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(Radiology. 2000;216:517-523.)
© RSNA, 2000

Experimental Studies

Usefulness of MR Imaging-Derived Thermometry and Dosimetry in Determining the Threshold for Tissue Damage Induced by Thermal Surgery in Rabbits1

Nathan J. McDannold, BS, Randy L. King, BS, Ferenc A. Jolesz, MD and Kullervo H. Hynynen, PhD

1 From the Department of Radiology, Division of MRI, Brigham and Women’s Hospital, Harvard Medical School, 221 Longwood Ave, Longwood Medical Research Center, 007c, Boston, MA 02115 (N.J.M., R.L.K., F.A.J., K.H.H.), and the Department of Physics and Astronomy, Tufts University, Medford, Mass (N.J.M.). Received September 8, 1999; revision requested October 29; revision received November 29; accepted January 12, 2000. Supported by National Cancer Institute grants CA 46627 and P01 CA 67165 and contract 282-97-0080 from the U.S. Department of Health and Human Services. Address correspondence to N.J.M. (e-mail: njm@bwh.harvard.edu).


   ABSTRACT
TOP
 ABSTRACT
INTRODUCTION
 MATERIALS AND METHODS
RESULTS
DISCUSSION
 REFERENCES
 
PURPOSE: To investigate in vivo the feasibility of using magnetic resonance (MR) imaging–derived temperature and thermal dose measurements to find the threshold of thermal tissue damage.

MATERIALS AND METHODS: Sonications were delivered in rabbit thigh muscles at varying powers. Temperature-sensitive MR images obtained during the sonications were used to estimate the temperature and thermal dose. The temperature, thermal dose, and applied power were then correlated to the occurrence of tissue damage observed on postsonication images. An eight-element phased-array transducer was used to produce spatially flat temperature profiles that allowed for averaging to reduce the effects of noise and the voxel size.

RESULTS: The occurrence of tissue damage correlated well with the MR imaging–derived temperature and thermal dose measurements but not with the applied power. Tissue damage occurred at all locations with temperatures greater than 50.4°C and thermal doses greater than 31.2 equivalent minutes at 43.0°C. No tissue damage occurred when these values were less than 47.2°C and 4.3 equivalent minutes.

CONCLUSION: MR imaging thermometry and dosimetry provide an index to predict the threshold for tissue damage in vivo. This index offers improved online control over minimally invasive thermal treatments and should allow for more accurate target volume coagulation.

Index terms: Magnetic resonance (MR), thermometry, 44.121412 • Ultrasound (US), experimental studies, 44.1298 • Ultrasound (US), focused, 44.1298 • Ultrasound (US), therapeutic, 44.1298


   INTRODUCTION
TOP
 ABSTRACT
 INTRODUCTION
MATERIALS AND METHODS
RESULTS
DISCUSSION
 REFERENCES
 
One of the main reasons focused ultrasound surgery and other noninvasive thermal therapies have not been used widely in the clinical setting has been the difficulty in determining the relative success or failure of the treatments. However, with the development of image guidance, especially with magnetic resonance (MR) imaging, these treatments offer many promising clinical uses. MR imaging has been shown to be useful in guiding and monitoring thermal surgery (1–6). It can be used to define the tissue to be treated, to measure induced temperature changes (7–9), to ensure proper targeting by localizing subthreshold heating (10), and to image the tissue after the thermal treatment (3,11–13).

Although MR imaging can depict tissue damage created by thermal therapies, in some cases it may be impossible to do so. For example, it may be hard to distinguish thermally damaged tissue in tumors from the tumor itself (11) and the edema that surrounds it. It also may not be desirable to cause tissue coagulation at all, as in hyperthermia treatments, for example. Thus, having another method beyond posttreatment imaging to determine the extent of tissue damage is necessary for optimal thermal treatments.

Three methods commonly are used to predict whether tissue damage induced by thermal therapies will occur. The first method uses the power output of the treatment device and the exposure length to predict tissue damage (14–18). The second method postulates a critical temperature, Tc, above which tissue damage occurs (19). The third method uses the entire temperature history of the tissue to estimate the thermal dose (20).

Temperature and thermal dose measurements have the advantage of being independent of inhomogeneities in the acoustic, optic, or electric properties of tissue and rates of perfusion. Thermal dose measurements have the advantage of being independent of the length and shape of the temporal heating profile. Thermal dose typically has been used in hyperthermia treatments, in which the heating profile is long and stays near 43.0°C, but also has been suggested for use during short-duration, high-temperature, focused ultrasound surgery (21). Others have investigated the use of MR imaging–derived dosimetry for predicting the size of lesions produced during focused ultrasound surgery (22–24). The purpose of this study was to investigate whether MR imaging–derived thermometry and dosimetry of thermal exposures can be used to quantify the threshold of tissue damage in vivo.


