameters (pulse width and pulse energy) with massive computer simulations. Other imaging modalities such as optical Doppler tomography and optical coherence angiography relied on blood flow or blood flow induced optical speckles to acquire information of port-wine stain lesions. However, the lack of blood flow right after laser surgery could not confirm full photocoagulation of lesion vessels. It was hypothesized that partially coagulated lesion vessels could remain refractory after laser treatment. Another limitation of these optical imaging modalities is their shallow imaging depth of <NUM>-<NUM>. In contrast, this invention takes simple experimental approaches to address the SP clinical problems.

The disclosed techniques, methods and apparatus of this invention are based on the physical principle of photoacoustic effect and its temperature-dependence. In early literature of photoacoustic techniques, the temperature-dependent photoacoustic effect was utilized in a range of temperature related measurements including measuring flame temperature and measuring solid thermal diffusivity. Esenaliev et al. reported real-time optoacoustic monitoring of temperature in ex vivo canine tissues in <NUM>. Larin et al. reported optoacoustic laser monitoring of cooling and freezing of ex vivo canine liver in <NUM>. Shah et al. reported photoacoustic temperature monitoring of ex vivo porcine tissue in <NUM>. Oraevsky et al. described optoacoustic imaging methods for medical diagnosis and real time optoacoustic monitoring of change in tissue properties, and an improved temperature calibration method in <CIT>, <CIT>, and <CIT>. In a continuous-wave laser thermal therapy described by Oraevsky et al. , tissue temperature varies very slowly. The continuous-wave laser has no negative effects on an asynchronous photoacoustic temperature-sensing process. However, short surgical light pulses of a SP surgery system heat up a surgical target within its short pulse duration and the surgical target cools down quickly. Accurately measuring a dynamic temperature rise due to energy deposition of a short surgical laser short laser pulse within a live tissue is challenging. It has not been demonstrated in scientific literature yet. Technically, it requires the temperature-sensing light pulse to be synchronized to the surgical light pulse with an exact short time delay. Additionally, the strong surgical light pulses generate strong photoacoustic signals upon absorption by light absorbers in 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 methods described by Oraevsky et al. and others in prior art are valid for quasi-continuous wave surgical pulses, but invalid for dynamic temperature rise created in SP surgeries with short surgical light pulses.

In summary, the prior art is deficient in methods to address clinical problems in SP surgeries such as laser tattoo removal and laser treatment of vascular malformation. The surgery techniques, apparatus and methods are disclosed below to fill the gaps between the science theory of SP and clinical practices.

As an example, <FIG> shows an example of a revolutionary, pulsed light SP surgical system wherein the inclusion of an acoustic detector differentiates it from a conventional SP light surgical system in prior arts. This SP pulsed light surgical system comprises a tunable pulsed light source <NUM> to produce light beams <NUM> under the control of its control system <NUM>; and a patient interface <NUM> operable to be in contact with a tissue <NUM>. In one implementation of <FIG>, the light beams <NUM> comprise only a surgical pulsed light beam. The pulse width of the pulsed light beam is less than <NUM> seconds, or less than <NUM> seconds, or less than <NUM> seconds in order to efficiently excite photoacoustic signals. The patient interface <NUM> comprises a light delivery unit <NUM>, an acoustic detector <NUM>, and an interface medium <NUM>. The light delivery unit <NUM> shapes the light beam profile, delivers the light beam with an articulated arm, adjusts the light beam diameter and transmits the light beam through the interface medium <NUM> to a tissue <NUM> surface. The interface medium <NUM> is preferable to be transparent to light beams and conductive to ultrasound. The pulsed light beam <NUM> excites photoacoustic waves <NUM> that propagate through the interface medium <NUM> and are detected by the acoustic detector <NUM>. The detected photoacoustic signals <NUM> are digitized, analyzed in the control system <NUM> for the control of the tunable pulsed light beam <NUM>. It is noted that a tunable light source in this document broadly means a light source with tuning capabilities in its central wavelength, or light pulse width, or light pulse energy or a combination of them. Tunable light source itself is not difficult to make. However, the missing part is how the central wavelength and other surgical light pulse parameters should be tuned according to surgical targets in tissue. The inclusion of an acoustic detector is exactly the missing part in prior arts of SP surgeries. The inclusion of an acoustic detector makes sense to the utilization of a tunable light source in SP surgeries for the first time. Both the tunable light source and the acoustic detector make an optimized SP surgery possible, a goal that has been desired for decades.

