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WO2025238199A1 - Procédé de mesure de la température d'une zone à l'intérieur d'un objet et dispositifs associés - Google Patents

Procédé de mesure de la température d'une zone à l'intérieur d'un objet et dispositifs associés

Info

Publication number
WO2025238199A1
WO2025238199A1 PCT/EP2025/063509 EP2025063509W WO2025238199A1 WO 2025238199 A1 WO2025238199 A1 WO 2025238199A1 EP 2025063509 W EP2025063509 W EP 2025063509W WO 2025238199 A1 WO2025238199 A1 WO 2025238199A1
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WO
WIPO (PCT)
Prior art keywords
temperature
ultrasound
area
value
echo
Prior art date
Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
Pending
Application number
PCT/EP2025/063509
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English (en)
Inventor
Rémi Souchon
Omar GACHOUCH
Current Assignee (The listed assignees may be inaccurate. Google has not performed a legal analysis and makes no representation or warranty as to the accuracy of the list.)
Institut National de la Sante et de la Recherche Medicale INSERM
Centre Leon Berard
Universite Claude Bernard Lyon 1
Original Assignee
Institut National de la Sante et de la Recherche Medicale INSERM
Centre Leon Berard
Universite Claude Bernard Lyon 1
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Application filed by Institut National de la Sante et de la Recherche Medicale INSERM, Centre Leon Berard, Universite Claude Bernard Lyon 1 filed Critical Institut National de la Sante et de la Recherche Medicale INSERM
Publication of WO2025238199A1 publication Critical patent/WO2025238199A1/fr
Pending legal-status Critical Current
Anticipated expiration legal-status Critical

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Classifications

    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B8/00—Diagnosis using ultrasonic, sonic or infrasonic waves
    • A61B8/52—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves
    • A61B8/5207—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves involving processing of raw data to produce diagnostic data, e.g. for generating an image
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B8/00—Diagnosis using ultrasonic, sonic or infrasonic waves
    • A61B8/52—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves
    • A61B8/5215—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves involving processing of medical diagnostic data
    • A61B8/5223—Devices using data or image processing specially adapted for diagnosis using ultrasonic, sonic or infrasonic waves involving processing of medical diagnostic data for extracting a diagnostic or physiological parameter from medical diagnostic data
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01K—MEASURING TEMPERATURE; MEASURING QUANTITY OF HEAT; THERMALLY-SENSITIVE ELEMENTS NOT OTHERWISE PROVIDED FOR
    • G01K11/00—Measuring temperature based upon physical or chemical changes not covered by groups G01K3/00, G01K5/00, G01K7/00 or G01K9/00
    • G01K11/22—Measuring temperature based upon physical or chemical changes not covered by groups G01K3/00, G01K5/00, G01K7/00 or G01K9/00 using measurement of acoustic effects
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01K—MEASURING TEMPERATURE; MEASURING QUANTITY OF HEAT; THERMALLY-SENSITIVE ELEMENTS NOT OTHERWISE PROVIDED FOR
    • G01K11/00—Measuring temperature based upon physical or chemical changes not covered by groups G01K3/00, G01K5/00, G01K7/00 or G01K9/00
    • G01K11/22—Measuring temperature based upon physical or chemical changes not covered by groups G01K3/00, G01K5/00, G01K7/00 or G01K9/00 using measurement of acoustic effects
    • G01K11/24—Measuring temperature based upon physical or chemical changes not covered by groups G01K3/00, G01K5/00, G01K7/00 or G01K9/00 using measurement of acoustic effects of the velocity of propagation of sound

Definitions

  • the present invention relates to a method for measuring the temperature of an area inside an object.
  • the present invention also deals with a corresponding calculator and device for measuring said temperature.
  • ultrasound scanners Being low-cost and portable devices, ultrasound scanners seem to be a good alternative to offer non-invasive monitoring treatment response.
  • the specification describes a method for measuring a temperature of an area of an object, the method comprising :
  • - a step of determining the temperature of the area based the value of the echo signal stretching factor and on another value, the value being the temperature change value when the temperature change value is known or a reference value when the temperature change value is unknown.
  • the method for measuring might incorporate one or several of the following features, taken in any technically admissible combination:
  • the step of obtaining comprises comparing the first ultrasound echo signal and the second ultrasound echo signal by a comparison technique.
  • the comparison technique enables to deduce the time delay between the first ultrasound echo signal and the second ultrasound echo signal, the step of obtaining further comprising obtaining the value of the echo signal stretching factor based on the time delay.
  • the comparison technique enables to deduce the time delay between the first ultrasound echo signal and the second ultrasound echo signal, the step of obtaining further comprising converting the time delay into a displacement and obtaining the value of the echo signal stretching factor based on the displacement.
  • the temperature change value is comprised between 0,1 °C and 5,0 °C.
  • the step of modifying is achieved by using a pulsed heat source.
  • the duration of the step of modifying is comprised between 10 -5 s and 1 s.
  • the temperature is obtained by applying a function on the echo signal stretching factor, the function being parametrized by coefficients derived from the temperature change value.
  • the method further comprises a step of measuring a reference stretching factor of the area and the reference value being the reference stretching factor of the area.
  • the method further comprises a step of estimating the speed of sound in the area based on the determined temperature.
  • the method further comprises a step of controlling the value of the temperature change value by comparing a value obtained at the end of the step of obtaining with an expected range of values for the echo signal stretching factor.
  • the method further comprises a step of emitting a warning in case the obtained value is outside the expected range of values for the echo signal stretching factor.
  • the object is a living subject.
