phase imaging on a .2-t mr scanner: application to temperature monitoring during ablation procedures

Original Research

Phase Imaging on a .2-T MR Scanner: Application to Temperature Monitoring During Ablation Procedures

Shantanu Sinha, PhD Thomas Oshiro, BS Usha Sinha, PhD Robert Lufkin, MD

Proton phase shift imaging methods with keyholing were developed to rapidly monitor temperature during MR-guided radiofrequency (RF) interventional procedures on a .2-T open configuration scanner. Tempera- ture calibration was performed on thermally controlled gel phantom and ex vivo bovine liver samples. Keyhol- ing methods were implemented for rapid imaging and tested both in simulation experiments and in the gel phantom. Phase drifts from extraneous sources were monitored and compensated for using reference phantoms. Sequence parameters TE, TR, and flip angle (FA) were optimized for maximum temperature sensitivity and minimum noise. Reduction of phase noise from coupling of the magnetic field to external perturbations using navigator-echo-based correction schemes were also investigated. The extraneous phase drifts from the magnet could be minimized by keeping the electromagnet on continuously. Navigator echo corrected keyholed FLASH sequences (TE = 30 msec, TR = 60 msec, FA = 40”. 64 X 128 matrix) were used to monitor the R F lesioning process in gel phantoms yielding images every 4 seconds with a temperature sensitivity of .015 ppm/”C. RF ablation in the bovine tissue was monitored using navigator-echo-corrected keyholed fast low angle shot (FLASH] sequences (TE = 30 msec, TR = 100 msec, FA = 40”, 128 X 256 matrix) with a temporal resolution of 13 seconds and a temperature sensitivity of .007 ppm/”C. The results indi- cate that monitoring of an RF ablation procedure by mapping temperature with sufficient temporal resolution is possible using phase images of FLASH sequences on a .2-T open scanner.

Index terms: Temperature mapping- Radiofrequency ahlation - Phase im aging * Keyholing * lntenrentional MRJ

JMRl 1997: 7:918-928

From the Department of Radiological Sciences, UCLA School of Medicine, Box 951721, 10833 Le Conte Avenue, Lns Angeles, CA 90024. E-mail: L

sinha~m~l.rad.ucla.edu. Received August 13. 1996: ralsion requested Jan- uary 23, 1997: revision received May 7: accepted May 13. Presented at the annual mqeting of the International Sociely for Magnetic Resonance in Med- icine, New York, 1996. Address reprint requests to S.S.

0 ISMRM. 1997

Abbreviations: FA = flip angle, FLASH = fast low angle shot, FOV i field of view. FWHM = full width at half maximum, W = keyhole factor. PRF =

pmton resonance frequency. ROI = region of interesl, RF = rddiofmquency, SNR = signal-to-noise ratio, TSNR = signal-to-noise ratio of the temperature map.

THE INTRODUCTION OF the open magnet configuration with wide angle accessibility to the patient and real time monitoring of images within the scanner room has sparked considerable interest in MR-guided interventional procedures. Such procedures, including MR-guided biop- sies, hyperthermia, cryoablation, and ablation using laser, radiofrequency, and focused US, have been performed both in conventional 1.5-T scanners (closed configuration] and at low fields (.5 T, open and closed configurations] (1-61. A critical part of such MR-guided ablation procedures is the ability to monitor spatially localized changes in temperature using heat-sensitive MR pulse sequences. Four different NMR properties of tissues have shown potential as parameters sensitive to temperature changes. These are the spin-lattice relaxation time T1, the molecular diffusion coefficient D, the proton density p, and the proton resonance frequency (PRF) (1,7-14). The PRF parameter has been shown to have excellent potential for monitoring temperature changes during hyperthermia in a closed configuration 1.5 T system (10-141. The shifts of the PRF can be measured conveniently from phase images of gradient-echo sequences,

The objective of this study was to map the temperature profile of lesions created by radiofrequency (RF) ablation procedures using phase images of gradient echoes. Be- cause RF interventional procedures take place in relatively short times (- 1-4 minutes), monitoring by dynamic MRI must be performed at a fairly rapid rate. Therefore, we incorporated a keyholing technique along with phase imaging to reduce the scan time of the dynamic scans by sampling reduced k-space. The effect of keyholing on the spatial and contrast resolution of the heat lesions on phase images was studied by simulation of a typical heat lesion in a full resolution phase image and comparison of this image with those reconstructed from reduced raw data matrices. The temperature dependence of the phase shift is linearly proportional to the external field; the result is that this technique is less sensitive by a factor of approximately 7 at .2 T compared with that at 1.5 T. It

~~

91 8

was therefore necessary to optimize different sequence parameters such as TE, TR, and flip angle (FA) at this reduced field strength to maximize the sensitivity to temperature while minimizing the uncertainty in temperature measurements (SD). The success of this method, which involves spatial and temporal mapping of phase, depends critically on the temporal stability of the external magnetic field flux density. The magnetic field stability of an open configuration electromagnet system can be expected to be reduced compared with a closed configuration superconducting system. This reduced stability is due to both the coupling of the fringe fields of the open configuration to extraneous magnetic field fluctuations as well as the drift of the main magnetic field B, due to the fluctuations of the current through the electromagnet. We have therefore investigated the field stability of the open magnet and implemented methods to improve the stability of the phase images. These included the use of reference phantoms at constant temperature as well a s navigator-echo-corrected pulse sequences. Two additional factors had to be ad- dressed to acquire images during the RF ablation procedure: (4 the electronic interference between the MR receiver system and the RF generator and (b) the susceptibility-induced artifacts of the RF ablation probe, which prevented visualization in the immediate vicinity of the probe. We present in this report the results of phase imaging at .2 T for temperature measurements, with a temporal resolution of 4 seconds, during an RF ablation procedure both in a gel phantom as well as in an ex vivo bovine liver sample.

