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DOI: 10.1148/rg.276075097
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RadioGraphics 2007;27:1809-1818
© RSNA, 2007

REFRESHER COURSES

From the RSNA Refresher Courses

MR Imaging in Hyperthermia1

James R. MacFall, PhD and Brian J. Soher, PhD

1 From the Department of Radiology, Duke University Medical Center, Box 3808, Erwin Road, Durham, NC 27710. Presented as a refresher course at the 2007 RSNA Annual Meeting. Received May 4, 2007; revision requested May 16 and received June 21; accepted June 25. Supported by grant P01-CA42745 from the National Institutes of Health. All authors have no financial relationships to disclose. Address correspondence to J.R.M. (e-mail: james.macfall{at}duke.edu).

There is growing clinical evidence that the combination of radiation therapy and hyperthermia, when delivered at moderate temperatures (40°–45°C) for sustained times (30–90 minutes), is of benefit with regard to palliative relief of cancer, tumor response, local control, and survival. Adequate measurement of the temperature distribution achieved with the hyperthermia is a key element in successful therapy. Thermal dosimetry, even invasive dosimetry, is a complex topic when applied to the heterogeneous tissue of a tumor and associated organ systems. Imaging in hyperthermia therapy is performed primarily for estimation and control of temperature. Magnetic resonance (MR) imaging has unique parameter dependences that make it possible to monitor hyperthermia therapy by detection of proton resonant frequency changes or diffusion coefficient changes. In addition, MR imaging can be used to assess vascular parameters that not only allow selection of suitable patients for therapy but may also allow demonstration of response to therapy. Finally, as the use of thermally sensitive liposomes for delivery of chemotherapeutic agents is developed, MR imaging may allow determination of local drug dose.

© RSNA, 2007

There is growing clinical evidence that the combination of radiation therapy and hyperthermia, when delivered at moderate temperatures (40°–45°C) for sustained times (30–90 minutes), is of benefit with regard to palliative relief, tumor response, local control, and survival (1–3). Positive phase III clinical trials comparing radiation therapy with or without hyperthermia have been reported for several sites, including recurrent breast cancer on the chest wall (4), melanoma (5), advanced head and neck cancer (6), esophageal cancer (7), cervical cancer (8), and glioblastoma multiforme (9). In some trials where invasive temperatures were measured during treatment, retrospective analysis has shown that descriptors of temperatures achieved were significantly correlated with beneficial outcome (10–14). Other phase III trials showed no significant correlation between time-temperature relationships and outcome, despite the positive benefit seen by adding hyperthermia to radiation therapy (6,9). There were important limitations in the design of these clinical trials: (a) thermometry data were sparse, being based on invasive measurements, and (b) there was no a priori attempt to control the thermal dose delivered.

With respect to the first point above, a number of research groups have demonstrated that magnetic resonance (MR) imaging is effective and accurate for noninvasively assessing temperature changes in tissues associated with absorption of nonionizing radiation (15–21). Concerning the second point, two phase III trials (22–24) have demonstrated that controlled delivery of thermal dose is clearly related to treatment outcome. These trials did not have the limitations listed above. Nevertheless, extensive invasive thermometry and real-time manual manipulation of treatment settings were performed, which was inefficient and lacked the accuracy of data that reflected the true temperature distribution in the tumor (25). Extensive invasive thermometry cannot be done in a routine fashion, particularly in deep-seated tumors, and is a confounding factor in any trial designed to demonstrate clinical benefit from thermal therapy. The positive results of these two clinical trials clearly provide compelling rationale for continued development of noninvasive thermometry in hyperthermia. Hyperthermia can be an effective adjunctive therapy, but adequate attention must be paid to ensuring that the target temperatures have been achieved for an adequate period of time.

