Ultrasound-Guided MRI: Preliminary Results Using aMotion Phantom

Matthias Gunther* and David A. Feinberg

In principal, both ultrasound (US) imaging and MRI can be per-formed simultaneously within the magnetic field of the scannerdue to their different physical natures (i.e., sound waves vs. elec-tromagnetic radiation). One potential application is to use USinformation for dynamic spatial organ localization to improve MRimaging, similar to MR navigator echoes. In this work, preliminaryresults are presented for US-guided MRI using position and ori-entation information extracted from US data to update the imageslice position of an SSFP sequence in real time. Effective prospec-tive motion compensation is shown for a phantom whose sinusoi-dal displacement can be fully corrected. Possible applications ofsimultaneous US-MRI acquisitions include real-time US-guidedcardiac MRI or interventional procedures. Magn Reson Med 52:27–32, 2004. © 2004 Wiley-Liss, Inc.

Key words: hybrid imaging; ultrasound; navigation; magneticresonance; MRI

Ultrasonography or ultrasound (US) is a diagnostic modal-ity that acquires tomographic images of the body interiorusing sound waves in the low megahertz range. Such USsignals and the electromagnetic MR signals, in theory, canbe detected independently and without interference. Con-sequently, the US scanner might be operated in the mag-netic field of the MR scanner without the two modalitiesadversely affecting one another (1). Several different com-binations to hybridize these two techniques, where onemight be used to enhance the other, can be envisioned. Forexample, improved biopsy localizations (2) might beachieved by combining the real-time speed of US with theMR contrast enhancement of tumors in fusion imageswhile a person remains in the magnet. An alternativeapplication (3), related to the present experiment, is toattempt to improve the image quality of cardiac imaging orto overcome motion artifacts using US information, similarto the use of MR navigator echoes.

There are a few fundamental differences between navi-gator echoes and the potential use of US information forspatial localization. A typical navigator signal is created bya selective RF pulse excitation of a limited volume yield-ing a 1D line image. Using real-time US, cine 2D imaging ofa region of interest (ROI) can be acquired. A 2D US imageand MR pulse sequence could be acquired simultaneouslyin the same ROI, with no time penalty in the MR imageacquisition, disturbance of a steady-state free precession(SSFP) sequence, or MR signal saturation effects from the

US. A true 3D motion trajectory, in principle, could betracked by combining multiple US images acquired simul-taneously at different orientations using multiple USprobes. We present preliminary in vitro investigationswhere cine 2D US imaging was performed in the MRscanner to obtain positional information on a movingstructure which in turn was used to continuously adjustthe slice position of the MR signal acquisition during theSSFP sequence.

MATERIALS AND METHODS

In general, the experimental setup consists of four majorcomponents (see Fig. 1): 1) a standard clinical US-scanner;2) a PC with A/D card for US data acquisition and object-tracking; 3) a standard clinical 1.5 T MR scanner; 4) amotion phantom and probe mounting device.

A standard clinical US scanner (VST Master, Diasonics)including a cardiac transducer (2.25 MHz) was used todetect the position of the structure to be tracked. The USscanner was modified to access the US raw data and thecorresponding line number information directly.

The US probe was attached to a plastic water tank witha mounting device built with a K’NEX construction kit (seebelow). Since this mounting structure remained fixed withrespect to the water tank, the position and orientation ofthe probe could be measured manually. This informationwas used to transform the actual position of the trackedobject from US coordinates to MR coordinates.

Considerable efforts were made to shield the US system,especially the cable to the probe, to avoid electronic inter-ference with the MR coils. The US probe itself was coveredwith metal-coated Mylar (Birthday balloon, Betallic),which is thin enough to permit the passage of ultrasoundpulses yet thick enough to shield the electromagnetic RF-radiation at frequencies that can be detected by the MRreceiver coil.

A computer (Pentium 3 733 MHz, Windows 2000,Adlink 20 MHz A/D Card NuDAQ PCI-9810 with 10 bitresolution) sampled the US raw data including the corre-sponding line numbers, which are obtained from the USscanner via coax cables, in real time. The current linenumber is acquired via the user port. No scan conversionis performed but the tracking is based on the raw data.In-house-developed software (using the algorithm de-scribed below) detects the position of the moving object(i.e., balloon, heart) and transforms the US coordinates tothe corresponding coordinates of the MR scanner. The dataof the calculated position and orientation are sent to theMR system via a TCP/IP connection using named pipes.

