WO2022166982A1 - 一种消融计算方法及消融计算系统 - Google Patents

一种消融计算方法及消融计算系统 Download PDF

Info

Publication number
WO2022166982A1
WO2022166982A1 PCT/CN2022/075489 CN2022075489W WO2022166982A1 WO 2022166982 A1 WO2022166982 A1 WO 2022166982A1 CN 2022075489 W CN2022075489 W CN 2022075489W WO 2022166982 A1 WO2022166982 A1 WO 2022166982A1
Authority
WO
WIPO (PCT)
Prior art keywords
temperature
map
phase difference
ablation
phase
Prior art date
Application number
PCT/CN2022/075489
Other languages
English (en)
French (fr)
Inventor
刘文博
韩萌
旷雅唯
吴朝
Original Assignee
华科精准（北京）医疗科技有限公司
Priority date (The priority date 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 date listed.)
Filing date
Publication date
Application filed by 华科精准（北京）医疗科技有限公司 filed Critical 华科精准（北京）医疗科技有限公司
Priority to CN202280006671.1A priority Critical patent/CN116324459A/zh
Publication of WO2022166982A1 publication Critical patent/WO2022166982A1/zh

Links

Images

Classifications

    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/44—Arrangements or instruments for measuring magnetic variables involving magnetic resonance using nuclear magnetic resonance [NMR]
    • G01R33/48—NMR imaging systems
    • G01R33/58—Calibration of imaging systems, e.g. using test probes, Phantoms; Calibration objects or fiducial markers such as active or passive RF coils surrounding an MR active material
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/44—Arrangements or instruments for measuring magnetic variables involving magnetic resonance using nuclear magnetic resonance [NMR]
    • G01R33/48—NMR imaging systems
    • G01R33/54—Signal processing systems, e.g. using pulse sequences ; Generation or control of pulse sequences; Operator console
    • G01R33/56—Image enhancement or correction, e.g. subtraction or averaging techniques, e.g. improvement of signal-to-noise ratio and resolution
    • G01R33/561—Image enhancement or correction, e.g. subtraction or averaging techniques, e.g. improvement of signal-to-noise ratio and resolution by reduction of the scanning time, i.e. fast acquiring systems, e.g. using echo-planar pulse sequences
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01R—MEASURING ELECTRIC VARIABLES; MEASURING MAGNETIC VARIABLES
    • G01R33/00—Arrangements or instruments for measuring magnetic variables
    • G01R33/20—Arrangements or instruments for measuring magnetic variables involving magnetic resonance
    • G01R33/44—Arrangements or instruments for measuring magnetic variables involving magnetic resonance using nuclear magnetic resonance [NMR]
    • G01R33/48—NMR imaging systems
    • G01R33/54—Signal processing systems, e.g. using pulse sequences ; Generation or control of pulse sequences; Operator console
    • G01R33/56—Image enhancement or correction, e.g. subtraction or averaging techniques, e.g. improvement of signal-to-noise ratio and resolution
    • G01R33/561—Image enhancement or correction, e.g. subtraction or averaging techniques, e.g. improvement of signal-to-noise ratio and resolution by reduction of the scanning time, i.e. fast acquiring systems, e.g. using echo-planar pulse sequences
    • G01R33/5615—Echo train techniques involving acquiring plural, differently encoded, echo signals after one RF excitation, e.g. using gradient refocusing in echo planar imaging [EPI], RF refocusing in rapid acquisition with relaxation enhancement [RARE] or using both RF and gradient refocusing in gradient and spin echo imaging [GRASE]
    • Y—GENERAL TAGGING OF NEW TECHNOLOGICAL DEVELOPMENTS; GENERAL TAGGING OF CROSS-SECTIONAL TECHNOLOGIES SPANNING OVER SEVERAL SECTIONS OF THE IPC; TECHNICAL SUBJECTS COVERED BY FORMER USPC CROSS-REFERENCE ART COLLECTIONS [XRACs] AND DIGESTS
    • Y02—TECHNOLOGIES OR APPLICATIONS FOR MITIGATION OR ADAPTATION AGAINST CLIMATE CHANGE
    • Y02A—TECHNOLOGIES FOR ADAPTATION TO CLIMATE CHANGE
    • Y02A90/00—Technologies having an indirect contribution to adaptation to climate change
    • Y02A90/30—Assessment of water resources