  • the specification also relates to a calculator adapted to:
  • the specification also concerns a device for measuring a temperature of an area of an object, the device for measuring comprising :
  • an imaging ultrasound unit adapted to collect a first ultrasound echo emitted by the area in response to a first ultrasound excitation signal
  • a temperature modifying unit adapted to modify the temperature of the area by a temperature change value
  • an imaging ultrasound unit adapted to collect a second echo ultrasound emitted by the area in response to a second ultrasound excitation signal
  • a calculator adapted to obtain an echo signal stretching factor of the area based on the first ultrasound echo and the second ultrasound echo, and to determine the temperature of the area based on the value of the echo signal stretching factor and on another value, the value being the temperature change value when the temperature change value is known or a reference value when the temperature change value is unknown.
  • FIG. 1 is a schematic representation of a device for measuring the temperature of an area
  • FIG. 2 is a graph representing, on the Y axis, the variation of the speed of sound in an area inside an object, plotted against the temperature of the area, on the X axis,
  • FIG. 3 is an example of flowchart illustrating the carrying out of a method for measuring
  • a device 10 for measuring temperature is represented on figure 1.
  • the object 14 is a human being.
  • the object 14 is an animal, and more preferably a mammal.
  • the device 10 is a device adapted to measure the temperature of an area 12 inside an object 14.
  • the device 10 is thus adapted to carry out a method for measuring a temperature of an area 12 inside an object 14, as will be described later in this specification.
  • the area 12 is a zone inside the object 14, from which ultrasound echoes can be received.
  • the area 12 is, for instance, a zone of interest in a soft biological tissue.
  • Prostate, liver, kidney or muscle are examples of biological tissue.
  • the speed of sound c (speed of propagation inside the biological tissue) varies with the temperature T of the biologic tissue, as illustrated on figure 2.
  • Figure 2 illustrates the variation of the speed of sound c with the temperature T of canine liver samples.
  • the function f(T) is supposed to be known and can be obtained by any known ultrasonic tissue characterization method, such as the method disclosed in the article “Temperature dependence of ultrasonic propagation speed and attenuation in excised canine liver tissue measured using transmitted and reflected pulses” published by II. Techavipoo, T. Varghese, and colleagues in the Journal of Acoustical Society of America (volume 115, issue 6) in June 2004, for example.
  • the device 10 for measuring comprises an imaging ultrasound transducer 16, a heating unit 18 and a calculator 20.
  • the imaging ultrasound transducer 16 is adapted to send ultrasound waves to the area 12 and to collect the echoes emitted by the area 12 in response to the ultrasound waves.
  • the imaging ultrasound transducer 16 is thus adapted to provide an ultrasound image of the area 12.
  • the ultrasound transducer 16 is here any transducer adapted to apply ultrasound waves in a controlled manner to fulfill the requirement of health safety of the object 14.
  • the heating unit 18 is adapted to heat the area 12 in a region named heating region hereinafter.
  • Such heating unit 18 is a pulsed heating unit in so far as it generates a temperature increment by applying a pulse.
  • the heating unit 18 is another ultrasound transducer 16 adapted to apply ultrasound pulses, which generates the temperature increase.
  • This other ultrasound transducer 16 may also be adapted to apply a therapy to the user.
  • the device 10 for measuring therefore comprises two ultrasound transducers, one for imaging and another one for heating.
  • the device 10 for measuring only comprises one ultrasound transducer, the imaging ultrasound transducer 16 being also used to heat the area 12.
  • the heating unit 18 is a radiofrequency ablation probe.
  • the calculator 20 is an electronic circuit designed to manipulate and/or transform data represented by electronic or physical quantities in registers of the calculator 20 and/or memories into other similar data corresponding to physical data in the register or memory memories, other types of display devices, transmission devices or storage devices.
  • the calculator 20 is produced in the form of a programmable logic component, such as an FPGA (Field Programmable Gate Array), or even an integrated circuit, such as an ASIC (Specific Integrated Circuit).
  • a programmable logic component such as an FPGA (Field Programmable Gate Array)
  • an integrated circuit such as an ASIC (Specific Integrated Circuit).
  • the method when the method is carried out in the form of one or more software programs, that is to say in the form of a computer program, also called a computer program product, it is also capable of being recorded on a medium, not shown, readable by computer.
  • the computer-readable medium is, for example, a medium capable of storing electronic instructions and of being coupled to a bus of a computer system.
  • the readable medium is an optical disk, a magneto-optical disk, a ROM memory, a RAM memory, any type of non-volatile memory (for example FLASH or NVRAM) or a magnetic card.
  • a computer program comprising software instructions is then stored on the readable medium.
  • figure 3 is a flowchart showing an example of carrying out a method for measuring.
  • the method for measuring is a method for measuring the temperature T of the area 12.
  • the method for measuring comprises a first step of collecting S50, a step of increasing S52, a second step of collecting S54, a step of obtaining S56 and a step of determining S58.
  • a first ultrasound excitation signal is emitted by the imaging ultrasound transducer 16 towards the area 12.
  • the first ultrasound excitation signal is typically a broadband signal with duration between 0.1 and 1 microsecond, and with central frequency between 1 MHz and 40 MHz.
  • Said first ultrasound excitation signal induces an ultrasound wave, propagating towards the area 12.
  • the emitted ultrasound wave is focused on the area 12.
  • Such focusing can be achieved using any known means of focusing an ultrasound wave, for example using time delays between independent piezoelectric elements of an imaging transducer.
  • Such focusing can be achieved using multi-angle plane or wide beams, as described by G. Montaldo et al. in the article entitled “Coherent plane-wave compounding for very high frame rate ultrasonography and transient elastography”, IEEE Trans Ultrason Ferroelectr Freq Control. 2009; 56(3): 489-506.