METHODS

Temperature Sensitivity of Phase Images The phase image of a gradient-echo sequence permits

the measurement of the phase distribution of the protons in the image, relative to a reference phase. Because the phase distribution reflects the frequency distribution of the protons, a change in temperature will be seen in the phase images as a change in phase. It has been shown that the temperature dependence of the phase can be expressed as follows:

where d@(x, y) is the phase difference induced by a temperature change of dT(x, y] at a pixel location (x, y); y is the gyromagnetic ratio of proton; C is the constant of proportionality; Be, is the external magnetic field flux; and TE is the echo time of the gradient-echo sequence (10). It can be seen from Equation (11 that the phase change for a given change in temperature is proportional to the external magnetic field as well as the TE of the sequence. Gradient echoes wilh echo times in the range of 6.8 to 30 msec have been used to generate temperature-sensitive phase images at 1.5 T (1 0-1 4). To achieve the same sensitivity at .2 T as at 1.5 T, it is necessary to use TEs in the range of 50 to 140 msec. However, it is also necessary to take into account the noise in the derived temperature maps to choose the optimum TE.

Accuracy of Temperature Derived from Phase Images The accuracy of the phase image depends on the sig-

nal-to-noise ratio (SNR] of the magnitude image [13,14]. This dependence can be evaluated by an examination of the signal and noise in the complex plane. The noise in

the phase image is related to the SNR of the magnitude image as follows:

where Nph, is the noise in the phase image and S,, and Nmq are the signal and noise of the magnitude image, respectively (13). The uncertainty in temperature measurements is then given by:

Temperature maps usually are derived from subtracted phase images performed to eliminate phase wraps. The temperature uncertainty is then decreased by a factor of d2 when determined from subtracted phase images.

Estimation of the Optimum Sequence Parameters: TE, TR, and FA

The signal strength of the fast low angle shot [FLASH) sequence is dependent on tissue parameters as well as sequence parameters: TE, TR, and FA, 0. The dependence of SNR of the temperature map (TSNR) on TE can be defined as follows:

TSNR 0~ Temperature Sensitivity/Nph,, 141

which, using Equations [ 11 and [2] and the FLASH signal intensity equation, reduces to:

TSNR TE/tan-' [l/exp(-TE/T2*)], (51

where temperature sensitivity and Nphase are defined by Equations [ 11 and [2], respeclively. The optimum value of TE will be that which maximizes the expression on the right side of Equation 151.

The other sequence parameters, TR and FA, do not affect the temperature sensitivity but can still influence the TSNR by their effect on the noise in the phase images. The TR influences the image magnitude signal, with long TRs (compared with T11 yielding higher values of the image signal intensity. However, our objective was also to minimize scan times to achieve a high temporal resolution for monitoring temperature during the RF ablation procedure. Therefore, the minimum TR permitted by a sequence was chosen rather than the optimum TR, as dictated by signal strength considerations. At any TE and TR, the FA was determined from the Ernst relationship to maximize the magnitude image signal because this in turn minimizes the phase noise (Equation [2]).

MR Image Acquisition A Siemens vertical field water-cooled electromagnet

OPEN MR scanner (Siemens Medical Systems, Iselin, NJ) , operating at .2 T with a maximum gradient amplitude of 10 mT/m and a gradient rise time to maximum amplitude of 1,000 msec, was used to acquire all of the MR images in both phantom and ex vivo bovine liver samples. The phantom consisted of four reference cylinders of 2.5 cm in diameter encased in insulating foam that surround a central gelatin cylinder of 10 cm. Reference cylinders were made of .2 g/l CuSO, and 1.5% agarose gelatin and the central cylinder consisted of 1.5% agarose gelatin and .2 mM NiC1,. The reference phantoms, maintained at room temperature, were used to correct for phase drifts arising from extraneous sources apart from the variation of temperature, the parameter of interest (10). A copper-

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b

a d

C

Figure 1. Images of the gel phantom: magnitude image showing the central phantom and the four reference phantoms (a), phase image at temperature T (b], reference phase image at ambient temperature (c), and phase image after phase wrap and phase drift corrections [a]. The reference phantoms, maintained at ambient temperature, are not visible in the last image.

constantan therniocouple placed within the imaged slice allowed in vivo monitoring of temperature for our calibration runs using a digital thermometer (Omega Instru- ments, Stamford, CT]. Susceptibility artifacts were minimized by placing the thermocouple wires vertically along the direction of the B, field.

Estimation of Phase Noise Gel phantom experiments were carried out a t ambient

temperature to verify the relationship of the magnitude SNR to the phase SNR (Equation [2]). The TE of a gradient-echo sequence was varied between 34 and 260 msec (TR fixed at 300 msec and FA fixed at 40'). The SD of a region of interest (ROI) in the subtracted phase image was compared with the signal of the same ROI in the corresponding magnitude image (noise in the magnitude images remained constant for the images acquired at different TEs) .

Calibration of Phase Change with Temperature Several experiments were performed to determine the

dependence of phase change in gradient-echo phase images on temperature. Phase changes were calibrated against temperatures indicated by the thermocouple. Temperature runs involved heating the central gel phantom (in a microwave oven) described above to a temperature of approximately 80°C and then monitoring the temperature both by the thermocouple as well as the MR phase images during the cooling period. The temperature dependence of tissue was measured by monitoring the phase changes in gradient-echo phase images of a bovine liver sample. Temperature calibrations of the bovine liver sample were performed by loading the sample in a vial,

which was then placed in a temperature-controlled water bath (at an initial temperature of 6OOC). A thermocouple placed in the liver sample (in the imaging plane) was used to monitor the temperature. For each sequence and choice of sequence parameters, temperature sensitivity and temperature accuracy were determined.

Monitoring and Correction of Phase Dnf% The Siemens OPEN scanner uses an electromagnet in

an open C configuration. Therefore, main magnetic field variations arise from (a) variations in the current through the electromagnet, which is pronounced during the warm-up period when the magnet is switched on initially, and (b) coupling of the magnet's fringe fields to external magnetic field perturbations. B, magnetic field fluctuations will result in fluctuations of the resonant frequency, which will then reflect as variations of phase in the phase images. It is important to separate the contributions to phase change from extraneous sources other than from temperature to accurately map the temperature. We tracked drifts in the magnetic field B, using a FLASH sequence with TR = 100 msec, TE = 40 msec, FA = 40", field of view (FOV) = 200 mm, and a 256 X 256 matrix. Images were acquired at 5-minute intervals over a period of 2 hours on several days. Phase drift correction was performed by using reference phantoms maintained at room temperature. The details of the postprocessing of the phase images are given below. The field drift was monitored after allowing a period of at least 24 hours after powering up the magnet.

