Abstract:
An apparatus for deriving tissue temperature from thermal strain includes a thermal strain measuring module. The module uses ultrasound ( 156, 158 ) to measure thermal strain in a region, within a subject, that surrounds a location ( 166   a,    166   f ) where a temperature sensor is disposed. Also included is a temperature measurement module configured for, via the sensor, measuring a temperature at the sensor while the sensor is inside the subject. Further included is a patient-specific thermal-strain-to-temperature-change proportionality calibration module. The calibration module is configured for calibrating (S 238 ) a coefficient and for doing so based on a measurement of a temperature parameter at that location derived from output of the temperature measurement module and on a measurement of thermal strain at that location obtained via the strain measuring module. The coefficient is usable, in conjunction with a thermal strain measurement derived from another location within the region, in evaluating (S 242 ), for that other location, another temperature parameter.

Description:
FIELD OF THE INVENTION 
       [0001]    The present invention relates to using ultrasound thermal strain measurements to determine temperature and, more particularly, to such use in conjunction with a thermal sensor utilized internally. 
       BACKGROUND OF THE INVENTION 
       [0002]    Thermal ablation techniques are an excellent alternative to major surgery, which can pose a risk even with the most experienced surgeon. These techniques are minimally invasive, requiring only needles (radiofrequency (RF), cryotherapy and microwave ablation) or a non-invasive heat source such as by using ultrasound, e.g., high-intensity focused ultrasound (HIFU). In most of the procedures, the cancerous tissue is heated to above 60° Celsius (C) and subject to necrosis. 
         [0003]    Radiofrequency ablation (RFA) is currently the only FDA approved minimally invasive heating therapy in the United States. RF ablation uses a probe with an active electrode tip through which a 460-500 kilohertz (KHz) alternating current is conducted. The current propagates through the body to the grounding pads placed either on the back or the thigh of the patient. The current causes ionic agitation and frictional heating. Heat is then dissipated through thermal conduction to ablate the tumor. RFA is frequently used to treat liver cancer. There are about 500,000 new cases of metastatic liver cancer in the western world and about 1 million new cases for primary liver cancer worldwide (83% of which are in developing countries). RFA and microwave ablation therapies are also gaining tremendous popularity in China due to the large number of liver cancers reported (e.g., 433,000 new cases in 2009 alone). Current treatment protocols use the simplistic spherical ablation volume predicted from the device manufacturers&#39; specifications. The actual treatment volumes greatly deviate from the prediction, resulting in large recurrence rates (approx. 35%). 
         [0004]    RF ablation is typically performed under ultrasound, computed tomography (CT) or magnetic resonance imaging (MRI) guidance. Follow up is done with a CT scan or MRI within a month to assess effectiveness of ablation and again at 3 month intervals along with tumor markers to detect residual disease or recurrence. One common reason for the high recurrence rates is the inability to monitor and control ablation size to adequately kill the tumor cells. Real-time feedback is accordingly provided to the clinician by means of a temperature map of the ablated zone. This can currently be achieved with reasonable accuracy with MR based temperature imaging. However, MRI is expensive and may not be readily available. Ultrasound is another modality that is commonly used for image guidance during placement of the needle. Due to its ease of use and availability it is a preferred method for monitoring the lesions. However, the only way it is currently used for monitoring treatment is by visualizing the hyperechoic lesions on a B-mode image. Low contrast exists between normal and ablated tissue. Visual artifacts arise from gas bubbles. Thus, the visualization currently afforded by ultrasound is only approximate and not a good indicator of the treatment efficacy. Also reliance on gas bubbles for echogenicity encounters the problem that bubble formation mainly occurs at temperatures elevated above those needed for the ablation, potentially resulting in unnecessary cell damage and prolongation of the procedure. 