Definitions

  • the present application relates to the field of medical devices, and more particularly, to an ablation calculation method and an ablation calculation system.
  • Magnetic resonance temperature imaging Magnetic Resonance Temperature Imaging, MRTI
  • MRTI Magnetic Resonance Temperature Imaging
  • One of the current magnetic resonance temperature imaging methods is the temperature measurement method based on the proton resonance frequency (Proton Resonance Frequency, PRF) displacement.
  • the distribution of tissue susceptibility is uneven, and objective environmental factors such as tissue movement caused by respiration/blood flow pulsation have a greater impact.
  • Common errors caused include phase unwrapping errors, errors caused by rapid changes in magnetic susceptibility, and errors caused by motion, etc. It is easy to cause the problem that the temperature map finally obtained is quite different from the actual temperature, which makes the temperature map lose its reference meaning.
  • the present application provides an ablation calculation method and a related ablation calculation system.
  • an ablation calculation method comprising:
  • the ablation (for each pixel) is calculated from the temperature map.
  • the temperature difference map is obtained as follows: the phase map at any time is subtracted from the phase map at the reference time to obtain the phase difference map at this time, and the phase difference map corresponding to at least one echo time is selected as With reference to the phase difference map, calibrate the phase difference maps to be calibrated corresponding to other echo times based on the reference phase difference map, and obtain the calibrated phase difference map.
  • the echo time corresponding to the reference phase difference map is smaller than the calibrated phase difference map.
  • the temperature difference map at this moment is calculated using the reference phase difference map and the calibrated phase difference map.
  • the echo time corresponding to at least one reference phase difference map does not exceed one of the following: 18ms, 17ms, 16ms, 15ms, 14ms, 13ms, 12ms, 11ms, 10ms, 9ms, 8ms, 7ms, 6ms, 5ms or 4ms.
  • the phase difference map to be calibrated is unwrapped according to the phase periodicity, and the calibrated phase difference map is obtained.
  • the ablation calculation method of the present invention further includes the step of eliminating the phase drift caused by the magnetic resonance system (for example, the B0 drift error); further, in the step of eliminating the phase drift caused by the magnetic resonance system, a number of physical temperature stable An area of unchanged and uniform tissue is used as a thermal reference point, and the phase shift is performed by subtracting the average phase difference of the thermal reference point from each phase difference image or by subtracting the average temperature change of the thermal reference point from the temperature change map Correction.
  • the phase shift is performed by subtracting the average phase difference of the thermal reference point from each phase difference image or by subtracting the average temperature change of the thermal reference point from the temperature change map Correction.
  • the ablation calculation method of the present invention further includes the step of correcting the error caused by the magnetic susceptibility, and the step of correcting the magnetic susceptibility is performed on the phase difference map or the temperature difference map, and the step includes:
  • the steps to perform susceptibility correction on the temperature difference map include:
  • a first temperature map is obtained according to the reference phase difference map, and a corresponding second temperature map is obtained according to the calibrated phase difference map,
  • the steps for performing susceptibility correction on the phase difference map include:
  • the ablation calculation method of the present invention further includes the step of correcting motion-induced phase errors by linear least squares fitting at each pixel using at least two sets of phase maps corresponding to different echo times. The resulting phase error is removed.
  • the ablation calculation method of the present invention further includes the step of obtaining a weighted temperature map, where the weighted temperature map is obtained by weighting a temperature map obtained by at least one reference phase difference map and a temperature map obtained by at least one calibrated phase difference map.
  • the echo time corresponding to at least one calibrated phase difference map is not less than 20ms, 19ms, 18ms, 17ms, 16ms, 15ms, 14ms, 13ms or 12ms.
  • the pixel ablation is calculated using the following formula:
  • E a is the activation energy
  • A is the frequency factor
  • R is the universal gas constant
  • T( ⁇ ) is a function of temperature (°C) and time ⁇
  • t is the current time
  • a set threshold eg 1
  • a storage medium is provided, and program codes are stored on the storage medium, and when the program codes are executed, the ablation calculation method of the present invention is implemented.
  • an ablation calculation system which includes an ablation calculation module, and the ablation calculation module can execute the ablation calculation method of the present invention. Further, the ablation calculation module may further include:
  • an information acquisition module which is used for receiving or acquiring magnetic resonance information, the magnetic resonance information at least including a phase map corresponding to the echo time obtained by scanning the target to be measured by using a gradient echo sequence containing i different echo times,
  • the i is a positive integer greater than or equal to 2;
  • a temperature difference calculation module which is used to select at least two sets of phase maps corresponding to different echo times to obtain corresponding temperature difference maps
  • a temperature map calculation module which is used to obtain a temperature map according to the temperature difference map, and the temperature map may be a weighted temperature map;
  • an ablation calculation module which is used for calculating the ablation condition of each pixel according to the temperature map.
  • phase drift correction module for performing the step of eliminating phase drift caused by the magnetic resonance system
  • a magnetic susceptibility error correction module for performing the steps of correcting errors caused by magnetic susceptibility
  • a motion error correction module for performing the steps of correcting motion-induced phase errors.
  • a laser interstitial hyperthermia instrument which includes: a host, a laser ablation device, and an optical fiber assembly, wherein the host includes a processor and is loaded with program codes, which are implemented when the program codes are executed.
  • the ablation calculation method of the present invention is provided, which includes: a host, a laser ablation device, and an optical fiber assembly, wherein the host includes a processor and is loaded with program codes, which are implemented when the program codes are executed.
  • the method of the present invention uses a non-retrospective algorithm or an iterative algorithm, which has a small amount of computation, saves computing time, and can quickly obtain temperature and ablation results.
  • Fig. 1 is the amplitude, phase map and temperature map of the magnetic resonance data obtained in the in vitro and in vivo environment of the prior art
  • FIG. 2 is a schematic flowchart of an ablation calculation method provided by an embodiment of the present application.
  • FIG. 3 is an acquired phase diagram, a phase difference diagram and a temperature diagram provided by an embodiment of the present application.
  • FIG. 5 is a schematic diagram of an experimental device provided by an embodiment of the application.
  • FIG. 6 is a partial magnification of a temperature map of laser interstitial hyperthermia in an in vitro pork experiment provided by an embodiment of the present application;
  • Fig. 7 is a schematic diagram of the change of temperature with time during the experiment of gel tissue mimic (Fig. 7(a)) or isolated pork (Fig. 7(b)) according to an embodiment of the present application;
  • FIG. 8 is a representative temperature map of dog O1 in an in vivo experiment provided by an embodiment of the application.
  • FIG. 9 provides laser interstitial hyperthermia ablation results for 3 representative dogs (dogs 01-03) according to an embodiment of the present application.
  • FIG. 10 is a schematic diagram of the ablation estimation predicted by the ablation calculation method for the post-ablation T2w image and all the other six cases (dogs 04-09) of different ablation laser doses provided by an embodiment of the application.
  • Magnetic resonance thermography can guide a variety of energy-delivery treatments, such as laser interstitial hyperthermia, focused ultrasound therapy, and radiofrequency ablation, to monitor target tissue temperature and ablation effects.
  • the present invention uses magnetic resonance temperature imaging-guided laser interstitial hyperthermia as an example to illustrate the method of the present invention.
  • Magnetic resonance temperature imaging-guided laser interstitial hyperthermia is a minimally invasive treatment method, which is a method for the treatment of surgically challenging patients. Tumors by location (anatomical or functional) create new options. This method can rapidly coagulate tissue and induce tumor cell necrosis through protein denaturation by applying a temperature of 50-80°C or higher for tens of seconds.
  • laser interstitial hyperthermia can target tumors more precisely, reduce discomfort and infection risk, and shorten patient hospital stays.
  • concurrent magnetic resonance thermography plays an important role for more efficient ablation of tumor cells and better protection of healthy surrounding cells and key structures.
  • Most laser interstitial hyperthermia ablation procedures rely on thermometry based on proton resonance frequency shift.
  • the existing magnetic resonance temperature imaging method is greatly affected by factors such as the environment, which is likely to cause a problem that the temperature map finally obtained is greatly different from the actual temperature.
  • the inventor found through research that the main sources of errors in the acquired temperature map were phase errors caused by unwinding and dislocation, magnetic susceptibility errors and phase errors caused by motion.
  • the magnetic susceptibility causes a reduction in image amplitude and a corresponding error in the image phase, which corrupts the reconstructed temperature map in and around the heating center.
  • Errors in reconstructing the temperature map can lead to erroneous estimation of the ablation area, which can lead to changes in treatment efficacy and thermal damage to critical tissues. Therefore, accurate temperature imaging is critical for the efficacy and safety of laser interstitial hyperthermia, especially when laser interstitial hyperthermia is applied to tight ablation regions in brain tissue.
  • thermometry method based on the proton resonance frequency shift is based on the fact that the resonance frequency of hydrogen protons varies with the temperature in water molecules.
  • the change in the local magnetic field with temperature can be described as:
  • ⁇ is the proton resonance frequency coefficient that varies with temperature, which is taken as 0.008-0.015ppm/°C here.
  • the corresponding resonant frequency change of water protons affected by temperature can be expressed as:
  • ⁇ T represents the temperature change
  • ⁇ f represents the resonant frequency change
  • ⁇ represents the gyromagnetic ratio
  • B 0 represents the static magnetic field strength
  • Gradient echo sequences are sequences used in thermometry based on proton resonance frequency shifts. According to formula (3), the longer the gradient echo sequence is, the larger the phase difference may be caused by the same temperature change, indicating that higher temperature sensitivity can be obtained.
  • CSF Cerebrospinal Fluid
  • thermograms measured by thermometry based on proton resonance frequency shifts due to the movement of cerebrospinal fluid in the brain.
  • the magnitude and phase signals of the CSF are often altered on pulsed gradient echo sequences by the normal dynamic motion of the CSF, which can confound temperature estimates.
  • CSF motion can also cause pixels in and around the ventricle to shift, causing errors in phase contrast images.
  • the in vivo temperature map shows pseudo-hyperthermia in the third ventricle due to cerebrospinal fluid movement. Temperature errors are more pronounced on pulsed gradient echo sequences with shorter echo times because they are less tolerant of the intensity of the phase shift introduced by the cerebrospinal fluid flow in (3).
  • ⁇ 0 represents the local magnetic field change caused by magnetic susceptibility
  • a heated center with a sharp temperature change shows severe signal loss on the order of longer echo times.
  • the intra-voxel spin phase shift is caused by local magnetic field inhomogeneities caused by temperature and susceptibility changes.
  • the magnetization artifact caused by laser heating is an important cause of the error.
  • phase errors around the heating center translate into pseudo-low temperatures on magnetic resonance thermography.
  • gradient pulse sequences with as short an echo time can provide better temperature sensitivity and signal-to-noise ratio, which is the current dilemma.
  • an embodiment of the present application provides an ablation calculation method, which includes:
  • the ablation of each pixel is calculated according to the temperature map.
  • the ablation calculation method obtains i groups of phase maps based on gradient echo sequences containing i different echo times, selects at least two sets of phase maps corresponding to different echo times to obtain corresponding phase difference maps, and according to the temperature difference map Get a temperature map.
  • the inventor found through research that the echo time of the gradient echo sequence is proportional to the size of the magnetic susceptibility artifact, so the phase map obtained by the gradient echo sequence corresponding to the smaller echo time is subject to the magnetization caused by heating.
  • the influence of the rate change is minimal, and the image data still maintains the correct phase, so the temperature map can be obtained based on the i groups of phase maps and phase difference maps obtained from gradient echo sequences containing i different echo times to reduce the final
  • the error of the temperature map is for the purpose of improving the accuracy of the obtained temperature map.
  • the ablation calculation method is not a retrospective algorithm or an iterative algorithm, and the calculation amount is small, and an almost real-time temperature map can be provided, which has high reference significance.
  • the embodiment of the present application provides an ablation calculation method, as shown in FIG. 2 , including:
  • S101 use a gradient echo sequence containing i different echo times to scan the target to be measured, and obtain i groups of phase maps corresponding to the echo times, where i is a positive integer greater than or equal to 2;
  • the minimum value and the maximum value of the echo time in the gradient echo sequence can be determined according to actual requirements.
  • the minimum value of the echo time in the gradient echo sequence can be the minimum value that can be obtained by the magnetic resonance temperature imaging device, and the maximum value of the echo time in the gradient echo sequence generally does not exceed the echo of the object to be measured.
  • the upper limit of the value range of time For example, in the case of head imaging, the value range of the echo time of the optional gradient time sequence is 3-30ms, and the specific value of the echo time included in the gradient echo sequence should be within the range of the echo time. within the value range.
  • the acquired phase diagram is shown in FIG. 3( a ).
  • the phase difference map is calculated by a complex subtraction procedure. Complex subtraction avoids problematic phase wrapping.
  • the acquisition of gradient echo sequence information containing i different echo times can be read or received from a server or other storage device, or it can be acquired in real time according to the settings of the staff.
  • the specific method is not limited, and it depends on the actual situation.
  • S102 Select at least two sets of phase maps corresponding to different echo times to obtain corresponding temperature difference maps
  • a drift correction of the static magnetic field intensity of the phase map and the phase difference map may also be included, so as to eliminate errors caused by the static magnetic field intensity.
  • S104 Calculate the ablation situation of each pixel according to the temperature map, and the ablation may use various methods for calculating ablation based on temperature and time parameters, for example, using the following formula:
  • E a is the activation energy
  • A is the frequency factor
  • R is the universal gas constant
  • T( ⁇ ) is a function of temperature (°C) and time ⁇
  • t is the current time
  • a set threshold eg 1
  • the specific steps of obtaining the temperature difference map include:
  • phase difference map is the time before the energy (such as thermal energy, light energy, radiofrequency ablation, cryoablation) is transmitted to the target tissue.
  • energy such as thermal energy, light energy, radiofrequency ablation, cryoablation
  • the phase difference map is calibrated to obtain a calibrated phase difference map, and the echo time corresponding to the reference phase difference map is smaller than the echo time corresponding to the calibrated phase difference map;
  • the temperature difference map at this moment is calculated using the reference phase difference map and the calibrated phase difference map.