  • Synthetic Aperture Ultrasound Imaging a technique where all array elements are fired one by one, as described in the article by Jensen et al., “ Synthetic Aperture Ultrasound Imaging”, Ultrasonics. 2006; 44 Suppl 1 :e5- 15.
  • the size of the focusing region of the ultrasound wave induced by the first ultrasound excitation signal is smaller than (strictly inferior to or equal to) the size of the heated region.
  • the first ultrasound excitation signal interacts with the area 12 and the area 12 emits a first ultrasound echo.
  • This first ultrasound echo is collected by the imaging ultrasound transducer 16.
  • Said ultrasound echo is a signal that varies in time.
  • the heating unit 18 heats the area 12, more specifically the heating region.
  • Such heating increases the temperature of the area 12 by a temperature increment 8T
  • the temperature increment 8T is comprised between 0.1 °C and 5.0 °C.
  • the duration of the step of increasing S52 is comprised between 10 -5 s and 1 s.
  • this short duration between the first step of collecting S50 and the second step of collecting S54 is advantageous, because it allows an accurate estimation of the deformation of the echo signal.
  • a second ultrasound excitation signal is emitted by the ultrasound imaging ultrasound transducer 16 towards the area 12.
  • the second ultrasound excitation signal should be identical to the first ultrasound excitation signal.
  • the second signal and the first signal must be highly correlated, i.e. the cross-correlation coefficient between them should be close to +1 or -1.
  • its shape can be identical but its amplitude can be different, because changes in amplitude do not change the absolute value of the correlation coefficient.
  • pulse inversion An example of changing amplitude while preserving correlation is pulse inversion, where the second signal is inverted compared to the first signal, as described by Shen et al. in the article “Pulse Inversion Techniques in Ultrasonic Nonlinear Imaging” (Journal of Medical Ultrasound 2005; 13(1): 3-17).
  • pulse inversion negative correlation is obtained, and the delay corresponds to the location of the minimum correlation.
  • coded excitation on one or the other of the signals, or on both signals, as described by Weng et al. in the article “Coded Excitation for Ultrasonic Testing: A Review” (Sensors 2024, 24(7), 2167; doi: 10.3390/s24072167).
  • typical codes used for coded excitation include chirp encoding, or Barker codes, or complementary Golay codes.
  • the second ultrasound excitation signal interacts with the area 12 and the area 12 emits a second ultrasound echo.
  • This second ultrasound echo is collected by the ultrasound imaging ultrasound transducer 16.
  • the calculator 20 obtains a value of an echo signal stretching factor s of the area 12 based on the first ultrasound echo and the second ultrasound echo.
  • the echo signal stretching factor s characterizes the reaction of the area 12 to the temperature increase.
  • a negative echo signal stretching factor corresponds to a compression of the echo signal in the time domain while a positive echo signal stretching factor corresponds to a stretching of the echo signal in the time domain
  • the echo signal stretching factor can also be named echo strain, thermal strain or apparent strain.
  • the step of obtaining S56 comprises comparing the first ultrasound echo and the second ultrasound echo.
  • Such comparing can be achieved by applying cross-correlation techniques on both ultrasound echoes.
  • the comparison of the ultrasound echoes can be achieved by comparing the spectrum of both ultrasound echoes (comparison in the frequency domain).
  • the first technique is applied.
  • a two-steps correlation technique implying two successive operations may be used, instead of using the standard crosscorrelation.
  • a coarse estimation is obtained, using cross-correlation on two adjacent windows (for instance, 2 mm length each, 0% overlap) at the focus position.
  • This coarse estimation is an integer number of temporal samples. It was used to shift and “realign” the windows. After coarse realignment, there still exists is a residual lag ( ⁇ 0.5 time sample) that will be estimated in the second operation.
  • crosscorrelation is used again on the coarsely realigned windows, and a subsample interpolation technique is used to estimate the residual delay, with sub-sample accuracy.
  • the subsample interpolation technique can be a cosine interpolation.
  • amplitude modulation correction may be used to further improve estimation accuracy, as described by Lindop and colleagues in their article “Estimation of Displacement Location for Enhanced Strain Imaging” (in IEEE Trans Ultrason Ferroelectr Freq Control 2007; 54(9): 1751-1571).
  • AMC is designed to compensate the bias on time-delay estimates, which is induced by intra-window variations in RF signal amplitude. Bias correction is based on an accurate estimation of the center of mass of the displacement at the window.
  • the AMC technique is applied on the first and second ultrasound echo signals.
  • the echo signal stretching factor is the gradient of the displacement d, along the direction of propagation z of the ultrasound imaging beam.
  • the stretching factor is obtained directly from the time delay r, without the conversion to displacement.
  • the stretching factor is given by the following equation : dr
  • the calculator 20 determines the temperature T of the area 12 based on the temperature increment 8T and the value of the echo signal stretching factor s.
  • the calculator 20 determines the temperature T by calculating the following formula: sc/6T + o-i
  • the value c 0 1540 m/s can be used.
  • T T - - s c- max 2a 2 8T
  • T c _ max is the maximum temperature reached by the parabola f(T).
  • T T - - s c- min 2a 2 8T
  • T c _ min is the minimum temperature reached by the parabola f(T).
  • the parabola is notably inverted in the case of fat tissue.
  • the echo signal stretching factor s is given by:
  • the function to be applied to the echo signal stretching factor s is a function parametrized by coefficients derived from the temperature increment 8T.
  • the calculator 20 obtains the temperature T of the area 12 based on the temperature increment 8T and the value of the echo signal stretching factor s.
  • the Applicant has carried out experimental validation in tissue-mimicking gel phantoms and in biological tissue samples ex vivo, with good results.
  • Such method for measuring therefore enables obtaining the temperature T of the area 12 over a wide range of temperatures based on thermal strain imaging.
