MAGNETIC RESONANCE IN MEDICINE 5, 384-389 (1987)

Localized Imaging Using Stimulated Echoes

PAMELA MILLS,* WIL CHEW,? LAWRENCE LITT,*’? AND MICHAEL MOSELEY~

Departments of TRadiology and *Anesthesiology, University of California, San Francisco, California 941 43

Received April 6, 1987; revised July 27, 1987

Stimulated-echo localized spectroscopy was combined with phase- and frequency-en- coding gradients to obtain “zoom” or magnified images of specific organs in situ. The technique requires neither surface coils nor an imaging coil arrangement that exclusively isolates the target organ. This technique can be readily applied to conventional spectrometer imagers that have limited computational capabilities. 0 1987 Academic Press, Inc.

INTRODUCTION

In conventional 2D spin-echo imaging, the NMR signal is obtained from nuclei everywhere throughout the sensitive volume of the imaging coil (1). In many appli- cations this poses no significant disadvantages since images of the entire sensitive volume are desired. In some special applications, such as fetal or pituitary imaging, most of the sensitive volume is not of interest. The production of high-resolution images of the tissue of interest requires the use of decreased pixel sizes, and this can be achieved in conventional spin-echo imaging by increasing the number of phase- encoding steps and the block size. Unfortunately, such an approach requires sig- nificant computational power that is not always readily available on conventional spectrometer imagers. Furthermore, a straightforward reduction in the field of view in the phase-encoding and frequency-encoding directions results in aliasing of the signal outside the region of interest (ROI) into the image. Alternative approaches to higher resolution and decreased pixel size involve exciting the ROI alone. Efficient inner volume imaging and localized imaging can be approached in various ways (2- 4 ) . Surface coil imaging, which has been used for regions of interest near the body’s surface, is difficult because the NMR signal weakens considerably for nuclei that are further away from the coil, and because artifacts from the inhomogeneous B, field of the surface coil modulate the proton density (2). Inner volume imaging utilizing a homogeneous B1 field was proposed by Feinberg et al. as part of an effort to eliminate motion artifacts (3) . The pulse sequence used in that approach employed multiple spin echoes and resulted in loss of signal from nuclei having a short T2. In this com- munication we combine gradient spatially localized spectroscopy (using stimulated echoes) with phase- and frequency-encoding gradients to obtain ZOO^" (i.e., mag- nified) images in which the NMR signal comes predominantly from the volume of interest.

0740-3 194187 $3.00 384 Copyright 0 1987 by Academic Press, Inc. All rights of reproduction in any form reserved.

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METHODS AND RESULTS

Stimulated-echo localized spectroscopy, which utilizes three frequency selective pulses in the presence of three orthogonal gradients (4-6), localizes the resultant stim- ulated echo to a orthorhombic ROI. Images of the region are obtained by placing a phase-encoding gradient after the first excitation pulse and a frequency-encoding gra- dient during signal acquisition. The stimulated-echo sequence for a slice transverse to the magnetic field is shown in Fig. 1. The strengths of the phase- and frequency- encoding gradients are adjusted in order to define the image’s field of view (FOV). Phase alternation of the second and third pulses with the receiver eliminates image artifacts from additional echoes, as discussed later.

Images of the brain of an anesthetized 4-kg New Zealand white rabbit were obtained. Anesthesia was induced with an intramuscular injection of ketamine (100 mg) and xylazine (20 mg), and maintained with 1% isoflurane in oxygen. The rabbit was in- tubated, paralyzed with pancuronium ( 1 mg/h), and mechanically ventilated. The head was placed in a 90-mm i.d. home-built birdcage imaging coil tuned to 85.56 MHz, the proton frequency of our 2-T GE-CSI spectrometer imager (7). The rabbit brain was positioned in the center of the gradient field. A TI-weighted coronal image of a slice through the head, shown in Fig. 2, was obtained using a spin-echo sequence with the following parameters: 64 phase encode steps X 256 frequency encode points, number of scans (NA) = 4, TE of 40 ms, TR of 340 ms, FOV of 90 X 90 mm, slice thickness of 5 mm, and total imaging time 5.8 min. Figure 3 shows a pixel by pixel enlargement of the rabbit brain from Fig. 2.

The stimulated-echo localized imaging sequence was used to obtain the 64 X 256 image of the rabbit brain that is shown in Fig. 4. The following parameters were used to obtain this image: 5 mm slice thickness, FOV 45 X 45 mm, TE = 30 ms, TM = 40 ms, TR = 580 ms, NA = 64, and total imaging time of 40 min. (We have followed the conventional terminology as follows: TE in the spin-echo image is the time from the excitation pulse to the center of the echo. TE in the stimulated echo is the time between the first and second pulses that are indicated in Fig. 1. TM is the delay between the second and third pulses.)