The MRI scanner was a 1.5 T whole-body system (So-nata, Siemens, Erlangen, Germany) operating with a dualcomputer sequence control system, for hardware control

Advanced MRI Technologies, Sebastopol, California.Grant sponsors: NIH, National Center for Research Resources; Grant number:R43RR17474.*Correspondence to: Dr. rer. nat. Matthias Gunther, Advanced MRI Technol-ogies, 652 Petaluma Ave., Sebastopol, CA 95472. E-mail: matthias.guenther@AdvancedMRI.comReceived 28 October 2003; revised 12 March 2004; accepted 15 March 2004.DOI 10.1002/mrm.20140Published online in Wiley InterScience (www.interscience.wiley.com).

Magnetic Resonance in Medicine 52:27–32 (2004)

© 2004 Wiley-Liss, Inc. 27

and user interface. An in-house-developed program re-ceives the positional information and sends it to the MRinterface computer and then to the sequence. No hardwaremodifications were needed on the MR scanner.

Existing MR sequences had to be modified to be able toadjust the slice position and orientation during runtimeaccording to the phantom location identified by the track-ing program. We implemented a fully balanced SSFP pulsesequence (trueFISP) with real-time update capabilities.The sequence automatically uses the shortest possible rep-etition time for a given parameter set (field of view (FOV),sampling bandwidth, matrix size). The current slice posi-tion and orientation is read from a set of variables, whichis updated as soon as new positional data from the UStracking system is available. The sequence then calculatesthe necessary gradient and RF-pulse commands and sendsthe information to the MR control system. Ideally, thecontrol information of a single line should be processed atonce to provide the shortest latency time and take advan-tage of the most recent position information available.However, the MR acquisition has to be continuous or theexecution of the sequence will stop. Since the MR controlsystem has to process the information provided by thesequence, the necessary data has to be present at a certaintime before it can be played out. The execution time theMR control system needs to process the sequence data isalmost independent of the amount of control informationwithin a single block. This means that to ensure properexecution of a sequence, control information for multiplelines has to be grouped together before sending it to theMR control system, which increases the latency.

We two sets of parameters: one to acquire images withlow spatial but high temporal resolution, and one withhigher in-plane resolution but increased acquisition time.The parameters were (TE / TR / matrix size / acquisitiontime): Set 1: 1.5 ms / 3.0 ms / 64 � 128 / 192 ms; Set 2:1.9 ms / 3.8 ms / 192 � 256 / 730 ms.

To examine the feasibility of the proposed technique, amotion phantom was developed that is capable of perform-ing cyclic motion at different frequencies (�1 Hz and�1/6 Hz), which are comparable to cardiac and respiratorycycles in humans. The US and phantom mountings are

assembled from a plastic toy construction set called “BigBall Factory” (K’NEX Industries) as proposed in Ref. 4. Atank filled with Gd-doped water was lined with bubblewrap on the inside to scatter the ultrasound pulse whichattenuates artifacts from multiple reflections at the tankwall. The heart phantom consists of a water-filled balloonsurrounded by plastic rods, which stabilize it and, at thesame time, are of similar size as coronary arteries. Thisballoon is moved by a long rod (1.4 m), which is attachedto an electric motor via a gear wheel-based transmissionoutside the magnet.

The transmission allows a quick switching between thetwo rotational speeds (cycle duration 1 and 6 sec, respec-tively). Other speeds can be obtained with minor modifi-cations. The main motion of the phantom is along onedirection (head–foot), which is parallel to the axis of themagnet, but there is also a small perpendicular componentof motion “x” (left–right). Similarly, an attempt was madeto minimize the “y” vertical axes of motion since theimaging plane of the ultrasound transducer was chosenalong the “z-axis,” parallel to the patient table.

Active contours (5,6) were used to describe the structureto be followed in the acquired US images. A contour con-sists of points connected by lines to delineate the objectshape. Only global state transformations but no modifica-tion of individual points of the active contour is allowed.These transformations are constrained to translation andscaling in the x- and y-direction as well as rotation. There-fore, all possible forms of a given contour can be describedusing a five-dimensional state space. Further improve-ments can be made by incorporating more degrees of free-dom (DOF) to account for structural changes, such as de-formations. Nonetheless, the present work incorporatesonly the five basic transformations.

