LASER THERAPY DEVICE FOR THE THERAPY OF LIVING TISSUE

Description

[0001] The invention relates to a device for laser therapy of living tissue comprising a pulsed laser light source and automatic control of the emitted power. The invention relates in particular to an ophthalmological laser device for the treatment of the fundus (back of the eye).

[0002] Lasers are used to treat a wide variety of retinal diseases. During laser irradiation of the fundus (back of the eye), the light-absorbing layers—the retinal pigment epithelium (RPE) and the choroid—are primarily heated, from where the heat spreads to the adjacent layers. This heat distribution depends not only on the absorbed laser energy but also on the pulse duration and the area of ​​irradiation. The typical area per laser spot on the retina is 100–500 µm in diameter, with an irradiation time of usually 20–500 ms. The shorter the irradiation time, the less heat diffuses into the volume outside the irradiated region.

[0003] Depending on the therapeutic goal, different temperatures are desirable to avoid undertreatment (i.e., the absence of a therapeutic effect) or overtreatment (i.e., unintended collateral damage). Within the specified time range, approximately 43–50 °C is required for sublethal stimulation of the retina (e.g., hyperthermic tissue stimulation for the prophylactic treatment of dry age-related macular degeneration, AMD), 50–70 °C for subthreshold photocoagulation, or 70–90 °C for mild photocoagulation (e.g., diabetic maculopathy), and 90–120 °C for strong photocoagulation (panretinal irradiation for diabetic retinopathy or retinal detachments). Selective thermomechanical disruption of the RPE (e.g.,For central serosa retinopathy, nano- to microsecond pulses are used, with temperatures > 140 °C occurring only in the RPE (selective retinal therapy, SRT) and leading to the selective generation of microbubbles on the intracellular melanosomes. From a pulse duration of 2 µs, thermal necrosis has been observed in addition to disruption. Generally, pulse durations between 2 and 50 µs are considered the transition range between thermomechanical and thermal effects.

[0004] Due to the significant fluctuations in light transmission through the eye and the highly variable light absorption of the RPE (20-80% in the visible spectral range), the aforementioned temperature increases cannot be achieved with fixed laser settings. Therefore, measuring and controlling the temperature increase during irradiation is highly desirable.

[0005] For example, an optoacoustic method for temperature measurement is known from US 2003 / 032949 A1, in which retinal absorbers are thermoelastically excited by means of repetitively short-pulsed laser radiation (pulse duration in the nanosecond range, repetition rate typically 1 kHz), which is superimposed on the therapeutic laser radiation, and the emitted pressure transients at the cornea of ​​the eye can be detected by means of an ultrasound transducer. Due to the temperature dependence of the Grüneisen coefficient, the pressure amplitudes increase with the retinal temperature and, after prior calibration, indicate the current mean temperature increase at the retina. Strictly speaking, temperature determination in this way is only possible as long as the target tissue is merely reversibly heated, but not yet thermally transformed and permanently damaged.Nevertheless, concepts have been developed to assess tissue damage caused by a laser pulse sequence ("burst") and, in particular, to perform temperature measurements during a first subsequence of the burst in order to control the laser during the application of a second subsequence of the same burst, see, for example, Publication US 2010 / 292763 A1.

[0006] The goal is for the treating physician to trigger the therapeutic laser only once for each laser spot to be treated, assuming that the intended tissue damage occurs precisely where it is needed, without causing excessive collateral damage. As a rule, this "gentle" laser therapy results in only faint or even invisible lesions on the retina. Therefore, the physician has no direct and objective means of monitoring the effectiveness of the treatment.

[0007] An obvious disadvantage of the procedure described in US 2003 / 032949 A1 is the need to use a separate measuring laser in addition to the therapy laser, which increases the complexity and cost of the laser system.

[0008] Currently, continuously emitting high-power lasers with emission powers on the order of 10 W in the green spectral range are under development for retinal applications. In the present context, emission powers on the order of 5 W are included in the invention. It is expected that in the future, lasers with emission powers on the order of 100 W and nearly constant power over extended periods will also become available for medical purposes. When using such high powers, the total irradiation time per laser spot must be in the lower microsecond range (2-50 µs) or even below and must be very precisely controlled to avoid causing severe, uncontrollable damage to the retina. In particular, when using such lasers, a single pulse may be therapeutically effective instead of a laser pulse sequence.

