- Email: [email protected]
Factors influencing contrast in fast spin-echo MR imaging
Magneric Resonance Imaging. Vol. 10, pp. Printed in the USA. All rights reserved.
497-51
I,
1992
0730-725X192 $5.00 + .C@ Copyright 0 1992Pergamon Press Ltd.
l Original Contribution
FACTORS
INFLUENCING
CONTRAST
IN FAST SPIN-ECHO
MR IMAGING
R.T. CONSTABLE, A.W. ANDERSON, J. ZHONG, AND J.C. GORE Department of Diagnostic Radiology, Yale University School of Medicine, New Haven, CT 06510, USA Multi-echo pulse sequences for producing T.-weighted images in much reduced imaging times have recently been developed for romine dhricaf use. A number of recent articles have described the contrast obtained with fast spinecho (FSE) sequences and have generally indicated that they depict tissues very similarly to conventional spin-echo (SE) imaging. There are, however, some important differences in contrast between some tissues in FSE images. This work presents a detailed study of the contrast obtained with FSE imaging sequences and examines the image sequence and tissue parameters which influence contrast. The use of multiple refocusing pulses produces several subtle effects not seen in conventional SE imaging sequences, and in this study the precise nature and extent of such effects are described. The relative contributions to image contrast of magnetization transfer, the decoupling of J-modulation effects, the production of stimulated echoes and direct saturation effects, of diffusion and of the effects of the differential attenuation of different spatial frequencies, are each quantified. The mechanisms responsible for the brighter fat signal seen in FSE images, as well as the loss of signal from some other tissues, are explained. Computer simulations, phantom experiments, and clinical images are all used to support the conclusions.
Keywords: Magnetic resonance; Fast imaging; Contrast; NMR.
than in single or multi-slice conventional SE imaging. Figure 1 demonstrates these features clearly, showing single and multi-slice conventional SE and FSE images of the brain of a normal volunteer. Table 1 and Fig. 1 demonstrate that there are effects which alter contrast in FSE imaging that are either absent or of greatly reduced influence in conventional SE imaging. The investigation presented here examines the effects on
The image contrast (namely, the variations in NMR signal between different tissues) is not the same in fast spin-echo (FSE) imaging as in conventional spin echo (SE) imaging, even when the effective-TE values are the same. Although FSE images show contrast very similar to conventional SE images,‘,’ there are several subtle but important differences which indicate that different contrast mechanisms may be influential. A proper understanding of these differences may be important in optimizing the design of fast sequences. Some typical absolute signal intensities measured in single and multi-slice SE and FSE magnitude images of the brain of a normal volunteer are summarized in Table 1, where several differences in tissue signals, comparing the two imaging sequences, are apparent. (Imaging was performed on a 1.5-T GE Signa with the transmit and receive gains the same in all cases.) The primary features are that the absolute signal intensity for brain tissue is reduced (by 25-30’70) in multi-slice FSE imaging compared with conventional multi-slice SE imaging, although it is increased (by 15-20’70) in single slice versions, and that in both single and multislice FSE imaging the fat is typically much brighter
RECENED
Address
2/S/92; ACCEPTED 5/23/92. correspondence to Dr. R. Todd’Constable,
Table 1. Absolute
signal intensities Sequence
Gray matter White matter CSF Fat
SE
SE
(1 slice)
(7 slices)
FSE (1 slice)
FSE (7 slices)
324 226 603 192
311 218 581 201
377 262 623 464
240 148 568 435
imaging parameters: TE = 112msec, TR = 2500 msec, FSE: T = 14 msec, ETL = 16.
partment of Diagnostic Radiology, Yale University School of Medicine, 333 Cedar Street, New Haven, CT 06510.
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(A)
(B)
63
Fig. 1. Axial brain images of a normal volunteer. TE = 112 msec, TR = 2500 msec. (A) and (C) are single echo spin-echo images with (A) obtained in a seven slice acquisition and (C) obtained in single slice mode. (B) and (D) are FSE images (ETL = 16, 7 = 14 msec), obtained in (B) seven slice and (D) single slice acquisitions. Note that the subcutaneous fat surrounding the brain is very bright in the FSE images (B,D) but dark in conventional SE imaging (A,C), and that the brain tissues are much darker in the multi-slice FSE image (B).