  • the value of the echo time of the reference phase difference map is less than or equal to 18ms, preferably, the value of the echo time of the reference phase difference map does not exceed 17ms, 16ms, 15ms, 14ms, 13ms, 12ms, 11ms, 10ms, 9ms, 8ms or 7ms, more preferably, the value of the echo time of the reference phase difference map does not exceed 6ms, 5ms or 4ms.
  • the phase map corresponding to the minimum echo time in the gradient echo sequence can be used as the reference phase map, and the reference phase difference map can be obtained based on the reference phase map, so as to minimize the heating caused by the phase map as much as possible.
  • the effect of magnetic susceptibility changes.
  • the minimum echo time in the gradient echo sequence may be the minimum value that the magnetic resonance temperature imaging device can take.
  • the phase difference map corresponding to at least one echo time at this moment is selected as the reference phase difference map, and the calibration of the phase difference maps corresponding to other echo times includes the following steps:
  • the ablation calculation method further includes:
  • S105 the step of eliminating the phase drift caused by the magnetic resonance system, the step of eliminating the phase drift caused by the magnetic resonance system is performed on the phase difference map or the temperature difference map.
  • the steps to remove phase drift caused by the magnetic resonance system on the phase difference map include:
  • the steps to cancel the phase drift caused by the magnetic resonance system on the temperature difference map include:
  • Correction is made by subtracting the average temperature difference at any of the thermal reference points from the temperature difference map.
  • the ablation calculation method further includes:
  • S106 the step of magnetic susceptibility correction, the step of magnetic susceptibility correction is performed on the phase difference map or the temperature difference map;
  • the steps to perform susceptibility correction on the temperature difference map include:
  • the temperature value of the first temperature map in the pixel can be used to replace the temperature value of the second temperature map, or the The temperature value of the adjacent pixel in the second temperature map replaces the temperature value of the pixel in the second temperature map, or an approximate temperature is fitted based on the temperature value of the adjacent pixel and the temperature value of the first temperature map to replace the temperature value of the second temperature map.
  • the steps for performing susceptibility correction on the phase difference map include:
  • the ablation calculation method further includes:
  • S107 the step of correcting the phase error caused by the motion on the phase difference map or the temperature difference map
  • the steps to correct motion-induced phase errors on the phase difference map include:
  • Motion-induced phase error is removed by a linear least squares fit at each pixel using the reference phase difference map and the calibrated phase difference map;
  • the steps to correct motion-induced phase errors on the temperature difference map include:
  • Motion-induced phase errors are removed by a linear least squares fit at each pixel using the first temperature map and the second temperature map.
  • Figure 4(c) shows the phase difference (first row) and relative temperature change ( second line).
  • the phase error ⁇ (x,y) bias introduces a large temperature bias, but is correctly eliminated after a linear least squares fit.
  • steps S105 , S106 and S107 can be performed on the phase difference layer surface or on the temperature difference map. That is, the step of eliminating the phase drift caused by the magnetic resonance system is performed on the phase difference map and/or the temperature difference map, and the step of magnetic susceptibility correction is performed on the phase difference map and/or the temperature difference map. The step of correcting motion-induced phase errors is performed on a phase difference map and/or a temperature difference map.
  • Fig. 3(c) shows the phase difference map of unwrapping and drift correction
  • Fig. 3(d) shows the temperature map
  • Fig. 3(e) shows the susceptibility-corrected image
  • Fig. 3(f) shows the motion error Corrected image.
  • the obtaining a temperature map according to the temperature difference map includes:
  • S1031 Calculate the temperature using the reference phase difference map and the calibrated phase difference map, and weight the calculated temperature to obtain a temperature map of the target to be measured;
  • a weighted average of the reference phase difference map and the calibrated phase difference map is used to obtain an average temperature difference, and a temperature map of the object to be measured is calculated according to the average temperature difference.
  • the weighting may be various weighting methods, such as average weighting, or the temperature map may be a temperature map corresponding to a single echo time, that is, the weighting coefficient of the temperature map is 1, and the weighting coefficients of other phase temperature maps is 0.
  • step S1031 it may further include:
  • S108 Perform multiple interpolation processing on the temperature map of the object to be measured, and calculate the boundary of the ablation region by using the temperature map of the object to be measured after the interpolation processing.
  • the purpose of performing multiple interpolation processing on the temperature map of the target to be measured is to obtain a smoother ablation region boundary, and the specific number of difference processing may be 2 or 3 times.
  • E a is the activation energy
  • A is the frequency factor
  • R is the universal gas constant
  • T( ⁇ ) is a function of temperature (°C) and time ⁇
  • t is the current time.
  • the gel tissue simulant (gel phantom) was heated using a laser ablation system including a 10W, 980nm diode laser and a cooled laser applicator system.
  • Figure 5 shows the insertion of an ablation fiber and two fiber-optic temperature probes into the gel tissue simulant, with a gel-filled reference tube fixed around as an insulating reference.
  • the ex vivo experiments of pork and porcine brain were performed using the same scanning parameters as the tissue mock experiments. Two experiments were performed on each type of tissue (gel, pork, pig brain), one with several laser cycles and the other with continuous heating and cooling. The root mean square error between the MR measurement temperature and the fiber measurement temperature was calculated as a measure of temperature accuracy.
  • post-ablation images including T1 plus gadolinium (T1+Gd) contrast images, fluid-attenuated inversion recovery (FLAIR), diffusion-weighted MR (distorted, through FSL 5.0 by augmentation or corrected by EPSI method) and T2-weighted images.
  • FIG. 3 and FIG. 4 illustrate an example of temperature calculation of the ablation calculation method provided by the embodiment of the present application.
  • Fig. 3(a) one time obtained during laser hyperthermia, the coil combined phase image is first obtained by a multi-TE echo sequence;
  • Fig. 3(b) the phase difference map is then obtained, the white arrows indicate the phase around the heating center The phase wrapping occurs on the graph;
  • Figure 3(c) shows the phase difference map after phase unwrapping and B0 drift correction.
  • Fig. 3(d) The temperature map calculated from Fig. 3(c) according to the PRF migration method.
  • White arrows highlight susceptibility-induced errors.
  • Figure 3(e) Temperature map after susceptibility correction.
  • White arrows show errors caused by residual CSF motion.
  • an exemplary method flow for a representative pixel includes: Step 1, obtaining a phase difference map and a phase map obtained by unwrapping the reference phase difference map (TE1), as shown by black arrows, under the condition of rapid temperature change , the phase difference map corresponding to some echo times is wrapped. Step 2, obtain the phase unwrapping map, step 3, drift correction of static magnetic field strength (magnetic resonance system), that is, B0 drift correction is to reduce system fluctuation, step 4), phase error correction caused by magnetic susceptibility, use the shortest echo Time (TE) corrects for temperature errors caused by susceptibility changes (black arrows) over longer echo times. Step 5 Motion-induced phase error correction.
  • the graphs in the first row and the second row are the phase difference versus time and the corresponding temperature versus time, respectively.
  • the motion-induced phase errors (black arrows) are nearly the same for multiple echo times, thus leading to more pronounced temperature errors on shorter TEs.
  • the result of correcting the motion error shows a smoother phase and temperature curve.
  • Figure 6 shows a representative temperature map of an ex vivo pork experiment during laser interstitial hyperthermia.
  • Six representative images were selected from 300 frames (3s/frame) acquired during thermal cycling (#50 for the 50th frame, #146 for the 146th frame, and so on).
  • the first row and the second row respectively use the traditional phase unwrapping method and the multi-echo time-based phase unwrapping method (multi-TE unwrapping) proposed in the embodiments of the present application.
  • multi-TE unwrapping multi-echo time-based phase unwrapping method
  • the technical principle is as follows:
  • the phase unwrapping method of the prior art is applied to the time dimension for phase jump detection. If the phase difference map of the current frame is wrongly unwrapped, all subsequent frames will be affected.
  • the ablation calculation method proposed by the present invention is performed on the basis of multiple echo dimensions, thus avoiding interference from previous frames.
  • the third row is a phase-unwrapped and susceptibility-corrected single echo time-temperature map with damaged pixels around the heating center recovered correctly.
  • the last row is the result of combining multiple echo time data using the ablation calculation method provided in the embodiment of the present application.
  • the final magnetic resonance thermography was shown to be more uniform in temperature at the hot spot.
  • Figure 7 shows the temperature as a function of time during the gel tissue mimic (Fig. 7(a)) or ex vivo pork (Fig. 7(b)) experiments, measured by two thermometric fibers (red lines) and Calculated by the method provided in the embodiment of the present application (dashed black line).
  • the temperature-time behavior of the proton resonance frequency (PRF) calculation in the cooling phase is similar to that of the fiber-optic temperature probe (also known as the thermometer
  • the measured temperature-time measurements of the fiber optic match very well.
  • Table 2 lists the root mean square error (RMSE) values between the MR calculated values and the thermometric fiber measurements, which represent the temperature accuracy of the proposed algorithm.
  • RMSE root mean square error
  • Experiment 1 performed several laser cycles of heating, while Experiment 2 had successive heating and cooling stages.
  • the results showed that the RMSE error of gel, pork or pig brain tissue was less than 0.5°C in most cases.
  • the probe (left) represents the left temperature measuring fiber, and the probe (right) represents the right temperature measuring fiber.
  • RMSE root mean square error
  • experimental experimental L(R)
  • TE single echo time
  • our proposed method integrates the information of multiple echoes, so the obtained temperature map removes both CSF-induced errors and magnetic susceptibility-induced errors, showing a more uniform and symmetrical heating region.
  • Figure 9 shows the results of laser interstitial thermotherapy ablation of 3 representative dogs (dogs 01-03).
  • the first column is a T2w image after thermal ablation and after gadolinium administration (post - T2).
  • the location of the ablation fiber can be clearly shown on T2w MRI.
  • the second column shows the final estimated ablation lesions of the proposed method at a given laser dose. These estimated ablation lesions are superimposed on the post-ablation T2w in red (darker color on grayscale). on the image.
  • the area of the ablation zone estimated by the method of the present invention is displayed in the upper left corner of the ablation image.
  • the third column is a representative temperature map obtained when laser heating is most intense.
  • the last three columns are post-ablation FLAIR, DWI, and T1w images after gadolinium administration, respectively. They all show sharp transition zones between dead and living tissue.
  • Figure 10 shows the post-ablation T2w images for all other six cases with different ablation laser doses (dogs 04-09) and the final algorithm predicts the damage estimate.
  • Calculated ablation (second row) matches well with post-ablation evaluation (first row).
  • the ablation area estimated by the algorithm is shown in the upper left corner of the T2w image. Depending on the duration of laser heating, the laser ablation area ranged from less than 30 square millimeters to nearly 90 square millimeters.
  • the ablation calculation method proposed in the present invention combines the advantages of different echoes to obtain better temperature map measurement results. Moreover, the ablation calculation method of the present invention can significantly improve the robustness and signal-to-noise ratio of magnetic resonance thermal imaging, thereby avoiding damage to healthy tissue due to misestimation of low temperature.
  • the method of the present invention also has excellent cerebrospinal fluid flow error suppression and can provide accurate temperature measurements in or around the ventricle. Compensation for errors caused by cerebrospinal fluid motion is clinically important for laser interstitial hyperthermia in the treatment of periventricular brain lesions. Furthermore, the proposed algorithm is online compatible and does not require iterative computations, making it well suited for magnetic resonance thermography as very close to real-time temperature maps are required.
  • the magnetic resonance temperature imaging system provided by the embodiments of the present application will be described below, and the magnetic resonance temperature imaging system described below can be referred to in correspondence with the ablation calculation method described above.
  • the embodiment of the present application further provides an ablation calculation system, which includes an ablation calculation module, and the ablation calculation module can execute the ablation calculation method of the present invention.
  • the ablation calculation module may further include:
  • an information acquisition module which is used for receiving or acquiring magnetic resonance information, the magnetic resonance information at least including a phase map corresponding to the echo time obtained by scanning the target to be measured by using a gradient echo sequence containing i different echo times,
  • the i is a positive integer greater than or equal to 2;
  • a temperature difference calculation module which is used to select at least two sets of phase maps corresponding to different echo times to obtain corresponding temperature difference maps
  • a temperature map calculation module which is used to obtain a temperature map according to the temperature difference map, and the temperature map may be a weighted temperature map;
  • an ablation calculation module which is used for calculating the ablation condition of each pixel according to the temperature map.
  • phase drift correction module for performing the step of eliminating phase drift caused by the magnetic resonance system
  • a magnetic susceptibility error correction module for performing the steps of correcting errors caused by magnetic susceptibility
  • a motion error correction module for performing the steps of correcting motion-induced phase errors.
  • a laser interstitial hyperthermia instrument which includes: a host, a laser ablation device, and an optical fiber assembly, wherein the host includes a processor and is loaded with program codes, which are implemented when the program codes are executed.
  • the ablation calculation method of the present invention is provided, which includes: a host, a laser ablation device, and an optical fiber assembly, wherein the host includes a processor and is loaded with program codes, which are implemented when the program codes are executed.
  • the embodiment of the present application also provides another magnetic resonance temperature imaging system, including:
  • a data transmission module which is configured to receive the magnetic resonance sequence image and judge the integrity of the image
  • a temperature calculation module which is configured to select the sequence, calculate the phase difference, calibrate the phase difference, and calculate the temperature
  • a temperature display module which is set to display the temperature in the form of a pseudo-color map or an isotherm
  • an ablation calculation module which is configured to calculate and display the ablation result
  • the time for the system to perform a complete calculation does not exceed 1 s.
  • the time for the system to perform a complete calculation is preferably no more than 0.5s, most preferably no more than 0.1s.
  • an embodiment of the present application also provides a magnetic resonance temperature imaging system, including: a memory and a processor;
  • the memory is used to store program codes
  • the processor is used to call the program codes
  • the program codes are used to execute the ablation calculation method described in any one of the above embodiments.
  • an embodiment of the present application further provides a storage medium, where a program code is stored on the storage medium, and when the program code is executed, the ablation calculation method described in any of the foregoing embodiments is implemented.
  • the embodiments of the present application provide an ablation calculation method and a related device.
  • the ablation calculation method is not a retrospective algorithm or an iterative algorithm, and the calculation amount is small, and an almost real-time temperature map can be provided. D.