[0009] The extremely rapid heating of the tissue makes it impossible to use an optoacoustic method like US 2003 / 032949 A1 for temperature control, because the pressure transients generally require a travel time of up to approximately 20 µs from the retina to a sound sensor in a contact lens. The measurement signal would arrive too late to regulate the heating laser pulse.

[0010] Alternatively, one could rely on optical detection of the transient volume change of the irradiated tissue through reversible thermal expansion or contraction under laser light heating. Such methods have been tested, for example, in US 2013 / 102894 A1 and US 2015 / 011983 A1, and utilize interferometers and measurement light sources specifically designed for interferometry, typically reflected into the beam path of the therapeutic laser. However, the disadvantages here also lie primarily in the complexity and cost of the laser system.

[0011] The work by Xiaohua Feng et al., "Self-temperature regulation of photothermal therapy by laser-shared photoacoustic feedback", Vol. 40, No. 19, October 1, 2015, Optics Letters, pp. 4492-4495, describes a laser device with photoacoustic (= optoacoustic) feedback for temperature control of the target using only one heating laser. This laser operates in pulse-modulated mode, with long heating pulses and linearly chirped sequences of short pulses for temperature measurement via pressure transients using an ultrasonic transducer. This enables automatic adjustment of the irradiation dose to the target by modulating the pulse durations of the heating laser, allowing the target to reach and maintain predetermined temperatures with high precision. In this work, the target is a phantom, the laser emission power is 1 W, and the heating laser pulses have pulse durations in the millisecond to second range.The chirp sequence takes 40 microseconds, and there is no delay between the sequence and the ultrasound transducer. These simplified conditions do not allow for direct applicability to the problem of retinal therapy.

[0012] The idea of ​​using the heating laser light itself for temperature determination is pursued here, as already described in US 2010 / 292763 A1. Publication DE 10 2009 016184 A1 describes an approach with alternately applied short measuring laser pulses and much longer heating laser pulses from a common laser source, specifically for laser therapy of the fundus under temperature monitoring and control.

[0013] The work on photoacoustic imaging by Fei Gao et al., "Single laser pulse generates dual photoacoustic signals for differential contrast photoacoustic imaging", Scientific Reports | 7: 626 | DOI:10.1038 / s41598-017-00725-4, published online on April 4, 2017, reveals the possibility of photoacoustic imaging of biological tissue using only two pressure transients, triggered and detected at the beginning and end of a long heating laser pulse. Without stress confinement but while maintaining thermal confinement, the pressure transients at the beginning and end of the heating laser pulse exhibit opposite polarity and different amplitudes, which can provide information about the energy density deposited in the tissue during the pulse duration.

[0014] Against this background, the invention aims to first propose a laser therapy device for treating living tissue that enables self-regulating, gentle therapy exclusively with heating laser pulses, i.e., with laser pulses of relatively high power and pulse duration. The second objective of the invention is to further develop the device to enable an objective assessment of the effects.

[0015] The first problem is solved by a laser therapy device for treating living tissue comprising a pulsed laser light source emitting laser light with an emission power in the range of 5 W to 100 W, a device for guiding the laser light into an applicator, a triggering device for initiating tissue irradiation with the applicator, a detection device for measuring the time-dependent volume changes caused by laser absorption in the tissue, a computing unit for evaluating the detected volume changes and for outputting control commands to a control unit for controlling the power of the laser light emitted into the tissue, characterized in that a. Each triggering of tissue irradiation causes the control unit to apply a first heating laser pulse with a first power and a first pulse duration and at least one second heating laser pulse with a second power and a second pulse duration to the tissue at a predetermined time pulse interval, wherein b. the detection device optically or acoustically detects the volume changes caused by the rising and falling power edges of the first heating laser pulse and supplies the measured values ​​to the processing unit; c. the processing unit, based on the measured values ​​of the volume change and taking into account at least the predetermined slopes of the power edges of the first heating laser pulse, determines an estimated value for the temperature increase of the tissue during the irradiation of the first heating laser pulse; and d.The processing unit generates a command from the estimated value to the control unit, which instructs the control unit to adjust the second power and / or the second pulse duration of the second heating laser pulse so that the irradiation of the second heating laser pulse heats the tissue to a predetermined target temperature.