DISCUSSION

Figure 4 shows the 45 X 45 mm coronal image of the rabbit brain obtained with the spatially localized imaging sequence. This image is posterior to the pituitary gland,

time +TE+TM+TE+-

FIG. 1 . Stimulated-echo localized image pulse sequence.

echo

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FIG. 2. A 64 X 256 spin-echo coronal image of the rabbit head.

FIG. 3. Software zoomed image of the rabbit brain.

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FIG. 4. A 64 X 256 localized stimulated-echo image of the rabbit brain.

and in a plane containing the posterior hypothalamus and the anterior midbrain. The third ventricle is well delineated and the CSF fluid appears as a dark outline emanating from the third ventricle. The corpus callosum is seen as a faint outline below the cerebral cortex. The internal carotid arteries are observed on either side of the midline just below the brain. Peripheral fat can be seen surrounding the sphenoid sinus and the posterior aspect of the sella turcica. The shadowing of the left hemisphere is a result of an improper alignment between the center of the gradient field and the region of space selected for the image. A comparison of this image with the spin-echo image demonstrates that the gradient-zoomed image shows better resolution of substructure. The posterior hypothalamic region, the carotid arteries, and the cavernous sinus are more clearly resolved in the localized image. A pixel by pixel enlargement of the ROI from the spin-echo image (Fig. 3) does not improve resolution of the substructure. Furthermore the pixels themselves are observed in the fourfold pixel enlargement rather than improved resolution.

It is, of course, possible to decrease the pixel size by a straightforward increase in both the number of phase-encoding steps and the number of data points acquired. However, such a straightforward approach suffers from two limitations: (1) a significant increase in imaging time and (2) the need for computer facilities capable of processing large 2DFT data sets. The minimum total imaging time (for comparable signal-to- noise) increases in direct proportion to the decrease in pixel area. For example, a four- fold reduction in pixel area could be achieved with a fourfold increase in spin-echo imaging time. For the image of the rabbit brain, this poses no substantial increase in

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imaging time; however, as the region of interest becomes significantly less than the coil diameter, the total imaging time becomes unwieldy. A more substantial limitation to very small pixel area is posed by the computer facilities of our CSI 2-T spectrometer imager which cannot process large 2D data sets (even of the size 128 X 5 12). The ease of implementing stimulated echo magnification without resorting to increased acqui- sition time and improved computing capabilities underscores its practical importance.

It is possible to estimate the minimum pixel definition achievable in a 2-T imager equipped with standard (3 G/cm) gradient coils. Theoretically the minimum pixel size has been reached when the static field inhomogeneities are comparable to the gradient times the pixel width. The observed homogeneity of our 2-T imager is 0.2 ppm/cm. With 3 G/cm gradient coils this corresponds to a minimum pixel size of 0.0 13 mm (receiver bandwidth o f f 15 kHz). (Practical limitations in our case, which arise from imaging volumes of 300-2100 cm3 and typical proton linewidths after shimming of 100 Hz, result in a minimum pixel size of 0.08 mm.) These dimensions pose no limitations for our applications and approach the dimensions of NMR mi- croscopy (8, 9).

An important concern in localized imaging is the loss in S/N as the FOV is decreased. It is well known that improvement in SIN is achieved by making a coil whose size is comparable to the dimensions of the tissue imaged (10, 11). For the type of inner volume imaging we have described this is clearly impossible. Furthermore, in order to obtain coronal and sagittal slices the coil length must be substantially larger than the slice thickness. Without modifying our coil design a minimum decrease in SIN of 4-fold is expected due to the 4-fold reduction in volume of the localized image. An additional loss of one-half in the SIN ratio is expected with stimulated-echo imaging as compared with spin-echo imaging. Neglecting further losses in signal, each phase- encoding step should require at least 64 times the number of scans as a spin-echo image, requiring an unacceptably long imaging time to obtain comparable signal-to- noise. However, sufficient SIN is observed in our localized image with a 16-fold increase in the number of scans and the total imaging time is therefore reasonable. A further decrease in the FOV may require unacceptably long imaging times and necessitate dedicated coils.