   MATERIALS AND METHODS
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
RESULTS
DISCUSSION
 REFERENCES
 
Ultrasound
The ultrasound fields were generated by spherically curved, air-backed, phased-array transducers. The transducers had eight elements cut in a sector-vortex configuration (25–27). Each had a radius of curvature of 8 cm and a diameter of 10 cm. Two separate transducers were used over the course of the experiments with resonant frequencies of 1.1 and 1.6 MHz. The transducers were powered by a multichannel amplifier system that was constructed in-house (28).

The sector-vortex array typically is phased in several modes. In mode zero, each element of the transducer is in phase and is equivalent to a single-element transducer. This mode was used at a power high enough to localize the focus but low enough not to induce tissue damage (10). In mode four, alternating elements of the array are 180° out of phase. In this mode, the ultrasound field is canceled out at the natural focus, and eight peaks are created that are arranged radially about the direction of the ultrasound beam in the shape of a ring. As the ultrasound is delivered, conduction fills in the center of this ring and produces a spatially flat temperature profile. This mode was used in the experiments to create a large focal volume so that the induced temperature rise could be averaged over a section of tissue. A hydrophone scan of the focal plane in mode four is shown in Figure 1. The sonications were 30 seconds in duration. The characteristics of the ultrasound field were measured with a previously described method (1) and are summarized in Table 1.



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  		Figure 1. Hydrophone scan of the beam intensity pattern created by the sector-vortex phased-array transducer in mode four. A 0.75-mm-diameter hydrophone with a step size of 0.20 mm was used. The scan was obtained at the focal plane, perpendicular to the direction of the ultrasound beam.

 

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  		TABLE 1. Characteristics of the Sector-Vortex Transducers
 
Animal Experiments
Five to 10 sonications were delivered 1.5–2.0 cm into each thigh muscle of 13 male New Zealand white rabbits (Milbrook Breeding Labs, Amherst, Mass). Before the experiments, the hair on each thigh was removed with animal hair clippers and depilatory lotion. The animals were anesthetized with a mixture of sodium xylazine hydrochloride (12 mg per kilogram of body weight per hour; Xylaject; Phoenix Pharmaceuticals, St. Joseph, Mo) and ketamine hydrochloride (48 mg/kg/h; Ketalar; Abbott Laboratories, North Chicago, Ill). The anesthetic was injected as an intramuscular bolus every hour. The rabbits were provided housing, food, and veterinary care according to National Institutes of Health and Harvard Medical School guidelines (29).

The rabbits’ body temperatures were monitored constantly with copper-constantan thermocouples that were constructed in-house. One thermocouple was inserted rectally. In some experiments, another thermocouple was implanted in the thigh muscle a few centimeters away from the location of the sonications. The temperature in the thigh was used as the baseline temperature in the experiments. When the thermocouple in the thigh was not used, the rectal thermocouple temperature was used as the baseline.

Experimental Setup
The transducer was mounted in a computer-controlled, mechanically driven, MR imaging–compatible positioning system (GE Medical Systems, Milwaukee, Wis). An earlier version of the positioning system has been described (1,2). The rabbit’s thigh was placed in a hole in a plastic tray positioned above the transducer. The hole was the opening of a bag of deionized, degassed water. The bag rested on a polyvinyl chloride membrane above the transducer that also was submerged in degassed water. To provide proper coupling, a layer of degassed water was poured between the membrane and the plastic bag. A receive-only surface coil with a diameter of 12.7 cm was attached beneath the plastic tray below the thigh to improve the signal-to-noise ratio.

To maintain its body temperature, the rabbit was placed on a plastic mat that circulated heated water. A plastic coil that circulated warm water encircled the water bag to ensure further that the thigh was kept warm during the experiments and to avoid temperature gradients in the tissue. The temperature in the water was monitored with a thermocouple and was kept as close to the body temperature of the rabbit as possible. The whole apparatus was put into a clinical 1.5-T MR imager (Signa; GE Medical Systems). Figure 2 shows a diagram of the experimental setup.



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  		Figure 2. Diagram shows the experimental apparatus. MRI = MR imaging, RF = radio frequency.