Thermal dosimetry, even invasive dosimetry, is a complex topic when applied to the heterogeneous tissue of a tumor and associated organ systems. The dose is generally calculated by using invasive thermal probes such as the Luxtron (Mountain View, Calif) fiberoptic probe at five to 10 locations. The temperatures from the probes are evaluated, for example, every minute as a function of the current T90 (the temperature exceeded by 90% of the measured temperature points), and thermal doses are summed together for each minute of treatment to obtain the total dose in cumulative equivalent minutes at 43°C for the T90 temperature (CEM43°CT90) (1). As mentioned above, this methodology is challenged by the low density of temperature samples that can practically be achieved. Thus, there is a natural desire to try to achieve sampling densities closer to the pixel density of a digital image of the tumor volume. The development of image-based thermometry methods has the potential to automate and simplify hyperthermia treatment.

Regardless of the potential direct benefit of hyperthermia therapy alone or in conjunction with other therapies, another emerging role for tissue heating and thermal control is the use of temperature-sensitive liposomes to encapsulate chemotherapeutic agents such as doxorubicin to allow high-dose delivery to a tumor (http://clinicaltrials.gov/show/NCT00441376). Preclinical (26) and clinical work with doxorubicin-containing thermally sensitive liposomes also points to a strong need for improved thermal dosimetry. The doxorubicin-containing thermally sensitive liposomes show optimal release kinetics in the range between 40° and 42°C, and deviation above or below this range leads to less drug release and delivery. One concept being tested is temporal modulation of the temperature distribution across the tumor volume in order to achieve more uniform drug delivery. The implementation of volumetric thermal dosimetry and control is essential for clinical implementation of such strategies.

Also, as a mature imaging modality, MR imaging offers opportunities for more than just temperature measurement. It can provide data for treatment planning, dynamic control of treatment delivery (27), and posttreatment assessment of tissue damage (28).

Thus, while there are a variety of other imaging methods such as computed tomography, positron emission tomography, and ultrasonography that are used to stage and monitor cancer therapy, MR imaging has emerged as providing unique advantages for hyperthermia therapy. Hence, this article concentrates on MR imaging and reviews the methods for measurement of temperature with MR imaging during hyperthermia therapy. It is this key feature of MR imaging, the ability to both monitor temperature throughout a volume and obtain useful morphometric and functional information from tumor and normal tissues, that makes it the modality of choice to image regional hyperthermia therapy delivery.

Proton Resonant Frequency Shift Technique
The fundamental thermal sensitivity of certain MR imaging parameters has been recognized for some time (29–32), but it was not thought to be a feasible method for hyperthermia thermometry because of implied incompatibilities between heating devices and the MR system. There were also concerns raised about hysteresis effects caused by tissue changes during heating (33).

The work by Delannoy et al (34) at the National Institutes of Health in 1990 was a landmark accomplishment because it showed for the first time that a clinically relevant heating device could be operated in the bore of a clinical MR imaging unit. While this work was done by using the apparent diffusion coefficient (ADC), Ishihara et al (35) performed key measurements to show the temperature dependence of the proton resonant frequency shift (PRFS) of water, which was known in the NMR literature but had not been developed for use in imaging. Also, MR temperature imaging has been demonstrated for thermal ablative procedures (28,36), where the temperatures needed to achieve tissue destruction are relatively high compared with hyperthermic temperatures.

Further developments (15–20,32,37–40) have led to the delivery of clinical hyperthermia treatments in human patients while acquiring temperature data by using MR imaging (21,41,42).

When a subject or water phantom is placed in an MR imaging unit of magnetic field strength B0, the water protons acquire a resonant frequency (fL) described by the Larmor expression: fL = {gamma}B0, where {gamma} is the gyromagnetic ratio. The PRFS method relies on the fact that this resonant frequency of protons will shift by 0.01 ppm/°C. This is a result of the bond length between the hydrogen and the oxygen in the water getting longer as temperature increases; thus, the chemical shielding of the proton is reduced, giving increased chemical shift. For a 1.5-T imaging system where fL = 63.85 MHz, the proton in a water molecule will thus change its resonant frequency by 0.6385 Hz/°C.