  • the method is adapted to measure temperature inside solid matter, and especially in soft biological tissues and can be carried out in vivo or ex vivo.
  • the method can be carried at any time.
  • such method for measuring enables obtaining an instantaneous temperature measurement at any time anywhere.
  • the temperature increment is unknown, and the method needs to be calibrated.
  • the step of determining S58 takes into account a reference value instead of the temperature increment.
  • the reference value is a reference stretching factor of the area 12 and and the method further comprises a step of measuring said reference stretching factor of the area 12 .
  • Such technique corresponds to a calibration technique.
  • the calibration technique is performed by combining: - a priori knowledge of the temperature T c max at which the speed of sound exhibits a maximum (or a minimum), and
  • the known temperature can be the room temperature in the case of an object, or 37°C when measuring temperature in the human body.
  • the reference echo stretching factor s ref is measured using the exact same procedure, with a first step of collecting, a step of increasing, and a second step of collecting. The presence of such step applies to any measurements, which are mentioned hereinafter, when detailing the calibration technique.
  • the stretching factor s is always equal to zero, by definition, regardless of the value of 8T. As explained above, if one assumes that other parameters such as attenuation do not change with temperature, then the relation between the stretching factor s and the temperature T is a linear function.
  • 8T is unknown and attenuation and other parameter may vary with temperature, resulting in changes in 8T as temperature varies.
  • the linear relation no longer holds, and the relation between s and T becomes extremely complex.
  • Such experimental determination can be carried out by implementing the following operations: obtaining a calibration table, said calibration table being for instance obtained in a preliminary experiment, in which a calibration table is prepared by measuring the stretching factor as a function of temperature, for different temperatures, for example using samples in a water bath. storing this preliminary calibration data in a table s ca i(T) that contains the values s for the different temperatures that have been tested. measuring the temperature inside an object representative of the object on which the method is carried out. For instance, for an in-vivo organ, the representative object is an ex-vivo organ of the same type.
  • the method is easier to implement since no MRI device 10 is needed and the computation to be realized are simple.
  • Such knowledge of the temperature can be used to follow an ultrasound therapy, by notably monitoring the heating induced by the therapy. This could ensure a better efficiency of the therapy and a better fulfillment of the security requirements.
  • Each function can also designate a corresponding table providing with the values.
  • the table is stored in the memory and the calculator 20 reads the result value in the table instead of calculating the value.
  • the method further comprises a step of estimating the speed of sound c in the area 12 based on the measured temperature T.
  • the calculator 20 applies the function f on the measured temperature T.
  • the result of the function f is the speed of sound c in the area 12.
  • the method comprises a step of controlling the value of the temperature increment.
  • the calculator 20 controls whether or not the obtained value is within the expected range.
  • the fact that the human body has a temperature around 37°C implies that the echo signal stretching factor s should be comprised between 1.10 3 and -2.10’ 3 .
  • the controller When the obtained value is outside the expected range, the controller emits a warning.
  • Such warning may, for instance, consists in the emission of a specific sound.
  • Such warning indicates generally that the temperature has raised more or less than expected.
  • the present method for measuring can also be used for temperature decrease with the same formula, the increase having a negative value.
  • the step of increasing S52 is a step of modifying the temperature of the area 12 by a temperature change value, the temperature change value being a temperature decrease or a temperature increment.
  • a temperature modifying unit 16 is used.
  • Figure 4 Speed of sound vs. temperature in canine liver samples at 5 MHz.
  • the process is illustrated at various temperatures that have been chosen arbitrarily (20, 40, 60 and 90°C). At each temperature, the additional heat pulse increases temperature by the same amount ST (horizontal arrow). The resulting variation in speed of sound Ac is shown by the vertical arrow.
  • FIG. 5 Simulated pressure distribution in a canine liver model: (left) Pressure field of the HIFII heat source, (right) Corresponding temperature distribution, immediately after 135 ms heating, in a medium at 50°C.
  • Figure 6 (left) Density map for the pre-heating simulation, (center) Speed of sound map for the pre-heating simulation, and (right) for the post-heating simulation.
  • Figure 7 Illustration of the experimental setup.
  • the sample temperature was controlled by setting the water bath temperature at the desired level.
  • the therapy transducer was only used to deliver the small additional heat pulse. For temperatures up to 40°C, only one water tank was used. For temperatures 50°C and above, two water tanks were used, one to heat the samples (tank on the left), and one maintained at 37°C to perform the measurements (tank on the right), because the transducers would have been damaged otherwise. Images were acquired with a L7-4 transducer connected to a Verasonics Vantage system.
  • Figure 10 Post-processing steps of data acquired from experimental measurements. For the samples measured at 23, 40, 60 and 80°C. (upper part) The first row represents the delay maps of the 24th frame. The dashed black boxes represent the region used to estimate the undesirable deformations, (middle part) The second row represents the displacement maps after an average filter, and subtraction of the undesirable deformation, (dowb part) The last row represents the thermal strain maps after an average filter. The square represents the focus position where the thermal strains are extracted.
  • Figure 11 Spatially-averaged thermal strains (Frames 20-120) extracted from the map (figure 10) at the focus position, for the sample measured at 23°C.
  • the origin of time was set to frame #20, i.e. immediately after the end of the burst.
  • the red box indicates the frames where an averaging was applied.
  • the high variability in that time window is unexplained. It may be attributable to motion induced by radiation force during the HIFII exposure. After that transient regime (i.e. for t > 250 ms), the thermal strain shows lower variability, but it slowly decays towards zeros, resulting in a time-dependent bias. This behavior was attributed to cooling of the heat pulse.