Landscapes

  • Physics & Mathematics (AREA)
  • High Energy & Nuclear Physics (AREA)
  • Condensed Matter Physics & Semiconductors (AREA)
  • General Physics & Mathematics (AREA)
  • Health & Medical Sciences (AREA)
  • General Health & Medical Sciences (AREA)
  • Nuclear Medicine, Radiotherapy & Molecular Imaging (AREA)
  • Radiology & Medical Imaging (AREA)
  • Engineering & Computer Science (AREA)
  • Signal Processing (AREA)
  • Magnetic Resonance Imaging Apparatus (AREA)

Abstract

一种消融计算方法及消融计算系统，该消融计算方法包括：使用含有i个不同回波时间的梯度回波序列对待测目标进行扫描，得到与回波时间对应的相位图，i为大于或等于2的正整数(S101)；选取至少两组对应不同回波时间的相位图获得对应的温度差图(S102)；根据温度差图获得温度图(S103)；根据温度图计算消融情况(S104)；消融计算系统包括能够执行消融计算方法的消融计算模块。

Description

一种消融计算方法及消融计算系统
本申请要求2021年2月8日提交的发明名称为“一种消融计算方法及消融计算系统”的中国专利申请号202110184184.1的优先权，其内容出于所有目的通过引用整体并入本文。
技术领域
本申请涉及医疗器械领域，更具体地说，涉及一种消融计算方法及消融计算系统。
背景技术
磁共振温度成像(Magnetic Resonance Temperature Imaging，MRTI)可以实现无创、实时、在体监测被试物体内部温度分布及变化，在微创和无创热疗，例如磁共振间质热疗、聚焦超声治疗等治疗的消融监测过程中具有重要的用途。
当前的磁共振温度成像方法之一为基于质子共振频率(Proton Resonance Frequency，PRF)位移的测温法，在实践过程中发现，基于质子共振频率位移的测温法受到磁共振线圈磁场均匀性，组织磁化率分布不均匀，呼吸/血流搏动等引起的组织运动等客观环境因素的影响较大，引起的常见误差包括相位解包裹误差，磁化率急速变化导致的误差，运动引起的误差等，容易造成最终获取的温度图与实际温度相差较大的问题，这使得该温度图失去了参考意义。
如何对根据磁共振数据进行更加准确的消融计算，是本领域仍需解决的技术问题。
发明内容
为解决上述技术问题，本申请提供了一种消融计算方法及相关的消融计算系统。
在第本发明的第一个方面，提供了一种消融计算方法，该方法包括：
使用含有i个不同回波时间的梯度回波序列对待测目标进行扫描，得到与所述回波时间对应的相位图，所述i为大于或等于2的正整数；
选取至少两组对应不同回波时间的相位图获得对应的温度差图；
根据所述温度差图获得温度图；
根据所述温度图计算(各像素的)消融情况。
进一步地，该方法中，获得温度差图是这样进行的：使用任一时刻的相位图减去基准时刻的相位图得到该时刻的相位差图，选取至少一个回波时间对应的相位差图作为参考相位差图，基于参考相位差图对其他回波时间对应的待校准相位差图进行校准，得到经校准的相位差图，参考相位差图对应的回波时间小于其所校准的相位差图对应的回波时间；使用参考相位差图和经校准的相位差图计算该时刻的温度差图。
其中，至少一个参考相位差图对应的回波时间不超过以下之一：18ms，17ms，16ms，15ms，14ms，13ms，12ms，11ms，10ms，9ms，8ms，7ms，6ms，5ms或4ms。
更进一步地，该方法中，上述校准如下进行：
根据相位差图与回波时间成正比例的关系，基于回波时间和参考相位差图，计算得到待校准相位差图的估计值；
使用待校准相位差图的估计值，根据相位周期性对待校准相位差图进行解包裹，得到经校准后相位差图。
可选地，本发明的消融计算方法还包括消除磁共振系统引起的相位漂移的步骤,(例如B0漂移误差)；进一步地，消除磁共振系统引起的相位漂移的步骤中，选取若干物理温度稳定无变化且组织均匀的区域作为热参考点，通过从每个相位差图像减去所述热参考点的平均相位差或者温度变化图中减去所述热参考点的平均温度变化，进行相位漂移校正。
可选地，本发明的消融计算方法还包括校正磁化率引起的误差的步骤，磁化率校正的步骤在相位差图或温度差图上进行，该步骤包括：
在温度差图上进行磁化率校正的步骤包括：
根据所述参考相位差图得到第一温度图，根据所述经校准相位差图得到对应的第二温度图，
判断所述第二温度图中每个像素对应的温度值与所述第一温度图中相应像素对应的温度值的差值的绝对值是否超过预设温度阈值，如果是，则对所述第二温度图中相应像素对应的温度值进行校正；
在相位差图上进行磁化率校正的步骤包括：
判断所述经校准相位差图中每个像素对应的相位差值与所述参考相位图中相应像素对应的相位差值的差值绝对值是否超过预设相位差阈值，如果是，则对经校准相位差图中相应像素对应的相位差进行校正。
可选地，本发明的消融计算方法还包括校正运动引起的相位误差的步骤，该步骤通过使用至少两组对应不同回波时间的相位图在每个像素处的线性最小二乘拟合将运动引起的相位误差去除。
可选地，本发明的消融计算方法还包括获取加权温度图的步骤，加权温度图是通过至少一个参考相位差图获得的温度图和至少一个经校准的相位差图获得的温度图进行加权获得。其中，至少一个经校准的相位差图对应的回波时间不小于20ms，19ms，18ms，17ms，16ms，15ms，14ms，13ms或12ms。
以上可选的消除磁共振系统引起的相位漂移的步骤，校正磁化率引起的误差的步骤，校正运动引起的相位误差的步骤，和获取加权温度图的步骤单独、部分或者全部选择作为本发明的消融计算方法的一部分的方案都属于本发明内容的范围。
可选地，本发明的消融计算方法中，像素的消融是使用如下公式计算的：
Figure PCTCN2022075489-appb-000001
其中，E a表示活化能，A是频率因子，R是通用气体常数，T(τ)是温度(℃)与时间τ的函数，t是当前时间，Ω值超过设定阈值(例如1)的像素视为已消融。本领域技术人员所知道的基于温度进行消融计算的其他方法和参数也可以作为替代方案，作为本发明的一部分。
在第本发明的第二个方面，提供了一种存储介质，存储介质上存储有程序代码，程序代码被执行时实现本发明的消融计算方法。
在第本发明的第三个方面，提供了一种消融计算系统，其包含消融计算模块，所述消融计算模块能够执行权本发明的消融计算方法。进一步地，消融计算模块可以进一步包含：
信息采集模块，其用于接收或者采集磁共振信息，磁共振信息至少包含使用含有i个不同回波时间的梯度回波序列对待测目标进行扫描得到的与所述回波时间对应的相位图，所述i为大于或等于2的正整数；
温度差计算模块，其用于选取至少两组对应不同回波时间的相位图获得对应的温度差图；
温度图计算模块，其用于根据所述温度差图获得温度图，温度图可以是加权温度图；
消融计算模块，其用于根据所述温度图计算各像素的消融情况。
相位漂移校正模块，其用于执行消除磁共振系统引起的相位漂移的步骤，
磁化率误差校正模块，其用于执行校正磁化率引起的误差的步骤，
运动误差校正模块，其用于执行校正运动引起的相位误差的步骤。
在第本发明的第四个方面，提供了一种激光间质热疗仪，其包括：主机、激光消融设备和光纤组件，主机包含处理器处理器加载有程序代码，程序代码被执行时实现本发明的消融计算方法。
本发明的实施方案的创新点包括一下一个或更多个：
1、消除了部分回波时间对应的回波序列的相位差图解包裹时发生错误的问题；
2、消除了因为升温导致导致的部分区域磁化率异常导致温度异常，无法显示温度和消融情况的问题；
3、消除了脑脊液，心脏搏动等运动引起的异常温度误差，例如脑室中的异常温度。
4、本发明的方法使用非追溯性算法或迭代算法，运算量较小，节省了计算时间，能够快速的得到温度和消融结果。
附图说明
为了更清楚地说明本申请实施例或现有技术中的技术方案，下面将对实施例或现有技术描述中所需要使用的附图作简单地介绍，显而易见地，下面描述中的附图仅仅是本申请的实施例，对于本领域普通技术人员来讲，在不付出创造性劳动的前提下，还可以根据提供的附图获得其他的附图。
图1为现有技术在离体和在体环境中获得的磁共振数据的幅值、相位图和温度图；
图2为本申请的一个实施例提供的一种消融计算方法的流程示意图；
图3为本申请的一个实施例提供的获取的相位图、相位差图和温度图；
图4为本申请的另一个实施例提供的一种消融计算方法的部分流程示意图；
图5为本申请的一个实施例提供的实验装置示意图；
图6为本申请的一个实施例提供的离体猪肉实验的激光间质热疗温度图的局部放大；
图7为本申请的一个实施例提供的在凝胶组织模拟物(图7(a))或离体猪肉(图7(b))实验期间，温度随时间的变化示意图；
图8为本申请的一个实施例提供的体内实验中的狗01的代表性温度图；
图9为本申请的一个实施例提供的对3只代表性狗(狗01-03)的激光间质热疗消融结果；
图10为本申请的一个实施例提供的消融后的T2w图像以及所有其他六种不同消融激光剂量的情况(狗04-09)的消融计算方法预测的消融估计示意图。
具体实施方式
磁共振温度成像可以引导多种能量输送型治疗手段，例如激光间质热疗、聚焦超声治疗、射频消融等，监控目标组织温度和消融效果。本发明以磁共振温度成像引导的激光间质热疗作为例子说明本发明的方法，磁共振温度成像引导的激光间质热疗是一种微创治疗手段，为治疗位于手术上具有挑战性的部位(解剖学或功能性)的肿瘤创造了新的选择。该方法通过施加温度在50～80℃或更高的温度达数十秒，可以快速凝固组织并通过蛋白质变性诱导肿瘤细胞坏死。与开放式手术相比，激光间质热疗可更精确地靶向肿瘤，并减少不适感和感染风向，并缩短患者住院时间。在激光间质热疗的消融过程中，并发磁共振热成像对于更有效地消融肿瘤细胞以及更好地保护健康的周围细胞和关键结构起着重要的作用。大多数激光间质热疗消融程序取决于基于质子共振频率位移的测温法。
但是，正如背景技术所述，现有的磁共振温度成像方法受到环境等因素的影响较大，容易造成最终获取的温度图与实际温度相差较大的问题。
在激光间质热疗的消融过程中，发明人通过研究发现获取的温度图中错误的主要来源是解卷错位导致的相位误差、磁化率误差和运动导致的相位误差。随着消融激光剂量的改变，磁化率会导致图像振幅减小以及图像相位中的相应误差，从而破坏加热中心及其周围的重建温度图。重建温度图的错误可能会导致消融区域的估计错误，从而可能导致治疗效果的变化以及对关键组织的热损伤。因此，准确的温度成像对于激光间质热疗的有效性和安全性至关重要，尤其是当激光间质热疗应用于脑组织中较紧的消融区域时。
发明人通过进一步研究发现，基于质子共振频率位移的测温法基于以下事实：氢质子的共振频率随水分子中的温度而变化。对于含水组织，局部磁场随温度的变化可描述为：
Figure PCTCN2022075489-appb-000002
其中，α是随温度变化的质子共振频率系数，在这里取0.008-0.015ppm/℃。受温度影响的水质子的相应共振频率变化可以表示为：
Δf＝αγB 0·ΔT；  (2)
其中，ΔT表示温度变化，Δf表示共振频率变化，γ表示旋磁比，B 0表示静态磁场强度。