In order to fully utilize the disclosed methods below for optimized SP surgical outcomes, it is desirable to utilize a more advanced tunable light source <NUM>, which can produce a surgical light pulse or a temperature-sensing light beam or both under the control of its control system <NUM>. For most of implementations of <FIG>, the temperature-sensing light beam comprises temperature- sensing light pulses that have a tunable time delay relative to surgical light pulses. However, the temperature-sensing light beam could be an intensity modulated light beam to excite photoacoustic waves when depth-resolved tissue information is not required for an application, or a chirped intensity-modulated light beam to allow a very low-resolution depth discrimination. The pulse width of the temperature-sensing light pulses is less than <NUM> seconds, or less than <NUM> seconds, or less than <NUM> seconds in order to efficiently excite photoacoustic signals. The central wavelength and pulse width of the temperature-sensing light beam could be fixed for some implementations. The temperature-sensing light pulse energy is of subtherapeutic level. For the tunable light source <NUM>, its tuning capabilities should match with the needs of a specific SP surgery application. The most practical implementation of such a more advanced tunable light source is to integrate a tunable surgical light source unit and a tunable temperature-sensing light source unit into a single package with a shared power supply subsystem, a shared cooling subsystem and a shared control system <NUM>. The light beams of the light source units need to be combined, and sent out from the same light 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.

Yet another implementation of <FIG> could be a surgical planning system that provides optimized surgical laser parameters for other conventional pulsed light SP surgical systems. Such a surgical planning system comprises a tunable pulsed light source that can produce a subtherapeutic 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 operable to be in contact with the target tissue. The patient interface comprises a light delivery unit, an acoustic detector, and an interface medium. The light delivery unit shapes beam profiles of light beams, delivers light beams with an articulated arm, adjusts diameters of light beams, and transmits light beams through the interface medium to a tissue surface. The light beams excite photoacoustic waves that propagate through the interface medium and are detected by the acoustic detector. The detected photoacoustic signals are digitized, analyzed by the control system for determining optimal surgical light pulse parameters to be used by another conventional pulsed light SP surgical system. The advantage of such a surgical planning system is that its laser pulse repetition rate could be much higher than a surgical system and the time for acquiring optimized surgical laser parameters is much shorter.

The key of this invention is the inclusion of an ultrasonic detector in a SP surgery system. A pulsed light SP surgical system of <FIG> could be generalized as a pulsed radiation SP surgical system by replacing the surgical light pulse with any form of radiation (radio waves, microwaves, infrared light, visible light, Ultraviolet, X-rays, Gamma rays) pulse that heats up lesions or extraneous contrast agents attached to lesions in tissue, replacing the temperature-sensing light pulse with any form of pulsed or modulated radiation (radio waves, microwaves, infrared light, visible light, Ultraviolet, X-rays, Gamma rays) that can be absorbed by lesions in tissue or extraneous contrast agents attached to lesions in tissue, and effectively excite photoacoustic waves, and replacing the light delivery unit in the patient interface with a radiation beam delivery unit for delivering a radiation beams to tissue. All techniques and methods for <FIG> apply to the generalized radiation SP surgical system. The examples presented below are mostly based on pulsed laser surgical systems because laser tattoo removal and laser treatment of vascular malformation are major concerns of this invention.

<FIG> show examples of patient interfaces with different configurations to facilitate optimized SP laser surgery. <FIG> shows a schematic example of a patient interface comprising a single element ultrasonic transducer <NUM>, and an interface medium <NUM> in acoustic contact with tissue <NUM>. Because a light delivery unit that should be shown in the patient interface is no different from that of a conventional laser SP system, the light delivery unit is skipped in <FIG> for simplicity. Laser beams <NUM> (a surgical laser beam, or a temperature-sensing laser beam, or both) can be selectively delivered to the tissue <NUM> surface according to the requirements of the methods. In some implementations, laser beams <NUM> could only comprise a surgical laser beam. The single element ultrasonic transducer <NUM> is positioned to detect photoacoustic waves without blocking laser beams <NUM>. The transducer <NUM> could be a traditional ultrasonic transducer or one based on optical detection techniques. The interface medium <NUM> allows the transmission of laser beams and photoacoustic waves. The interface medium <NUM> could be saved in a more simplified configuration where the single element ultrasonic transducer is in direct acoustic contact with tissue at a tissue surface area immediately next to a tissue surface area illuminated by the laser beams. One advantage of this simplified configuration is that it may simultaneously allow the delivery of skin cooling agent through free space. Spectroscopic photoacoustic signals are acquired by the single element ultrasonic transducer <NUM> from a one-dimensional, depth-resolved space in the tissue, and are digitized and analyzed by the control system <NUM> of <FIG>.