tissue contrast of using multiple refocusing pulses3 in NMR imaging. FSE imaging sequences, based on the rapid acquisition with relaxation enhancement (RARE) technique first introduced by Hennig et a1.,4v5produce multiple, closely spaced echoes, each separately phase encoded to represent a different line in the spatial frequency domain or k-space, for a given slice within a single TR interval. Such multiple echo trains have different sensitivities to some effects than single echo sequences, and these are a primary source of contrast differences between FSE and conventional SE imaging. However, the precise roles of individual mechanisms have not been clearly separated before. The mechanisms for altered contrast in FSE imaging may include an altered effective spatial point spread function (PSF) for each tissue type6 which arises because different phase-encoding steps are collected at differ-
ent echo times; the effect of multiple refocusing pulses on removing the modulation due to J-Coupling7-9 in tissue; multi-slice effects including magnetization transfer from off-resonance saturation as well as direct saturation from the overlap of slice profiles;‘O-l3 the influence of stimulated echoes’4*‘5 produced by imperfect pulses; and the effects of diffusion through imaging gradients. w’* Each of these mechanisms is examined both theoretically and experimentally. Phantom and clinical images are used to support the conclusions. The different mechanisms are treated in the sections which follow with each section examining the role of a separate contrast mechanism. Two mechanisms which are not examined and which are not felt to impact significantly on image contrast are chemical exchange effects and the transient approach to equilibrium within an extended echo
Contrast in fast spin-echo MRI 0 R.T.
train. Chemical exchange, as well as diffusion between adjacent areas of different susceptibility, are not believed to be of significant influence in affecting transverse relaxation at fields of 1.5 T or below,” even though T, is sensitive in exchanging systems to the number and spacing of 180” radiofrequency (RF) pulses in a Carr-Purcell-Meiboom-Gill (CPMG) sequence.20 The approach to equilibrium affects the amount of signal recorded at the first few echoes in an echo train,21,22 which can affect the PSF but should not significantly alter tissue contrast, particularly if long effective TEs are used. An analysis of motion/flow effects is beyond the scope of this study although in the spine, cerebrospinal fluid (CSF) flow may have a significant impact on contrast. While emphasis is given to explaining the differential enhancement of the fat signal in FSE images when compared to conventional SE images, the contrast behavior of other tissues, such as gray and white matter and CSF, is also examined. The discrepancies in tissue contrast between FSE and SE imaging are explained in terms of the above effects. Each of these mechanisms impacts on all tissues, but the magnitude of each particular effect is dependent, in part, on tissue properties such as the number of resonance peaks and the coupling between them, (in the case of J-Coupling), the relaxation times, Tl and T2, of the individual tissues, (as in the PSF), and the role of cross-relaxation (as in magnetization transfer effects). The magnitude of each effect is quantified for a number of specific tissues. THE EFFECTIVE
PSF
Several reports have been published6,21-26 describing the effect of collecting different phase encode steps at different TE times or under non-steady state conditions. In FSE, k-space is separated into NJETL segments (NY = # phase encode steps) and each segment is assigned to a particular echo in the echo train. One k,-line per segment is collected in each TR. The lines for adjacent k-space segments are collected on adjacent echoes within the echo train. The effective TE is determined by the echo number (1 I echo # I ETL) at which the segment containing the 0th order phase encode line is located. The value of TE for the k, = 0 step uniquely determines the integrated signal for the reconstructed object. The sampling strategy for traversing k-space in FSE has been previously described in detail.1*6 The effective PSF is dependent upon the T2 weighting assigned to the different spatial frequencies according to the echo time at which each spatial frequency is measured. In most cases, the PSF leads to either slight edge enhancement, if the high spatial frequencies are collected at the early echoes, or slight
499
CONSTABLE ET AL.