Navigation-echo-based correction for gradient-echo sequences.-The open C-arm configuration of the .2 T al- lows coupling of external magnetic fields via the fringe field of the magnet. This effect is most pronounced in gra-

920 * JMRl September/October 1997

a

- 2 450 .- STDphase=91 2'atan(9.7/Smagn) 400

2 350 300

3 250 m

E 200 I

3 100

2 150 a

50

0 (I]

0 0 0 0 0 0 0 0 0 0 Z W B S 8 8 R 8 8 Smag (arbitrary units)

TSNR= 3.5klE

atan(ll(5WExpi-TEl52))) R"2 = 0.992

b o- 0 50 100 150 200 250

TE (W

dient echoes acquired at long TEs, where view-to-view phase variations [arising from external field variations) cause banding artifacts in the image. Navigator-echo- based schemes have been suggested as a method of correction for phase variations that may arise from different sources: gross patient motion, physiologic motion, and coupling to varying external magnetic fields (1 5). Obvi- ously, the major contributor to phase noise in phantom measurements is the last factor. In the corrected gradient-echo sequences used here, phase correction was achieved by collecting a non-phase-encoded echo (called the navigator echo) after the normal echo collection. The phase variations that arise during k-space traversal was corrected by the second echo, yielding images that are not as sensitive to phase variations from external magnetic fields during a scan. The geometry parameters of the scan were maintained the same as the uncorrected sequences for all of the temperature calibration runs.

Postprocessing (Phase Unwrapping and Drift Correction)

Phase images were processed off line on a SUN Sparc2 workstation (SUN Microsystems, Milpitas, CA). Postpro- cessing of the phase images involved two steps: (a) phase unwrapping and (b) correction for the phase drift from sources other than temperature. Phase images from the host scanner were scaled from 0 to 4,096, corresponding to phase angles of 0" and 360°, respectively. In terms of this scaling, the phase unwrapping was performed by subtraction of the phase image at any temperature from a reference phase image (10). An alternate method of phase unwrapping by complex subtraction of the temperature images from the reference image also has been proposed recently ( 1 1).

After phase unwrapping, each phase image was then corrected for the phase changes that occur during a heating or cooling cycle from effects other than temperature. These effects were corrected by using the four tubes, which were maintained at ambient temperature. These reference tubes were used in both phantom and the ex vivo bovine liver calibration studies. Correction involved subtraction of the average phase of the four reference tubes at any temperature from that of the sample being heated (either the gel phantom or the bovine sample).

Kegholing or Partial k-Space Acquisition Image acquisition times (at .2 T) for temperature map-

ping using phase images are in the range of 20 to 24 seconds [TE - 30-60 msec, TR - 100 msec, and 256 phase encode levels). Qpical ablation periods using RF or laser sources are 1 to 2 minutes long, and a 24-second scan may not provide adequate temporal resolution for moni-

Figure 2. Phase noise measurements in a gel phantom at ambient temperature. Var- iation of phase noise with magnitude image signal intensity (a). TSNR (defined as (TE)/ tat-l(l/S-,J with TE (b]. The phase noise was measured as the SD of an ROI (6 X 6 pixels) in the subtracted phase images. The magnitude signal, S-, was measured with an ROI of the same size placed a t an identical location in the corresponding magnitude image.

toring the ablation procedure. We investigated keyholing techniques to reduce scan times. Keyholing refers to partial k-space sampling and recently has been proposed to increase temporal resolution during a dynamic study after injection of a contrast agent (16-18). A full resolution reference image is acquired before administration of the contrast agent, and the dynamic study consists of ac- quiring a number of lower resolution images, where only the central k-space lines are collected. Reconstruction of the postcontrast images is performed by using the high k-space lines of the precontrast high resolution image in addition to the acquired central k-space lines.

Postprocessing for Partial k-Space Acquisition In the current imaging protocol, the full data matrix set

was first acquired at ambient temperature. A partial k- space sampled image also was acquired at ambient temperature, and this was taken as the reference image for the subtraction. Subsequent images during the temperature cycle were acquired with lower resolution. All of the images, including the reference image, were then reconstructed after substituting the high k-space lines from the full data matrix set into the reduced k-space data set of the dynamic scan. Phase information was then ex- tracted using the same steps for removing phase wraps and phase drifts, as outlined in the section on postprocessing described earlier.

Simulation Experiments for Keyholing Simulation studies have been performed earlier to as-

sess the effect of keyholing on the spatial and contrast resolution in magnitude images (18). Because different factors affect magnitude and phase images as well a s the noise in these images, an independent simulation of the effects of keyholing in phase images was undertaken in- stead of extrapolating the results from magnitude images. Furthermore, phase images are complicated both by the presence of phase wraps and phase drifts, which are, in our method, removed by subtraction from a reference image and from the average phase of four reference tubes, respectively. We designed studies using a simulated heat lesion to study the effect of keyholing on phase images. A typical lesion generated by an RF ablation procedure depends on the power and duration of the RF generator, diameter, and length of the exposed tip of the RF probe (19). For an RF probe with dimensions of 1- mm diameter and 5-mm exposed tip, a probe tip temperature of 70°C maintained for a duration of 2 minutes generates a lesion with radius of 3 mm and length of 7 mm (20). The temperature distribution within the lesion is hottest close to the RF tip and falls off as lc4 away from the probe renter, where r is the distance from the probe

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Temperature (“C) Figure 3. The variation of phase in a gel phantom with temperature as determined from thermocouple measurements using FLASH sequences with a TE of 40 and 60 msec and a fixed TR of 100 mser. The temperature sensitivity and temperature uncertainty for both plots are listed in Table 1.

center. We assumed a rectangular lesion (10 X 7 mml with a temperature of 80°C at the center, decreasing to 27°C at the edges of the lesion, which is a representative size and temperature distribution for RF ablation procedures (20).