<FIG> shows another schematic example of a patient interface comprising a linear arrayed ultrasonic transducer <NUM> and an interface medium <NUM> in acoustic contact with tissue <NUM>. Because a light delivery unit that should be shown in the patient interface is no different from that of a conventional laser SP equipment, the light delivery unit is skipped in <FIG> for simplicity. Laser beams <NUM> (a surgical laser beam, or a temperature-sensing laser beam, or both) can be selectively delivered to the tissue <NUM> surface according to the requirements of the methods. The linear arrayed ultrasonic transducer <NUM> is positioned to detect photoacoustic waves without blocking laser beams <NUM>. The linear arrayed transducer <NUM> could be a traditional ultrasonic transducer or one based on optical detection techniques. The interface medium <NUM> allows the transmission of laser beams and photoacoustic waves with minimum loss. The interface medium <NUM> could also be saved in a more simplified configuration where the linear arrayed ultrasonic transducer is in direct acoustic contact with tissue at a tissue surface area immediately next to a tissue surface area illuminated by the laser beams delivered through free space. One advantage of this simplified configuration is that it may simultaneously allow the delivery of skin cooling agent through free space. Spectroscopic photoacoustic signals are acquired by the linear arrayed ultrasonic transducer <NUM> from a two-dimensional, depth-resolved space in the tissue, and are digitized and analyzed by the control system <NUM> shown in <FIG>.

<FIG> shows another schematic example of a patient interface comprising a linear arrayed ultrasonic transducer <NUM>, an interface medium <NUM> in acoustic contact with tissue <NUM>, and a rotational stage <NUM> that mounts the linear arrayed ultrasonic transducer <NUM> and the interface medium <NUM> and rotates around the central axis of the illuminated oval area on tissue surface for the acquisition of a three-dimensional tissue information. The arrows <NUM> show the rotational direction of the rotation stage <NUM>. Because a light delivery unit that should be shown in the patient interface is no different from that of a conventional laser SP system, the light delivery unit is skipped in <FIG> for simplicity. Laser beams <NUM> (a surgical laser beam, or a temperature sensing laser beam, or both) can be selectively delivered to the tissue <NUM> surface according to the requirements of the methods. Laser beams <NUM> are delivered with a flexible multimode fiber or a flexible fiber bundle and accessory optics (not shown in <FIG>) to allow rotation. The linear arrayed ultrasonic transducer <NUM> is positioned to detect photoacoustic waves without blocking laser beams <NUM>. The linear arrayed transducer <NUM> could be a traditional ultrasonic transducer or one based on optical detection techniques. The interface medium <NUM> allows the transmission of laser beams and photoacoustic waves with minimum loss. The interface medium <NUM> could also be saved in a more simplified configuration where the linear arrayed ultrasonic transducer is in direct acoustic contact with tissue at a tissue surface immediately next to a tissue surface area illuminated by the laser beams delivered through free space. One advantage of this simplified configuration is that it may simultaneously allow the delivery of skin cooling agent through free space. Spectroscopic photoacoustic signals are acquired by the linear arrayed ultrasonic transducer <NUM> from a three-dimensional, depth-resolved space in the tissue, and are digitized and analyzed by the control system <NUM> shown in <FIG>.

<FIG> shows another schematic example of a patient interface comprising a single element ultrasonic transducer <NUM>, an acoustic wave reflector <NUM>, and an interface media <NUM> in acoustic contact with tissue <NUM>. Because a light delivery unit that should be shown in the patient interface is no different from that of a conventional laser SP system, the light delivery unit is skipped in <FIG> for simplicity. Laser beams <NUM> (a surgical laser beam, or a temperature- sensing laser beam, or both) can be selectively delivered to the tissue <NUM> surface according to the requirements of the methods. The single element ultrasonic transducer <NUM> is positioned to detect photoacoustic waves reflected by the acoustic reflector <NUM> without blocking laser beams <NUM>. The usage of an acoustic reflector <NUM> allows the laser beams and the transducer on the same side of tissue without blocking each other. The single element ultrasonic transducer <NUM> could be a traditional ultrasonic transducer or one based on optical detection techniques. The interface medium <NUM> allows the transmission of laser beams and photoacoustic waves with minimum loss. Spectroscopic photoacoustic signals are acquired by the single element ultrasonic transducer <NUM> from a one-dimensional, depth-resolved space in the tissue, and are digitized and analyzed by the control system <NUM> shown in <FIG>.

<FIG> shows another schematic example of a patient interface comprising a linear arrayed ultrasonic transducer <NUM>, an acoustic reflector <NUM>, and an interface media <NUM> in acoustic contact with tissue <NUM>. Because a light delivery unit that should be shown in the patient interface is no different from that of a conventional laser SP system, the light delivery unit is skipped in <FIG> for simplicity. Laser beams <NUM> (a surgical laser beam, or a temperature sensing laser beam, or both) can be selectively delivered to the tissue <NUM> surface according to the requirements of the methods. The linear arrayed ultrasonic transducer <NUM> is positioned to detect photoacoustic waves reflected by the acoustic reflector <NUM>. The usage of an acoustic reflector <NUM> allows the laser beams and the transducer on the same side of tissue without blocking each other. The linear arrayed ultrasonic transducer <NUM> could be a traditional ultrasonic transducer or one based on optical detection techniques. The interface medium <NUM> allows the transmission of laser beams and photoacoustic waves with minimum loss. Spectroscopic photoacoustic signals are acquired by the linear arrayed ultrasonic transducer <NUM> from a two- dimensional, depth-resolved space in the tissue, and are digitized and analyzed by the control system <NUM> shown in <FIG>.