 
Temperature Imaging
Temperatures were measured by exploiting the temperature dependence of the proton-resonant frequency (9). This method recently was shown to be the most sensitive of the current MR imaging thermometry methods commonly used (30). Changes in the proton-resonant frequency were estimated (N.J.M.) by measuring changes in phase and dividing by 2{pi} times the time in which the phase developed (the echo time of the image). A fast spoiled gradient-echo sequence was used to acquire the phase maps (31). The following parameters were used: 50.6/25 (repetition time msec/echo time msec); flip angle, 30°; bandwidth, 3.1 kHz; field of view, 16 cm; section thickness, 3 mm; matrix size, 256 x 128; imaging time per image, 6.7 seconds. An echo time of 25 msec was used because the optimal signal-to-noise ratio in fast spoiled gradient-echo phase-difference imaging occurs when the echo time is equal to T2*, and the T2* of muscle is around 25 msec (31). The small temperature gradients and the relatively long sonication time allowed for the long imaging time required when the optimal echo time was used.

The temperature dependence of the proton-resonant frequency shift in rabbit skeletal muscle is 0.00909 ppm/°C and is linear above the tissue coagulation threshold temperature (24). With an echo time of 25 msec and a magnetic field strength of 1.5 T, the temperature dependence of the sequence was 11.0°C/radian. Twelve sets of magnitude, real, and imaginary images were acquired, which yielded a total imaging time of 80 seconds. A reference phase map was acquired 7 seconds before the application of the ultrasound. The phase-difference images were constructed from this reference image by using a complex phase-subtraction method (31).

To avoid phase wraparound, the images were subtracted pairwise, and then the phase was summed along the time dimension. For the nth phase subtraction {Delta}{phi}n:

where {phi}i was the phase at the ith time point. This method prevented phase wrapping as long as the temperature did not cause a phase change greater than {pi} radians in the time of one image acquisition ({Delta}T <= 34.5°C in 6.7 seconds).

The images were acquired in a plane perpendicular to the direction of the ultrasound beam to reduce the effects of averaging of the temperature over the 3-mm section thickness. The position of the focus was found by localizing the maximum voxel value in the region of the focus. The temperature at each voxel was averaged with its eight nearest neighbors to reduce the effects of noise.

Thermal Dose Calculation
As proposed by Chung et al (22), the Sapareto and Dewey equation (20) was used to estimate the thermal dose delivered to the tissue from the time-temperature profiles gathered by means of MR imaging thermometry:

where t43 is the thermal dose in equivalent minutes at 43.0°C, T is the mean temperature during time {Delta}t, and R is a constant that equals 0.50 above 43.0°C and 0.25 below 43.0°C. The temperature profiles of each voxel were interpolated linearly with a step size of {Delta}t = 0.1 second.

Postsonication Imaging
After the sonications, T2- and T1-weighted fast spin-echo images were acquired for about 2 hours to determine whether there was a tissue response to the applied sonications. The parameters used in these images are shown in Table 2. The T1-weighted images were acquired before and after injection of a bolus of gadopentetate dimeglumine (Magnevist; Berlex Laboratories, Wayne NJ) at a concentration of 0.3 mmol/kg in each rabbit’s ear vein. After the injection, images were obtained until the enhancement reached its maximum. Contrast material was administered shortly after the last sonication in each thigh, and again roughly 2 hours after the last sonication.


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  		TABLE 2. Imaging Parameters for the T2- and T1-weighted Fast Spin-Echo Sequence Used to Characterize the Tissue Response to Sonications
 
In normal tissue, thermal lesions have been shown to appear as dark regions surrounded by a bright rim on T2-weighted images and as regions of no enhancement surrounded by regions of high enhancement on contrast material–enhanced T1-weighted images, and these images have been shown to correlate well with postmortem examination results (11,13,24,32). Because of the high contrast between normal tissue and tissue treated with focused ultrasound, no grading scheme was used to determine whether the tissue was damaged. The same author who calculated the temperature and dose information at each sonication evaluated the postsonication images for tissue damage. To avoid bias during this analysis, after the completion of all the experiments the images of each sonication were displayed in random order.


   RESULTS
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
RESULTS
DISCUSSION
 REFERENCES
 
Temperature and Thermal Dose Measurements
Figures 3 and 4 show examples of how the MR imaging–derived temperature measurements were analyzed. Figure 3 is an example of time series of temperature-sensitive images that show the development of a 19 W/30 sec sonication in mode four created with the 1.6-MHz array. The ring produced by the sector-vortex geometry is evident in the early stages of the ultrasound delivery, as well as how this ring fills in over time to yield a flat temperature profile.