If MR images are made at two different temperatures, the familiar magnitude image will not show a change, but we can also make an image where the intensity of each pixel is proportional to the image phase angle {varphi}, which depends on the pixel frequency fp (which is close to fL but shifted by temperature and field inhomogeneity) and the echo time TE, as {varphi} = fpTE. With an echo time of 20 msec, the image phase will thus change by 4.6° of image phase for each degree Celsius of temperature change at 1.5 T. It is important to remember that the method can measure only temperature change and the size of the effect depends on the echo time and magnetic field strength. While the image phase has a limited range of –180° to +180° of phase (if the real and imaginary parts of the image are available), this does not limit the method in practice if temperature updates are performed often enough that the phase change between updates does not reach the limit. The total temperature change over time is then given by summing the individual changes.

As an example of the PRFS method, Figure 1 shows a cylindrical gel phantom that is inside of a radiofrequency heating device that has four H-shaped antennae on the outside with a water coupling path between the antennae and the phantom. The phantom can be heated by energizing the antennae with, typically, 50 W of 140-MHz radiofrequency energy. This causes a temperature rise of about 1°C per minute, since the phantom is designed to have conductivity and other material parameters adjusted to approximate those of tissue. The phantom has thin catheter tubes along its length that allow MR-compatible temperature probes (Luxtron) to be inserted to measure the temperature directly. Figure 2 is a graph of the image phase as a function of the probe temperature as the phantom is heated. The inset is a phase image obtained along the axis of the phantom (coronal) that shows the positions of the temperature probes.


Figure 1
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  		Figure 1.  Photograph shows a gel-filled phantom (large white arrow) inside a cylinder with four radiofrequency antennae on the outside. The small white arrow indicates one of the H-shaped antennae. There is a water-filled space (black arrow) between the antennae and the phantom.

 

Figure 2
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  		Figure 2.  Graph shows the MR image phase as a function of directly measured temperature at six different locations in a gel phantom surrounded by a water bolus, as shown in Figure 1. The inset is a phase image obtained along the axis of the phantom (coronal view); the small white inhomogeneities are the probe locations.

 
A major difficulty with the PRFS method is that of image phase "drift" over the time of a series of temperature measurements that may span a hyperthermia therapy delivery session of between 1 and 2 hours. The image phase may drift enough to make the PRFS method unusable. An obvious approach is to provide reference materials in the image that do not change phase with temperature in order to monitor and correct the phase drift. Mineral oils provide a suitable such material, since the resonant frequency of oil changes a negligible amount when heated. A difficulty with such a scheme is that the references must be outside of the region where the temperature is to be estimated. Just measuring the phase drift outside of the object does not necessarily determine the drift inside the object. However, it turns out that making simple assumptions such as that the drift is slowly varying over space can provide very good estimates of the drift in places where it is not measured. Use of such reference materials also illustrates the potential of measuring absolute temperature, since the fat-water frequency difference should be a function of absolute temperature, which may be especially valuable in regions that have mixtures of fat and water such as breast tissue.

Figure 3 shows an axial image obtained across a cylindrical radiofrequency heating device and phantom similar to that in Figure 1. The phantom contains temperature probes for direct temperature measurement; the water coupling bolus contains tubes filled with a mineral oil reference material. The graph in Figure 4a shows the phase temperature change over a period of 65 minutes for a number of ROIs in the phantom and references. During the first 35 minutes, the phantom remained at constant temperature. Radiofrequency heating was then initiated for the next 20 minutes, followed by 10 minutes with the radiofrequency heating stopped. Although the phantom stayed at a constant temperature during the initial unheated period, the calculated temperature drifted downward uniformly for all regions. The drift downward in the gel phantom during heating was masked by the actual temperature rise occurring.