An energy term is calculated for the contour by sum-ming the inverse exponentials of the intensity of the USimage at the location of each point. The energy of a contouris an inverse measure for the probability that the contourmatches the desired structure within the US data. Here,the lower the energy the higher the matching probability.The CONDENSATION algorithm (7) was used to calculatethe most likely contour. It samples the state space numer-

FIG. 1. Flow chart illustrating the combined use ofcine ultrasound imaging with MR-imaging, Thesecond (or more) US probe is optional, since ini-tially only one probe was used. The US image issent to a computer to extract the positional andorientational information which is used to updatethe imaging parameters of the MR sequence.

28 Gunther and Feinberg

ous times in a random manner, but following a determin-istic diffusion equation, which describes the possiblemovement of the structure. The most likely contour isestimated by calculating the sum of all samples weightedby their probability. This method is very stable even forimages with low signal-to-noise ratios. Furthermore, theCONDENSATION algorithm provides an adjustable butfixed execution time necessary for constant update rates(12 ms for the described hardware and an average numberof contour points, i.e., 20).

To further improve the accuracy of the tracking algo-rithm a simple form of extrapolation of the actual positionypred(t) was implemented. This allows compensation ofthe latency between US measurement and MR slice acqui-sition. We used a simple second-order extrapolation byLagrange’s classical equation utilizing three recent mea-surement points yn(tn), which were known to lie within atleast one-fourth of a period of the motion cycle:

ypred�t� � �t � t2� � �t � t3�/�t1 � t2�/�t1 � t3� � y1

� �t � t1� � �t � t3�/�t2 � t1�/�t2 � t3� � y2

� �t � t1� � �t � t2�/�t3 � t1�/�t3 � t2� � y3. [1]

The latency of the system was estimated using three dif-ferent methods: 1) a theoretical summation of the execu-tion times for each step; 2) measurement without latencycorrection, using the amplitude of the residual motionthereof and utilizing Eq. 3 of the Appendix to calculate theapparent delay; 3) measurement with several latency cor-rection delays (0, 50, 80, 100, and 150 ms). These measure-ments were performed with the first MR imaging parame-ter set described above to yield a high temporal resolution.

RESULTS

Figure 2 shows the phantom and probe mounting deviceon the patient table of the MR magnet. On the right-handside the electric motor including the transmission can be

seen. The plastic rod, which transfers the rotational mo-tion to the balloon phantom, is in the center of the image.On the left the water tank containing the balloon restsinside the MR head coil. In front of the head coil the probemounting with the attached US transducer can be seen.The US scanner and computer setup are not visible. Notethat the phantom and patient table are moved to the centerof the magnet during the experiments.

Images were acquired using the SSFP sequence. As ex-plained in Materials and Methods, instead of sending asingle line to the MR control system, multiple lines have tobe sent in a block. In most cases 25 ms is sufficient toensure that there are always data in the queue. For atypical repetition time of an SSFP sequence of 3 ms, eightlines have to be grouped. This is still adequate to reducemotion artifacts which occur during acquisition. The im-provement in image quality is demonstrated in Fig. 3,where the slice position was updated from US informationwithin the SSFP sequence while maintaining steady state.The left image was acquired without tracking, showing astar-like motion artifact from the plastic rods on the water-filled balloon. By enabling the tracking, the balloon andthe rods stay in the middle of the image in a cine acquisi-tion and are depicted clearly. Note that the water tank isblurred now since it appears to move relative to the bal-loon. For cardiac imaging this is unlikely to represent aproblem, as the displacement of the heart is not as large asin the phantom and no high signal structures are close by.

The execution time for each step is presented in Table 1,yielding an estimated theoretical overall delay betweenthe measurement of the phantom position by the US scan-ner and the updating of the actual slice information on theMR scanner of about 30 ms. Due to MR scanner con-straints, there is an additional delay of 20–40 ms (usually8 times TR of SSFP sequence) until the MR slice is actuallyacquired. Thus, the delay expected from known sourceslies between 50 and 70 ms.

Figure 4 displays the position of the balloon vs. timewith and without latency correction. A single line of each

FIG. 2. Photo of motion phantom and US probemounting. The water tank is filled with Gd-dopedwater and lined internally with conventional air-bubble plastic wrap to attenuate reflection of USsignal. The heart phantom consists of a water-filledballoon surrounded by four plastic rods and ismoved by the lateral plastic rods connected to thegears and motor at the end of the patient table. TheUS probe cable is shielded with conventional alu-minum foil, whereas the probe window itself isshielded by thin metal-coated Mylar foil.