可以在复杂的磁共振成像的相位中观察到由于温度变化引起的共振频率的变化。对于给定的梯度回波序列的间隔时间TE，可以根据相位差Δφ计算相对温度变化ΔT，该方程可表示为：
Figure PCTCN2022075489-appb-000003
梯度回波序列是基于质子共振频率位移的测温法中使用的序列。根据公式(3)可知，梯度回波序列越长，相同的温度变化可能导致相位差越大，表明可以获得更高的温度灵敏度。
参考图1，随着梯度回波序列的回波时间增加，相位对比和相位包裹都增加，这表明在稍后的回波时间内，温度灵敏度更高，相位解缠程序更多。图1中，在(a)(离体，猪脑)和(b)(体内)中通过本申请实施例使用的含有4个不同回波时间的梯度回波序列获得的第一至第四回波的幅值(上排)和相位(第二行)。使用传统的PRF算法根据每个TE(回波时间)设置计算温度图(下排)。在较长的回波时间内会出现更多的相位包裹，因为图像对比度也相应增加。请注意，强烈的激光加热会由于磁化率的变化而导致信号损失，并且还会转化为加热中心周围像素的相位和温度误差。在体内实验中，请注意脑脊液(Cerebrospinal Fluid，CSF)运动(motion)会在MRTI上引起不合适的高温，这在前一个回波中更明显，因为较短的TE对所引入的相似相位误差的耐受性较低。
例如，由于大脑中脑脊液的运动，扫描间运动可能是基于质子共振频率位移的测温法测量的温度图中的一个大问题。脑脊液的大小和相位信号经常通过脑脊液的正常动态运动在脉冲梯度回波序列上更改，这可能会混淆温度估计。脑脊液运动还可能导致心室内和周围心室的像素移动，从而导致相差图出现错误。如图1b所示，体内温度图显示由于脑脊液运动，第三脑室内的伪高温。温度误差在较短回波时间的脉冲梯度回波序列上更为明显，因为它们对(3)中的脑脊液流动引入的相移强度的容忍度较小。
实际上，水质子的局部磁场也应考虑磁化率x 0，公式(1)变为：
Figure PCTCN2022075489-appb-000004
其中，σ χ0表示由磁化率引起的局部磁场变化。
进一步研究发现，激光加热会在激光尖端周围的GRE成像中引起明显的磁化伪影。仍然参考图1，温度急剧变化的加热中心(如图1(a)中箭头所示)在较长的回波时间数量级上显示严重的信号损失。体素内自旋相移是由温度和磁化率变化引起的局部磁场不均匀引起的。
激光加热引起的磁化伪影，尤其是较长回波时间的梯度回波序列对应的图像中的磁化伪影，是造成误差的重要原因。仍然参考图1，在离体或体内实验中，加热中心周围的相位误差转化为磁共振热成像上的伪低温。通常情况下，在成像过程汇总，建议使用具有尽可能短的回波时间的梯度脉冲序列以最大程度地减小磁化率伪影。但是，更长的回波时间的梯度脉冲序列可以提供更好的温度灵敏度和信噪比，这是当前面临的两难选择。
为了兼顾温度灵敏度、信噪比和低误差，本申请实施例提供了一种消融计算方法，该方法包括：
使用含有i个不同回波时间的梯度回波序列对待测目标进行扫描，得到与所述回波时间对应的相位图，所述i为大于或等于2的正整数；
选取至少两组对应不同回波时间的相位图获得对应的温度差图；
根据所述温度差图获得温度图；
根据所述温度图计算各像素的消融情况。
所述消融计算方法基于含有i个不同回波时间的梯度回波序列获得i组相位图，选取至少两组对应不同回波时间的相位图获得对应的相位差图，并根据所述温度差图获得温度图。发明人经过研究发现，梯度回波序列的回波时间与磁化率伪影的大小成正比关系，因此与较小回波时间对应的梯度回波序列获得的相位图受到的由于加热而导致的磁化率变化的影响最小，其图像数据仍保持正确的相位，因此可以基于含有i个不同回波时间的梯度回波序列获得的i组相位图以及相位差图进行温度图的获取，以降低最终获得的温度图的误差，提高获得的温度图的准确性的目的。
进一步的，所述消融计算方法并非追溯性算法或迭代算法，运算量较小，可提供几乎实时的温度图，具有较高的参考意义。
本申请实施例提供了一种消融计算方法，如图2所示，包括：
S101：使用含有i个不同回波时间的梯度回波序列对待测目标进行扫描，得到与所述回波时间对应的i组相位图，所述i为大于或等于2的正整数；
在步骤S101中，所述梯度回波序列中回波时间的最小取值和最大取值均可根据实际需求而定，一般情况下，为了尽量降低磁化率变化导致的磁化率伪影，所述梯度回波序列中回波时间的最小取值可以取磁共振温度成像设备能够取到的最小值，所述梯度回波序列中回波时间的最大取值一般不超过对待测目标成像的回波时间的取值范围上限。例如，在对头部成像来说，其可选的梯度时间序列的回波时间的取值范围为3～30ms，所述梯度回波序列中包含的回波时间的具体取值均要在该取值范围内。
参考图3，具体参考图3(a)，获取的相位图如图3(a)所示。然后如图3(b)所示，通过复数相减程序计算出相差图。复相减法可以避免有问题的相位包裹。
含有i个不同回波时间的梯度回波序列信息的获取可以是从服务器或其他存储设备中读取或接收，也可以是根据工作人员的设定实时获取的，本申请对获取梯度回波序列的具体方法并不做限定，具体视实际情况而定。
S102：选取至少两组对应不同回波时间的相位图获得对应的温度差图；
可选的，在步骤S102和步骤S103之间，还可以包括对相位图和相位差图的静态磁场强度漂移校正，以消除静态磁场强度引起的误差。
S103：根据所述温度差图获得温度图。
S104：根据所述温度图计算各像素的消融情况，消融可以使用基于温度和时间参数计算消融的各种方法，例如使用如下公式计算的：
Figure PCTCN2022075489-appb-000005
其中，E a表示活化能，A是频率因子，R是通用气体常数，T(τ)是温度(℃)与时间τ的函数，t是当前时间，Ω值超过设定阈值(例如1)的像素视为已消融。
下面对本申请实施例提供的消融计算方法的各个步骤的可行执行方式进行描述。
在上述实施例的基础上，在本申请的一个实施例中，获得温度差图的具体步骤包括：
使用所述不同时刻中任一时刻的相位图减去基准时刻的相位图得到该时刻的相位差图，基准时刻为对目标组织传输能量(例如热能、光能、射频消融、冷冻消融)之前的任意时刻，优选地为进行能量传输之前不久的时刻，例如即将进行能量传输的时刻；选取在该时刻至少一个回波时间对应的相位差图作为参考相位差图，对其他的回波时间对应的相位差图进行校准，得到经校准的相位差图，所述参考相位差图对应的回波时间小于其所校准的相位差图对应的回波时间；
使用所述参考相位差图和所述经校准的相位差图计算该时刻的温度差图。
可选的，参考相位差图的回波时间的取值小于或等于18ms，优选地，参考相位差图的回波时间的取值不超过17ms，16ms，15ms，14ms，13ms，12ms，11ms，10ms，9ms，8ms或7ms，更优选地，参考相位差图的回波时间的取值不超过6ms，5ms或4ms。
可选的，可使用梯度回波序列中的最小回波时间对应的相位图作为参考相位图，并基于参考相位图获得参考相位差图，以尽可能最小化相位图受到的由于加热而导致的磁化率变化的影响。如前文所述，所述梯度回波序列中的最小回波时间可以为磁共振温度成像设备能够取到的最小值。
所述选取在该时刻至少一个回波时间对应的相位差图作为参考相位差图，对其他回波时间对应的相位差图进行校准包括以下步骤：
使用所述参考相位差图和待校准的回波时间对应的相位差图，根据相位差与回波时间成正比例的关系，基于回波时间和参考相位差图的相位差，计算得到所述待校准的回波时间对应的相位差图的相位差的估计值；，然后使用所述估计值，根据相位周期性对所述待校准的相位差进行解卷积，得到经校准后的相位差。
在上述实施例的基础上，在本申请的另一个实施例中，所述消融计算方法还包括：
S105：消除磁共振系统引起的相位漂移的步骤，所述消除磁共振系统引起的相位漂移的步骤在相位差图或温度差图上进行。
在相位差图上进行消除磁共振系统引起的相位漂移的步骤包括：
选取多个热参考点(Region ofInterest，ROI)，通过从每个相位差图减去所述热参考点的平均相位差；
在温度差图上消除磁共振系统引起的相位漂移的步骤包括：
在温度差图中减去任一所述热参考点的平均温度差，进行校正。
在上述实施例的基础上，在本申请的另一个实施例中，所述消融计算方法还包括：
S106：磁化率校正的步骤，磁化率校正的步骤在在相位差图或温度差图上进行；
在温度差图上进行磁化率校正的步骤包括：
根据所述参考相位差图得到第一温度图，根据所述经校准相位差图得到对应的第二温度图；
可选地，在所述第一温度图和各个所述第二温度图中确定预设区域；
判断所述第二温度图中每个像素对应的温度值与所述第一温度图中所述预设区域中相应像素对应的温度值的差值的绝对值是否超过预设温度阈值，如果是，则对所述第二温度图中相应像素对应的温度值进行校正；校正可以有多种方法，例如可以使用该像素中第一温度图的温度值替换第二温度图的温度值，或者使用第二温度图中相邻像素的温度值替换第二温度图中该像素的温度值，或者基于相邻像素的温度值和第一温度图的温度值拟合一个近似温度替代第二温度图的温度值；
在相位差图上进行磁化率校正的步骤包括：
在所述参考相位差图和经校准的相位差图中确定预设区域；
判断所述经校准相位差图中每个像素对应的相位差值与所述参考相位图中所述预设区域中相应像素对应的相位差值的差值绝对值是否超过预设相位差阈值，如果是，则对经校准相位差图中的相位差进行校正，校正方法与前文类似，不再重复。
在上述实施例的基础上，在本申请的又一个实施例中，所述消融计算方法还包括：
S107：在相位差图或温度差图上进行的校正运动引起的相位误差的步骤；
在相位差图上进行的校正运动引起的相位误差的步骤包括：
通过使用所述参考相位差图和所述经校准相位差图在每个像素处的线性最小二乘拟合将运动引起的相位误差去除；
在温度差图上进行的校正运动引起的相位误差的步骤包括：
根据所述参考相位差图得到第一温度图，根据所述经校准相位差图得到对应的第二温度图；
通过使用所述第一温度图和所述第二温度图在每个像素处的线性最小二乘拟合将运动引起的相位误差去除。
仍然参考图4(c)，图4(c)示出了在没有(左图)和有(右图)运动误差校正的情况下作为时间的函数的相差(第一行)和相对温度变化(第二行)。对于较短的回波时间，相位误差Δφ(x,y) bias会引入较大的温度偏差，但在线性最小二乘拟合之后可正确消除。
如前文所述，步骤S105、S106和S107均即可在相位差图层面上进行，也可在温度差图上进行。即所述消除磁共振系统引起的相位漂移的步骤在相位差图和/或温度差图上进行，所述磁化率校正的步骤在相位差图和/或温度差图上进行。所述校正运动引起的相位误差的步骤在相位差图和/或温度差图上进行。
仍然参考图3，图3(c)表示解包裹和漂移校正的相差图，图3(d)表示温度图，图3(e)表示磁化率校正后的图像，图3(f)表示运动误差校正后的图像。
在上述实施例的基础上，在本申请的又一个实施例中，所述根据所述温度差图获得温度图包括：
S1031：使用所述参考相位差图和经校准的相位差图计算温度，并对计算得到的温度进行加权以获得待测目标的温度图；
或
使用对所述参考相位差图和经校准的相位差图进行加权平均以获得平均温度差，并根据所述平均温度差计算所述待测目标的温度图。
在步骤S1031中，加权可以是各种加权方法，例如平均加权，或者温度图可以是单独的一个回波时间对应的温度图，即该温度图的加权系数为1，其他相位温度图的加权系数为0。
在步骤S1031之后还可包括：
S108：对所述待测目标的温度图进行多次插值处理，并利用插值处理后的所述待测目标的温度图计算消融区域边界。