[0016] The sub-claims are aimed at advantageous designs, in particular also at further training for impact assessment.

[0017] Previously, short, low-energy laser pulses were typically used to determine temperature; these pulses, in themselves, were not intended to contribute significantly, if at all, to heating the tissue. The general aim was to avoid altering the tissue's state during the temperature measurement process.

[0018] The device according to the invention now dispenses with applying short laser pulses or bursts of such laser pulses, but instead applies exclusively long heating laser pulses with medium to high laser power, each of which results in a temperature increase of several degrees Celsius over its pulse duration.

[0019] To treat a single laser spot - especially on the retina - only very few pulses are needed, preferably just two.

[0020] The first heating laser pulse, which heats a laser spot, should preferably not yet cause permanent tissue damage (e.g., denaturation), but rather only induce sufficiently rapid volume changes through thermal expansion of the tissue at the beginning and contraction at the end of the heating laser pulse, which can be detected by an optical or acoustic detection device. The detection devices generate electrical volume change signals that can be transmitted to a processing unit and evaluated by it.

[0021] The detection device can, for example, be designed to optically detect thermally induced movements of light-scattering or reflecting tissue layers within the irradiated tissue. It may include an interferometer and at least one photodetector that measures and outputs time-varying light intensity. Optical coherence tomography (OCT) or a high-coherence light source (single longitudinal mode) with speckle detection can be used for this purpose. The therapy laser itself can also emit light with sufficiently high coherence, so that a small portion of the therapy light can simultaneously serve as measurement light. The determined movements—such as the change in the distance between two scattering tissue layers—allow for a direct conclusion to be drawn about the desired volume change.

[0022] According to the invention, it is preferred that the detection device is designed to detect pressure waves generated thermally in the irradiated tissue and comprises at least one ultrasound transducer that detects a pressure transient and outputs it as measured values. Sufficient time is available for the ultrasound waves to propagate from the heated tissue to the ultrasound transducers both during the long pulses and between the pulses.

[0023] According to the invention, it is provided that the second heating laser pulse, which is applied only after the detection and evaluation of the volume change signals, is controlled with the information from the first heating laser pulse.

[0024] Preferably, the first and second pulse durations lie within the interval of 2 to 50 microseconds. However, other pulse durations, especially shorter ones, are not excluded. The distinguishability—the separability—of the volume change signals at the beginning and end of the pulse depends on the detection method; optical detection allows this even for very short pulses. Acoustic detection is generally simpler in terms of equipment.

[0025] The pulse interval is usually defined as the time difference between similar features of two successive laser pulses, for example, the rising pulse edges or the pulse maxima. However, with few pulses of long duration, it is more advantageous to define the pulse interval as the duration of the pause between two pulses, i.e., the time difference between the falling edge of one heating laser pulse and the rising edge of the next. Preferably, the time interval between the first and at least the second heating laser pulse should be predetermined to be greater than the thermal relaxation time of the tissue.

[0026] Prior knowledge regarding thermal relaxation time exists in the literature, e.g., Brinkmann and Birngruber, "Selective Retina Therapy," in Z. Med. Phys 17 (2007), pp. 6-22. There, it can be seen that the thermal relaxation time depends on the size and structure of the object heated with laser light and is approximately 64 µs, particularly for an assumed homogeneous RPE absorber layer 4 µm thick. The authors further conclude: "If the aim is to heat all RPE cells uniformly without significantly heating the surrounding tissue, pulse durations of approximately 5-30 µs are suitable. However, if one wants to generate pronounced peak temperatures at the melanosomes, pulse durations < 5 µs should be chosen. With pulse durations > 50 µs, the cell-limited temperature increase is lost." This also results in the aforementioned preferred interval for pulse durations.

[0027] It is particularly preferred that the time interval between pulses is chosen to be so large that the temperature increase of the tissue achieved with the first heating laser pulse has already been largely eliminated by heat dissipation before the second pulse is triggered, i.e. that the tissue has returned to the base temperature of the tissue before the start of irradiation, typically 37 °C.