Figure 3A
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  		Figure 3a.  (a) Axial MR image of the phantom in Figure 1 shows the positions of the catheters that contain the direct temperature measurement probes (small arrows). The large arrows indicate the locations of the mineral oil–filled tubes that were placed outside the phantom in the water bolus. The lines in the water bolus are the folds of the plastic membrane that contains the water, and the bright inhomogeneities on the outside are due to the conductive antennae for heating placed on the outside of the outer cylinder. (b) Image obtained by summing all of the temperature changes measured with the PRFS method, which are indicated by an orange-toned intensity scale, shows the circularly symmetric heating pattern at the end of the heating.

 

Figure 3B
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  		Figure 3b.  (a) Axial MR image of the phantom in Figure 1 shows the positions of the catheters that contain the direct temperature measurement probes (small arrows). The large arrows indicate the locations of the mineral oil–filled tubes that were placed outside the phantom in the water bolus. The lines in the water bolus are the folds of the plastic membrane that contains the water, and the bright inhomogeneities on the outside are due to the conductive antennae for heating placed on the outside of the outer cylinder. (b) Image obtained by summing all of the temperature changes measured with the PRFS method, which are indicated by an orange-toned intensity scale, shows the circularly symmetric heating pattern at the end of the heating.

 

Figure 4A
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  		Figure 4a.  Graphs of the temperature change at the different regions of interest (ROIs) in Figure 3b calculated with the PRFS method over time. During the 65 minutes shown, the phantom remained unheated and at a constant temperature for the first 35 minutes, followed by 20 minutes of heating and then 10 minutes of cooling. (a) Graph shows data that have not been corrected for phase drift, which is clearly demonstrated by the PRFS measurements in the oil references (the constantly decreasing values). (b) Graph shows the data corrected for drift by using spatial interpolation and the oil references. The data are seen to be comparable with the direct temperature measurements (black lines).

 

Figure 4B
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  		Figure 4b.  Graphs of the temperature change at the different regions of interest (ROIs) in Figure 3b calculated with the PRFS method over time. During the 65 minutes shown, the phantom remained unheated and at a constant temperature for the first 35 minutes, followed by 20 minutes of heating and then 10 minutes of cooling. (a) Graph shows data that have not been corrected for phase drift, which is clearly demonstrated by the PRFS measurements in the oil references (the constantly decreasing values). (b) Graph shows the data corrected for drift by using spatial interpolation and the oil references. The data are seen to be comparable with the direct temperature measurements (black lines).

 
The graph in Figure 4b, which was obtained by using the ROIs to sample the oil reference phase values and a low-order interpolation spatially, shows the temperature measured with the temperature probes (black lines) and the temperature estimated with the PRFS technique, corrected for drift. After correction, the gel was shown to have a constant temperature during the nonheated period and to have temperature changes consistent with the measured temperature changes during the heat periods. There is an apparent overcorrection in the locations near the edge of the phantom during the period just before the heating. This may indicate the degree of validity of the spatial interpolation method or assumptions. The orange-toned image in Figure 3b is a cumulative temperature change map formed at the end of the heating period and showing the typical circular heating pattern in the phantom.

An example of this method for a subject who underwent hyperthermia therapy in a clinical trial is shown in Figure 5 (the data were obtained under an institutional review board–approved protocol with informed consent). The subject’s tumor (soft-tissue sarcoma of the calf) is visible in the upper right quadrant of Figure 5a. The graph shows the PRFS temperature estimates in the ROIs defined by the color blocks in Figure 5c and 5d. ROI 1 is in the tumor, near the invasive fiberoptic probe (not visible), ROI 2 is in normal muscle, and ROIs 3 and 4 are in the oil reference materials. The phases and amplitude of the applied radiofrequency heating were adjusted to deliver the highest heating to the tumor region.