         [0005]    Another proposed ultrasound technique for ablation monitoring is ultrasound thermometry. Ultrasound thermometry can potentially enable mapping the temperature distribution during thermal therapies in 3D spatial and temporal dimensions. Through the concept of thermal dose (derived from the time history of temperature rise), the extent of the ablation zone can be determined over the entire volume. Hence, ultrasound thermometry provides significant advantages over temperature measurements obtained from a single or a few thermocouples that provide only a sparse sampling of the ablation zone. The underlying principle of ultrasound thermometry is that the speed of sound in the tissue changes as a function of temperature which manifests as apparent shifts (displacement) in ultrasound echoes. The resulting temperature induced strain (derived by differentiating the displacement along the direction of the ultrasound beam) is nominally proportional to the temperature rise in the range up to 50° C. The proportionality constant (thermal strain to temperature coefficient) is typically estimated through a calibration performed in a water bath wherein a known temperature rise that produces the corresponding thermal strain is noted. One such study discloses calibration curves for different body tissue types. Varghese, T., Daniels, M. J., “Real-time calibration of temperature estimates during radiofrequency ablation”, Ultrasonic Imaging, 26(3):185-200 (2004) (hereinafter “Varghese”). The curves, which each relate temperature rise to thermal strain, are each seen to be essentially linear over a hypothermia temperature range which extends up to 50° C. Accordingly, a proportionality constant can be derived for each tissue type. U.S. Patent Publication No. 2013/0204240 to McCarthy discloses an integrated catheter tip (ICT) that includes a thermocouple. The ICT is used for hyperthermia therapy. Readings from the thermocouple are used to measure temperature adjacent to the ICT. A radiometer is also used in the measurement, because heating is caused by microwave energy and because a more complete picture of the temperatures in the treatment region is desired. 
       SUMMARY OF THE INVENTION 
       [0006]    The above-described Varghese method of proportionality factor calibration is feasible in laboratory studies and not in a clinical situation. Indeed, one could use calibration curves for a particular tissue type available from the literature. However, such values are only approximate with a high standard deviation arising from the difference in the method and local variations in the tissue composition. Even for a given tissue type, the temperature dependence of ultrasound propagation speed significantly varies, based on tissue composition, e.g., water content and fat content. The composition, for a given patient, can locally vary even within same organ such as the liver. Hence, for a given patient and subject an in situ estimate of the proportionality factor, i.e., the temperature-strain coefficient, affords greater accuracy in knowing the local temperatures throughout the intended ablation region  160 . The accurately measured temperatures can be inputted into a thermal model to accurately predict temperatures in the ablation regime. For the in situ estimate, it is proposed herein below to obtain a reliable “ground truth” temperature value in vivo at the site of thermal treatment, as via a thermocouple onsite. The thermocouple may be at the tip of a tine of an RF ablation electrode. Applications of the inventive technology also extend to hyperthermia therapy. In McCarthy, for example, in which a thermocouple is used in hyperthermia therapy, ultrasound thermometry would offer an economical and safe alternative to microwaves for the regional temperature monitoring. Using the patient-specific coefficient proposed herein makes the ultrasound-thermometry-based monitoring more accurate. 
         [0007]    In an aspect of what is proposed herein, an apparatus for deriving tissue temperature from thermal strain includes a thermal strain measuring module. The module uses ultrasound to measure thermal strain in a region, within a subject, that surrounds a location where a temperature sensor is disposed. Also included is a temperature measurement module configured for, via the sensor, measuring a temperature at the sensor while the sensor is inside the subject. Further included is a patient-specific thermal-strain-to-temperature-change proportionality calibration module. The calibration module is configured for calibrating a coefficient and for doing so based on a measurement of a temperature parameter at that location derived from output of the temperature measurement module and on a measurement of thermal strain at that location obtained via the strain measuring module. The coefficient is usable, in conjunction with a thermal strain measurement derived from another location within the region, in evaluating, for that other location, another temperature parameter. 
         [0008]    In the ablation context and operationally, the clinician performs a test shot or heating to a few degrees and ultrasound data is collected. Ultrasound strain estimates are obtained over the entire intended ablation region, and the patient specific coefficient is determined. With this coefficient, temperature estimates are obtained over the region. Since the normal temperature of the human body is 37° C., the temperature estimates are below 50° C., i.e., in the hyperthermia range. A model is now used to predict temperatures in the ablative range. The input to the model is ultrasound determined temperature estimates, and ablation device parameters like power and impedance. The model is then run with various combinations of thermal conductivities and electrical conductivities. This is done as an optimization to best match an output temperature distribution with that obtained by the test shot. The optimization operates on the equations below: 
         [0000]    
       
         
           
             
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         [0000]    where k is the thermal conductivity, p is the density, C is the specific heat, a is the electrical conductivity. 