Ultrasound-Guided MRI 29

image is shown for the whole time series (time alongabscissa). Initially, the tracking is switched on and theballoon appears stationary. Further to the right, tracking isswitched off and the motion of the balloon relative to theMR system becomes apparent. Nonetheless, some residualmotion of the balloon remains even with tracking enabled(top image in Fig. 4). The time delay between US dataacquisition and actual update on the MR slice positionleads to a residual motion with (almost) a �/2 phase shift,as shown in the Appendix. This phase shift can be clearlyseen in the top image of Fig. 4. Using Eq. 3 of the Appendixand a measured relative amplitude of 0.12 � 0.02 for theresidual motion the apparent delay can be estimated to bearound 125 � 20 ms.

Measurements were performed with different latencycorrections as described above. The same small amount ofresidual motion was found for 50 and 100 ms delay, witha minimum at 80 ms. The bottom image in Fig. 4 wasacquired with a latency correction of 80 ms, which re-moves the residual motion almost entirely.

DISCUSSION

Previously, ultrasound has been used inside the MR scan-ner as a therapeutic modality to perform thermal surgery(8,9), where it had been found that simultaneous MR andUS imaging was impossible. However, for such an appli-cation much higher transmitter intensities than commonly

used for diagnostic purposes are employed. US waveshave been used simultaneously within the MR scanner forapplications at lower power levels (10,11), although re-sults for combined MR and US imaging were never pub-lished. At diagnostic transmitter gains only minor interfer-ences between the two imaging modalities are expecteddue to their different frequencies and the intrinsic physi-cal differences of their signals (i.e., sound waves vs. elec-tromagnetic radiation). Combining the data from the twoimaging modalities in real time allows for exchange ofphysiologic information and permits modification and im-provement in the MRI data acquisition. For example, thespeed of real-time ultrasound imaging could be combinedwith the tissue characterization during contrast enhance-ment, the resolution and large FOV of MRI to track devicesor needles during interventional procedures simulta-neously, as published recently in a sequential setup (2). Inthis work, we used the US data to track the displacementof a motion phantom and to adjust the MR imaging sliceaccordingly. In the future this will be applied to cardiactracking.

The interference level between both modalities is verylow. There is no sacrifice of quality for the US imagesexcept for distortions on the monitor display due to thestrong magnetic field of the MR scanner. This is a displayproblem but no degradation of the US data itself. Someinterference can be seen within the MR images due toelectronic activity but we are confident these noise sourcescan be reduced substantially by improving the shielding.

Several publications deal with US equipment within theMR scanner, e.g., focused US for surgery purposes (8,9),US exposimetry to image the impact of US waves on tissue(10), or fetal cardiotocography to examine the influence ofMRI on human fetuses (11).

The images shown in Figs. 3 and 4 used the US spatialinformation to adjust the MR slice position within an SSFPsequence, thus following the moving object to keep itsmagnetization in steady state. The improvement in imagequality by reducing the motion artifacts is evident.

FIG. 3. Comparison between untracked(left) and tracked (right) object, with andwithout US tracking, respectively. On theleft side with US tracking disabled, theedges of the water tank are sharp, while theballoon is blurred due to motion artifact.Note the star-like motion artifact of the rodssurrounding and moving with the balloon.On the right image, the US tracking is en-abled and produces a detailed image of theballoon and the rods. The water tank now isblurred since the spatial parameters of theMR sequence are continuously updated tofollow the moving balloon.

Table 1Average execution time for each step during the real-time UStracking

US image acquisition (56 Hz) 18 msFeature extraction: Tracking algorithm 16 ms(acquisition during tracking) Net 18/2 � 16 ms 25 msTransfer coordinates to MR host computer via named

pipes �0.1 msHost computer to MR control computer 4 msTotal cycle time of US coordinate update: 29.1 ms

30 Gunther and Feinberg

A particle filter algorithm was used for tracking theposition of the motion phantom. A great advantage of thistype of optimization technique is the fixed execution time,which can be controlled in terms of tracking accuracy byadjusting the number of samples drawn. This allows thesystem to run in real time with a constant update rate.Furthermore, the tracking algorithm is relatively robust,especially in cases where the tracked object is (partially)occluded.