[0028] While the first heating laser pulse serves to acquire information about the tissue in the selected laser spot, according to the invention, the at least second heating laser pulse causes the tissue to be heated to a target temperature for a therapeutic effect. The second heating laser pulse is typically more energetic than the first heating laser pulse; that is, the product of laser power and pulse duration is typically larger for the second heating laser pulse than for the first.

[0029] It is considered advantageous that the first and second heating laser pulses have the same pulse duration and that the laser power is controlled only for the second heating laser pulse. With previously developed high-power laser diodes, this can be readily achieved by controlling the direct current. This eliminates the need for expensive optical components for electronically controlled shading or deflection of the laser light, such as an acousto-optic modulator. However, the use of light modulators to vary the applied power is also within the scope of the present invention.

[0030] Provided that the emission power of the laser light source exhibits a sufficiently steep rising and falling edge during the switching on and off of a heating laser pulse, with a smooth, as constant as possible plateau in between, only two significant volume change signals are actually generated, provided that the energy irradiated during the edge duration, δt, corresponds to the typical pulse energies of probe pulses from previous detections. For photoacoustic detection, for example, probe pulses with an energy of 5 µJ are known to be sufficient. At a laser power of P max = 10 W, this corresponds to an edge duration of δt = 1 µs, in which the applied laser power increases linearly from zero to its maximum value and then decreases again.

[0031] A particular advantage of the invention is that the evaluation of the volume change resulting from the first heating laser pulse, used to determine an estimated temperature increase, can be applied in exactly the same way to at least the second heating laser pulse. This allows for an estimated value for the actual therapeutically effective target temperature achieved by the laser after the treatment of a laser spot. This estimated value can be displayed to the treating physician in real time and, moreover, recorded (i.e., stored electronically) and compared with the target temperature desired by the physician. The user thus has immediate control over the effect and, if desired, access to a stored therapy protocol, which can also be used as electronic proof of treatment or to document a therapy history.

[0032] Applying heating laser pulses in the microsecond time regime with laser powers in the watt range to absorbing tissue results in readily detectable, rapid volume changes due to tissue expansion (dV1) and contraction (dV2) during the power changes dP1 and dP2 of the laser pulse. This occurs at the rising (dP1) and falling (dP2) edges of the laser pulse, provided the power changes are rapid enough, with short edge durations dt1 and dt2, respectively. The heating laser pulse edges thus determine the measurement signals to be acquired, namely the tissue response directly to the switching on and off of the laser. For pulses with a constant plateau profile, considering the slopes of the laser power at the pulse edges is sufficient for signal evaluation.

[0033] The expansion rates dV1 / dt1 and contraction rates dV2 / dt2 can be directly measured using optical interferometric methods by tracking positional displacements of tissue layers, such as the tissue surface. These rates are highest during the edges of the heating laser pulse and allow for the estimation of the tissue's coefficient of thermal expansion, which, in turn, can be determined from previous studies as a temperature-dependent material property over a wide temperature range.

[0034] Simultaneously, the rapid volume changes trigger pressure waves in the irradiated tissue, which propagate through the tissue and can be registered as a time-dependent pressure signal (pressure transient) by an ultrasound sensor. In this case, the previously known temperature dependence of the Grüneisen coefficient is used to directly infer the temperature from the pressure transients. Specific measured quantities dD1 and dD2 could include, for example, the maximum amplitude of the pressure transient, the integral of the pressure transient over a time interval of increased amplitude, or the frequencies of the pressure transients.

[0035] Experimentally, using acoustic detection, it was found that the ratio of pressure measurement signals from the end of the laser pulse to the beginning can be used as a measure of tissue heating during the laser pulse, provided the power changes and edge durations at both pulse ends are equal in magnitude. However, this is generally not the case. Nevertheless, the laser pulse edges can either be measured with each pulse or—given sufficient laser stability from pulse to pulse—derived from pre-tabulated data, which the laser light source manufacturer could provide.

[0036] If the edge profiles differ at the beginning and end of the heating laser pulse, the ratio of the pressure measurement signals can be normalized to the edge slopes, for example according to: α = dD 2 ΔP 2 dD 1 ΔP 1

[0037] The size ΔP This can be considered the maximum slope of the laser flanks. ΔP = dP dt t = t max or alternatively, the average slope ΔP = 1 δt ∫ 0 δt dP dt t ′ dt ′ averaged over the flank duration δt be introduced. Other standards in this sense are also possible.