<FIG> shows another schematic example of a patient interface comprising a linear arrayed ultrasonic transducer <NUM>, an acoustic reflector <NUM>, an interface media <NUM> in acoustic contact with tissue <NUM>, and a rotational stage <NUM> that mounts the linear arrayed ultrasonic transducer <NUM>, the acoustic reflector <NUM> and the interface medium <NUM>, and rotates around the axis of the laser beams. The arrows <NUM> show the rotational direction of the rotation stage <NUM>. Because a light delivery unit that should be shown in the patient interface is no different from that of a conventional laser SP system, the light delivery unit is skipped in <FIG> for simplicity. Laser beams <NUM> (a surgical laser beam, or a temperature- sensing laser beam, or both) can be selectively delivered to the tissue <NUM> surface according to the requirements of the methods. The linear arrayed ultrasonic transducer <NUM> is positioned to detect photoacoustic waves reflected by the acoustic reflector <NUM>. The usage of acoustic reflector <NUM> allows the laser beams and the transducer on the same side of tissue without blocking each other. The linear arrayed ultrasonic transducer <NUM> could be a traditional ultrasonic transducer or one based on optical detection techniques. The interface medium <NUM> allows the transmission of laser beams and photoacoustic waves with minimum loss. Spectroscopic photoacoustic signals are acquired by the linear arrayed ultrasonic transducer <NUM> from a three-dimensional, depth-resolved space in the tissue, and are digitized and analyzed by the control system <NUM> shown in <FIG>.

<FIG> shows another schematic example of a patient interface comprising a Dammann grating <NUM>, a <NUM>-D arrayed ultrasonic transducer <NUM>, and an interface medium <NUM> in acoustic contact with tissue <NUM>. Because a light delivery unit that should be shown in the patient interface is no different from that of a conventional laser SP system, the light delivery unit is skipped in <FIG> for simplicity. Laser beams <NUM> (a surgical laser beam, or a temperature-sensing laser beam, or both) can be selectively delivered to the tissue <NUM> surface according to the requirements of the methods. The laser beams are transformed into <NUM>-D arrayed laser beams <NUM> by a Dammann grating <NUM>. <NUM>-D arrayed laser beams pass through the free space not being taken by the <NUM>-D arrayed ultrasonic transducer <NUM> and the interface medium <NUM> before reaching the tissue <NUM>. The excited photoacoustic waves are detected by the <NUM>-D arrayed ultrasonic transducer <NUM>. The <NUM>-D arrayed ultrasonic transducer <NUM> could be a traditional ultrasonic transducer or one based on optical detection techniques. The interface medium <NUM> allows the transmission of laser beams and photoacoustic waves with minimum loss. Spectroscopic photoacoustic signals <NUM> are acquired by the <NUM>-D arrayed ultrasonic transducer <NUM> from a three-dimensional, depth-resolved space in the tissue, and are digitized and analyzed by the control system <NUM> shown in <FIG>.

<FIG> shows another schematic example of a patient interface comprising a dichroic mirror <NUM>, a scanning lens <NUM>, a <NUM>-D Galvo scanner <NUM>, and a Fabry-Perot sensor <NUM> in acoustic contact with a tissue <NUM>. Because a light delivery unit that should be shown in the patient interface is no different from that of a conventional laser SP system, the light delivery unit is skipped in <FIG> for simplicity. Laser beams <NUM> (a surgical laser beam, or a temperature sensing laser beam, or both) can be selectively delivered to the tissue <NUM> surface according to the requirements of the methods. Both the dichroic mirror <NUM> and the Fabry-Perot sensor <NUM> are transparent for the laser beams <NUM> (a surgical laser beam, or a temperature-sensing laser beam, or both). An ultrasonic- wave-detection laser beam <NUM> is scanned by the <NUM>-D Galvo scanner <NUM> and the scanning lens <NUM>, and reflected by the dichroic mirror <NUM> to the Fabry- Perot sensor <NUM> for the detection of photoacoustic waves. The Fabry-Perot sensor <NUM> comprises of two layers of dielectric mirrors and an acoustic-wave-sensing layer between two dielectric mirrors. The ultrasonic-wave-detection laser beam <NUM> could be in the form of <NUM>-D arrayed laser beams or <NUM>-D arrayed laser beams in other implementations. An optical system that sends out the ultrasonic-wave-detection laser beam <NUM>, detects the reflected ultrasonic-wave-detection laser beam <NUM> modulated by photoacoustic waves is skipped from <FIG> for simplicity. Spectroscopic photoacoustic signals could be acquired from a one-dimensional or a two- dimensional or a three-dimensional, depth-resolved space in the tissue, and be digitized and analyzed by the control system <NUM> shown in <FIG>.