Since there is an inherent delay between position mea-surement of the tracked structure and the actual update ofthe imaging slice location, motions with large displace-ments cannot be compensated for entirely. In this work,we used a second-order extrapolation of the actually mea-sured position to predict the location of the tracked struc-ture at the time when the imaging will take place. Tominimize the error of this prediction the extrapolationtime has to be short with respect to the used measurementdata points. In our case this constraint was best fulfilled forthe slower motion representing the breathing cycle. Thecurrent setup is not capable of following the motion phan-tom during a simulated heart cycle due to the latency ofabout 80 ms. The estimate of the apparent latency usingthe equations of the Appendix yielded 125 � 20 ms, whichis slightly longer than the actually measured delay. Sincethis estimate is based on the relative amplitude of theresidual motion, a large systematic error is unavoidabledue to the low resolution of the underlying images. Be-sides this uncertainty there might be another reason for themismatch between estimated and measured delay. Thelarge apparent latency could result from image distortionsrather than from a temporal delay. We are using a dedi-cated cardiac MR scanner with a short gradient coil inset.Consequently, the gradients become quickly nonlineareven for relatively short distances away from the magnetisocenter. The effect can be clearly seen in Fig. 4, wherethe residual motion shows a deviation from the expectedsinusoidal form precisely when the balloon approachesthe outer regions of the water tank. For cardiac imagingthis nonlinearity poses no problem, since the patients arepositioned such that their hearts are centered in the mag-net and the motion of the heart will be much smaller thanthe motion of our phantom.

With the given hardware, an update can be performedevery 30 ms but with a latency of about 80 ms. About

25–30 ms of the latency originate from the limited process-ing speed of the MR control system as described in Mate-rial and Methods. This latency could probably be reducedby performing some calculations (e.g., geometrical gradi-ent pulse transformations) up front. However, this latency,which could be termed MR scanner reaction time, poses aproblem for all real-time feedback navigation techniques.Ten to 20 ms of the latency can be attributed to the acqui-sition of the US image. By reducing the FOV the imagingtime can be reduced further. Unfortunately, instead ofincreasing the frame rate the clinical US scanner used inthis work automatically switches to a mode where eachline is repeated multiple times, limiting the actual framerate to about 50 Hz. Most of the remaining latency seems tostem from the filter and amplifier chain, which processeseach US line before it can be digitized by the PC describedabove. Thus, major improvements are possible by switch-ing to different US hardware (e.g. Ref. 12) to allow moreflexible control over the US imaging process. With thesemodifications an update rate of at least 50 Hz with areduction of latency from 80 ms to an estimated 20 msseems feasible in the near future.

Finally, it must be noted that the present experimentwas conducted to establish proof of principle, in vitro, bytracking a sinusoidal phantom motion, which was easy tocontrol and to experimentally validate. Such motion doesnot normally exist in vivo, particularly not in diseasedhearts with arrhythmic contractions. Only weak assump-tions were made about the sinusoidal nature of the motion,although the periodic nature of the motion might benefitthe extrapolation algorithm. However, the transition of thepresented results to nonperiodic motion should bestraightforward, especially with reduced latency, thus re-ducing the uncertainty of the extrapolation.

While US is routinely used to image nonperiodic heartmotion, there are several technical issues which will needto be overcome before US-guided MRI might improve car-diac imaging. For instance, the appropriate positioning ofthe US probe in a sonographic window could interferewith the RF coil placement, although the substernal posi-tion looks promising. Safety studies using US with differ-ent pulse sequences need to be conducted. The presentexperiment only evaluated US guidance in an SSFP se-quence, which is commonly used in cardiac MRI. How-ever, evaluation of other imaging sequences and actual in

FIG. 4. Position of the balloon vs. time. Left:tracking is enabled; Right: no tracking isapplied. Only a single line of each image isshown for the whole time series visualizingthe motion of the phantom. Some residualmotion is apparent with tracking enabled(top) originating from two imperfections ofthe prototype system. One source of errorresults from no compensation for spatialdistortions from nonlinearities of the gradi-ent system. The larger error source is due tothe inherent delay between US image mea-surement and actual update of the position.At the bottom, results are shown with addi-tional latency correction where the residualmotion error is removed almost entirely.

Ultrasound-Guided MRI 31

vivo cardiac MR imaging is yet to be established. As the USacquisition is independent of the specific imaging se-quence, there is no apparent limit to combine it with fasterimaging techniques such as partial Fourier or parallel im-aging except for the issue of probe and coil placement.