Three phase images were acquired with a gradient-echo sequence (TE = 34 msec, TR = 69 msec, FA = 40”). All three images were acquired with 256 phase-encoding lines and represent the high resolution images, similar to the precontrast images in dynamic contrast studies (1 6). The first phase image was used as the high resolution image and the second phase image was used as the reference image for subtraction. A lesion with the size and temperature distribution given above was simulated in the third phase image. All three images were two-dimensional inverse Fourier transformed to obtain the raw data lines (2 1). The simulation was performed using the inner k lines (192, 128, 64 or 32) from the second and third images with the remaining higher k-space lines substi- tuted from the first image. A comparison of the lesions in the keyholed images with the original simulated lesion was made for different keyhole factors (defined as the ratio of lines acquired in the high resolution image to that acquired in the low resolution images).

Spatial and contrast distortions in keyholed images were quantified as follows. ‘The full width at half maximum (FWHM) of two-dimensional temperature profiles along the phase-encode axis was used as the index of spatial distortions. The percentage spatial distortion was defined as follows:

161

where FWHM,, and FWHM,, are the FWHM of the simulated lesion profile in the full resolution and keyholed images, respectively. The average of the four highest temperature points within the lesion was used to quanhfy contrast distortions. The percentage contrast distortion was defined as follows:

where T,, and Tkl are the average of the four highest temperatures in the simulated lesion in the full resolution and keyholed images, respectively.

RF Ablation Both the agar gel phantom and the ex vivo bovine liver

sample were used to investigate temperature mapping by phase imaging during the course of RF ablation using a 480-kHz RF generator (Radionics Inc., Burlington, MA). Attempts were made to decrease the interference between the RF generator and imaging electronics by incorporation of a specially designed bandpass filter at the output of the generator. We initially attempted to use the “MR- compatible” RF probe from Radionics, which was con- structed of an outer sheath of an MR-compatible Lufkin needle (EZ-E-M Inc., Westbury, NY) with a built-in temperature sensor. The phase shift images acquired at long TEs (necessary for accurate temperature monitoring) are, however, extremely susceptible to magnetic field inhomogeneities. Therefore, the commercially available MR- compatible RF probes could not be used because of the prohibitive amount of susceptibility artifacts, which precluded visualization of the region around the probe [al- though these needles, in our experience, work satisfac- torily in magnitude images). We investigated other possible materials for the present feasibility study and found that a probe made of a solid copper 18-gauge wire, insulated for most of its length but exposed at the tip for approximately 1 cm, with a copper-constantan thermocouple attached externally along its length yielded susceptibility artifact free phase images.

A reference image with a 128 X 128 matrix was acquired with the RF probe and thermocouple in place within the gelatin. A navigator-echo-corrected sequence was used with the following parameters: TE = 30 msec, TR = 60 msec, FA = 40°, FOV = 200 mm, slice thickness = 5 mm. The probe was connected to the RF generator and the thermocouple was connected to the digital monitor. Care was taken to ensure that there was no change in the positions of the probe, thermocouple, and phantom after the initial image was acquired. Repeated images were acquired, during and immediately after the RF ablation, at the same location using a low resolution image matrix (64 X 128) with other sequence parameters maintained identical as the high resolution image, yielding a temporal resolution of 4 seconds per image. The bovine liver sample was imaged using keyholing with a setup similar to that for the gel phantom. A high resolution 256 X 256 matrix was acquired (TE = 30 msec, TR = 100 msec, FA = 40°, FOV = 200 mm, slice thickness = 10 mm) for the bovine liver acquisition. Low resolution images (128 X 256) were acquired during ablation, resulting in a temporal resolution of 13 seconds per image for the tissue sample. In both studies, the thermocouple was used to monitor the temperature near the RF tip and RF power was adjusted to maintain -90°C temperature at the tip for 1 minute. Phase unwrapping was performed using a reference image, as in the calibration studies. The drift correction was made using phase data from ROIs remote from the RF probe, which was presumed to be at the ambient temperature, due to the sharp decrease in temperature with increasing distance from the RF probe. Calibration studies yielded the temperature coefficient of phase change for each type of phantom and this value was used to convert the phase changes during RF ablation to temperature units. We did not use the thermocou-

922 JMRl September/October 1997

Temperature Calibration of the Gel Phantom (From Phase Images of FLASH Sequences)

TE Sample Acquisition Sensitivity Correlation Average SD

30 Navigator Gel 16 1.42 .997 2.68 30 Keyhole [128/256) Gel 16 1.38 .993 3.69

40 None Gel 16 1.87 .964 4.69 40 Navigator Gel 16 1.84 .998 I .52 60 None Gel 16 2.65 ,963 2.42 60 Navigator Gel 16 2.68 .996 1.03

m e Time (seconds] (Degrees/'C) Coefficient ("CI (msec) Additional Features

Navigator

30 I

Temperature ("C) figure 4, Calibration of the phase change with temperature in a bovine liver sample using a navigator-echo-corrected FLASH sequence [TE = 60 msec, TR = 100 msec). Phase changes were measured in an ROI of 3 X 3 pixels.

ple data because it measured temperature at only one point in the lesion and the lesion had a fairly sharp temperature profile (note: the thermocouple was only used to measure the temperature near the tip to ensure that sufficient current was being maintained to create a lesion). Thermal profiles were obtained in gel phantoms during both the heating (RF power on) and cooling (RF power off) cycles. For bovine tissue, thermal profiles were obtained only during the cooling cycle. The additional noise in the phase images from the RF generator precluded any meaningful measurements during the heating cycle in the case of the bovine liver sample. The lower T2* of the bovine liver sample lead to decreased SNR compared with that of the gel phantom. . RESULTS

Figure 1 shows the gel phantom and reference tubes (Fig. la, magnitude image) and typical phase images at an elevated temperature (Fig. lb), at the ambient temperature (Fig. lc), and after phase unwrap and drift correction (Fig. Id). Phase unwrap was performed using the reference phase image acquired at time t = 0 (Fig. lb). Ambient temperature gel phantom data were used to verify the relationship between the noise in the phase images to the TE of the sequence. The T2* of the gel phantom

0 0

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0 - -60- No Drift Correction

Drift Correction using 0

Reference Phantoms O .