对所述待测目标的温度图进行多次插值处理的目的是为了获得更平滑的消融区域边界，差值处理的具体次数可以是2或3次。
在计算消融区域边界的过程中，具体利用如下公式：
Figure PCTCN2022075489-appb-000006
其中，E a表示活化能，A是频率因子，R是通用气体常数，T(τ)是温度(℃)与时间τ的函数，t是当前时间。Ω值超过设定阈值(例如1)的像素视为已消融。
下面结合具体实验对本申请实施例提供的消融计算方法进行验证。
使用包括10W，980nm二极管激光器和冷却的激光施加器系统的激光消融系统对凝胶组织模拟物(凝胶体模)进行加热。相位图像是在3TMR扫描仪(Ingenia，Philips Healthcare，Best，荷兰)使用16条接收线圈使用多回波时间梯度回波序列获取的：翻转角＝30°，TE＝6/12/18/24ms，TR＝22ms，矩阵＝176×176，FOV＝200 x 200mm 2，切片厚度＝5mm，3s/图像。
如图5所示，还将两个MR兼容的光纤温度探头插入到组织模拟物中，探头尖端位于靠近消融光纤的位置，以获取各点的凝胶温度。由于加热过程中光纤探头受到消融光纤的影响，因此温度计仅监视冷却阶段。图5中。具体地，图5中显示在凝胶组织模拟物中插入了消融光纤和两个光纤温度探头，填充凝胶的参考管固定在周围作为绝缘参考。
使用与组织模拟物实验相同的扫描参数进行猪肉和猪脑的离体实验。对每种类型的组织(凝胶，猪肉，猪脑)进行了两次实验，其中一种进行了若干次激光循环加热，另一种进行了持续的加热和冷却。计算出MR测量温度和光纤测量温度之间的均方根误差，作为温度精度的测量值。
杜宾狗的体内实验已获得清华大学伦理审查委员会的批准。九只成年杜宾狗接受了激光间质热疗。加热过程在3T MR扫描仪(Ingenia，Philips Healthcare，Best，荷兰)上通过32条接收头线圈使用多回波时间梯度回波序列进行监控。
经过消融手术后，为了获得有关消融区实际范围的详细信息，获取了消融后图像，包括T1加钆(T1+Gd)对比图像，流体衰减反转恢复(FLAIR)，扩散加权MR(失真，通过FSL 5.0通过增补或通过EPSI方法校正)和T2加权图像。
仍然参考图3和图4，图3和图4说明了本申请实施例提供的消融计算方法的温度计算的一个实例。图3(a)，在激光热疗期间获得的一个时间，首先通过多TE回波序列获得线圈组合相位图像；图3(b)，然后获得相位差图，白色箭头指示在加热中心周围的相位图上发生的相位包裹；图3(c)显示经过相位解包裹和B0漂移校正的相位差图。图3(d)根据PRF偏移方法从图3(c)计算得到的温度图。白色箭头突出显示了磁化率引起的误差。图3(e)磁化率校正后的温度图。白色箭头显示了残留的CSF运动引起的误差。图3(f)运动校正的温度图。
图4中，关于代表性像素的一个示例性方法流程包括：步骤1，获取相差图以及参考相位差图(TE1)解包裹获得的相位图，如黑色箭头所示，在温度快速变化的情况下，一些回波时间对应的相位差图发生了包裹。步骤2，获取相位解包裹图，步骤3，静态磁场强度(磁共振系统)漂移校正，即B0漂移校正是为了减少系统波动，步骤4)，磁化率引起的相位误差校正，使用最短的回波时间(TE)校正了磁化率变化(黑色箭头)在较长的回波时间上引起的温度误差。步骤5运动引起的相位误差校正。第一行的图和第二行图分别是随时间变化的相位差和相应的温度随时间的变化。对于多个回波时间，运动引起的相位误差(黑色箭头)几乎相同，因此在较短的TE上导致更明显的温度误差。校正运动误差的结果显示出更平滑的相位和温度曲线。
组织模拟物和离体实验结果：
图6示出了激光间质热疗期间离体猪肉实验的代表性温度图。从热循环期间获取的300帧(3s/帧)中选择六个代表性图像(#50表示第50帧，#146表示第146帧，依次类推)。第一行和第二行分别是使用传统的相位解包裹方法和本申请实施例中提出的基于多回波时间的相位解包裹方法(多TE解包裹)。使用现有技术的相位展开方法，由于激光热量引起的磁化率变化，温度图上的像素会严重损坏，即使不再使用激光也无法恢复。技术原理如下：将现有技术的相位展开方法应用于时间维度进行相位跳变检测，如果当前帧的相差图被错误地解包裹，则随后的所有帧都会受到影响。另一方面，本发明提出的消融计算方法是在多回波维度的基础上进行的，因此避免了来自先前帧的干扰。第三行是经过相位解包裹和磁化率校正的单个回波时间温度图，加热中心周围的损坏像素已正确恢复。最后一行是使用本申请实施例提供的消融计算方法进行的多回波时间数据组合结果。显示出最终的磁共振热成像在热点的温度更均匀。
图7示出了在凝胶组织模拟物(图7(a))或离体猪肉(图7(b))实验期间随时间变化的温度，分别由两个测温光纤测量(红色线条)和本申请实施例提供的方法计算得出的(虚线黑色线条)。在多次加热(图7(a))或单次加热(图7(b))的情况下，质子共振频率(PRF)计算的温度-时间行为在降温阶段与光纤温度探头(也成为测温光纤)测量的测得的温度-时间非常匹配。表2列出了MR计算值和测温光纤测量值之间的均方根误差(RMSE)值，这些值代表了所提出算法的温度精度。实验1进行了几次激光循环加热，而实验2则是连续的加热和冷却阶段。结果表明，在大多数情况下，凝胶，猪肉或猪脑组织的RMSE误差均小于0.5℃。图7中，探头(左)表示左侧测温光纤，探头(右)表示右侧测温光纤。
表2.本申请实施例提供的方法在光纤测量的温度和MR计算的温度之间的比较。
Figure PCTCN2022075489-appb-000007
缩写：RMSE，均方根误差；实验，实验L(R)，左侧(右侧)的光纤温度探头。
图8显示了体内实验中的狗01的代表性温度图。需要注意的是，消融区域位于靠近第三脑室和侧脑室的位置。选择在激光消融过程中获取的100帧(3s/帧)图像叠加在消融后T2w磁共振热成像上。从上至下是通过现有技术算法分别根据单个回波时间(TE)数据(TE＝6ms和TE＝24ms)计算出的温度图，并使用本发明所提出的算法根据多(联合)TE回波序列计算出的温度图。第一行(TE＝6ms)显示了第三侧脑室和侧脑室内的伪高温，表明短TE计算温度严重受CSF流动伪影的影响。CSF诱发的第三脑室内伪影(白色箭头指示)仍存在于第二行(TE＝24ms)，但被所提出的联合TE回波序列算法很好地抑制了。第二行显示，与较短的TE(TE＝6ms)相比，更长的TE(TE＝24ms)可以提供更平滑的边界和更好的温度SNR，但是如上所述，由于磁化率的变化，加热中心周围的像素会损坏。另一方面，我们提出的方法整合了多个回波的信息，因此获得的温度图同时消除了CSF引起的误差和磁化率引起的误差，显示出更均匀和对称的加热区域。
图9显示了对3只代表性狗(狗01-03)的激光间质热疗法消融结果。第一列是热消融后并施用钆后的T2w图像(术后-T 2)。在T2w MRI上可以清楚地显示出消融光纤的位置。第二列显示了在给定的激光剂量下，而并本发明提出的方法的最终估计的消融损伤，这些估计的消融损伤以红色(灰度图上为较深颜色)叠加在消融后的T2w图像上。从而证明了本发明提出的方法计算的估计消融与消融后MRI之间的良好一致性。本发明方法估计的消融区面积显示在消融图像的左上角。第三列是当激光加热最强烈时获得的代表性温度图。最后三列分别是施用钆后的消融后FLAIR，DWI和T1w图像。它们都显示出死组织与活组织之间的尖锐过渡区。
图10显示了所有其他六种不同消融激光剂量的案例(狗04-09)的消融后T2w图像以及最终算法预测损伤估计。消融计算值(第二行)与消融后评估值(第一行)匹配良好。算法估计的消融区域显示在T2w图像的左上角。根据激光加热持续的时间，激光消融面积范围从小于30平方毫米到近90平方毫米。
上面的实验结果表明，可以使用本申请实施例提供的消融计算方法校正由于加热激光器本身引起的质子共振频率温度图中磁化率引起的误差。我们首先提出通过质子共振频率移位方法将多回波时间梯度回波脉冲序列应用于磁共振热成像。相比单回波序列，多梯度回波序列可提供更多信息，而无需额外的扫描时间，并且为相位展开和伪影消除提供了新 方法。
较短的回波时间可以忍受磁化率伪像，但对噪声敏感，而较长的回波时间则具有较好的温度敏感性和信噪比，但受磁化率伪像的影响很大。本发明提出的消融计算方法融合了不同回波的优势，以获得更好的温度图测量结果。而且，本发明的消融计算方法可以显着提高磁共振热成像的鲁棒性和信噪比，从而避免了由于误估了低温而对健康组织造成的损害。
本发明的方法还具有出色的脑脊液流动误差抑制能力，并且可以在心室内或周围提供准确的温度测量。补偿脑脊液运动引起的错误对于激光间质热疗治疗脑室周围脑部病变在临床上很重要。此外，提出的算法是在线兼容的，不需要迭代计算，因此非常适合于磁共振热成像，因为需要非常接近实时的温度图。
下面对本申请实施例提供的磁共振温度成像系统进行描述，下文描述的磁共振温度成像系统可与上文描述的消融计算方法相互对应参照。
相应的，本申请实施例还提供了一种消融计算系统，其包含消融计算模块，所述消融计算模块能够执行权本发明的消融计算方法。进一步地，消融计算模块可以进一步包含：
信息采集模块，其用于接收或者采集磁共振信息，磁共振信息至少包含使用含有i个不同回波时间的梯度回波序列对待测目标进行扫描得到的与所述回波时间对应的相位图，所述i为大于或等于2的正整数；
温度差计算模块，其用于选取至少两组对应不同回波时间的相位图获得对应的温度差图；
温度图计算模块，其用于根据所述温度差图获得温度图，温度图可以是加权温度图；
消融计算模块，其用于根据所述温度图计算各像素的消融情况。
相位漂移校正模块，其用于执行消除磁共振系统引起的相位漂移的步骤，
磁化率误差校正模块，其用于执行校正磁化率引起的误差的步骤，
运动误差校正模块，其用于执行校正运动引起的相位误差的步骤。
在第本发明的第四个方面，提供了一种激光间质热疗仪，其包括：主机、激光消融设备和光纤组件，主机包含处理器处理器加载有程序代码，程序代码被执行时实现本发明的消融计算方法。
本申请实施例还提供了另一种磁共振温度成像系统，包括：
数据传输模块，其设置成接收磁共振序列图像，并判断图像完整性；
温度计算模块，其设置成用于选择序列、计算相位差、校准相位差、计算温度；
温度显示模块，其设置成将温度以伪彩图或等温线的模式展示；
消融计算模块，其设置成计算消融结果并进行显示；
其中，所述系统进行一次完整计算的时间不超过1s。
在本申请的一些实施例中，所述系统进行一次完整计算的时间优选不超过0.5s，最优选不超过0.1s。
相应的，本申请实施例还提供了一种磁共振温度成像系统，包括：存储器和处理器；
所述存储器用于存储程序代码，所述处理器用于调用所述程序代码，所述程序代码用于执行上述任一实施例所述的消融计算方法。
相应的，本申请实施例还提供了一种存储介质，所述存储介质上存储有程序代码，所述程序代码被执行时实现上述任一实施例所述的消融计算方法。
综上所述，本申请实施例提供了一种消融计算方法及相关装置，所述消融计算方法并非追溯性算法或迭代算法，运算量较小，可提供几乎实时的温度图，具有较高的参考意义。
本说明书中各实施例中记载的特征可以相互替换或者组合，每个实施例重点说明的都是与其他实施例的不同之处，各个实施例之间相同相似部分互相参见即可。
对所公开的实施例的上述说明，使本领域专业技术人员能够实现或使用本申请。对这些实施例的多种修改对本领域的专业技术人员来说将是显而易见的，本文中所定义的一般原理可以在不脱离本申请的精神或范围的情况下，在其它实施例中实现。因此，本申请将不会被限制于本文所示的这些实施例，而是要符合与本文所公开的原理和新颖特点相一致的最宽的范围。

