A simple approach to T2 imaging in MRI

Mogneric Resonance Imaging, Vol. 6, pp. 64-646, Printed in the USA. All rights reserved.

1988

0730-725X/88 $3.00 + 40 Copyright 0 1988 Pergamon Press plc

l Original Contribution

A SIMPLE APPROACH TO T2 IMAGING IN MRI J.M. POPE AND N. REPIN* Dept. of Biophysics, School of Physics, The University of New South Wales, P.O. Box 1, Kensington N.S.W. Australia 2033, *Dept. of Radiology, The Royal North Shore Hospital, St. Leonards, N.S.W. Australia 2065 A simple method for obtaining images whose contrast depends only on T2 is described and tested both on phantoms and in vivo. The method works reliably and effectively under clinically realistic operating conditions using standard imaging protocols. It can result in a substantial reduction in imaging times for T2 weighted images. Keywords:

T2 imaging; MRI; Contrast; Surface coil images.

the acquisition

INTRODUCTION NMR relaxation times Tr and T2are of paramount importance in determining tissue contrast in ‘H NMR imaging.‘,* It has long been recognized that these parameters are also sensitive to differences between normal and pathological conditions.3-5 Indeed the ability of spin-echo MRI to identify lesions relies largely on such differences, notably in the spin-spin relaxation time T2. However T2 weighted images are a complex function of both spin-spin and spin-lattice relaxation times T2 and Tlas well as spin density and pulse sequence timing parameters TR and Tn.‘*6 Further the spin lattice relaxation time T,in particular is I

I

a

of additional

raw data,

thus greatly

increasing imaging times. We have adopted a less sophisticated approach to T2 imaging which involves little additional computation and negligible increase in imaging times. The method involves simply dividing out the Tland spin density contributions to image contrast, yielding images which depend substantially only on T2. The method has been tested both on phantoms and volunteers. THEORY

For both partial saturation (PS) and spin echo (SE) imaging sequences the signal intensity in a given pixel is given by (1)

N(H)e-TE’G

I

1_,~-TR/TI~-TR/Tz _Cos~(e-T~/G _e-T~/Q) 1(1) (1 - e- TR’T1)sinf3

time, Tn the echo time and 0 the flip angle. For the case TR >> T2 this can be written where TR is the repetition

a function of magnetic field strength, making comparison of images obtained on different imaging systerns problematic. An obvious way of overcoming many of these difficulties is to extract from the raw image data, maps of the individual parameters T, , T2 and spin-density N(H). However such calculated images involve substantial additional computation and generally require

I a N(H)e-TE’T2

[

1 (2)

(1 - e- TR’T1)sinO 1 _ cOsee-TR,T,

a result which applies also to small flip angle rapid imaging sequences which employ gradient ethos, pro-

RECEIVED l/13/88; ACCEPTED 4/S/88. Acknowledgment-The authors would like to acknowledge helpful discussions with other members of the MRI unit at the Royal North Shore Hospital, notably Dr. W.

Sorby and Dr. J. Roche, who also facilitated access to the imaging machine. Address correspondence to J.M. Pope, Dept. of Biophysics, School of Physics, The University of New South Wales, P.O. Box 1, Kensington N.S.W. Australia 2033. 641

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vided T2 is replaced by T;, the time for the free induction signal to de-phase in the residual gradients of the static field. For multi-echo (ME) sequences the ratio of the signal from a late echo at TE, to that from the first echo at TE < TE, is therefore given by

Table 1. Comparison of relaxation times (in milliseconds) for CuSO, solutions measured on a Bruker SXP spectrometeer operating at 60 MHz with values obtained on a GE Signa imaging system at 63.9 MHz Spectrometer cuso‘j (mM)

that is the contributions to signal intensity from spin density, flip angle and T,divide out, being common to both ethos in the same sequence. If such a division is carried out on a pixel by pixel basis then, contrast in the resulting image will depend only on T2(or T; in the case of gradient echo images). It is not strictly a calculated T2image in the conventional sense because the pixel values are not linear in T2.Compared with a simple two-point fit calculated T2 image it requires less computation and the image contrast is more closely related to that obtained in a conventional T2weighted image, without the additional complication of a functional dependence on T,and N(H). RESULTS