, , I I , I I , , ? , I I I I I I I I , I I I

-100 -

-120 0

0 20 40 60 80 100 1 Time (min)

0

Figure 5. Magnetic field drifts measured by phase shifts (converted to temperature units) for a period of 2 hours. The magnet was on continuously for 48 hours before these drift measurements. The drift of the central phantom is shown before applying any correction (a) and after correction using the average phase of the four reference phantoms (+).

was calculated from a multi-echo data set to be 52 msec. Figure 2a is a plot of the magnitude signal versus the SD of the phase image (the noise in the magnitude image was constant for all TEs). The SD of the phase image varies as t a n - l ( l / S - ) , as shown by the fit (Fig. 2a). A plot of TSNR as defined by TE/tan-l(l/S,,) (Equation [5]), using the signal of an ROI in the magnitude image, S,, at each TE, is shown in Fig. 2b.

Figure 3 is a plot of the phase variation in a gel phantom versus temperature, with the temperature monitored independently with a thermocouple. The phase was measured at the same location in all images in an ROI (6 X 6 pixels) placed in the center of the gel phantom. FLASH sequences at two echo times (40 and 60 msec) with the fixed TI3 of 100 msec were investigated. Other sequence parameters were as follows: FOV = 200 mm, slice thickness = 10 mm, matrix size = 128 X 128. Table 1 lists the temperature sensitivity, temperature uncertainty, and correlation offit for the different sequences. The temperature uncertainty was calculated as the SD of the 6 X 6 pixel ROI in the subtracted phase image (converted to temperature units using the measured sensitivity of the sequence). The temperature sensitivity of the frequency shift for the NiCI,-doped gel phantom was calculated to be ,015 ppm/"C (average over all of the measurements

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35 40 45 50 55 60 65 70 75 80 85 Temperature ("C)

Figure 6. Phase variations with temperature as measured by a navigator-echo-based gradient-echo sequence at TE = 30,40, and 60 msec. The sequence parameters were similar to those used to generate the plots of Fig. 5. (See Table 1 for the temperature sensitivity and temperature uncertainty of these mcasure- ments.)

listed in Table 1). Figure 4 shows the calibration of the temperature phase variation for the bovine tissue. The images were processed with a smaller RO1 (3 X 3 pixels) to ensure that measurements were made in a region where the temperature was uniform. The geometry scan parameters were the same as for the phantom studies, and a TE of 60 msec (navigator echo corrected) with a TR of 100 msec was used. The temperature sensitivity of the frequency shift for the bovine liver sample was calculated to be .007 ppm/"C.

Figure 5 shows the magnetic field drifts with time in an ROI placed in the central gel phantom as monitored by the phase images of a FLASH sequence (TJ3 = 40 msec, TR = 100 msec, FA = 40°, FOV = 200 mm) acquired every 5 minutes. Measured phase shifts are converted to temperature units using the calibration data for the gel phantom. The drift was monitored for 2 hours after the magnet was allowed to stabilize for a 48-hour period. The average phase value of the reference tubes was used for the drift correction (Fig. 5). The SD of the phase in the ROI (averaged over all of the measurements) in temperature units was 1.44"C and arises from the noise in the phase images [Equations [2] and [3]). Figure 6 shows the phase changes with temperature in a gel phantom as measured by the navigator-echo-based sequences (at TE = 30, 40, and 60 msec) with sequence parameters identical to that of the conventional FLASH sequences (Fig. 3) . The sensitivity and SD of each sequence are reported in Table 1.

A temperature calibration run also was performed with a keyholed sequence to investigate the effect of keyholing on the sensitivity and SD of temperature measurements (Fig. 7). Data from a nonkeyholed sequence acquired under similar conditions also is shown for comparison (com- pare rows 1 and 2 of Table 11. Figure 8 shows the results of the simulation study to investigate the effects of keyholing on reproducing the spatial and temperature profile of a typical heat lesion generated by RF ablation. Table 2

0 TE=30 full resolution

TE=30 keyholed -20

40 45 50 55 60 65 70 75 80 05 Temperature ("C)

Figure 7. Verification of the sensitivity and SD of phase measurements using a keyholed sequence (m): 128 phase levels were acquired during the temperature run and the image was keyholed to 256 levels using data from a reference image at ambient temperature. Data from an identical nonkeyholed sequence is included to facilitate comparisons (0).

lists the spatial and contrast deviations of the lesion in the keyholed images compared with the simulated lesion.

Gel phantom images acquired during and immediately after an RF ablation procedure with a 64-level keyholed sequence and a temporal resolution of 4 seconds are shown in Figure 9. The three-dimensional temperature profiles obtained during the ablation (Figs. 9a through 9c) and immediately after the generator was powered down (Figs. 9d through 90 also are shown to the right of each image of the dynamic scan.

DISCUSSION The postprocessing of the data, consisting of phase un-

wrapping by subtraction from a reference image and phase drift correction using the average phase of constant temperature reference tubes, was adequate to yield accurate phase data (Fig. 1). Figure Id shows that after drift correction, only the central phantom (which was heated) registers a phase change, whereas the reference tubes do not register any change because they were maintained at ambient temperature.

It is important to consider the noise in the phase images while attempting to optimize the value of TE for temperature measurements. Simple geometric considerations of the noise in the complex signal and the noise in the phase image leads to an arctan (l /Sr, ,J relationship, confirmed by experiment (Fig. 2a). Theoretic considerations of the SNR of the temperature map (Equation [5]) yields an optimum value of TE for phase mapping at a value approximately equal to the T2* of the tissue under investigation. However, this value of TE may not be appropriate for other reasons, such as the increase in acquisition time and increased susceptibility artifacts [discussed below in greater detail). These two factors have to be considered in RF ablations, and thus, we calibrated sequences at lower values of TE than given by T2*.