         [0009]    The model is then re-initialized with the determined k and σ and run with these values to predict ablative temperatures. The above test shot and subsequent model initialization procedure can be completed in 3-4 minutes. Now the clinician is ready to begin the ablation procedure. In this mode, as the therapy progresses, real-time power and impedance profiles are passed on to the model from the ablation generator. These profiles are part of a database of various temperature profiles with different values of electrical conductivity and thermal conductivity, the profiles having been generated a priori, even before the patient is on the table. Each of the profiles pertains to a particular output power of the RF ablation generator and impedance in the electrical flow from the RF ablation generator, through the electrode and onto the pads in completing the circuit. The profile links the output power and impedance to temperature increments throughout the region. The model calculates the current ablation temperature throughout a three-dimensional (3D) volume at each time step for the power and impedance input from the generator, and a thermal dose contour progresses as the therapy progresses. This progression is visualized on the screen in real-time. At the discretion of the clinician, or via automatic image matching to the intended ablation region, the therapy is stopped as the contour covers the tumor boundary with a margin. A more complex model could have heterogeneous zones of k and σ and not just one k and σ for the entire tissue. An example of a thermal model is provided in commonly-owned International Publication No. WO 2014/076621 to Anand et al. 
         [0010]    Details of the novel technology for patient-specific ultrasound thermal strain to temperature coefficient calibration are set forth further below, with the aid of the following drawings, which are not drawn to scale. 
     
    
     
       BRIEF DESCRIPTION OF THE DRAWING 
         [0011]      FIG. 1  is a schematic and conceptual diagram exemplary of patient-specific ultrasound thermal strain to temperature coefficient calibration in accordance with the present invention; and 
           [0012]      FIG. 2  is flow chart of a particular variation on methodology performable with the structures shown in  FIG. 1 . 
       
    
    
     DETAILED DESCRIPTION OF EMBODIMENTS 
       [0013]    FIG. depicts, by illustrative and non-limitative example, an apparatus  100  for deriving tissue temperature from thermal strain. The apparatus  100  includes an RF ablation generator  102 , and energy exchange and sensing device  104 , RF grounding pads  106 , and an ultrasound imaging system  108 . 
         [0014]    The RF ablation generator  102  includes a temperature measurement module  110  and a communication module  112 . 
         [0015]    The energy exchange and sensing device  104  includes an ablation needle  114  and a needle holder  116 . 
         [0016]    Included in the ultrasound imaging system  108  are an ultrasound imaging probe  118 , a thermal-strain measuring module  120 , a patient-specific thermal-strain-to-temperature-change proportionality calibration module  122 , a therapy monitoring module  124 , a display  126 , and a user control interface  128 . 
         [0017]    An RF ablation electrode  130  is incorporated within the ablation needle  114  and comprises one or more tines  132 . Each tine  132  has, at a distal end, a tip. Offset slightly in from the tip is a thermocouple  138  or other thermal sensor. 
         [0018]    All of the modules, and other data processing elements, may be implemented in any known and suitable combination of hardware, software and/or firmware. 
         [0019]    Also, instead of an ablation needle, a catheter may deliver the electrode  130 . 
         [0020]    In addition, instead of an electrode for ablation, another ablation technique that uses internal temperature sensors may be employed, such as microwave ablation via microwaves delivered by the energy exchange and sensing device  104 . 
         [0021]    Non-ablation applications such as hyperthermia-based therapy in which ultrasound thermometry is used to monitor temperature are also within the intended scope of what is proposed herein. 
         [0022]    The imaging probe  118  may be trans-thoracic and an internal probe such as a transesophageal echocardiography (TEE) probe. 
         [0023]    Shown on the display  126  in  FIG. 1  for illustrative purposes is a B-mode ultrasound image  140 . Overlaid on the image  140  is a temperature map  142 . 