[0038] The steeper the changes in laser light power, the faster the tissue volume changes. This results in higher expansion and contraction rates and higher pressure amplitudes.

[0039] Regarding acoustic detection, the value that can be determined, in other words, relates the energies of the triggered pressure waves to the incident laser energies for the pulse end edge and pulse beginning edge, respectively, can be calculated. From the ratio α and the known temperature dependence of the Grüneisen coefficient of the tissue, the resulting temperature increase of the tissue up to the pulse end can be derived.

[0040] Particularly at pulse durations shorter than the tissue's thermal relaxation time, a temperature gradient arises in the beam propagation direction z, corresponding to the absorption and scattering of the radiation µa(z) and µs(z). The value α therefore only yields an average weighted temperature value, which depends on all irradiation and tissue parameters.

[0041] To approximate the temperature increase, the temperature-dependent behavior of the Grüneisen coefficient and the initial temperature before irradiation (body temperature, room temperature) must be known. The temperature dependence of the Grüneisen coefficient can be described, for example, by a quadratic equation of the form Γ T = C ∗ T 2 − T 0 2 − 2 T m T − T 0 for tissues in the temperature range of 0–100 °C. In the case of porcine retinal tissue in a cuvette, the maximum of the parabola was T m = 105.4°C, and the zero crossing T0 = -21°C approximated (from: Kandulla et al, Journal of Biomedical Optics 11(4) 2006), where C is a proportionality constant.

[0042] One way to calculate the average temperature increase in the tissue, T Helz , The solution involves calculating this directly using a fit function. The average temperature increase results from the parabolic approximation above. T Heιz ‾ = T m ± αB + T m − T 0 2 with B = T B 2 − T 0 2 − 2 T m ⋅ T B − T 0

[0043] At T B This refers to the base temperature (e.g., room temperature or body temperature). To calculate the average temperature increase... T Helz on the often more interesting maximum temperature increase T max To be able to conclude that the temperature is highest at the location of the highest temperature, a constant f can be determined. T max = T Heιz ‾ − T B ⋅ f + T B

[0044] The constant f generally depends on the irradiation geometry and duration, as well as on the thermal tissue parameters. In the case of only an extensive homogeneous absorber without scattering, when irradiated with pulse durations below the thermal relaxation time, the following results: f = 2. In general, f For specific sample types or, in this case, tissue classes, the values ​​can be determined using theoretical models or experimentally (e.g., using thermocouples, temperature-dependent fluorescent nanoparticles, etc.).

[0045] To control the irradiation process to a defined target temperature, T ziel , At least two heating laser pulses are applied, with the first heating laser pulse having an energy E 1, which does not yet allow the target temperature to be reached and preferably also does not cause thermal damage to the tissue. From the temperature measured after the first pulse T Heiz can the energy E2. The second heating laser pulse can be adjusted by setting the power and / or the pulse duration so that the target temperature is reached after the end of the second pulse, for example by T Ziel = T Heιz ‾ − T B ∗ E 2 E 1 + T B

[0046] Multiple pulses can also be applied to gradually approach the target temperature.

[0047] As long as the heat energy deposited in the tissue through laser absorption cannot leave the irradiated volume via heat transport mechanisms (thermal confinement), it is essentially irrelevant whether the increased energy E2. This is achieved by increasing the power or the pulse duration of at least one second heating laser pulse. Initially, only the products of power and pulse duration are relevant. However, it is considered advantageous not to increase the second pulse duration, but rather to keep it the same as the first and increase only the power. Heat transfer is time-dependent, and the transport processes during the second heating laser pulse are naturally most similar to those of the first heating laser pulse at the same pulse duration. Furthermore, current control of the laser light source shows that switching on and off can lead to different laser power profiles at different pulse durations. If one wants to forgo a light modulator to control the emission of the continuously activated laser and instead operate the laser with current pulses, one cannot necessarily assume that the pulse shape will be preserved at different pulse durations.This can complicate the evaluation of the measured values ​​and is easily avoidable if all pulse durations are chosen to be the same.

[0048] It is advantageous to choose a pulse interval large enough to allow the tissue to cool down to its baseline temperature as completely as possible between pulses. Alternatively, any residual heat buildup can also be factored into the calculation.