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  		Figure 3. Temperature-sensitive phase subtractions of fast spoiled gradient-echo MR images (50.6/25) obtained during a 30-second sonication delivered in rabbit thigh. The temperatures measured from this time series of images were used to calculate the thermal dose. The images were acquired perpendicular to the direction of the ultrasound beam. The dimension of each subplot is 1.9 x 1.9 cm. s= seconds.

 


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  		Figure 4. Graphs show MR imaging-derived temperature and thermal dose profiles for the sonication shown in Figure 3. Top: Mean temperature rise versus time of nine voxels centered at the focus. Middle: Mean thermal dose versus time of nine voxels centered at the focus. Bottom: Spatial temperature profile through the center of the focus at peak temperature rise. The error bars in the top and bottom graphs are SDs.

 
Figure 4 shows the temperature and dose profiles for this sonication measured from the images in Figure 3. The top and middle sections of Figure 4 show the time development of the mean temperature and thermal dose of 9 voxels centered at the focus for this sonication. The bottom section of Figure 4 shows the spatial temperature profile through the focus at peak temperature rise. The flat temperature profile at the peak temperature rise and the relatively small temperature gradients are evident in these plots. A total of 138 such sonications were delivered in the course of all of the experiments. Forty-seven sonications were excluded from the results because of animal motion or visible noise in the images or because the sonications were on top of a blood vessel or muscle tissue boundary. Lesions on top of blood vessels were excluded because of flow-induced phase changes on the fast spoiled gradient-echo images.

The noise in the image subtractions was calculated by taking four 3 x 3-voxel regions away from the ultrasound focus for each sonication. The regions were chosen to be away from blood vessels, muscle boundaries, and fat. The mean and SD of these voxel regions corresponded to a temperature change of 0.8°C ± 0.3 (n = 3,916). The mean and SD of all of the individual voxels in these regions corresponded to -0.1°C ± 1.0 (n = 35,244). There was no noticeable increase in noise away from the focus when the ultrasound was turned on. At peak temperature rise, the mean SD of a 3 x 3-voxel region centered at the focus was 1.0°C ± 0.8 (n = 91). This corresponded to 8.4% ± 4.8 of the mean temperature rise, which was 12.4°C ± 2.9.

The thermocouple implanted in the thigh was not used during 31 of the 91 sonications analyzed in the experiment. In these cases, the rectal temperature was used for the baseline temperature. The mean difference between the rabbit body temperature measured with the thermocouple in the thigh and with that in the rectum was 0.4°C ± 1.0; the maximum difference was 2.5°C.

Postsonication Images
Figure 5 shows examples of the T2-weighted and contrast-enhanced T1-weighted images of lesions created by the focused ultrasound heating. These lesions were caused by 13.9-, 15.4-, and 21.6-W sonications. The lesions were characterized by a dark region surrounded by a bright ring on T2-weighted images and by a region of no contrast enhancement surrounded by a ring of strong enhancement on the T1-weighted images. The size of the lesions did not vary much during the 2 hours after the delivery of the sonications. However, the signal intensity increased on the T2-weighted images for some of the lesions during this time.



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  		Figure 5. Postsonication MR images of the lesions produced in rabbit thigh muscle with 30-second sonications delivered with an eight-element sector-vortex array in mode four. The sonications were delivered at a frequency and power of (left column) 1.6 MHz/13.9 W, (middle column) 1.6 MHz/15.4 W, and (right column) 1.1 MHz/21.6 W. Top row: T2-weighted images (2,000/25). Bottom row: Subtractions of T1-weighted images (500/17) obtained before and after contrast enhancement. These images were all obtained perpendicular to the direction of the ultrasound beam.

 
Correlation of Tissue Changes with Temperature, Dose, and Power
Figure 6 shows the correlation of tissue response as seen in the postsonication imaging with thermal dose, temperature, and acoustic power. This information gauged how well these parameters could be used to predict the creation of thermal tissue damage. The thermal dose and temperature correlated well with tissue changes, whereas the power did not. All of the locations with thermal doses less than 4.3 equivalent minutes at 43.0°C and temperatures less than 47.2°C did not show any tissue changes, and all of the locations with thermal doses above 31.2 equivalent minutes at 43.0°C and temperatures above 50.4°C did. In between these ranges, the tissue response appeared to be indeterminate. No correlation was observed between the applied power and the occurrence of tissue damage around the threshold exposure.



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  		Figure 6. Graphs show the correlation of the creation of lesions to MR imaging-derived (top) thermal dose, (middle) temperature, and (bottom) applied power.