CONCLUSION

A method was presented which allows simultaneous ac-quisition of US and MR images independent of each other.Our results demonstrate how real-time US imaging can beperformed inside an MRI scanner simultaneously duringthe MRI data acquisition. Results for US-guided MRI on amotion phantom were shown where the US signal acts asa navigator to adjust the slice position and orientation ofan MR sequence in real time. One advantage of US trackingcompared to MR navigator signals is that US does notsaturate the magnetization, so no additional signal atten-uation occurs while imaging the same tissue volume withUS and MRI. The current setup permits measuring andcorrecting for the displacement of a periodically movingphantom every 30 ms. However, further refinements ofhardware and US acquisition are anticipated and couldyield US positional information within a 2 ms TR of theSSFP sequence. This method can be applied to all se-quence types. A long-term goal of this work is for contin-uous 2D or 3D ultrasound tracking of the heart’s positionduring the MR image acquisition to improve spatial reso-lution in MRI cardiac and coronary imaging. With UStracking, the MRI acquisition time could be extended tolonger durations with or without breathholding. It is yet tobe seen if 2D and 3D ultrasound tracking of organs leads toimproved tracking accuracy compared to 1D MR navigatorecho techniques.

APPENDIX

Dependency of Tracked Position for Varying DelaysBetween Measurement and Update

Assume that there is a delay between the measurement ofthe cyclic (frequency f [1/s]) changing object position andthe subsequent update of the MR sequence. With the cyclicfrequency � 2•�•f the deviation �P of the estimatedposition from the true position can be expressed as:

�P�t,� � sin� � t� � sin� � �t � ��

� 2 � sin� �

2 � � cos� � �t �

2��� � 2 � sin� �

2 � � sin� � �t �

2� ��

2�.

For small values of the phase difference will be close to�/2, i.e., if the original motion is sinusoidal the apparent

residual motion of the tracked object will behave like acosine. In general, the phase difference between the orig-inal and the residual motion can be expressed as:

Phase difference ��

2�

�

2. [2]

The amplitude of the residual motion will be:

Relative amplitude �2 � sin� �

2 � � 2 � sin�� � f � �. [3]

If the latency equals the time of one period 1/f (i.e., thecorrection occurs exactly one wavelength behind) the am-plitude of the residual motion will be zero. This holds trueonly for perfectly periodic motion patterns. More complexmotion can be described by a sum of sinusoidal function,thus, using the same formalism as above to describe la-tency behavior of motion-compensated systems.

ACKNOWLEDGMENT

The authors thank Lawrence E. Crooks for help on theshielding of the ultrasound system.

REFERENCES

1. Lunsford LD. Modern stereotactic neurosurgery. Boston: MartinusNijhoff; 1988.

2. Piron CA, Causer P, Jong R, Shumak R, Plewes DB. A hybrid breastbiopsy system combining ultrasound and MRI. IEEE Trans Med Imag2003;22:1100–1110.

3. Feinberg DA, Gunther M. Simultaneous MR and ultrasound imaging:towards US-navigated MRI. In: Proc 11th Annual Meeting ISMRM,Toronto, 2003. p 381.

4. Huber ME, Stuber M, Botnar RM, Boesiger P, Manning WJ. Low-costMR-compatible moving heart phantom. J Cardiovasc Magn Reson 2000;2:181–187.

5. Kass M, Witkin A, Terzopoulos D. Snakes: active contour models. In:Proc. 1st Int. Conf. Computer Vision, 1987. p 259–268.

6. Cohen LD. On active contour models and balloons. Computer VisGraph Image Proc Image Understanding 1991;53:211–218

7. Isard M, Blake A. CONDENSATION—conditional density propagationfor visual tracking. Int J Comput Vis 1998;29:5–28.

8. 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.

9. Cline HE, Schenck JF, Hynynen K, Watkins RD, Souza SP, Jolesz FA.Magnetic resonance-guided thermal surgery. Magn Reson Med 1993;30:98–106.

10. Walker CL, Foster FS, Plewes DB. Magnetic resonance imaging ofultrasonic fields. Ultrasound Med Biol 1998;24:137–142.

11. Shakespeare SA, Moore RJ, Crowe JA, Gowland PA, Hayes-Gill BR. Amethod for foetal heart rate monitoring during magnetic resonanceimaging using Doppler sonography. Physiol Meas 1999;4:363–368.

12. Richard WD, Zar DM, LaPresto EL, Steiner CP. A low-cost PCI-bus-based ultrasound system for use in image-guided neurosurgery. Com-put Med Imag Graph 1999;23:267–276.

32 Gunther and Feinberg