Figure 5A
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  		Figure 5a.  Estimates of temperature change (starting at body temperature) during hyperthermia therapy in a patient with a soft-tissue sarcoma of the lower leg. (a) Axial image of the calf and the tumor shows the anatomy. The leg is surrounded by a water bolus and the heating antennae are on the outside of the cylinder, as shown in Figure 1. (b) Graph shows the temperature changes during the 70 minutes of therapy. The different lines correspond to the ROIs, which are shown as color blocks in c and d. 1 = ROI in the tumor near the invasive fiberoptic temperature probe (black line on graph), 2 = ROI in normal muscle, 3 and 4 = ROIs in the oil references. The drift correction has been applied to the graph, as demonstrated by the near-constant values of the oil references. (c, d) Anatomy images obtained at the beginning (c) and near the end (d) of therapy show the regional temperature changes by means of an overlaid orange-toned temperature scale.

 

Figure 5B
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  		Figure 5b.  Estimates of temperature change (starting at body temperature) during hyperthermia therapy in a patient with a soft-tissue sarcoma of the lower leg. (a) Axial image of the calf and the tumor shows the anatomy. The leg is surrounded by a water bolus and the heating antennae are on the outside of the cylinder, as shown in Figure 1. (b) Graph shows the temperature changes during the 70 minutes of therapy. The different lines correspond to the ROIs, which are shown as color blocks in c and d. 1 = ROI in the tumor near the invasive fiberoptic temperature probe (black line on graph), 2 = ROI in normal muscle, 3 and 4 = ROIs in the oil references. The drift correction has been applied to the graph, as demonstrated by the near-constant values of the oil references. (c, d) Anatomy images obtained at the beginning (c) and near the end (d) of therapy show the regional temperature changes by means of an overlaid orange-toned temperature scale.

 

Figure 5C
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  		Figure 5c.  Estimates of temperature change (starting at body temperature) during hyperthermia therapy in a patient with a soft-tissue sarcoma of the lower leg. (a) Axial image of the calf and the tumor shows the anatomy. The leg is surrounded by a water bolus and the heating antennae are on the outside of the cylinder, as shown in Figure 1. (b) Graph shows the temperature changes during the 70 minutes of therapy. The different lines correspond to the ROIs, which are shown as color blocks in c and d. 1 = ROI in the tumor near the invasive fiberoptic temperature probe (black line on graph), 2 = ROI in normal muscle, 3 and 4 = ROIs in the oil references. The drift correction has been applied to the graph, as demonstrated by the near-constant values of the oil references. (c, d) Anatomy images obtained at the beginning (c) and near the end (d) of therapy show the regional temperature changes by means of an overlaid orange-toned temperature scale.

 

Figure 5D
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  		Figure 5d.  Estimates of temperature change (starting at body temperature) during hyperthermia therapy in a patient with a soft-tissue sarcoma of the lower leg. (a) Axial image of the calf and the tumor shows the anatomy. The leg is surrounded by a water bolus and the heating antennae are on the outside of the cylinder, as shown in Figure 1. (b) Graph shows the temperature changes during the 70 minutes of therapy. The different lines correspond to the ROIs, which are shown as color blocks in c and d. 1 = ROI in the tumor near the invasive fiberoptic temperature probe (black line on graph), 2 = ROI in normal muscle, 3 and 4 = ROIs in the oil references. The drift correction has been applied to the graph, as demonstrated by the near-constant values of the oil references. (c, d) Anatomy images obtained at the beginning (c) and near the end (d) of therapy show the regional temperature changes by means of an overlaid orange-toned temperature scale.

 
The graph in Figure 5b shows that the invasive measure and ROI 1 (both in the tumor) agree quite well and that the drift correction method results in a constant value for the oil reference materials. The tumor heated up rapidly and showed modulation of the temperature due to blood flow changes alternated with radiofrequency power changes as the body attempted to cool the tumor region. The orange-toned images in Figure 5c and 5d show the temperature changes at the beginning (virtually zero) and after 60 minutes of heating.