<FIG> shows another schematic example of a patient interface comprising a vacuum structure <NUM>, a circular arrayed ultrasonic transducer <NUM>, and an interface medium <NUM> in acoustic contact with a tissue <NUM>. Because a light delivery unit that should be shown in the patient interface is no different from that of a conventional laser SP system, the light delivery unit is skipped in <FIG> for simplicity. Laser beams <NUM> (a surgical laser beam, or a temperature sensing laser beam, or both) can be selectively delivered to the tissue <NUM> surface according to the requirements of the methods. The circular arrayed transducer <NUM> could be a traditional ultrasonic transducer or one based on optical detection techniques. The circular arrayed ultrasonic transducer <NUM> is in acoustic contact with the wall of the vacuum structure <NUM>. The vacuum structure <NUM> is designed to suck part of the tissue into it in a way similar to a body cupping device. Interface medium <NUM> such as water could be injected into the bottom of the vacuum structure to immerse the tissue sucked into the vacuum structure. The laser beams <NUM> illuminate the tissue inside of the vacuum structure and excite photoacoustic waves. The circular arrayed ultrasonic transducer <NUM> could acquire spectroscopic photoacoustic signals from a two- dimensional, depth-resolved space in the tissue. The circular arrayed ultrasonic transducer <NUM> could be adjusted in elevational direction for acquiring 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 <NUM> shown in <FIG>.

<FIG> shows an example of wavelength tuning operation for optimizing SP laser treatment of unknown pigments. The optimal surgical laser wavelength should maximize the laser energy deposition ratios of unknown pigments to nature chromophores. The following procedure is designed for a laser SP surgery system whose surgical laser pulse width is short enough to effectively excite photoacoustic signals. However, if the surgical laser pulse is too long to effectively excite photoacoustic signals, a short temperature-sensing laser pulse generated by a more advanced dual-pulse (a surgical laser pulse followed by a delayed temperature-sensing laser pulse) laser system should be used to excite photoacoustic signals in the following procedure. First, a series of wavelength points that comprises characteristic peaks and valleys of oxygenated hemoglobin and deoxygenated hemoglobin is determined; Second, a tissue area is selected and the patient interface is operated to be in acoustic contact with the selected tissue area; Third, the acoustic detector is preferably configured in the mode of acquiring a <NUM>-D, depth-resolved photoacoustic tissue information with a single surgical laser pulse. It is noted that the acoustic detector could be configured to acquire <NUM>-D, depth-resolved tissue information in a simplified implementation and the following steps might need slight modifications; Fourth, the control system sends out multiple subtherapeutic surgical laser pulses for each wavelength point, and acquires photoacoustic signals detected by the acoustic detector; Fifth, an averaged <NUM>-D, depth- resolved tissue information is acquired after tomography reconstruction for each wavelength point; Sixth, unknown pigments are identified along with nature chromophores after their relative extinction coefficients are calculated and their relative extinction coefficient curves are fitted. If no nature chromophores are presented in the <NUM>-D, depth-resolved tissue space, known extinction coefficient curves of nature chromophores from literature could be used; Seventh, relative energy deposition ratio curves of unknown pigments to nature chromophores are calculated; Finally, the optimal surgical laser wavelengths are determined for different types of unknown pigments with an algorithm that puts different priority weights on different nature chromophores. Multiple laser treatments with optimized lasing wavelengths for different types of unknown pigments might be performed in series for the optimal laser treatment outcome. The same technique in <FIG> 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 using a wavelength-tunable, pulsed or modulated radiation beam to effectively excite photoacoustic waves from lesions in tissue or extraneous contrast agents attached to lesions in tissue.

For the more advanced dual-pulse (a surgical laser pulse followed by a delayed temperature-sensing laser pulse) laser SP surgical system, this invention provides a method for photoacoustic temperature sensing in live tissue including a non-invasive Gr<IMG>neisen parameter calibration procedure. This method overcome limitations of methods in prior art. This method detailed in <FIG> and <FIG> can non-invasively measure dynamic temperature of a surgical target in a tissue heated by a short surgical laser pulse after calibration. The calibration procedure is based on a hypothesis that the heating of a surgical target by laser pulses is a linear process and the maximum temperature rise of the surgical target is proportional to the laser energy deposited into the surgical target when the temperature is measured immediately after the surgical laser pulse.