Comparing the temperature sensitivities at three TEs (30, 40, and 60 msec; with navigator echo correction)

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a b C d e Figure 8. Image of a lesion simulated in a phase image, seen as a rectangular region of hyperintensity (a). Image of the lesion reconstructed using 192 (b], 128 (c). 64 (dl, and 32 [el raw data lines of the image and the remaining high resolution lines from an image without the simulated lesion. A zoomed region of the phase image centered on the rectangular lesion is shown here.

from Table 1, it can be seen that this factor increases approximately linearly with TE (TE = 30, 40, 60 msec: 3, 4 , 6 and corresponding temperature sensitivity (in phase- degrees/"(=) = 1.42, 1.84, 2.68: 3, 3.9, 5.7). This is in accordance with Equation 111. The SD of the phase images, expressed in temperature units, decreases with TE, in the range 30 to 60 msec (Table 1). From Equation [3] , it can be seen that the phase noise, in temperature units, has a minimum at approximately TE = T2*. The T2* of the gel phantom is -52 msec, and thus, a minimum in phase noise is expected around this value. This is verified by the experimentally observed SDs, in which all of the TEs used were less than or approximately equal to the TE at the theoretic minimum of phase noise. The SDs quoted here for the gel phantom are larger than that obtained for temperature mapping using phase images at higher field strengths of 1.5 T and 4.7 T ( 10.13). The rea- son for the increased phase noise at low field strengths may be attributed to the decreased signal intensity of the magnitude signal at the lower fields [magnitude signal intensity increases approximately linearly with the main magnetic field]. The frequency shift of the bovine liver sample is less than that of the gel phantom (.007 ppm/ "C compared with .015 ppm/"C for the gel). This may be because liver tissue may contain fat, the resonance frequency of which does not shift as much with temperature as does the water resonance (10). The measurement ROI may contain both types of tissue. leading to a reduction in the observed phase shift. The decreased frequency shifts of muscle and tumor compared with that of pure water have been reported earlier a t 1.5 T (14,22).

The choice of TE for the purpose of monitoring temperature changes depends on the application. For interventional procedures such as hyperthermia, in which the temperature changes occur slowly, temperature measurements can be performed at fairly long intervals. How- ever, for interventional procedures such as RF and laser ablation, it is important to have a high temporal resolution because the heat-lesioning procedure is completed in a fairly short time (- 1 4 minutes). Therefore, a gradient echo with a TE of 60 msec can be used in conjunction with hyperthermia applications, whereas shorter TE gradient-echo sequences may be necessary to monitor interventional procedures. As in the choice of TE of the sequence, the choice of TR is also dependent on the application. If speed is critical, it will be important to reduce TR, even at the expense of increased SD of the measurements. We chose a sequence with a TE of 30 msec and a TR of either 60 msec or 100 msec, with a combination of 50% keyholing to achieve a balance between temperature sensitivity, SD of temperature measurements, and scan times, to monitor temperature during RF ablations.

The magnetic field drift is critically dependent on the magnet equilibration condition, and it was found that when the magnet and scanner system were allowed to sta-

bilize for more than 24 hours, the phase drifts were much lower than when the system was switched on for only a few hours. In the latter condition, initial measurements of phase in the central phantom showed that the drift correction using the peripheral reference tubes was not adequate (phase drifts were greater by a factor of 3 compared with when the magnet was stabilized for more than 24 hours). All phase measurements reported here were therefore performed with the magnet left on continuously. The drift before correction) in a period of 2 hours in the open system (-120°C) is higher than that found at high fields (- 30°C) in a closed system (1 0). This is not surprising because superconducting systems are inherently more sta- ble than the electromagnet systems (23). In fact. the manufacturer-quoted field drift for our system is I. 1 ppm/ hour after a warm-up period of 30 minutes, which is nearly a factor of 10 greater than that quoted for superconducting systems [23]. The important aspect, however, is that this drift could be corrected using the reference phantoms (Fig. 5). The corrected phase drift over 2 hours was in temperature units, within the average SD of the phase measurements ( 4 1.44"C).

Gradient-echo sequences are sensitive to magnetic field inhomogeneities, especially at longer echo times. Further- more, the design of the open magnet is such that an in- teraction can exist as a result of the coupling of the main magnetic field with external magnetic fields (a fringe field coupling). If the external magnetic fields are fluctuating, then view-to-view phase changes during acquisition de- grade the image quality. In phase images, this factor could increase noise, ie, the SD of phase measurements. One method of correction involves the use of a navigator echo acquired at each level, the phase of which is used to correct for view-to-view phase variations (15). Therefore, we investigated gradient-echo sequences at TE = 30.40. and 60 msec with navigator-echo phase correction. From Table 1, it can be seen that the SD of the measured temperature is less for the navigator-corrected sequences when compared with a conventional sequence with the same parameters. This is also reflected in higher values of the correlation coefficient of the linear fit when comparing the navigator corrected with conventional FLASH sequences. This is as expected because the navigator-echo-based phase correction scheme reduces the view-to-view phase variations. The temperature sensitivity that is dependent on the TE, however, is not significantly different between the corrected and uncorrected sequences. The navigator- echo sequences may show an even greater advantage over uncorrected sequences in scanning a live subject in whom gross motion is a potential source of phase errors. Contri- butions to phase noise from cardiac motion, breathing, and vascular pulsatility will certainly increase this factor in subjects compared with stationary phantoms. However, evidence from the success of navigator-echo corrected-diffusion sequences (1 5), in which sensitivity lo motion is


I Table 2 I Spatial and Contrast Distortions in Simulated Keyholed Lesions

Number of YO Spatial Yo Contrast Acauired Levels Distortion Distortion

I 3.56 0 0 I _ _ - 192 I 128 64

3T7 4.2 14

2.6 10.3 7.4

32 52 43.9

more severe than in phase imaging, indicates that phase noise contribution from physiologic motion can be reduced by incorporation of navigator-echo-based schemes.

Keyholing techniques have been proposed to increase the temporal resolution of a dynamic scan to monitor contrast changes in amplitude images after administration of the contrast agent. The processing of phase images (to remove phase wraps and field drifts) is more involved than direct measurements of signal intensity changes from amplitude images during a dynamic study. There have been no earlier investigations of the effect of keyholing on phase images. We undertook a gel phantom study to confirm the accuracy of temperature measurements obtained with the keyholing technique. The temperature sensitivity of the full resolution and keyholed images are in good agreement (Table 1). The SD of the phase measurements increased by a factor of --g2 (keyhole sequence SD = +3.69", nonkeyhole sequence SD = +2.68"). This is because the magnitude image SNR of the keyholed images is reduced compared with a full resolution image by the square root of the ratio of the number of levels in the full resolution to that acquired in the keyholed image. This reduction in magnitude SNR is reflected in the SD of the phase measurements (Equation [211.