By varying the TM delay, stimulated echo imaging can be used to obtain direct T, measurements of short TI relaxation processes. However, this also decreases SIN unless the TM delay is fairly short. In the absence of rf phase cycling or gradients three additional spin echoes appear following the third excitation pulse. These additional echoes occur at time TM - TE, TM, and TM + TE and result in interference patterns in the image. Partial elimination of the interfering echoes can be obtained by appro- priately long TM delays (TM > 3TE) with a concomitant loss in S/N due to both TI and T2 relaxations. The addition of a gradient following the second excitation pulse dephases transverse magnetization without affecting the longitudinal magnetization that forms the stimulated echo. The advantage of this gradient is to allow shorter TM delays and thus improve the SIN of the stimulated echo.

Other gradient spatial localization techniques can be combined with phase- and frequency-encoding gradients to produce localized images (12-16). The utility of spa- tially localized images (albeit not magnified images) has already been demonstrated as an adjunct to localized spectroscopy ( 4 ) . An advantage of the localized, magnified,

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well-resolved image obtained with our stimulated-echo technique is that it utilizes neither subtraction techniques (as ISIS (12)) nor saturation of the nuclei outside the region of interest (as VSE (13)). Imperfect subtraction and imperfect saturation produce aliasing of the magnetization outside the region of interest into the desired zoom image just as a simple reduction of the FOV produces in a spin-echo image. Stimulated-echo localized imaging selectively excites only the magnetization in the region of interest.

The stimulated-echo localized imaging presented here can obtain inner volume images of 20 X 20 X 3000 mm in reasonable imaging times without modification of the imaging coils or an increase in the gradient strength. Furthermore, these localized images can be processed using the computational facilities of the NMR spectrometer imager and without substantial increases in acquisition time. For many applications, such as in situ fetal imaging, this may provide sufficient localization to enhance res- olution of the fetal structure. Further decreases in pixel size and enhancements in S/N will require modifications in hardware, such as those already in use in microscopic imaging applications (9).

ACKNOWLEDGMENTS

We acknowledge financial support from the NIH (Grant R23 GM34767) and helpful discussions wih John Kucharczyk on rabbit physiology.

REFERENCES

1. W. A. EDELSTEIN, J. M. S. HUTCHISON, G. JOHNSON, AND T. REDPATH, Phys. Med. Biol. 25, 751

2. M. E. MOSELEY, W. M. CHEW, M. C. NISHIMURA, T. L. RICHARDS, J. MURPHY-BOESCH, G. B. YOUNG,

3. D. A. FEINBERG, J. C. HOENNINGER, L. E. CROOKS, L. KAUFMAN, J. C. WATTS, AND M. ARAKAWA,

4. J. GRANOT, J. Mugn. Reson. 70, 488 (1986). 5. E. L. HAHN, Phys. Rev. 80, 580 (1950). 6. J. FRAHM, K. D. MERBOLDT, W. HANICKE, AND A. HAASE, J. Mugn. Reson. 64,81 (1985). 7. C. E. HAYES, W. A. EDELSTEIN, J. F. SCHENCK, 0. M. MUELLER, AND M. EASH, J. Mugn. Reson. 63,

8. L. W. HEDLUND, G. A. JOHNSON, J. P. KARIS, AND E. L. EFFMAN, J. Comput. Assist. Tomogr. 10,

9. G. A. JOHNSON, M. B. THOMPSON, S. L. GEWALT, AND C. E. HAYES, J. Mugn. Reson. 68, 129 (1986). 10. D. I. HOULT, Prog. NMR Spectrosc. 12,41 (1978). 11. F. W. WEHRLI, in “Magnetic Resonance of the Reproductive System” (S. McCarthy and F. Haseltine,

12. R. J. ORDIDGE, A. CONNELLY, AND J. A. B. LOHMAN, J. Mugn. Reson. 66,283 (1986). 13, W. P. AUE, S. MULLER, T. A. CROSS, AND J. SEELIG, J. Magn. Reson. 56, 350 (1984). 14. P. R. LUYTEN, A. J. H. MARIEN, B. SIJTSMA, AND J. A. DEN HOLLANDER, “SOC. Magn. Reson. Med.

15. J. GRANOT, “SOC. Magn. Reson. Med. Fifth Annual Meeting, Montreal,” p. 973, 1986. 16. D. M. DODDRELL, J. M. BULSING, G. J. GALLOWAY, W. M. BROOKS, J. FIELD, M. IRVING, AND H.

(1980).

T. M. MARSCHNER, L. H. PITTS, AND T. L. JAMES, Mugn. Reson. Imaging 3, 383 (1985).

Radiology 156, 743 (1985).

622 (1985).

948 (1986).

Eds.), Chap. 2, p. 23, SLACK Inc., NJ, 1987.

Fifth Annual Meeting, Montreal,” p. 997, 1986.

BADDELEY, J. Mugn. Reson. 70, 3 19 (1986).