  • Figure 13 Estimated temperature using the thermal strain as a function of temperature. The conversion from thermal strain to temperature was based on the thermal strains values already shown in figure 9. The linear fit (dashed line in figure 12) was used to convert from thermal strain values to temperature.
  • HIFII high intensity focused ultrasound
  • RF radiofrequency
  • MRI magnetic resonance imaging
  • thermal strain imaging is based on the detection of echo-shifts in response to thermal expansion and changes in speed of sound (SOS).
  • This variation of the signals measured before and after heating can be estimated in the frequency domain, from a shift of the central frequency, or in the time domain, with the estimation of a time-delay.
  • the initial temperature and speed of sound have to be known, and requires a calibration table for the tissues being investigated.
  • the estimation of the temperature is limited.
  • speed of sound and temperature seems to be parabolic ], such that the SOS reaches a peak around 50 - 70 °C, before to decrease at higher temperature.
  • This experimental section describes the method for determining a temperature based on thermal strain imaging, that seems to show a linear relation between thermal strain and temperature over a wider range of temperatures in simulation and tissue-mimicking material. Illustrating the potential of this new method to provide a convenient method for monitoring temperature changes before coagulation, during HIFII treatment and RF ablation.
  • the key to the new method is to change the temperature to be measured, temporarily and by a small amount.
  • a pulsed heat source designed to induce a local, reproducible, and temporary change of temperature 67 in the tissue being investigated (see figure 4).
  • the change in speed of sound Ac is approximately equal to 8T.dc/dT, where dc/dT is the local value of the partial derivative at temperature T..
  • thermal strain 5 induced by this pulsed heat source is also directly proportional to temperature, as in the following equation 4:
  • equation 5 can be written according to the form of equation 6, namely:
  • T c max is the maximum temperature reached by the parabola.
  • the above-described method makes no assumption on the initial temperature and does not require knowledge of the speed of sound.
  • the estimation of the temperature is then simply based on the measured thermal strain (Eq. 5).
  • Eq. 5 the measured thermal strain
  • the heat source is a HIFU transducer. It is driven by a short acoustic burst that lasts 150 milliseconds. The measurement is taken within milliseconds after the acoustic burst, so that thermal diffusion and perfusion can be neglected.
  • the method relies on the ability of the pulsed heat source to deliver a reproducible variation T.
  • the Applicant envisioned that using a HIFII transducer driven at a constant acoustic power and a constant exposure duration would achieve high reproducibility.
  • heat source simulations were performed to assess how reproducible this strategy is, especially when the background temperature changes.
  • imaging simulations were performed to assess the feasibility of the method.
  • the heat source was a bowl-shaped therapeutic transducer with a diameter of 30 mm, focused at 80 mm, and sinusoidal excitation with frequency 4.3 MHz, as apparent from table I hereinafter;
  • Two-dimensional (2D) ultrasound wave propagation simulations were first conducted at different temperatures, using the k-Wave MATLAB toolbox (http://www.k-wave.org).
  • the model geometry used for the simulations was composed of a homogeneous canine liver matrix (100 mm x 26 mm), and the material properties of the propagation medium were set at a density of 1050 kg/m 3 , and an attenuation coefficient of 0.7 dB/MHz/cm. To simplify the simulations, the attenuation coefficient was the same for each temperature studied, and considered as independent of temperature.
  • the simulations of heat diffusion in the homogenous medium were performed using the BioHeat Transfer Equation (BHTE) implemented in k-Wave. It had an initial homogeneous temperature distribution (from 20 to 90°C), a density of 1050 kg/m3, a thermal conductivity of 0.512 W/m/K, a specific heat of 3600 J/kg/K. To create a local increase of 2°C of the temperature at the focus position, the pressure at the surface of the transducer was set to 2.6 MPa, and the exposure duration was set to 135 ms.
  • BHTE BioHeat Transfer Equation
  • the transmit/receiver linear array of the acoustic imaging simulations represents a L7- 4 transducer (Verasonics Inc, Kirkland, WA, USA), configured with 64 elements to speed up computation of the simulations (element size: 0.25 mm).
  • the elements were excited individually with a broadband pulse (Frequency: 5 MHz, Bandwidth: 80%, Pressure at the surface: 1 MPa, Cycle: 2.5), focus at 25 mm.
  • the imaging transducer was placed opposite to the therapeutic transducer.
  • the figures present only a small portion of interest (315mm x 26 mm) within the simulation region (see figure 5). The focus of the therapeutic and imaging simulations was set at the same position.
  • the purpose was to create local changes in impedance to generate backscatter.
  • the speed of sound map (figure 6) was homogenous and set according to the temperature studied (from 20 to 90°C).
  • the speed of sound (figure 6) was calculated based on the temperature maps of the heat diffusion simulations (figure 5), using the fitted calibration table (figure 4).
  • the k-Wave raw data were processed with several post-processing steps.
  • the raw RF signals were beamformed (delay and sum) to get a single scan line, based on the transducer focus distance and apodization settings.
  • the loss of energy due to attenuation was compensated using time gain compensation, and a Gaussian filter was applied to reduce the noise outside the transmit frequency range.
  • the background temperature of the medium was homogeneous. For each temperature level, a total of 60 simulations were performed to evaluate the variability induced by the echo-shift estimation during the post-processing.
  • the speckle pattern was varied by changing the density distribution, as explained previously. No random noise was added.
  • the time delay between the pre- and post-heating data was calculated using a crosscorrelation technique, and converted to a displacement.
  • a two-steps implementation of the zero-lag cross correlation method was used.
  • a coarse estimation was obtained, using two adjacent windows (2 mm length each, 0% overlap) centered around the focus position. This coarse estimation is an integer number of temporal samples. It was then used to shift and “re-align” the windows.