The water self-diffusion coefficient, which is called the ADC in heterogeneous tissue, is temperature sensitive. The ADC of water (D, usually expressed in units of millimeters squared per second) in tissue changes with temperature through the following simple relationship:


Formula

where Ea is the water activation energy, k is the Boltzmann constant, and T is the absolute temperature. This relationship can be restated in terms of temperature change ({Delta}T) as follows:


Formula

Thus, the fractional change in temperature can be calculated from the fractional change in the diffusion coefficient. The value of Ea is about 0.21 eV for water, so this method yields about a 2.7% change in D for each degree of temperature change.

The ADC can be measured with most modern MR imaging systems by using echo-planar imaging pulse sequences. The process involves acquiring two images, each with different sensitivity to diffusion. The diffusion sensitivity is related to the strength and timing of the gradients and is denoted as the b value of the pulse sequence. The best signal-to-noise ratio in measuring an ADC value results from obtaining one image with diffusion weighting b = 0 and a second one with b = 1/D. If the first image intensity is denoted S(0) and the second is S(b), then the ADC is calculated from the natural logarithm of the image ratio:


Formula

Echo-planar imaging is a fast technique that is challenged by low resolution, low signal-to-noise ratio, and image distortion. In addition, the need to sensitize the images to diffusion leads to long echo times, which further reduce signal-to-noise ratio for tissues that have short T2 relaxation times. Fortunately, many tumors have longer T2 values, so this method can still work in such pathologic tissues. As an example of the noise and linearity of the method, Figure 6 shows the ADC fractional change versus temperature change in canine tissues being heated with radiofrequency energy.


Figure 6
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  		Figure 6.  Graph shows the dependence of the fractional change in ADC in canine tissue as a function of temperature change.

 
Although the ADC method thus appears to have adequate linearity with temperature, a more complicated analysis shows that it has lower sensitivity to temperature change than the PRFS method. As an example, for an echo-planar diffusion measurement at 1.5 T on a system with a maximum gradient strength of 23 mT/m with repetition time = 6 seconds, 4-mm section thickness, 90-kHz bandwidth, 24-cm field of view, 128 x 128 matrix, and b = 750 sec/mm2, the minimum echo time value is 109 msec, leading to a temperature noise (calculated from the image and diffusion noise) of 1.2°C. To measure temperature to within 1°C, longer imaging times or larger voxels would have to be used.

For the same object, by using the PRFS technique and a gradient-echo acquisition with repetition time = 34 msec, echo time = 20 msec, 16-kHz bandwidth, and a 256 x 256 matrix, the temperature noise was only 0.26°C. The PRFS method in this example thus has twice the image spatial resolution, a shorter 8-second imaging time compared with the longer 12-second imaging time of the ADC measurement (since two images must be acquired), and, most important, much lower temperature noise (by about a factor of 5!). Hence, the PRFS method has become the preferred method of MR temperature measurement.

To have successful hyperthermia therapy, the tumor must be "heatable." If a tumor has large sections of highly vascular tissue, it may be difficult to achieve local heating in these sections of the tumor without overheating poorly or normally vascularized tissue. Hence it is of value, when deciding whether a particular tumor is a good candidate for hyperthermia, to perform contrast material–enhanced or, even better, dynamic contrast-enhanced MR imaging to evaluate tumor vascularity (17,19,43). In addition to allowing the basic morphometry and vascularity of the tumor to be assessed, the vascular parameters (permeability, extravascular space) that can be obtained from an analysis of the dynamic contrast-enhanced data are potentially markers of tissue response to therapy.