blurring, if the high spatial frequencies are collected at late echoes. Significant effects are observed however, in proton density weighted (or short effective TE) images, for small objects with short T2, in which case the blurring can be severe and results in a loss of object peak signal intensity and contrast. In this context small means that the object spans only a few pixels in the phase encode direction since the blurring occurs only in that direction. A long thin object running perpendicular to the phase encode direction may be lost through this blurring mechanism. Alternatively, small object contrast is enhanced if the high order phaseencode steps are collected at the early echoes.6,26 As described by Constable and Gore,6 PSF effects can lead to an enhancement of the thin layer of fat surrounding the brain, but this should only enhance the signal in regions where the fat tissue is thin in the phase-encode direction. Additionally, this only occurs if the effective echo time is near the end of the echo train such that the high spatial frequency information is collected at early echoes and thus, is enhanced. In this situation, increased signal may be expected when thin sections of the subcutaneous fat tissue layer are perpendicular to the phase-encode direction, whereas in the regions where the tissue is aligned parallel to the phase-encode direction the signal should remain unchanged. In the FSE image shown in Fig. lB, the subcutaneous fat is noticeably brighter around the entire brain in comparison to the conventional SE image shown in Fig. 1A. In FSE sequences typically used for T2-weighted clinical studies with an effective TE of 80-120 msec, an inter-echo spacing (7) of 14-20 msec, and echo train lengths of 8-16 echoes, the PSF is not sufficiently altered to produce this type of edge enhancement. An edge/small object enhanced FSE image (r = 14 msec, ETL = 16, TE = 224 msec, TR = 2000 msec) is shown in Fig. 2A, where the fat is bright around the entire brain, while a blurred image (7 = 14 msec, ETL = 16, TE = 28 msec, TR = 2000 msec) is shown in Fig. 2B. The markedly different PSFs, primarily affect the sharpness of the image obtained as the extreme examples of Fig. 2 demonstrate. The uniformity of the fat signal in each image argues against PSF effects significantly altering contrast. Thus, the altered PSF produced in FSE imaging may account for slight changes in contrast for objects spanning only a few pixels in the phase-encode direction, but does not lead to the gross contrast changes seen in FSE when compared to conventional SE imaging. STIMULATED
ECHOES
Theory Stimulated echoes have also been suggested as a possible explanation for the bright fat signal seen on
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Fig. 2. Axial brain FSE images of a normal volunteer. The imaging parameters are (A) TE = 224 msec, (B) TE = 28 msec. In each image, r = 14 msec, ETL = 16, and TR = 2000 msec. Note the edge enhancement effects in (A) and the blurring due to PSF effects in (B). Edge enhancement occurs for objects which span only a few pixels in the phase encode direction. This
does not account for the bright fat observed in Figs. 1B and ID.
late echo FSE images. Stimulated echoes may comprise a significant amount of the signal present in an FSE image4 but previous studies have not made clear where these echoes alter tissue contrast. Whenever three or more RF pulses of less than 180” flip angle are applied serially, stimulated echoes will be produced. l4 Stimulated echoes arise from regions at the edges of slices where the selective RF pulses produce smaller nutations than at the center of each slice. In conventional SE imaging, stimulated echoes may cause image artifacts’4*‘5 and therefore are removed by additional dephasing gradients. In FSE imaging, rather than suppressing the stimulated echoes, the echoes are assigned the correct phase and are added to the regular spin echo.‘** In the following experiments, performed on a healthy volunteer, the contribution of stimulated echoes to the total signal was altered and contrast measurements were performed.
signals was altered by changing the amplitude of the slice selection gradient during each 180”RF pulse, thereby changing the fractional contribution of stimulated echoes to the total signal in a single slice FSE sequence. If stimulated echoes play a significant role in determining tissue contrast, then changing the contribution of the stimulated echo component in this way should significantly alter the tissue contrast. Four experiments were performed with four different amplitudes of the slice-selection gradient for the 180”RF refocusing pulses. The slice-selection gradient for the 90” RF pulse was held constant and none of the RF pulses were changed. It should be noted that changing the gradients should have no effect on regions at the center of each slice where the flip angle remains 180”) but only serves to alter the fraction of each slice that produces stimulated echoes.
Results and Discussion Experiments Stimulated echoes are produced because each refocusing pulse is significantly off-resonance towards the edge of each slice, the slice selection gradient acting to tip the effective field out of the transverse plane. In our experiments, the magnitude of the stimulated echo
In Fig. 3, FSE images are shown in which the magnitude of the slice selective gradient for the 180” RF pulses was varied. In Fig. 3A, the gradient is set to its normal value. In Figs. 3B and 3C, the gradient is reduced to l/2 and then l/3 of its normal value, respectively, thereby decreasing the contribution of stimu-
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CONSTABLE ET AZ.
Fig. 3. FSE images (TE,rr = 100 msec, TR = 2800 msec, ETL = 16, and r = 20 msec) are shown in which the contribution of stimulated echoes to the overall signal is altered by changing the magnitude of the gradient applied with the 180” refocusing RF pulses. In (A), the gradient is set to its normal value. In (B) and (C), the gradient is reduced to l/2 and then l/3 of its normal value, respectively, decreasing the contribution of stimulated echoes. The contrast is virtually identical in (A-C). In (D), the gradient is increased to twice the normal value which leads to sharp signal loss because of the much thinner slice refocused
lated echoes. It may be noted that the contrast is virtually identical in Figs. 3A-C. In Fig. 3D, the gradient is increased to twice the normal value, which leads to sharp but uniform signal loss because the slice that is refocused is now much thinner. Table 2 summarizes the intensity measurements from the images of Fig. 3 for gray and white matter, fat, and CSF. It may be concluded from these results that stimulated echoes are not responsible for the differences in contrast between FSE and SE sequences, as they have a negligible impact on contrast. This lack of change in tissue contrast confirms an earlier study2’ indicating that stimulated echoes, while contributing to the overall signal, do not strongly influence T2-weighted tissue contrast. It should be borne in mind however, that stimulated echo effects may account for the 10% increase in signal seen in FSE images over SE images in single slice
imaging. Furthermore, the slice profiles in FSE can be different from those produced in SE imaging because of the cumulative effect of multiple selective refocusing pulses, and thus absolute signal values are not ex-
Table 2. Absolute
signal intensities
Relative Z-gradient
Gray matter White matter CSF Fat Imaging parameters: ETL = 16.