Claims (12)

  1. 一种消融计算方法，其特征在于，包括：
    使用含有i个不同回波时间的梯度回波序列对待测目标进行扫描，得到与所述回波时间对应的相位图，所述i为大于或等于2的正整数；
    选取至少两组对应不同回波时间的相位图获得对应的温度差图；
    根据所述温度差图获得温度图；
    根据所述温度图计算消融情况。
  2. 根据权利要求1所述的消融计算方法，其特征在于，获得所述温度差图是这样进行的：
    使用任一时刻的相位图减去基准时刻的相位图得到该时刻的相位差图，选取至少一个回波时间对应的相位差图作为参考相位差图，基于所述参考相位差图对其他回波时间对应的待校准相位差图进行校准，得到经校准的相位差图，所述参考相位差图对应的回波时间小于其所校准的相位差图对应的回波时间；
    使用所述参考相位差图和经校准的相位差图计算该时刻的温度差图。
  3. 根据权利要求2所述的消融计算方法，其特征在于，所述校准如下进行：
    根据相位差图与回波时间成正比例的关系，基于回波时间和参考相位差图，计算得到待校准相位差图的估计值；
    使用所述估计值，根据相位周期性对所述待校准相位差图进行解包裹，得到经校准后相位差图。
  4. 根据权利要求1至3中任一项所述的消融计算方法，其特征在于，还包括消除磁共振系统引起的相位漂移的步骤。
  5. 根据权利要求4所述的消融计算方法，其特征在于，所述消除磁共振系统引起的相位漂移的步骤中，选取若干物理温度稳定无变化且组织均匀的区域作为热参考点，通过从每个相位差图像减去所述热参考点的平均相位差或者温度变化图中减去所述热参考点的平均温度变化，进行相位漂移校正。
  6. 根据权利要求1至5中任一项所述的消融计算方法，其特征在于，还包括校正磁化率引起的误差的步骤，磁化率校正的步骤在相位差图或温度差图上进行，该步骤包括：
    在温度差图上进行磁化率校正的步骤包括：
    根据所述参考相位差图得到第一温度图，根据所述经校准相位差图得到对应的第二温度图，
    判断所述第二温度图中每个像素对应的温度值与所述第一温度图中相应像素对应的温度值的差值的绝对值是否超过预设温度阈值，如果是，则对所述第二温度图中相应像素对应的温度值进行校正；
    在相位差图上进行磁化率校正的步骤包括：
    判断所述经校准相位差图中每个像素对应的相位差值与所述参考相位图中相应像素对应的相位差值的差值绝对值是否超过预设相位差阈值，如果是，则对经校准相位差图中相应像素对应的相位差进行校正。
  7. 根据权利要求1至6中任一项所述的消融计算方法，其特征在于，还包括校正运动引起的相位误差的步骤，该步骤通过使用至少两组对应不同回波时间的相位图在每个像素处的线性最小二乘拟合将运动引起的相位误差去除。
  8. 根据权利要求1至7中任一项所述的消融计算方法，其特征在于，所述温度图是通过至少一个参考相位差图获得的温度图和至少一个经校准的相位差图获得的温度图进行加权获得的加权温度图。
  9. 根据权利要求1所述的消融计算方法，其特征在于，所述消融是使用如下公式计算的：
    Figure PCTCN2022075489-appb-100001
    其中，E a表示活化能，A是频率因子，R是通用气体常数，T(τ)是温度(℃)与时间τ的函数，t是当前时间，Ω值超过设定阈值的像素视为已消融。
  10. 一种存储介质，其特征在于，所述存储介质上存储有程序代码，所述程序代码被执行时实现权利要求1至9中任一项所述的消融计算方法。
  11. 一种消融计算系统，其特征在于，包含消融计算模块，所述消融计算模块能够执行权利要求1至9中任一项所述的消融计算方法。
  12. 一种激光间质热疗仪，其特征在于，包括：主机、激光消融设备和光纤组件，所述主机包含处理器，所述处理器加载有程序代码，所述程序代码被执行时实现权利要求1-9任一项所述的消融计算方法。
PCT/CN2022/075489 2021-02-08 2022-02-08 一种消融计算方法及消融计算系统 WO2022166982A1 (zh)