The gel phantom calibration with a keyholed technique confirmed that the process of unwrapping and drift correction using a reference full resolution image did not introduce any significant difference in the temperature sensitivity measurements. The gel phantom has a dimen- sion of 10-cm diameter X 5-cm height and is heated uni- formly so that this study does not reflect the effect of keyholing on the spatial resolution. We investigated this latter effect by a simulation study with a heat lesion simulated in a phase image. Visual inspection of the keyholed images using 192, 128, 64, and 32 levels reveals that the lesion is well reproduced to 64 levels. The percentage spatial and contrast deviations (Table 2) are within 14Vo and 8%, respectively, down to 64 levels. The spatial and contrast distortions are unacceptably high, however, when 32 levels are used. This is in rough agreement with the theoretic calculation presented in ref. 18 for the spatial resolution of magnitude keyholed images used in a dynamic contrast study. Here, a keyhole factor (KF) is defined as the following ratio.

KF = number of levels in the reference image/number of levels in the keyholed image

The theory predicts that for objects 5 pixels in size or larger, minimal intensity distortions (less than 10%) are seen for a keyhole factor of 4. In the simulation study, the 10 X 7 mm lesion size in pixels is 12 X 9 pixels [the FOV of the high resolution image = 180 mm, matrix size = 256 X 256). Because the object of interest (ie, the lesion) is larger than 5 pixels, a value of 4 for the keyhole factor should cause less than 10% intensity distortions in the dynamic changes. Because the high resolution ma-

trix was acquired at 256 levels, it follows that at least 64 levels must be acquired in the dynamic study to keep the distortions within 10%. This also is the result of the simulation study, which shows that reducing the number of levels in the dynamic scan down to 64 levels (KF = 4) still produces images with acceptable quality (Fig. 8). Reduc- ing the number of levels in the dynamic scans to 32 produces unacceptable intensity distortions [KF = 8) compared with using KF = 4 (64 levels of the dynamic scan). The jump in the percentage distortions on going from KF = 4 to KF = 8 (Table 2) is larger than the theoretic estimation for amplitude images (18) and probably reflects the different factors that affect phase and amplitude images. These simulation studies thus enable one to set a lower limit on the number of levels that must be acquired in a keyholed study to accurately monitor heat lesions of size in the range of 10 X 7 mm.

In the current investigation, the choice of a copper needle as the RF probe material was essentially dictated by susceptibility considerations. However, it should be men- tioned that copper is not biocompatible and thus cannot be used in vivo. The requirements for a MR-compatible RF probe material are a s follows: (a) biocompatibility, (b) reasonably high strength and ductility, (c) good electric conductivity, and (d) susceptibility matching that of brain tissue, resulting in negligible magnetic field inhomogeneities a t the needlejtissue interface. Titanium and plat- inum are possible choices as electrode material because their susceptibilities are closest to that of brain tissue and meet the other requirements listed above (24). Cop- per needles were used here because the entire study was conducted in gel phantoms and ex vivo tissue only.

Several factors were involved in choosing an appropriate sequence to monitor the RF ablation process. The RF probe introduces magnetic field inhomogeneities that cause susceptibility artifacts, which are more pronounced at longer values of TE. Echo times in the range of 30 to 60 msec did not produce artifacts around the needle in phase images. The second factor in sequence selection was the scan time, because the RF ablation lasts from 1 to 4 minutes. To adequately monitor the temperature changes during the RF ablation procedure, it is necessary to have a temporal resolution of less than 15 seconds. This was accomplished with a keyholed sequence with 64 acquired levels and a TR of 60 msec, which reduced scan times to the order of 4 seconds, per- mitting adequate resolution during a 2-minute RF ablation procedure. The temperature profiles of the RF lesions shown during and immediately after the ablation clearly demonstrate the potential for using keyholed phase images of gradient-echo sequences at .2 T. The interference from the RF generator caused a deterioration of the image SNR, especially at high voltage settings of the generator. However, this was minimized by incorporation of the iso- lation filter at the output of the generator. Temperature measured by the thermocouple attached to the RF probe was in good agreement with the hottest temperature within the lesion temperature profile (Fig. 9). The temperature profiles obtained in the bovine liver sample were noisier than those obtained in gel phantoms. This is pri- marily due to the fact the tissue SNR is reduced by a factor of 2.5 compared with that of the gel phantom. Ad- ditional factors that lead to noisier profiles may be the inherent inhomogeneous composition of the tissue (mix- ture of fat and water protons).

Factors such a s the lower SNR of images, larger drift of the electromagnet, coupling to external magnetic fields, and susceptibility artifacts due to the RF probe render it difficult to make accurate temperature measurements

926 JMRl SeptemberlOctober 1997

100

40 80

70

60 50 4n

30

20

10

0

a d

1 DO 90

BU 70

bO su 40

30 ZU

10

0

b e

Irn 90

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6[1

SO

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30

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Figure 9. Gel phantom images at different time points of an FW ablation procedure. Images shown her? have been phase unwrapped and corrected for drifts. Images were acquired while the RF power was increased (a+) and after the RF generator was powered down (d-f). The three-dimensional temperature profiles are shown [to the right of the image) for the rectangular ROI (in white) shown super- posed on the subtracted phase image.

from phase images using a low field system. This paper addresses these factors and demonstrates feasibility at low fields in stationary gel phantoms and ex vivo tissue, albeit with a reduction in temperature accuracy compared with imaging at 1.5 T. An extension to in vivo imaging will introduce the additional factor of physiologic motion contributing to phase noise, but it is anticipated that the use of navigator echoes will reduce the view-to- view variations in phase to permit meaningful measurements of temperature.