  • the residual lag ( ⁇ 0.5 time sample) was estimated in the second step, using cross-correlation again, with the same window length and overlap, but only testing for three delays (-1 ; 0; and +1 time sample).
  • AMC amplitude modulation correction
  • Two measurement points were used to serve as auto-calibration, assuming a linear relation between thermal strain and temperature. The first one was at initial temperature (20°C) and the second one was at the maximum temperature reached by the parabola (60 °C) (see figure 4). The other measurement points (thermal strain from 30 to 90°C) were then converted to temperature using this calibration, and compared to ground truth values.
  • a L7-4 linear array transducer connected to a Verasonics Vantage system (Vantage 256, Verasonics., Kirkland, WA, USA) was used to acquire the beamformed RF signals pre- and post-heating.
  • the details of the imaging transducer are described on Table I.
  • a custom therapeutic transducer, with a PZT ceramic from Boston piezo-optics (Boston piezo-optics, Bellingham, MA, USA), connected to a generator and an amplifier were used to create a burst (time: 150ms, power: 20W) and increase locally the temperature at the focus position.
  • the details of the therapeutic transducer are reported on Table 1.
  • the imaging transducer and the therapy transducer were positioned opposite to each other, and the beam imaging technique was used to align them.
  • the beam imaging technique consists in using a conventional ultrasound scanner with no transmitted signal, and simultaneous reception on all elements of the probe, to capture and display the therapeutic ultrasound beam, in real time.
  • a homogeneous phantom cylinder of Zerdine, tissuemimicking material, (CIRS Inc, Norfolk, VA) was placed at the focal positions of both transducers and imaged in a water tank at different temperatures (23, 25, 30, and 40°C).
  • the two samples were immerged and heated in a second water tank, then they were both transferred to the first water tank. 30-60 seconds of waiting were carried out to let the water waves settle down, then both measurements (ultrasound and thermocouple) were recorded simultaneously.
  • the first method (one bath) was preferred because it induces a highly controlled and homogeneous temperature distribution within the samples. But for temperatures >40°C, the imaging probe would have been damaged. This is why the second bath was added, allowing to test samples at 50°C and above, at the cost of a less controlled temperature.
  • the time delays were calculated between the image acquired before the burst (image 5) and the images following the burst (images 20 to 120), by two-step correlation and amplitude modulation corrections, exactly as described in the post-processing simulation section [15,16,27],
  • the cross-correlations were used along the RF signals, starting 1-mm from the imaging transducer, with a window length of 1 mm and no spacing between the non-overlapping windows, creating a map of time delays.
  • This time delay map was converted to a displacement map, then filtered with a 3x3 pixels (0.9*3 mm) moving average filter to reduce the noise. Undesirable sample displacement and deformation were present.
  • the thermal strain map was obtained using the gradient of the filtered displacement map, and a second moving average filter (3x3 pixels, 0.9*3 mm) was applied to further reduce the noise. The thermal strains at the focus were then extracted from the map.
  • Two measurement points were used to serve as auto-calibration, assuming a linear relation between thermal strain and temperature. The first one was at initial temperature (23°C) and the second one was the thermal strain equal to zero (estimated to be at around 60 °C). The other measurement points (thermal strain from 30 to 88°C) were then converted to temperature using this calibration, and compared to values obtained from thermocouple measurements carried out on the phantom replica.
  • the dashed line is a linear fit based on thermal strain values at 20 and 60°C. This linear relation will be used in figure 9 to convert the other thermal strain values (i.e. those obtained from 30 to 90°C) to temperature.
  • Figure 9 shows the corresponding temperatures, estimated from the thermal strains and the linear fit.
  • Figure 10 shows the time delay maps computed on the 24th Frame, at 23, 40, 60 and 80°C.
  • Figure 10 central one shows the corresponding displacement maps, after smoothing and subtraction of the undesirable deformation (black dashed boxes).
  • Figure 10 (right) shows the corresponding thermal strain images after smoothing, along with the focus position where the thermal strains were extracted and spatially averaged (red square). The process was repeated for every image frame, resulting in a time-varying spatially-averaged thermal strain value (figure 11).
  • the final (spatially-averaged and temporally averaged) thermal strain value for the corresponding experiment was calculated by averaging over the 20 first frames after the burst (red square in figure 11). After that transient regime (i.e. for t > 250 ms), the thermal strain shows lower variability, but it slowly decays towards zeros, resulting in a time-dependent bias. This behavior was attributed to cooling of the heat pulse.
  • Figure 12 shows the thermal strains for temperatures from 23 to 88 °C, along with a linear fit (dashed line) based on thermal strain values at 23°C (initial temperature) and 53.5°C (zero-thermal strain).
  • Figure 13 shows the corresponding temperature estimates in the range 30-88 °C.
  • Thermal strains induced by a constant acoustic power bust was estimated in canine liver simulation and experimentally in Zerdine phantom samples, at different temperatures.
  • the thermal strains were derived from echo-shifts on beamformed RF data, generated by a small and local change in temperature.
  • the thermal strain was then directly converted to estimate the temperature, using an auto-calibration table built from the thermal strain at initial temperature (20°C) and the zero-thermal strain (around 60°C). This auto-calibration enables us to estimate the temperature without knowing the ST and a possible use of the new method in-vivo.
  • the results, both in simulations and in experiments demonstrated the linear relation between thermal stain and temperature, and the possible estimation of temperature over a wider range of temperatures compared to previous studies.
  • the reproducibility of the heat source was obtained by using the same exposure duration, same acoustic power, by positioning the phantom sample at the same distance of the therapeutic transducer, and by assuming constant attenuation, acoustic intensity, specific heat, and density, for every measurement, as shown for some tissues in previous studies.
  • these parameters are likely to vary with temperature, especially when coagulation occurs.