Priority Applications (1)

Application Number Priority Date Filing Date Title
CN202280006671.1A CN116324459A (zh) 2021-02-08 2022-02-08 一种消融计算方法及消融计算系统

Applications Claiming Priority (2)

Application Number Priority Date Filing Date Title
CN202110184184.1 2021-02-08
CN202110184184.1A CN114764133B (zh) 2021-02-08 2021-02-08 一种消融计算方法及消融计算系统

Publications (1)

Publication Number Publication Date
WO2022166982A1 true WO2022166982A1 (zh) 2022-08-11

Family

ID=82365215

Family Applications (1)

Application Number Title Priority Date Filing Date
PCT/CN2022/075489 WO2022166982A1 (zh) 2021-02-08 2022-02-08 一种消融计算方法及消融计算系统

Country Status (2)

Country Link
CN (2) CN114764133B (zh)
WO (1) WO2022166982A1 (zh)

Cited By (1)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
CN117249922A (zh) * 2023-11-17 2023-12-19 山东盈动智能科技有限公司 一种用于温度测试仪的温度校准方法及系统

Citations (10)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US5711300A (en) * 1995-08-16 1998-01-27 General Electric Company Real time in vivo measurement of temperature changes with NMR imaging
US20050065429A1 (en) * 2003-09-18 2005-03-24 Ge Medical Systems Global Technology Company, Llc Method for three plane interleaved acquisition for three dimensional temperature monitoring with MRI
CN1993628A (zh) * 2004-08-02 2007-07-04 皇家飞利浦电子股份有限公司 包括用相位图表示和被用作为相位参考的参考媒体的mri测温技术
US20090238429A1 (en) * 2008-03-19 2009-09-24 Joerg Roland Magnetic resonance tomography method and apparatus for correcting drift of the b0 field in a temperature map
CN103403567A (zh) * 2011-03-01 2013-11-20 皇家飞利浦有限公司 涉及图像比约束重建的加速mr温度测定绘图
CN107468251A (zh) * 2017-07-03 2017-12-15 中国科学技术大学 一种低场磁共振温度成像相位漂移的校正方法
CN108245158A (zh) * 2016-12-29 2018-07-06 中国科学院深圳先进技术研究院 一种磁共振温度测量方法及装置
CN108652627A (zh) * 2018-03-13 2018-10-16 安徽锐捷信息科技有限公司 一种磁共振温度成像方法及装置
CN110430834A (zh) * 2017-03-20 2019-11-08 皇家飞利浦有限公司 使用电性能断层摄影来定位消融组织
CN111714097A (zh) * 2020-06-30 2020-09-29 杭州佳量医疗科技有限公司 一种基于多梯度回波序列的双模态磁共振测温方法

Family Cites Families (7)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US7542793B2 (en) * 2002-08-22 2009-06-02 Mayo Foundation For Medical Education And Research MR-guided breast tumor ablation and temperature imaging system
CN105748151A (zh) * 2008-06-18 2016-07-13 工程服务公司 Mri兼容的具有校准人造模型和人造模型的机器人
US8326010B2 (en) * 2010-05-03 2012-12-04 General Electric Company System and method for nuclear magnetic resonance (NMR) temperature monitoring
DE102012208425B4 (de) * 2012-05-21 2013-12-12 Siemens Aktiengesellschaft Kontinuierliches Korrigieren von Phasenfehlern einer multidimensionalen, ortsselektiven Magnetresonanz-Messsequenz
CN107271937B (zh) * 2017-07-04 2019-07-23 南京拓谱医疗科技有限公司 一种三维多参数加权磁共振成像的同步采集和校准方法
CN110244245B (zh) * 2019-06-10 2021-09-07 苏州润蓝医疗科技有限公司 一种优化的磁场漂移矫正方法及装置
CN111374645A (zh) * 2020-03-24 2020-07-07 聚融医疗科技(杭州)有限公司 一种热消融实时监控的呼吸伪像校正方法及系统

Patent Citations (10)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
US5711300A (en) * 1995-08-16 1998-01-27 General Electric Company Real time in vivo measurement of temperature changes with NMR imaging
US20050065429A1 (en) * 2003-09-18 2005-03-24 Ge Medical Systems Global Technology Company, Llc Method for three plane interleaved acquisition for three dimensional temperature monitoring with MRI
CN1993628A (zh) * 2004-08-02 2007-07-04 皇家飞利浦电子股份有限公司 包括用相位图表示和被用作为相位参考的参考媒体的mri测温技术
US20090238429A1 (en) * 2008-03-19 2009-09-24 Joerg Roland Magnetic resonance tomography method and apparatus for correcting drift of the b0 field in a temperature map
CN103403567A (zh) * 2011-03-01 2013-11-20 皇家飞利浦有限公司 涉及图像比约束重建的加速mr温度测定绘图
CN108245158A (zh) * 2016-12-29 2018-07-06 中国科学院深圳先进技术研究院 一种磁共振温度测量方法及装置
CN110430834A (zh) * 2017-03-20 2019-11-08 皇家飞利浦有限公司 使用电性能断层摄影来定位消融组织
CN107468251A (zh) * 2017-07-03 2017-12-15 中国科学技术大学 一种低场磁共振温度成像相位漂移的校正方法
CN108652627A (zh) * 2018-03-13 2018-10-16 安徽锐捷信息科技有限公司 一种磁共振温度成像方法及装置
CN111714097A (zh) * 2020-06-30 2020-09-29 杭州佳量医疗科技有限公司 一种基于多梯度回波序列的双模态磁共振测温方法

Cited By (2)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
CN117249922A (zh) * 2023-11-17 2023-12-19 山东盈动智能科技有限公司 一种用于温度测试仪的温度校准方法及系统
CN117249922B (zh) * 2023-11-17 2024-03-08 山东盈动智能科技有限公司 一种用于温度测试仪的温度校准方法及系统

Also Published As

Publication number Publication date
CN114764133B (zh) 2023-08-08
CN116324459A (zh) 2023-06-23
CN114764133A (zh) 2022-07-19

Similar Documents

Publication Publication Date Title
Kolandaivelu et al. Noninvasive assessment of tissue heating during cardiac radiofrequency ablation using MRI thermography
Baron et al. In vivo T2‐based MR thermometry in adipose tissue layers for high‐intensity focused ultrasound near‐field monitoring
US9289154B2 (en) Techniques for temperature measurement and corrections in long-term magnetic resonance thermometry
US8368401B2 (en) Techniques for correcting measurement artifacts in magnetic resonance thermometry
de Senneville et al. Motion correction in MR thermometry of abdominal organs: a comparison of the referenceless vs. the multibaseline approach
US6559644B2 (en) MRI-based temperature mapping with error compensation
WO2011028505A1 (en) Techniques for correcting temperature measurement in magnetic resonance thermometry
US8024025B2 (en) T1-corrected proton resonance frequency shift thermometry
WO2022166982A1 (zh) 一种消融计算方法及消融计算系统
Botnar et al. Temperature quantification using the proton frequency shift technique: In vitro and in vivo validation in an open 0.5 tesla interventional MR scanner during RF ablation
Sinha et al. Phase imaging on a. 2‐T MR scanner: application to temperature monitoring during ablation procedures
Le Ster et al. RF heating measurement using MR thermometry and field monitoring: methodological considerations and first in vivo results
CN114910855B (zh) 一种磁共振温度成像方法及相关装置
de Senneville et al. MRI-guided high-intensity focused ultrasound sonication of liver and kidney
CN114814689B (zh) 一种磁共振温度成像方法
Shmatukha et al. Correction of proton resonance frequency shift temperature maps for magnetic field disturbances caused by breathing
Krahn et al. MRI-guided cardiac RF ablation for comparing MRI characteristics of acute lesions and associated electrophysiologic voltage reductions
Rempp et al. Prediction of cell necrosis with sequential temperature mapping after radiofrequency ablation
CN114754890A (zh) 一种磁共振温度成像方法与系统
JP4574781B2 (ja) 磁気共鳴装置及び温熱治療装置
US11644520B2 (en) Systems and methods for magnetic resonance based skull thermometry
Toupin et al. Combination of principal component analysis and optical-flow motion compensation for improved cardiac MR thermometry
CN114820838B (zh) 校正磁化率误差的磁共振温度成像方法
CN114764034A (zh) 一种减少运动误差的磁共振测温方法
De Landro et al. Analysis of cavitation artifacts in Magnetic Resonance Imaging Thermometry during laser ablation monitoring

Legal Events

Date Code Title Description
121 Ep: the epo has been informed by wipo that ep was designated in this application

Ref document number: 22749261

Country of ref document: EP

Kind code of ref document: A1

NENP Non-entry into the national phase

Ref country code: DE

122 Ep: pct application non-entry in european phase

Ref document number: 22749261

Country of ref document: EP

Kind code of ref document: A1