         [0024]    The apparatus  100  is operable in an coefficient calibration mode  144  (conceptually depicted in conjunction with a switchable arrow  146  in  FIG. 1 ) to, as shown by the formula  147 , calibrate a patient-specific ultrasound thermal strain to temperature coefficient  148  which serves as a proportionality factor between a thermal strain  150  and a temperature differential  151 . In coefficient calibration mode  144 , the RF ablation generator  102  is operated at a low power, keeping the temperature in the tissue below 50° C. It may be kept at 43° C., for instance, or in a range of 37° C. to 43° C. 
         [0025]    This is followed by a hyperthermia temperature-field determination mode  152 . In this mode  152 , the calibrated coefficient  148  is applied to thermal strain  150  that has been calculated for locations throughout the intended ablation region. 
         [0026]    A thermal model initialization mode  153  uses the determined temperature field and ablation device parameters to initialize the model for tissue properties discernable from comparing the temperature field to output temperatures of the model. 
         [0027]    The apparatus  100  is also operable in a body tissue ablation and concurrent model execution mode, or “tissue ablation mode”,  154  in which the RF ablation generator  102  is operated at a higher power, for ablation. The tissue is heated to above 55° C. and typically above 60° C. The model also operates ongoingly in the tissue ablation mode  154 . Ablation therapy is performed on a human, or animal, patient. 
         [0028]    Pulses  156  of ultrasound are emitted in the coefficient calibration mode  144 , and the return pulses  158  are analyzed to assess thermal strain in the intended ablation region  160 . Measurements of thermal strain  150  in the coefficient calibration mode  144  are taken at the thermocouples  138 , e.g., within a radius centered at the thermocouple of twice an ultrasonic spatial resolution (lateral or axial) of the apparatus  110 , and are used to calibrate the coefficient  148 . 
         [0029]    For an ablation needle  162 , each of one or more tines  164   a - g  has at its distal end  163  a respective thermocouple  138 . 
         [0030]    Partially or fully surrounding a location  166   a  of the thermocouple  138  for the tine  164   a  is a volumetric region  168   a  to be associated with a particular calibrated coefficient  148  that is to be computed. Likewise as an example,  FIG. 1  shows a second volumetric region  168   f  surrounding a location  166   f  of a respective thermocouple  138 . Although the regions  168   a ,  168   f  are portrayed as spherical, they can be any arbitrary shape. 
         [0031]    Although each region  168   a ,  168   f  is to be associated with a particular coefficient  148 , the value of the coefficient when computed for each of two different regions may turn out to be the same. They can be the same or almost the same if the tissue composition in the immediate vicinity of both respective locations  166   a ,  166   f  is the same or almost the same. A hypothetical tissue-composition-based divider  170 , which can actually be constructed by the user interactively onscreen, is shown in  FIG. 1 . Thus, the coefficient  148  can be expected to be calibrated to a different value for regions  168   a  on one side of the divider  170  than for the regions  168   f  on the other side of the divider. 
         [0032]    Regions  168   a ,  168   f  may overlap. Even if, for example, regions  168   a ,  168   f  are truncated at the divider  170 , regions on the same side of the divider may overlap. For the first region  168   a , for instance, other than the surrounded or thermocouple location  166   a , there is another location  172   a , and there are additional locations  174   a ,  176   a . When the temperatures at the additional locations  174   a ,  176   a  are estimated, i.e., in the hyperthermia temperature-field determination mode  152 , the coefficient  148  for the first region  168   a  is utilized. However, if the other location  172   a  is also within the adjoining region (not shown), a selection can be made between the regions sharing the location, or a combination such as an average of respective coefficients  148  can be computed. The average may be weighted by distance of the location  172   a  to the respective thermocouple locations  166   a ,  166   f  or, in the case of selection, selection can be made of the based on the closest thermocouple location. 