As an example of the value of contrast-enhanced MR imaging, Figure 7 shows contrast-enhanced MR images of two different tumors along with color-coded temperature maps obtained at the end of heat therapy, which show the hottest areas in red and the cooler areas in blue (note the somewhat smaller magnification compared with that of the MR images). The top MR image (Fig 7a) shows a strongly enhancing tumor rim consistent with high vascularity and a necrotic core that does not enhance. The thermal map (Fig 7b) shows that only cooler temperatures were achieved by the end of the heating, indicating that the tissue cooling from the vascular flow was enough to reduce the heating—even in the necrotic region.


Figure 7A
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  		Figure 7a.  MR images and temperature maps of two tumors with higher (a, b) and lower (c, d) vascularity. (a, c) Contrast-enhanced MR images show two sarcomas of the leg. (b, d) Corresponding temperature maps show temperature change from the beginning of heating by means of a color scale, which ranges from blue (0°–2°C) to red (>6°C). Note that the image magnification is different between the MR images and the temperature maps.

 

Figure 7B
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  		Figure 7b.  MR images and temperature maps of two tumors with higher (a, b) and lower (c, d) vascularity. (a, c) Contrast-enhanced MR images show two sarcomas of the leg. (b, d) Corresponding temperature maps show temperature change from the beginning of heating by means of a color scale, which ranges from blue (0°–2°C) to red (>6°C). Note that the image magnification is different between the MR images and the temperature maps.

 

Figure 7C
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  		Figure 7c.  MR images and temperature maps of two tumors with higher (a, b) and lower (c, d) vascularity. (a, c) Contrast-enhanced MR images show two sarcomas of the leg. (b, d) Corresponding temperature maps show temperature change from the beginning of heating by means of a color scale, which ranges from blue (0°–2°C) to red (>6°C). Note that the image magnification is different between the MR images and the temperature maps.

 

Figure 7D
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  		Figure 7d.  MR images and temperature maps of two tumors with higher (a, b) and lower (c, d) vascularity. (a, c) Contrast-enhanced MR images show two sarcomas of the leg. (b, d) Corresponding temperature maps show temperature change from the beginning of heating by means of a color scale, which ranges from blue (0°–2°C) to red (>6°C). Note that the image magnification is different between the MR images and the temperature maps.

 
The tumor shown in the lower MR image (Fig 7c) does not have such strong enhancement of the rim and has a larger necrotic region. The thermal map (Fig 7d) shows that higher temperatures were achieved in the vascular regions of this tumor compared with those achieved in the more vascular tumor and also shows a hot spot in the necrotic region. Careful evaluation of subjects as to whether their tumors are heatable has been an important part of conducting successful clinical trials of hyperthermia, since older, less successful trials may have entered subjects who were not suitable for such therapy.

As mentioned in the Introduction, an exciting new area of thermal therapy is to use heat-sensitive liposomes to encapsulate chemotherapeutic agents to allow targeted therapy through hyperthermia. These liposomes can have their properties adjusted to release their contents at 42°C, and thus local hyperthermia will cause most of the drug to be released in the heated region. By encapsulating MR contrast agents in the same liposomes, it may be possible to make MR images of the regions where the agent is released. By using quantitative MR imaging, the concentration of the MR contrast agent can be calculated, which may allow accurate assessment of the concentration of the drug deposited throughout the tumor (26,44–48).

Imaging in hyperthermia therapy is performed primarily to estimate and control temperature. MR imaging has unique parameter dependences that make this possible by detection of proton resonant frequency changes or diffusion coefficient changes. In addition, MR imaging can be used to assess vascular parameters that not only allow suitable selection of subjects for therapy but may also allow detection of response to therapy. Finally, as the use of thermally sensitive liposomes matures for delivery of chemotherapeutic agents, imaging may allow determination of local drug dose.

The authors gratefully acknowledge valuable discussions with and images from Mark Dewhirst, DVM, PhD, Paul Stauffer, MSEE, CCE, Oana Craciunescu, PhD, and Zeljko Vujaskovic, MD, PhD.

Abbreviations: ADC = apparent diffusion coefficient, PRFS = proton resonant frequency shift, ROI = region of interest

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