AND DISCUSSION

In order to evaluate the technique it was necessary first to investigate under what conditions our imaging system was capable of yielding accurate and reproducible relaxation times. A phantom was therefore constructed comprising six plastic bottles of 250 ml capacity containing water doped with CuS04 with concentrations in the range 0.1 mM to 100 mM. Samples of these same solutions in 10 mm NMR tubes were taken in order to compare relaxation times measured on the imaging system with values obtained on a conventional NMR spectrometer. The spectrometer employed was a Bruker SXP machine operating at 60 MHz which had been extensively modified and upgraded for automated measurement of relaxation times.’ Measurements of the spin-lattice relaxation time Tlwere made using either inversion recovery or saturation recovery sequences, while T2values were obtained using the Carr-Purcell spin-echo method.* In all cases at least twelve pulse delay times/echo times were employed to map out the relaxation and the measurements were repeated several times to check consistency. The results are summarized in Table 1. Also shown in Table 1 are relaxation times measured on the phantom in our GE Signa imaging system operating at 63.9 MHz (1.5 T). These measurements were all made using the headcoil with a multi-echo saturation recovery imaging sequence (ME) in multislice imaging mode with 5 mm slice thickness and four ethos at TE values of 20, 40, 60 and 80 ms. The sequence repetition time TR was stepped through five

0.1 1.0 3.0 10 30 100

TI 2501 960 439 163 55.7 15.2

Imager T2

f 24 f 16 * 5 Y!Z 2 k 0.4 f 0.2

461 263 193 115 38.1 14.0

f f f f 2 f

39 20 7 3 0.9 0.2

2344 951 404 129 44.6 12.0

TI

T2

k f f f f f

340 210 f 80 189 f 21 129 k 4 80 * 2.0 32.8 + 0.6 12.0 +

16 10 4 4 3.3 0.2

values between 125 and 2000 ms, giving five-point fits for the Tldata and four point fits for T2,a region of interest being selected from the final images in each bottle of the phantom. A comparison of T,values calculated from early and late ethos yielded no systematic differences. Likewise T2values calculated from data corresponding to different TR values agreed within experimental error. There was also no systematic dependence of relaxation times measured from center or end slices when more than one slice was imaged simultaneously, at least for slice spacings down to 1 mm. We also found no dependence of relaxation times on whether transmitter and receiver attenuation factors were set manually or using ‘auto prescan.’ Errors in the relaxation times for the imaging system in Table 1 represent the standard deviation of values obtained under different operating conditions and include two sets of measurements separated in time by 1 week. Comparison of the relaxation times from the imaging system with those obtained on the spectrometer indicate that the former are systematically shorter. This is also seen in Fig. 1, where the corresponding relaxation rates (l/T,, l/T,) are plotted against the concentration of CuS04. The solid line which approximates the spectrometer Tldata at high concentrations corresponds to the relation l/T,

= 0.658C + 0.027

where C is the CuS04 concentration in mM. In the case of Tlvalues the discrepancy is relatively small, although the higher frequency of the imager measurements indicates that these values should be slightly longer. Much greater differences were found in the T2 values, particularly those in excess of 100 ms, where both spectrometer and imager measured spinspin relaxation rates were found to deviate from a lin-

A

simple approach to Tz imaging 0

J. M. POPE AND N. REPIN

I

1

I

o-1

1-o

10

CuS04

CONCENTRATION

643

I 100

(mid)

Fig. 1. Comparison of relaxation rates for water protons in solutions of CuSO, of varying concentrations measured on an imaging machine at 63.9 MHz (square symbols) with those obtained on a spectrometer at 60 MHz (circles): l n l/T,; o q l/T,.

ear dependence on paramagnetic ion concentration at low concentrations due to effects of diffusion and convection in the bulk solutions, imperfect 180” pulses and field or phase instabilities. Similar systematic differences between imager and spectrometer measured T2 values have been reported by MacFall et a1,6 who attribute them to the effects of imperfect pulses and uncompensated eddy currents resulting from the field gradient pulses employed to provide spatial discrimination on the imaging system. Overall, however, the results confirm that the imager-measured relaxation times are reproducible to better than 10% under clinically realistic operating at conditions (multi-slice, auto-prescan, etc.) provided of course that pulse flip angles and transmitter and receiver attenuation factors are not altered during the course of the measurements. Furthermore, with the exception of the systematic deviations in T2 values for T2 > 100 ms mentioned above, the results appear surprisingly accurate. Having demonstrated the accuracy and reliability of relaxation times measured on the imaging system, we proceeded to employ the same phantom to test our

method of calculating T2 images. In order to do this we employed the standard command line image processing system (CLIPS) software supplied by GE, which allows simple arithmetic manipulation of images on a pixel by pixel basis. The results are illustrated in Fig. 2, which shows the raw images of our phantom derived from first and last ethos respectively at the top and the processed images below. The concentration of CuS04 in the bottles decreases from top left to bottom right in these images. The image at bottom left in Fig. 2 was obtained by simple division of the late echo image by the early echo one (after subtraction of an intensity offset introduced by the imaging machine and prescaling to avoid effects of integer arithmetic). Note that in both raw images, as a result of partial saturation effects, the signal intensities do not follow the order of concentrations of paramagnetic ions in the bottles, confirming that, as expected, contrast in these images is affected by Tl as well as T2 [Eq. (2)]. In the calculated image, in contrast, the intensities follow the order of T2 values. This image is however marred by noise arising from regions outside the