[0049] As already mentioned, a key advantage of the invention is that the actual target temperature achieved in the tissue can be measured. This is accomplished by the detection device recording the volume changes caused by the rising and falling edges of the second heating laser pulse, feeding these measurements to the processing unit, and the processing unit then calculating, displaying, and / or recording an estimated value for the temperature increase during the second heating laser pulse. Such a numerical, objective assessment of the effect of laser treatment is considered a significant advancement over previous therapeutic lasers.

[0050] Furthermore, high-power lasers in the range of 5 W to 100 W also open access to thermomechanical therapy approaches that focus on cell disruption through local vaporization. Here, too, treatment and even a certain degree of effect control are possible with the same laser device according to the invention, provided that only the software of the processing unit is extended.

[0051] Preferably, in this case—that is, when thermomechanical tissue damage is intended with a target temperature of, for example, more than 140 °C—the processing unit is configured to determine and record, starting from the beginning of the second heating laser pulse, the difference between the time of occurrence of a volume change on the rising edge of the heating laser pulse and the next time of occurrence of a volume change. If this measured difference is less than the second pulse duration, then the measurement signal indicating the volume change arrives prematurely and therefore cannot originate from the falling edge of the laser pulse. Rather, volume changes resulting from the formation of microbubbles in the tissue are detected. Blister formation is typically accompanied by pressure amplitudes that are up to two orders of magnitude larger than the purely thermally induced pressure amplitudes, provided the edges of the heating laser pulses are not too steep.These bubbles are therefore practically impossible to miss, and the moment they appear marks the beginning of bubble formation. Again, the time difference between the start of bubble formation and the end of the heating laser pulse is measurable and can be used as a measure of the thermomechanical treatment dose. This measure is directly evident from the time difference between the start of the laser pulse and the start of bubble formation, as determined by the processing unit, provided the pulse duration is known.

[0052] The invention will be explained in more detail below, also with the aid of figures. This will show: Fig. 1 a sketch of the laser system with computing unit and detector for volume change signals; Fig. 2 Experimental data on the slopes of laser light at different pulse durations and powers using a diode laser as an example; Fig.3a-3c Experimentally measured (time-shifted) pressure transients for heating laser pulses of comparable pulse duration with different power levels.

[0053] In Fig. 1 The sketch shows a diagram of the laser device according to the invention in an embodiment for the treatment of the retina. The means for guiding the light, the applicator, and the trigger for the applicator are not shown in this diagram.

[0054] A laser source (4) emits heating laser pulses (1) onto tissue (2). A detection device registers the change in tissue volume and generates electrical signals that are fed to a processing unit (6). The processing unit (6) can be a programmable microprocessor or a personal computer. It processes the raw data arriving from the detection device (3) – and stores it non-volatilely if necessary – and determines control commands for the laser (4) based on the processing results. A control device for the laser (4) is not shown separately, but is usually integrated into the laser (4) as a single unit. If, as is preferred here, the operating current of the laser (4) is varied to control the laser power, then the control device can be implemented as a simple analog circuit that converts the voltage values ​​directly from the processing unit (6) into currents.Alternatively, the control unit can also be integrated into the computing unit (6). The control unit can also be configured to digitally receive meta commands from the computing unit (6), such as values ​​for pulse duration and pulse power, and translate these into analog signals for controlling the laser (4).

[0055] According to the invention, the laser (4) has an emission power on the order of 5 W to 100 W. Here, by way of example, integrated into the laser (4) and not shown separately, means for temporally recording the emitted laser power of at least a portion of the laser light (1) are preferably provided. The detected measured values ​​of the laser power are supplied to the processing unit (6), which determines the slopes of the power edges for the laser pulses. Such beam control is common prior art. For very stable laser light sources, creating a table with slopes for laser pulses of different pulse durations and powers is also an option instead of continuously repeated measurements.