CONCLUSIONS This report describes the measurement of temperature

during and immediately after an RF ablation procedure using keyholed gradient-echo phase images with navigator-echo-based correction. Temporal resolution of temperature maps was 4 seconds in gel phantoms and 13 seconds in the tissue sample, achieved by keyholing down to 64 and 128 levels, respectively. The susceptibility-induced artifacts were reduced by a combination of a copper needle and a relatively shorter TE for the phase images. This study demonstrates the feasibility of monitoring temperature of a heat lesion in a bovine liver sample during an RF ablation procedure. This has immense potential for MR-guided interventional procedures because it provides accurate control via real-time temperature monitoring of the lesioning process.

Acknowledgments: The authors thank Dennis Atkinson, Siemens Medical Systems, Iselin, N J , for helpfiil discussions.

References 1. Cline HE, Hynenen K, Hardy CJ, Watkins R, Schenk JF, Jolesz

FA. MH temperature mapping of focussed ultrasound surgery. Magn Reson Med 1994: 31:628-636.

2. Hathout G, Lufkin R. Jabour B, Andrews J, Castro D. MR guided aspiration cytology in the head and neck at high field strength. J Magn Reson Imaging 1992; 2:93-94.

3. Matsumoto R, Oshio K, JolesL FA. Monitoring of laser and freezing-induced ablation in thc liver with TI-weighted MR imaging. J Ma@ Reson Imaging 1962; 2:555-562.

4. Cline HE, Schenck JF, Hynynen K, Watkins RD, Souza SP, Jolesz FA. MR-guided focused ultrasound surgery. J Comput Assist Tomogr 1992; 16:956-965.

5. Jolesz FA, Silverman SG, Kikiriis R, et al. Interactive image- guided interventions in an open-configuration MR imaging system (abstract]. Radiology 1994; 193(P):252.

6. Chung YC, Duerk JL, Lewin JS. Generation and observation of radiofrequency thermal lesion ablation for interventional magnetic resonance imaging (abstract]. In: Book of abstracts: International Society for Magnetic Resonance in Medicine 1996. New York: International Society for Magnetic Resonance in Med- icine, 1996; 1743. Hall AS, Prior MV, Hand JW, Young IR, Dickinson RJ. Obser- vation by MR imaging of in vivo temperature changes induced by radiofrequency hypothermia. J Comput Assist Tomogr 1990: 14:430-436.

8. Morvan D, Leroy-Willig A, Malgouyres A, Cuenod CA. Jehenson P, Syrota A. Simultaneous temperature and regional blood volume measurements in human muscle using an MRI fast diffusion technique. Magn Reson Med 1993; 29:371-377.

9. MacFall J, Prescott DM, Fullar E, Samulski TV. Temperature dependence of canine brain tissue diffusion coefficient measured in vivo with magnetic resonance echo-planar imaging. Int J Hyperthem 1995; 11:73-86.

10. De Poorter J, De Wagter C, De Deene Y, Thomsen C, Stahlberg F, Achten E. Noninvasive MRI thermometry with the proton

7.


resonance frequency [PRF] method: in vivo results in human muscle. Magn Reson Med 1995: 33:74-81.

11. Chung AH, Hynynen K, Colucci V, Oshio K, Cline HE, Jolesz FA. Optimization of spoiled gradient-echo phase imaging for in vivo localization of a focused ultrasound beam. Magn Reson Med 1996; 36:745-752.

12. De Poorter J, De Wagter C, De Deene Y, Thomsen C. Stahlberg F, Achten E. The proton-resonance-frequency-shift method compared with molecular diffusion for quantitative measurement of two-dimensional time-dependent temperature distribution in a phantom. J Ma@ Reson €3 1994; 103:234-241.

13. lshihara Y, Calderon A, Watanabe H, et. al. A precise and fast temperature mapping using water proton chemical shift. Magn Reson Med 1995; 34:814-823. Cline HE, Hynynen K, Schneider E, et al. Simultaneous magnetic resonance phase and magnitude temperature maps in muscle. Magn Reson Med 1996: 35:309-315.

15. de Crepigny AJ, Marks MP, Enzmann DR, Moseley ME. Navi- gated diffusion imaging of normal and ischemic brain. Magn Re- son Med 1995: 33:720-728.

16. Jones RA, Haraldseth 0. Muller TB, Rinck PA, Oksendal AN. K-space substitution: a novel dynamic imaging technique. Magn Reson Med 1993; 29:830-834. Plewew DB, Bishop JE, Soutar I, Cohen E. A segmented k- space imaging method for 3D dynamic breast MRI (abstract). In:

14.

17.

Book of abstracts: Society of Magnetic Resonance in Medicine 1993. Berkeley, C A Society of Magnetic Resonance in Medicine, 1993; 1251.

18. Spraggins TA. Simulation of spatial and contrast distortions in keyhole imaging. Magn Reson Med 1994; 31:320-322.

19. Organ LW. Electrophysiologic principles of radiofrequency lesion making. Appl Neurophysiol 1976; 39:69-76.

20. Cosman ER, Rittman WJ. Physical aspects of radiofrequency energy applications. In: Huang SKS, ed. Radiofrequency cathe- ter ablation of cardiac arrhythmias: Basic concepts and clinical applications. Mount Kisco: Futura, 1994: 13-23.

21. Gao J H , Xiong J, Lai S, et al. Improving the temporal resolution of functional MR images using keyhole techniques. Magn Reson Med 1996; 35:854-860.

22. MacFall, Samulski TV, Prescott DM, Fullar E. Thermal mapping using the MR image phase in vivo during hyperthermia (abstract). In: Book of abstracts: Society of Magnetic Resonance 1994. San Francisco, CA. Society of Magnetic Resonance, 1994; 1578.

23. Thomas SR. Magnets and gradient coils: types and character- istics. In: Bronskill M J , Sprawls P, eds. The Physics of MRI, 1993. New York: American Institute of Physics, 1993; 56-58.

24. Comte P. Electrode technology. Jn: Wieser HG, Elger CE, ed. Presurgical evaluation of surgeries, 1987. Berlin: Springer-Ver- lag, 1987: 109-111.

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phase imaging on a .2-t mr scanner: application to temperature monitoring during ablation procedures

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