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Fig. 2. Images of a phantom comprising six bottles of C&O4 doped water with concentrations 100 mM (top left) 30 mM, 10 mM, 3 mM, 1 mM, 0.1 mM (bottom right). Top images are unprocessed from early and late ethos at left and right respectively. The divided images below are shown with and without noise mask.

phantom where both raw images contained negligible signal intensity. To remove this effect a ‘noise mask’ was developed which sets signal intensity in the final processed image to zero wherever there is no appreciable signal in the raw images. The final image is shown at bottom right in Fig. 2. All the processing to achieve this result, including application of the noise mask, was performed automatically, taking less than 60 seconds for a 256 x 128 image. In order to assess the potential clinical value of the method we have tested the processing on images from patients and volunteers. Figure 3 shows headscans from a normal volunteer. The images were obtained with Ta values of 4000 ms (top), 1000 ms and 500 ms

(bottom). Images from the first ethos (Tn = 20 ms) appear in the first column, from the last ethos (Tn = 80 ms) in the center and the divided and processed images in the last column. Contrast in these latter images is very similar (although signal to noise is inferior in the short TR image) confirming that it reflects differences in T2 values only. The similarity of all the processed images to the Ta = 4000 ms, TE = 80 ms image is also gratifying, confirming that the method works effectively in vivo. It appears particularly effective in identifying tissues with long T2 values such as CSF. This is significant since the power of MRI to identify the presence of lesions depends in most instances primarily on their longer T2 values

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Fig. 3. Sagittal headscans from a normal volunteer, with TR = 4000 ms (top row), 1000 ms (center) and 500 ms (bottom). Images derived from early (Tn = 20 ms) and late (Tn = 80 ms) ethos are shown in each case at left and center respectively, with the corresponding processed images in the right hand column. All 5 mm slice thickness and two excitations.

compared with surrounding normal tissue. It is worth emphasizing that the total acquisition time for these images depends almost entirely on Ta, so that an important benefit of the technique is its ability to compute an image whose contrast depends only on T2,in a fraction of the time necessary to acquire such an image directly. Since the standard technique for identifying the presence of lesions in MRI is a long TR and Tn, heavily T2weighted image, this time saving may prove of considerable importance clinically. A further benefit of the image processing in surface coil imaging is demonstrated in Fig. 4, which shows a saggital section of the cervical spine. Again the primary images are at the top and the processed images below. In this case the image at bottom right represents the sum of processed images from three slices to improve signal to noise ratio. It is notable that the grading of image intensity with distance which is characteristic of surface coil images is removed in the processed images, giving a more uniform image intensity. Also the separation of spinal cord from CSF is much

clearer in the processed images consistent with T2 contrast only. The effect of image processing on Gd-DTPA contrast-enhanced images is shown in Fig. 5, in which normal images from a patient with glioma (post radiotherapy) are shown at left and the corresponding contrast enhancement images on the right. The top two images are unprocessed images from early (Tn = 30 ms) and late (Tn = 80 ms) ethos respectively, while the processed images are at the bottom. It is notable that the effect of the contrast enhancement is removed by the image processing, confirming that the primary effect of Gd-DTPA in MRI is on Tl . Presumably the effects of paramagnetic contrast on T2are less significant and insufficient to shorten them sufficiently to remove them from the range above ~200 ms where they are largely instrumentally determined. CONCLUSIONS We have demonstrated technique for calculating

a simple image processing images whose contrast de-

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Fig. 4. Surface coil image of the cervical spine. Unprocessed images with T, = 2000 ms and TE = 30 and 80 ms are shown at top left and right respectively with the corresponding processed image at bottom left. The image at bottom right is obtained by addition of three adjacent slices (slice thickness 3 mm and separation 1 mm).

pends only on spin-spin relaxation time T, and shown that it works effectively in-vivo under clinically realistic operating conditions using standard protocols. The method can substantially reduce imaging times for T, weighted images and improves image uniformity in surface coil imaging. The method has been employed to demonstrate that contrast enhancement in the presence of Gd-DTPA derives almost entirely from changes in TI of the affected tissues. Since T, values in-vivo are found to be field independent2v5 the method may facilitate the direct comparison of images obtained on different imaging machines.

4. 5.

REFERENCES 1. Mansfield, P.; Morris, P.G. NMR cine. New York: Academic Press; 2. Bottomley, P.A.; Foster, T.H. Pfeifer, L.M. Med. P&s. 11:425; 3. Hazelwood, C.F.; Nichols, B.L.

Fig. 5. Head images of a patient with glioma images on the left were obtained without contrast while those on the right lo-20 min after administration of Gd-DTPA. Top and centre rows correspond to early and late ethos respectively while the processed images are below.

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