<FIG> shows an example of calibrating a temperature-dependent relative logarithm function of Gr<IMG>neisen parameter of tissue. First, we start calibration process by measuring the equilibrium temperature T0 of the tissue; Second, we measure the photoacoustic signal of a surgical target excited by a temperature-sensing laser pulse with a constant laser pulse energy, perform a logarithm operation on the photoacoustic signal amplitude, and get a baseline signal; Third, we send a surgical laser pulse for heating the surgical target and measure the photoacoustic signal excited by the surgical laser pulse; Fourth, we send both the surgical laser pulse and the temperature-sensing laser pulse, measure the excited photoacoustic signal by dual pulses, subtract the photoacoustic signal excited by the surgical laser pulse from that of the dual pulses, calculate the amplitude of the photoacoustic signal excited by the temperature-sensing laser pulse, separate temperature-dependent part from other temperature-independent parts with a logarithm operation, subtract the baseline signal, and calculate a relative logarithm function value, logΓ(T0 + δT)- logΓ(T0) where G denotes the Gr<IMG>neisen parameter of the tissue and δT denotes the temperature rise caused by the heating laser pulse, for the temperature point of T0 + δT; Fifth, if the recorded photoacoustic signal amplitude does not show an abrupt increase due to the laser-induced cavitation at <NUM>, we wait until the tissue temperature returns to its original equilibrium temperature. Then we adjust the heating laser pulse energy to its ki times and return to the third step for acquiring another relative logarithm function value of logΓ(T0 + kiδT)- logΓ(T0) for the temperature point of T0 + kiδT. We should keep the increase of the surgical laser energy small in order to have a more accurate measurement of <NUM>. If a laser-induced cavitation is observed, we continue to the final step. In the final step, we have <MAT>.

Thus, we can calculate the absolute temperature rises (k0δT, k1δT,. , kmδT) by each surgical laser pulse and fit the function of logΓ(T)- logΓ(T0) between T0 and <NUM> where T denotes temperature. The equilibrium temperature could be the body temperature of a patient. It could also be an equilibrium temperature of an ex vivo tissue in an environment of known temperature. In practice, laser pulse energy fluctuates from pulse to pulse. Compensation with simultaneous laser pulse energy monitoring is necessary for the procedures above. As long as both the starting temperature and the temperature to be measured are between T0 and <NUM>, the calibrated relative logarithm function Gr<IMG>neisen parameter of tissue is valid for a temperature sensing operation as detailed below.

<FIG> shows an example of dynamic temperature sensing operation of surgical targets in a tissue after heating 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 profile of a surgical target along time can be accurately measured by repeating the following dynamic temperature sensing procedure. It is important that we start from a known body temperature and we know the starting point in the relative logarithm function of Gr<IMG>neisen parameter. First, we measure body temperature before laser surgical intervention; Second, we send a temperature-sensing laser pulse, measure the amplitude of the excited photoacoustic signal at body temperature and calculate its logarithm as a baseline signal; Third, we send only a surgical laser pulse and measure the photoacoustic signal excited by the surgical laser pulse; Fourth, we send both the surgical laser pulse and the temperature-sensing laser pulse and measure the excited photoacoustic signal; Fifth, we calculate the amplitude of the photoacoustic signal excited by the temperature-sensing laser pulse and separate temperature-dependent part from other temperature- independent parts with a logarithm operation; Sixth, we calculate the relative logarithm function value by subtracting the baseline signal; Finally, we determine the dynamic temperature at the end of the surgical laser pulse from the calibrated relative logarithm function of Gr<IMG>neisen parameter of tissue. For temperature sensing of a surgical target heated by a continuous-wave laser, it requires only two measurements of the photoacoustic signals excited by the temperature-sensing laser pulse at the body temperature and at a time point during CW laser surgery. More accurate measurement result is expected due to the more accurate, non-invasive calibration method in <FIG>. The same technique in <FIG> applies to pulsed laser (coherent light source) surgical systems, other non-coherent pulsed light-source surgical systems, high-intensity-focused- ultrasound therapy systems, and other general radiation SP systems using a surgical pulsed radiation beam to heat up lesions in tissue or extraneous contrast agents attached to lesions in tissue, and a pulsed or modulated temperature-sensing radiation beam to effectively excite photoacoustic waves from lesions in tissue or extraneous contrast agents attached to lesions in tissue.