[0056] An operating device (5) for user input of irradiation parameters into the processing unit (6) is preferably configured to display user-selectable intervals for irradiation parameters, the interval limits being hierarchically dependent on user inputs. This means that the operating device (5) can comply with certain safety rules and reject inputs that contradict these rules. When using a high-power laser with a high potential for damage, especially in retinal therapy, erroneous parameter settings by the user can thus be intercepted at an early stage. By way of example, and preferably, the operating device (5) can be configured to select the power of the first heating laser pulse based on the user input of the pulse duration in such a way that the first heating laser pulse does not cause permanent tissue damage.Furthermore, it is advantageous if the operating device (5) is designed to allow only values ​​greater than a thermal relaxation time of the tissue predetermined by the user input for the first pulse duration as user input for the time interval between the first and at least the second heating laser pulse. Hierarchical dependency here means that some user selection options are implemented with priority, while other selection options are considered secondary and are appropriately restricted by applying predetermined – e.g., programmed – safety rules, depending on the priority user settings.

[0057] The control element (5) can therefore include its own processor for data processing and a data storage device for non-volatile data storage.

[0058] The control element (5) can be configured such that user inputs include at least the target temperature and the laser pulse duration – for example, primarily – as well as, optionally and secondarily, the time interval between laser pulses and the initial power of the first heating laser pulse. Furthermore, the processor of the control element (5) can execute an encryption algorithm to encrypt data transmitted by the computing unit (6) together with the user inputs, thereby securing them inseparably and immutably. Data secured in this way can only be read together subsequently and can serve as an authenticated treatment record (also with a timestamp).

[0059] The variation of the ratio of the slopes, ΔP 2 / ΔP 1 , of the heating laser pulses has been experimentally investigated by varying the pulse duration and the laser power. Fig. 2 The figures shown for pulse durations of 5 µs and 50 µs at powers up to 14 W are for illustrative purposes only. They demonstrate that the ratios of the pulse slopes—as well as the slopes themselves—change with the applied power due to altered excitation conditions (including the laser pump current). It becomes clear that pulse-wise measurement of the slopes is advantageous for the evaluation described here if one wishes to avoid a cumbersome calibration of the laser light source. Furthermore, laser aging effects are possible, meaning that a single factory calibration is not always sufficient.

[0060] The Fig. 3a-3c Three examples of measured pressure transients at different laser powers are shown. The pressure transients are represented by the lower curves, and the derivatives of the heating laser pulse by the upper curves. For ease of comparison, all plots have been normalized so that the first maximum of the pressure transients corresponds numerically to the maximum slope ΔP1 of the rising edge of the first heating laser pulse. This makes it easy to see the contribution of a temperature increase to the pressure transient at the end of the heating laser pulse. The acoustic travel time from the sample to the sensor has been removed for clarity; that is, the pressure transients have been shifted by this travel time.

[0061] Fig. 3a shows pressure transients during irradiation with 614 mJ / cm²: The applied pulse with a pulse duration of 35 µs exhibits two slopes. ΔP 1 and ΔP Two pulses are generated, each triggering a pressure transient A1 and A2. The pressure transients are shown here on the same time axis as the heating laser pulses, i.e., shifted by the constant signal propagation time from the retina to the ultrasound transducer for better alignment. Energy measures dD1 and dD2 are assigned to the pressure transients, which here are determined, for example, from the maximum amplitudes of signals A1 and A2. From this, the ratio α is calculated, resulting in... T Helz The temperature at the end of the pulse reaches 55°C. Such a brief temperature increase generally does not damage living retinal tissue. The curves shown represent an example of the first heating laser pulse of the method according to the invention.

[0062] Fig. 3b shows pressure transients during irradiation with 1245mJ / cm²: When applying a pulse of twice the power compared to Fig. 3a The same signals branch off, now with increased values. The laser pulse edges rise faster, and the amplitudes of the pressure transients are significantly increased. Using the same calculations as for Fig. 3a This results in a temperature of 82°C at the end of the pulse, which is already within the usual range of the target temperature for photocoagulation. Fig. 3b This thus represents a possible second heating laser pulse of the invention, which leads directly to the therapeutic effect.