 

   DISCUSSION
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
RESULTS
DISCUSSION
REFERENCES
 
The results from this study show that the threshold for tissue damage induced by focused ultrasound correlates well with the MR imaging–derived thermal dose and temperature measurements but not with the applied power. The reason that power did not correlate was most likely because of differences in the body temperature; differing depths of tissue; and differences in the properties of the tissue from location to location, such as acoustic absorption and perfusion rate. Such a result indicates that acoustic intensity is not a good predictor of the threshold for thermal tissue damage in vivo. This method is likely to work equally well for monitoring and control of other thermal therapy exposures.

The indeterminate range of dose measurements (~27 minutes) and temperature measurements (~3°C) was expected because of the accuracy of the MR imaging temperature estimates and the baseline temperatures in the muscle. Another reason for this indeterminate region may have been because of inhomogeneities in the ultrasound field. Even though the sector-vortex array produced flat temperature profiles after some time, inhomogeneities in the temperature before this point could have caused nonuniform dose profiles. To produce truly flat dose profiles with the sector-vortex array, it would be necessary to switch temporally between different modes (26). However, it is expected that the histologic variations between animals and tissue locations were a substantial source of the variation observed in the study. The results from this study indicate that when all of these uncertainties are considered together, MR imaging–derived thermometry and dosimetry provide an index that can be used to estimate the threshold for tissue damage.

MR imaging–derived thermometry and dosimetry offer a distinct advantage over previous in vivo methods used to determine the threshold for tissue damage. A whole plane, or volume, can be imaged by means of MR imaging, whereas previous studies used invasive temperature probes with one or a few measurement points that were insensitive to spatial temperature inhomogeneities.

Results of previous studies (33–37) on the effects of hyperthermia have shown that the threshold for 100% necrosis in different tissues ranges 50–240 minutes. Meshorer et al (33) studied the tissue effects in porcine muscle that occurred after 30 minutes of heating. The acute effects observed 18–24 hours after therapy were that minimal necrosis occurred at 40°C–43°C, moderate damage occurred at 44°C–46°C, and severe damage occurred at temperatures greater than 46°C. These relationships between time and temperature convert to 0.5–30.0 minutes at 43°C for minimal necrosis, 60–240 minutes for moderate damage, and more than 240 minutes for severe damage. Those results agree well with the results found in this study.

According to the hyperthermia literature (21,24,38), this dose threshold of 240 minutes is adequate to coagulate all tissue and thus has been used as an indicator of tissue damage caused by focused ultrasound. The results from this study indicate that 240 minutes is a conservative estimate for predicting tissue damage in muscle tissue. However, more studies, especially in tumor tissue, are required to truly establish the treatment threshold for tumor surgery.

In conclusion, the results of this study show that MR imaging–derived thermometry and dosimetry can be used to assume that an adequate thermal exposure is delivered to coagulate the target tissue. With advances in MR imaging thermometry, the threshold presumably could be found with more precision. Knowledge of the threshold exposure necessary to cause tissue damage along with online temperature monitoring could allow for optimal thermal treatments. In addition, use of MR imaging thermometry could allow thermal exposures that are close to the thresholds to be explored for therapeutic purposes. An example of such exposure is increasing the blood vessel permeability for targeted drug or gene therapy (39).

Practical applications: Several ultrasound (40–44), laser (3,45,46), and radio-frequency (4,47) thermal surgery devices currently are being tested clinically. The ability to predict the induced tissue damage during treatments with these devices is important for several reasons. The foremost reason is to provide online information to the physician so that the desired target volume is exposed adequately. Second, the thermal imaging and dose calculation can increase safety by providing warning before the surrounding normal tissue is overexposed. Third, the ability to monitor the thermal dose removes the need to overexpose tissue. This reduces the treatment time and the danger of causing thermal buildup in the overlying tissue (48–50).


   ACKNOWLEDGMENTS
 
The authors thank Heather Martin, BS, for preparing the animals for the experiments and GE Medical Systems (Milwaukee, Wis) for supplying the positioning system and pulse sequence used in the experiments. The authors also thank TxSonics (who bought out this division of GE Medical Systems and now owns the equipment) for their support.


   FOOTNOTES
 
Author contributions: Guarantor of integrity of entire study, K.H.H.; study concepts and design, K.H.H.; definition of intellectual content, K.H.H.; literature research, N.J.M.; experimental studies, N.J.M., R.L.K.; data acquisition, N.J.M., R.L.K.; data analysis, N.J.M.; manuscript preparation, N.J.M.; manuscript editing and review, all authors.


   REFERENCES
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 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
RESULTS
DISCUSSION
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