In applications such as laser treatment of vascular malformations, it is desirable to effectively heat a surgical target with laser pulses whose laser pulse width matches to thermal relaxation time of the surgical target. Most energy of laser pulse will be confined to the surgical target instead of being spread to surrounding healthy tissues. Computer simulation with tissue models and a surgical target's dimension information is the only available method to estimate thermal relaxation time of a surgical target in tissue in the research field of laser treatment of vascular malformation. However, <FIG> shows an example of optimizing surgical laser pulse width by in vivo measurement of thermal relaxation time of a surgical target in tissue. In addition to the surgical laser pulse, a temperature-sensing laser pulse is required to perform the task of measuring thermal relaxation time of a surgical target. First, we measure body temperature and the photoacoustic signal of a surgical target at body temperature with a temperature-sensing laser pulse, and calculate the logarithm of the photoacoustic signal as a baseline signal; Second, we send a subtherapeutic surgical laser pulse and measure excited photoacoustic signal; Third, we send both the subtherapeutic surgical laser pulse and the temperature-sensing laser pulse with a precise delay time, measure excited photoacoustic signal of dual pulses, and calculate the amplitude of the photoacoustic signal excited by the temperature-sensing laser pulse; Fourth, we separate temperature-dependent part from other temperature-independent parts with a logarithm operation and calculate the relative logarithm function value by subtracting the baseline signal; Fifth, we determine the temperature of the surgical target at the precise delay time; Sixth, we wait until surgical target temperature returns to body temperature, and determine whether there are enough delay time points to fit the curve of the surgical target's temperature versus delay time. If there are no enough delay time points, we change the value of the delay time and repeat the procedures between the third step and the sixth step before we can fit the curve of temperature versus delay time and derive the thermal relaxation time of the surgical target. The optimal surgical laser pulse width is determined to be equal to the measured thermal relaxation time of the surgical target. The same technique in <FIG> applies to pulsed laser (coherent light source) SP surgical systems, other non-coherent pulsed light-source SP surgical systems, and other general radiation SP systems using a surgical pulsed radiation beam to heat up lesions in tissue or extraneous contrast agents attached to lesions in tissue, and a pulsed or modulated temperature-sensing radiation beam to effectively excite photoacoustic waves from lesions in tissue or extraneous contrast agents attached to lesions in tissue.

In a conventional laser SP surgery, surgical laser pulse energy is selected according to a clinician's past experiences. However, the inclusion of an acoustic detector makes it possible to optimize surgical laser pulse energy objectively for the first time. For the laser treatment vascular malformations, the optimal laser pulse energy would heat 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 would heat a selected surgical target to <NUM> and cause laser-induced cavitation.

<FIG> shows an example of surgical laser pulse energy optimization in a tunable laser during a laser photocoagulation SP surgery. We assume the laser wavelength is already optimized before laser power optimization. First, we select a surgical target with a <NUM>-D, depth resolved tissue information excited by a temperature-sensing laser pulse and acquired by an acoustic detector; Second, we optimize laser pulse width after measuring the surgical target' s thermal relaxation time; Third, we send a surgical laser pulse and measure excited photoacoustic signal; Fourth, we send both the surgical laser pulse and the temperature-sensing laser pulse, measure the photoacoustic signal excited by dual pulses, and calculate the amplitude of the photoacoustic signal excited by the temperature-sensing laser pulse; Fifth, we separate the temperature-dependent part with a logarithm operation, calculate a relative logarithm function value by subtracting the baseline signal acquired in previous steps, and determine the temperature of the surgical target; Sixth, we determine whether the surgical target reaches a predetermined temperature for photocoagulation. If not, we wait for the temperature to recover its original body temperature, increase the laser pulse energy and return to the third step. If the temperature reaches to a predetermined temperature for photocoagulation, we end the operation with the optimized laser pulse energy for photocoagulation of the surgical target. The same technique in <FIG> applies to pulsed laser (coherent light source) SP surgical systems, other non-coherent pulsed light-source SP surgical systems, high-intensity- focused-ultrasound therapy systems for optimizing ultrasound beam energy, and other general radiation SP systems using a surgical pulsed radiation beam to heat up lesions in tissue or extraneous contrast agents attached to lesions in tissue, and a pulsed or modulated temperature sensing radiation beam to effectively excite photoacoustic waves from lesions in tissue or extraneous contrast agents attached to lesions in tissue.

A distinct feature of laser photodisruption is the generation of acoustic shock-waves due to a laser-induced cavitation. We assume there is no need to further increase laser pulse energy once the acoustic shock-wave due to laser-induced cavitation is observed. For laser photodisruption in laser tattoo removal application, it is desirable to use a tunable surgical laser instead of a more advanced tunable dual-pulse (a surgical laser pulse followed by a delayed temperature-sensing laser pulse) surgical laser. <FIG> shows an example of in vivo surgical laser pulse energy optimization during a laser photodisruption SP surgery. The following procedure is designed for a laser SP surgery system whose surgical laser pulse width is short enough to effectively excite photoacoustic signals. We assume the laser wavelength is already optimized before laser pulse energy optimization. First, we select a surgical target with a <NUM>-D, depth resolved tissue information excited by the surgical laser pulse and acquired by an acoustic detector; Second, we measure the surgical target's thermal relaxation time and optimize surgical laser pulse width according to the surgical target' s thermal relaxation time; Third, we send a surgical laser pulse and measure excited photoacoustic signal; Fourth, we draw one point for the curve of photoacoustic signal versus laser pulse energy and determine whether there is an abrupt photoacoustic signal increase due to a laser-induced cavitation. If there is no abrupt photoacoustic signal increase, we can wait for the surgical target to recover its original body temperature, increase surgical laser pulse energy, and return to the third step. If an abrupt photoacoustic signal increase due to laser- induced cavitation is observed, the optimal laser pulse energy is achieved. The same technique in <FIG> applies to pulsed laser (coherent light source) SP surgical systems, other non-coherent pulsed light-source SP surgical systems, and other general radiation SP systems using a surgical pulsed radiation beam to heat up lesions in tissue or extraneous contrast agents attached to lesions in tissue, and excite photoacoustic waves from lesions in tissue or extraneous contrast agents attached to lesions in tissue.