[0063] The Fig. 3c The graph shows pressure transients during irradiation with 1488 mJ / cm²: This irradiation leads to blistering in the tissue. The second pressure transient begins even before the laser power decreases. It is therefore not due to thermoelastic relaxation upon cooling of the tissue, but rather to material vaporization. The double arrow 1 indicates the time difference between the onset of the acoustic signal and the end of the irradiation. Note the changed scales compared to the previous graph. Fig. 3a and Fig. 3b The amplitude, increased tenfold, is also an indication of microbubble formation and confirms the assumption. Since the irradiation power and the time required for microbubble formation are known, it is possible to determine the irradiance required to reach the bubble formation temperature of approximately 140°C. Similarly, the duration of the thermomechanical disruption until the end of the heating laser pulse can be used as a dose measure and, if necessary, adjusted and optimized for subsequent laser pulses by setting the power.

[0064] It should also be noted here that the curves of Fig. 3a-3c They were not recorded at the same laser spot. Therefore, the tissue absorption coefficients differ between the experiments.

Claims

1. Laser therapy device for the therapy of a living tissue (2) comprising a pulsed operated laser light source (4) emitting laser light (1) with an emission power of the order of magnitude of 10 W to 100 W, a device for guiding the laser light (1) into an applicator, a trigger device for triggering tissue irradiation with the applicator, a detection device (3) for measuring the time-dependent volume changes triggered by laser absorption in the tissue (2), a computing unit (6) for evaluating the detected volume changes and for issuing control commands to a control device for controlling the power of the laser light (1) irradiated into the tissue (2), wherein a. each triggering of the tissue irradiation causes the control device to apply a first laser pulse with a first power and a first pulse duration, and at least a second laser pulse with a second power and a second pulse duration at a predetermined time interval to the tissue (2), wherein b. the detection device (3) optically or acoustically records the volume changes caused by the rising and falling power flank of the laser pulse, and supplies the measurement to the computing unit (6); c. the computing unit (6), on the basis of the measurements of the volume change, taking into account at least the predetermined rises of the power edges of the first laser pulse, determines an estimated value for the temperature increase of the tissue (2) during the first laser pulse irradiation and characterised in that d. from the estimated value, the computing unit (6) generates a command to the control device, which causes the control device to set the second power and / or the second pulse duration of the second laser pulse so the irradiation of the second laser pulse heats the tissue to a predetermined target temperature.

2. Laser therapy device according to claim 1, characterised in that the control device is configured in such a way that the pulse durations of all laser pulses applied at the time of triggering are the same.

3. Laser therapy device according to any one of the preceding claims, characterised in that the detection device (3) records the volume changes caused by the rising and falling power edge of the laser pulse, and supplies the measurement to the computing unit (6) and the computing unit (6) determines, displays and logs an estimated value for the temperature increase during irradiation by the second laser pulse.

4. Laser therapy device according to any one of the preceding claims, characterised in that means are provided for the temporal recording of the emitted laser power which detects at least a portion of the laser light (1) and transmits measurements of the laser power to the computing unit (6), wherein for the laser pulses the computing unit (6) determines the rises of the power edges.

5. Laser therapy device according to any one of the preceding claims, characterised in that detection device (3) is designed to record pressure waves thermally triggered in the irradiated tissue and comprises at least one ultrasonic transducer that records a pressure transient and emits it as measuring values.

6. Laser therapy device according to any one of claims 1 to 4, characterised in that the detection device (3) can be designed, for example, for the optical detection of thermally triggered movements of light-scattering or reflecting tissue layers in the irradiated tissue and comprises an interferometer as well as at least one photodetector which measures a temporally changeable light intensity and emits this as a measuring values.

7. Laser therapy device according to any one of the preceding claims, characterised in that as of the start of the second laser pulse, the computing unit is designed to determine and log the difference between the time of occurrence of a change in volume at the rising edge of the laser pulse and the next time of occurrence of a change in volume.

8. Laser therapy device according to any of the preceding claims, characterised in that an operating device (5) for the user input of irradiation parameters into the computing unit (6) is designed to display to the user selectable intervals for irradiation parameters, wherein the interval limits depend hierarchically on user inputs.

9. Laser therapy device according to claim 9, characterised in that the control device (5) is designed to select the power of the first laser pulse based on the user input of the pulse duration in such a way that the first laser pulse does not cause lasting tissue damage.

10. Laser therapy device according to any one of claims 9 or 10, characterised in that the operating device (5) is designed to only allow as user inputs for the temporal pulse interval between the first and at least the second laser pulse, values that are greater than a thermal relaxation time of the tissue predetermined at least from the user input for the first pulse duration.