         [0033]      FIG. 2  is a flow chart exemplary of a procedure  200  for deriving tissue temperature from thermal strain  150 . The procedure  200  is performed serially through the above-described modes  144 ,  152 ,  153 ,  154 , transitioning mode-to-mode through the series automatically, without the need for user intervention. The needle holder  116  is attached fixedly to the probe  118  (step S 202 ). With the needle holder  116  attached, the probe  118  is positioned manually or via motorized movement to bring the tumor to be ablated into the field of view of the probe (step S 204 ). If the probe  118  has a 2D transducer array, the probe can, with the tumor within the field of view, be held motionless throughout the procedure  200 , either manually or by the motorized mechanism, for ablation of all tumorous body tissue within the field of view. Cyclical body motion, such as respiratory or cardiac, can be automatically and dynamically compensated through a combination of motion gating and ultrasound speckle-based motion tracking. With the probe  118  in place, the needle  114  can be manually advanced through the needle holder  116  and into the subject  180  under operator control by a distance, and at an orientation, that are readable from the needle holder. For example, the proximal end of the needle  114  can have graded markings that show how far the needle has been advanced. This information is entered via the user control interface  128  (step S 206 ). Accordingly, the tip of the needle  114  is at a known location in image space and is into or just short of the tumor. The clinician viewing the tumor interactively delimits and defines the intended ablation region  160  onscreen (step S 208 ). Under operator control, the one or more tines  164   a - g  are extended (step S 210 ). The tines  164   a - g  are stiff and extend invariantly into the body tissue, mainly or entirely tumorous, that is being pierced. Thus, the thermocouple locations  166   a ,  166   f  on the tines  164   a - g  and slightly offset from the tine tips are known. Alternatively, X-rays from a CT or fluoroscopy system registered to the ultrasound imaging system  108  can be employed to localize the locations  166   a ,  166   f . The coefficient calibration mode  144  is then initiated (step S 212 ). The RF ablation generator  102  is operated at a low power keeping the temperature in the tissue below 50° C. It may be kept at 43° C., for instance, or in a range of 37° C. to 43° C. The RF ablation generator  102 , in effect, sets the heating of the electrode  130  to a pre-designated temperature, or temperature range, that is below the maximum temperature of, for example, 50° C. (step S 214 ). Also, at this point, the RF generator begins self-checking the temperatures at all thermocouples  138  and regulates the temperatures ongoingly, in both the current coefficient calibration mode  144  and throughout the above-discussed ensuing modes  152 - 154 . The thermocouple temperatures are thus maintained to whatever is the current set temperature or temperature range (step S 215 ). A temperature reading is now taken by all thermocouples  138  at their respective locations  166   a ,  166   f  (step S 216 ). An ultrasound pulse  156  is issued in a current direction in volumetric space (step S 218 ). From a return echo pulse  158  in the same direction, an A-line is acquired and recorded (step S 220 ). Then, steps S 218  and S 220  are repeated in each direction for the intended ablation region  160 . In particular, a pulse  156  is issued (step S 222 ) and an A-line is acquired and recorded (step S 224 ). For a 2D ultrasound transducer array, the scan may proceed from an elevationally high row of scan lines progressively downward. Alternatively, a one-dimensional array can be pivoted mechanically for a similar scan. After the scan, the RF ablation generator then raises the temperature at the respective electrodes, and steps S 216 -S 224  are repeated (step S 226 ). The two A-lines of a current direction are cross-correlated (step S 228 ). The resulting offset is used to divide the two A-lines into segments such that a segment of one A-line is paired with what is, in view of the offset, a spatially close segment of the other A-line (step S 230 ). Segments of a pair are cross-correlated to fine tune the global offset to a local value, this being done for each pair (step S 232 ). The local values are the apparent displacements usable in computing thermal strain  150 . In particular, the local displacements are differentiated in the current, i.e., axial  182 , direction to yield the local value of the thermal strain  150  (step S 234 ). The local strain values are stored (step S 235 ). The local temperature differentials  151  are obtained by subtracting the temperature read in step S 216  from the temperature reading in step S 227  for each thermocouple  138  (step S 236 ). The coefficient(s)  148  are calibrated by, at the locations  166   a ,  166   f  of the thermocouples  138 , respectively evaluating the formula  147  with the local temperature differential  151  