The methods in <FIG> and <FIG> provide real time control of surgical laser pulse energy during laser SP surgeries. However, it is possible to accurately determine the required surgical laser pulse energy for a surgical target to reach a predetermined temperature without actually performing laser surgery. <FIG> shows an example of determining optimal surgical laser pulse energy to achieve a predetermined temperature without performing laser surgery. We assume the laser wavelength is already optimized. First, we select a surgical target with a <NUM>-D, depth resolved tissue information excited by the temperature-sensing laser pulse and acquired by an acoustic detector; Second, we optimize laser pulse width after measuring the surgical target's thermal relaxation time; Third, we send a subtherapeutic surgical laser pulse and detect its photoacoustic signal; Fourth, we send both a subtherapeutic surgical laser pulse and a temperature-sensing laser pulse, which is immediately after the end of the surgical laser pulse; Fifth, we measure 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, we separate temperature-dependent part with a logarithm operation, calculate a relative logarithm function value by subtracting the baseline signal acquired in previous steps; Finally, we calculate the temperature rise due to the subtherapeutic surgical laser pulse and determine the required surgical laser pulse energy for heating the target to a predetermined temperature. The same technique in <FIG> applies to pulsed laser (coherent light source) SP surgical systems, other non-coherent pulsed light-source SP surgical systems, high-intensity- focused-ultrasound therapy systems, and other general radiation SP systems using a surgical pulsed radiation beam to heat up lesions in tissue or extraneous contrast agents attached to lesions in tissue, and a pulsed or modulated temperature-sensing radiation beam to effectively excite photoacoustic waves from lesions in tissue or extraneous contrast agents attached to lesions in tissue.

Another potential usage of the acoustic detector is to provide in vivo temperature calibration for skin cooling devices that provide protections for skin epidermis layer in a laser SP surgery. Skin cooling is widely used in laser treatment of vascular malformation. Skin cooling can effectively prevent laser-induced cavitation in epidermis during laser tattoo removal as well. <FIG> shows an example of operations for an optimized laser SP surgery with skin cooling. First, we tune the surgical laser wavelength to maximize SP surgical effects; Second, we optimize the surgical laser pulse width according to the measured thermal relaxation time of the surgical target; Third, we optimize surgical laser pulse energy according to the expected surgical effects (photocoagulation or photodisruption); Fourth, we measure body temperature, send a temperature sensing laser pulse, measure the amplitude of the excited photoacoustic signal of an epidermis target, and calculate its logarithm value as a baseline signal; Fifth, we adjust skin cooling parameter, apply skin cooling, and send a delayed temperature-sensing laser pulse; Sixth, we measure the amplitude of the photoacoustic signal excited by the temperature-sensing laser pulse, calculate the relative logarithm function value by taking logarithm and subtracting the baseline signal; Seventh, we determine the temperature of the epidermis target and whether the epidermis target is cooled to a predetermined temperature. If not, we wait for temperature recovery and return to the fifth step. If yes, we perform optimized SP surgery. The same technique in <FIG> applies to pulsed laser (coherent light source) SP surgical systems, other non-coherent pulsed light- source SP surgical systems, and other general radiation SP systems using a surgical pulsed radiation beam to heat up lesions in tissue or extraneous contrast agents attached to lesions in tissue, and a pulsed or modulated temperature-sensing radiation beam to effectively excite photoacoustic waves from lesions in tissue or extraneous contrast agents attached to lesions in tissue.

The example described above is the most comprehensive implementation of the techniques disclosed herein.

Claim 1:
A selective photothermolysis device, comprising:
a tunable radiation source (<NUM>) configured to emit pulsed radiation and optionally a pulsed or modulated temperature-sensing radiation, and excite photoacoustic waves from one or more targets in a tissue (<NUM>);
a patient interface (<NUM>) comprising
a radiation delivery unit (<NUM>) configured to deliver said radiation to said tissue (<NUM>), and
an ultrasonic detector (<NUM>) for detecting said photoacoustic waves excited by said radiation from one or more targets in said tissue (<NUM>); and
a control system (<NUM>) that digitizes photoacoustic signals detected by said ultrasonic detector (<NUM>), analyzes said photoacoustic signals and controls the tunable radiation source to generate radiation pulses with an optimal central wavelength, and, optionally, an optimal pulse width, or an optimal pulse energy, or a combination of them for optimal surgical outcome.