  • the Applicant anticipates that it will not be possible to measure the temperature in regions where coagulation has occurred. However, it may not be an issue. Indeed, in most therapeutic applications, detecting coagulation is more important than measuring the temperature inside the coagulated area. Interestingly, the temperature Tc_max, for which the thermal strain becomes zero, also corresponds to the temperature at which coagulation occurs almost instantly, regardless of the duration of exposure. As a consequence, to ensure that coagulation has occurred, the Applicant has anticipated that it might be sufficient to deposit heat locally until the local thermal strain becomes zero. However, further work is needed to confirm this hypothesis.
  • the Applicant has anticipated that the present method can only be used in the temperature range wherein the attenuation does not change significant. This temperature range depends on the type of tissues. It has been showed non-constant acoustic attenuation for canine muscle, kidney, and prostate in the temperature range from 25°C to 95 °C . The case of canine liver was unclear, with nearly constant attenuation a first study, and varying attenuation reported in a posterior study.
  • radiofrequency (RF) ablation is the most common procedure for liver cancer and liver metastases.
  • RF radiofrequency
  • the treatment is limited by the lack of feedback regarding the temperature distribution. If attenuation turns out to be constant, the present method has the potential to provide temperature feedback during the procedure.
  • the reproducibility of the heat source can also be affected by the heterogeneity of the biological tissue being treated or by different biological tissues present on the propagation path.
  • the local heating ST cannot be predicted.
  • the two-point calibration method is still applicable.
  • Figure 14 Photos of the reference renal cortex sample at (a) room temperature (a) 20°C, (b) 40°C, (c) 60°C, and (d) 88°C. All four images are displayed with the same scale. The shrinkage of the sample is clearly visible.
  • Figure 16 Thermal strain in canine liver (C1 curve), muscle (C2 curve), kidney (C3 curve), and prostate (C4 curve) simulations for scenario #2, i.e. when the attenuation changes everywhere in the tissue, including on the propagation path of the acoustic power burst.
  • Figure 17 Post-processing steps of data acquired from experimental measurements. For one pig liver sample measured at 25, 60 and 80°C.
  • the first row represents the displacement maps (in meters) of the 24th frame after an average filter, and subtraction of the undesirable deformation (the dashed black boxes represent the region used to estimate the undesirable deformations)
  • the second row represents the thermal strain maps after spatial averaging.
  • the thermal strain is dimensionless. It is computed as the spatial gradient in the vertical direction of the displacement maps. The red square represents the focus position where the thermal strains are extracted.
  • the temperature was varied between 25 and 95 °C.
  • the intensity at the focus was calculated as described below.
  • the temperature increase induced by the acoustic burst was calculated using equation (3).
  • the corresponding thermal strain at the focus was calculated using equation (2).
  • the Applicant considers a focused treatment, wherein only the focus is being treated, while heating in neighboring tissues is negligible.
  • the Appplicant considers the case of a hypothetical treatment that affects uniformly a large area, encompassing the propagation path between the therapeutic transducer and its focus. This scenario also corresponds to the water bath experiment, described in the next section, where the entire sample was heated uniformly. Note that in existing therapeutic applications, the reality is expected to be somewhere between these two extreme cases. Hence these two cases are given as an illustration of the potential of the method, not as an actual prediction.
  • the details of the simulations are given below.
  • the simulation script was developed on Matlab (MATLAB R2018a, The MathWorks, Natick, MA, USA) and performed at different temperatures (from 25°C to 95°C).
  • the model geometry used for the simulations consisted of a homogeneous matrix representing canine liver, kidney, muscle, or prostate tissues.
  • the material properties of the propagation medium were set at a density of 1050 kg/m3 and a specific heat of 3600 J/kg/K.
  • the speed of sounds and attenuations at 4MHz were interpolated from the calibration tables of a parabolic dependence between speed of sound and temperature in all the tissues they investigated. Therefore, a second-order polynomial fit was applied over the speed of sound.
  • a is the attenuation
  • p the density
  • C the specific heat of the medium
  • I the acoustic intensity at the focus
  • t the duration of the acoustic burst.
  • the Appliant adjusted the attenuation only at the focus position.
  • the acoustic intensity, / 0 was then calculated using the following equation: where p is the peak acoustic pressure at the focus.
  • p is the peak acoustic pressure at the focus.
  • I was then calculated using the following equation:
  • I / o .exp ( - az) (5) where a is the attenuation, z the position of the focus, and I o the acoustic intensity at the surface of the transducer.
  • thermocouple would provide that information.
  • the Applicant wanted to avoid temperature measurement errors due to the viscous heating effect.
  • the Applicant decided not to use a control sample to simplify the experiment, which involves transferring samples from one tank to another.
  • the experiment setup for estimating temperature in biological tissues is similar to that used in our previous study in Zerdine phantom.
  • a L7-4 linear array transducer connected to a Verasonics Vantage system (Vantage 256, Verasonics., Kirkland, WA, USA) was used to acquire the beamformed RF signals pre- and post-heating.
  • An Olympus transducer (Accuscan A380, Olympus IMS, Japan) connected to a generator and an amplifier were used to create an acoustic burst (time: 100ms, power: 64W) and locally increase the temperature at the focus position.
  • the details of the Olympus transducer are reported on Table III.
  • the imaging transducer and the therapy transducer were positioned opposite to each other. Active beam imaging was used to visualize the sample and the beam simultaneously, in real time, and to align them.
  • the protocol was standardized based on the reference measurements, so the samples were kept for 25 min in the tank before acquisition. It allowed the samples to reach the temperature studied. For temperature 50°C and above, we waited for 30-60 seconds to let the water waves settle down.