and the local value of the thermal strain  150  (step S 238 ). The apparatus  100  now transitions to the thermal model initialization mode  153  (step S 240 ). The stored local strain values of all directions are respective multiplied by the calibrated coefficient  148  of the respective volumetric regions  168   a ,  168   f , or, for locations  172   a  in region overlap, optionally by an averaged coefficient (step S 242 ). The respective products, i.e., temperature differentials  151  that have been evaluated, are added to the corresponding, ambient starting temperatures, typically about 37° C., measured in step S 216  (step S 244 ). The resulting sums for the associated locations  172   a  constitute a hyperthermia temperature field  184  that, in the thermal model initialization mode  153 , is inputted into a thermal model  186  (step S 246 ). The thermal model  186  is then run with various combinations of thermal conductivities and electrical conductivities (step S 248 ). For the best match of the temperature field with the model-generated temperature field, the utilized thermal and electrical conductivities are determined (step S 250 ). The model  186  is re-initialized with these two parameters (step S 252 ). In the case of a model for liver tissue, typical model parameters are, for instance, an electrical conductivity of 0.148 Siemens per meter (S/m), a thermal conductivity of 0.465 watts per meter Celsius (W/mC), a density of 1060 kilograms per cubic meter (kg/m 3 ), a heat capacity of 3600 joules per kilogram Celsius (J/Ckg) and a perfusion rate of 6.4×10 −3 /second. 
         [0034]    In the tissue ablation mode  154 , real-time power and impedance profiles from the RF ablation generator  102  are time-step by time-step matched to current power and impedance values during ablation to extract respective temperature increments (step S 254 ). The increments are accumulated to yield in real time an ablation temperature field  188  (step S 256 ). Location-specific thermal dose measurements are ongoingly updated (step S 258 ). These measurements and/or current ablation temperatures can be thresholded to detect a stopping point for power production by the RF ablation generator  102  (step S 260 ). Thus, based on the calibrated coefficient  148 , monitoring is performed, during the provision of therapy, of temperature at one or more additional locations  174   a ,  176   a  within the region  168   a ,  168   f . Alternatively or in addition to the thresholding, one or more B-mode images  140  are acquired (step S 262 ) and color-coded temperature maps  142  corresponding to the real-time ablation temperature field  188  are overlaid over, or otherwise combined (e.g., alpha blended) with, the B-mode image(s) to form respective composite images  190  (step S 264 ). The clinician may accordingly visually judge when a stopping point for the heating has been reached and thus, via the user control interface  128 , halt power production by the RF ablation generator  102  (step S 266 ). Whether stopping is automatic or operator-initiated, the ultrasound imaging system  108  issues a command to the RF ablation generator  102  to halt heating via the RF ablation electrode  130  since ablation is now complete (step S 268 ). 
         [0035]    While the invention has been illustrated and described in detail in the drawings and foregoing description, such illustration and description are to be considered illustrative or exemplary and not restrictive; the invention is not limited to the disclosed embodiments. 
         [0036]    For example, instead of an overlaid temperature map, the map is displayable alongside the B-mode image. 
         [0037]    Other variations to the disclosed embodiments can be understood and effected by those skilled in the art in practicing the claimed invention, from a study of the drawings, the disclosure, and the appended claims. In the claims, the word “comprising” does not exclude other elements or steps, and the indefinite article “a” or “an” does not exclude a plurality. The word “exemplary” is used herein to mean “serving as an example, instance or illustration.” Any embodiment described as “exemplary” is not necessarily to be construed as preferred or advantageous over other embodiments and/or to exclude the incorporation of features from other embodiments. Any reference signs in the claims should not be construed as limiting the scope. 
         [0038]    A computer program can be stored momentarily, temporarily or for a longer period of time on a suitable computer-readable medium, such as an optical storage medium or a solid-state medium. Such a medium is non-transitory only in the sense of not being a transitory, propagating signal, but includes other forms of computer-readable media such as register memory, processor cache and RAM. 
         [0039]    A single processor or other unit may fulfill the functions of several items recited in the claims. The mere fact that certain measures are recited in mutually different dependent claims does not indicate that a combination of these measures cannot be used to advantage.