  • a 12-angle plane wave imaging sequence was used to acquire 120 frames (total duration 1.5s). The time between angle was 500 ps, and the resulting frame rate was 80 fps.
  • the thermal strains at the focus position were computed as explained before.
  • the time delays were calculated between the image acquired before the burst and the images following the burst.
  • the time delay map was created using a two-steps implementation of the zero-lag cross-correlation method along the RF signals (window length: 1-mm, 0% overlap).
  • the two-steps cross-correlation was composed of a coarse estimation and a residual estimation.
  • the estimation accuracy was improved by applying amplitude modulation corrections.
  • the time delay map was converted to a displacement map and filtered to reduce the noise (0.9*3 mm moving average filter).
  • the undesirable deformation from water movement and HIFII exposure were estimated and subtracted from the displacement map, based on windows positioned on both sides of the focus, at 4.5XX mm distance. Then, the thermal strain map was obtained using the gradient of the filtered displacement map and filtered a second time to reduce the noise (0.9*3 mm moving average filter). Finally, the thermal strains were extracted at the focus position.
  • Temperatures were estimated using a calibration table built from two measurement points, assuming a linear relation between thermal strain and temperature. The first point was the thermal strain at initial temperature (25°C). The second point was the zero-crossing of the thermal strain (i.e. the temperature for which the thermal strain is equal to zero), estimated from the experimental data. The estimated temperatures were then compared to the values obtained from thermocouple measurements conducted in the reference samples.
  • Figure 15 shows the thermal strain vs. temperature in the canine liver, muscle, kidney, and prostate simulations. This figure is for scenario #1 , wherein attenuation was changed only at the focus position.
  • Figure 16 shows the thermal strain vs. temperature in the canine liver, muscle, kidney, and prostate simulations. This figure corresponds to scenario #2, wherein the attenuation was changed at the focus position and on the propagation path of the acoustic power burst.
  • Table IV Time and size of reference samples at each temperature step.
  • Figure 17a shows the displacement delay maps computed on the 24th Frame, at 25, 60 and 80°C, after smoothing and subtraction of the undesirable deformation (black dashed boxes).
  • Figure 17b shows the corresponding thermal strain images after smoothing, along with the focus position where the thermal strains were extracted and spatially averaged (10 x 0.6 mm) (square).
  • Figure 18 shows the temporal average of the thermal strain vs. temperature in pig liver, shoulder muscle, renal cortex, and cardiac muscle from 25 to 88°C. Temporal averaging was performed over all 120 frames, corresponding to a time window of 1.5 s.
  • the dashed line is a linear fit based on thermal strain values at 25°C (initial temperature) and at 60°C (zero-thermal strain). This linear relation is used in Figure 19 to convert the other thermal strain values (i.e. those obtained from 30 to 88°C) to temperature.
  • Thermal strains induced by a constant acoustic power burst were simulated in canine liver, muscle, kidney, and prostate (figure 15). In a first case, the attenuation was assumed to change only at the focus position. The Applicant observed a linear relation between thermal strain and temperature up to 90 °C in liver, but not in muscle, kidney, and prostate. In these tissues, the slope increased for temperatures above 50-60°C, i.e. when coagulation occurred. The relation is no longer linear, but it is still possible to use the data to create an “attenuation-compensated” calibration table to convert thermal strain to temperature over the entire range from 25 up to 95°C.
  • Thermal strains and temperatures were also experimentally estimated in pig liver, shoulder muscle, renal cortex, and cardiac muscle (from 25 to 88°C). The results demonstrated a linear relation between thermal strain and temperature from 25°C to 80°C in liver tissues. In renal cortex samples, the linear relation was demonstrated from 25°C to 70°C. In shoulder and cardiac muscle tissues, the relation between thermal strain and temperature was linear from 25°C to 60°C. These experimental results are comparable to scenario #2 in the simulation results.
  • thermometry method The temperature range wherein this thermometry method can be used depends on the type of tissue treated. Indeed, this thermometry method can only be applied when the attenuation does not change significantly.
  • thermometry method is promising, and future research should focus on monitoring thermal treatments, such as HIFII and RF ablation in vivo.

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Abstract

La présente invention concerne le domaine de la mesure de la température d'une zone à l'intérieur d'un objet. La norme de référence pour mesurer les changements de température pendant les thérapies thermiques est l'imagerie par résonance magnétique, mais les scanners à ultrasons semblent être une solution de remplacement intéressante, car ce sont des appareils portables. Plusieurs procédés ont été développés pour estimer la température à l'aide d'ultrasons, en mesurant la contrainte thermique, le coefficient d'atténuation acoustique ou l'énergie de rétrodiffusion, mais les procédés utilisés doivent encore être optimisés. C'est la raison pour laquelle les inventeurs ont élaboré un procédé plus sensible. Ainsi, la présente invention concerne un procédé de mesure de la température d'une zone à l'intérieur d'un objet impliquant plusieurs échos ultrasonores obtenus à différents moments.
PCT/EP2025/063509 2024-05-16 2025-05-16 Procédé de mesure de la température d'une zone à l'intérieur d'un objet et dispositifs associés Pending WO2025238199A1 (fr)

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Citations (2)

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US20190262074A1 (en) * 2018-02-27 2019-08-29 Walter Kusumoto Ultrasound thermometry for esophageal or other tissue protection during ablation
EP3240611B1 (fr) * 2014-12-30 2020-09-30 Koninklijke Philips N.V. Étalonnage à ultrasons de contrainte thermique par rapport à la température spécifique à un patient

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Publication number Priority date Publication date Assignee Title
EP3240611B1 (fr) * 2014-12-30 2020-09-30 Koninklijke Philips N.V. Étalonnage à ultrasons de contrainte thermique par rapport à la température spécifique à un patient
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