strength
l/4
l/2
1
2
516 332 817 535
519 340 806 565
527 322 812 551
338 222 583 357
TE = 102
msec,
TR = 2800
msec, 7 =
17
msec,
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petted to be identical. Such slice profile changes will be reported separately. DIFFUSION
EFFECTS
We have also examined diffusion effects in FSE imaging as a possible mechanism for contrast changes. It is well understood’6-‘9 that the diffusion of water molecules through a nonuniform magnetic field can result in signal loss through intravoxel dephasing of the spins. It is also known that a train of multiple 180” RF pulses can serve to partially refocus diffusing spins if the echo spacing is short relative to the diffusion time. The loss of signal induced by the random motion of diffusing nuclei in a linear gradient led to the development of pulsed gradient spin-echo sequences (PGSE) for measuring diffusion coefficients. ” The attenuation of a spin echo in a PGSE experiment is given by, A(G)/A(O)
= exp[-y2G2h2(A
- a/3)0]
,
(1)
where A(G) and A (0) are the spin-echo amplitudes with and without a pulsed magnetic field gradient present, respectively, y is the gyromagnetic ratio, 6 is the duration of each segment of the pulsed field gradient, A is the separation between the centers of the pulsed field gradient lobes, and D is the diffusion coefficient of the material studied. ” Using Eq. (l), the effect of diffusion in both a conventional SE imaging sequence and a FSE sequence may be modelled. The FSE sequence may be modelled as a series of n separate SE imaging sequences, where 12= ETL. While the results of these calculations are dependent on the actual gradient moments in the two sequences, it is possible to derive some conclusions from the model. Consider, for example, the diffusion effects in the direction of the read gradient in the two sequences. If the initial dephasing lobe of the read gradient in the conventional SE sequence is applied following the 90” pulse, but before the 180’ pulse, then the model indicates that as the echo time increases the diffusion effects are slightly greater in conventional SE imaging than in FSE imaging. The amount by which the signal loss is greater in conventional SE imaging depends, of course, on the duration of the applied gradients and the diffusion coefficient of the material imaged. However, if the dephase lobe of the read gradient is moved to a position immediately prior to the readout gradient in the conventional SE imaging sequence, then the diffusion effects in this direction will be minimized and the FSE sequence will suffer the greater signal loss through diffusion effects. Similar results are found for the diffusion effects in the slice select and phase encode directions. Comparing the conventional SE se-
quence with the FSE sequence, diffusion effects for any tissue contribute to a differential change in signal intensity of less than 2% for that tissue, with the FSE sequence providing this slight signal gain through the refocusing of some diffusive losses by the multiple 180” refocusing pulses. That the signal gained through the refocusing of diffusion effects in FSE imaging of brain tissues is small, is reasonable, as the refocusing action of the multiple 180” pulses is partially offset by the addition of more gradients, which induce diffusion losses between the 90” pulse and the final readout gradient located at the effective TE. These “diffusion weighting” gradients are in fact the read gradients for the early phase-encoded echoes collected in a FSE imaging experiment. The relevant echoes considered here for diffusion effects are those echoes preceding the echo at which the zero phase encode step is located; thus preceding the echo which defines the effective TE. Before dismissing diffusion losses completely, it is essential to have an idea of D for fat protons. As corn oil mimics the behavior of fat in both conventional SE and FSE imaging (see below) we compared the diffusion coefficients of corn oil and water at room temperature. In diffusion experiments performed at 2.0 T, the water signal experienced a much greater decrease in signal intensity compared to oil as the gradient factor b, where b = y2G26’(A - 6/3) was increased. The diffusion coefficient for water was determined to be 2.9 x 10m3mm2/sec, well within the range of earlier published values, while the diffusion coefficient for corn oil was found to be 15.5 x 1O-5 mm2/sec; a factor more than 18 times lower than that of water. Based on these measurements and the above analysis, it may be concluded that diffusion effects play a minor role in altering contrast in FSE imaging and that these effects are not great enough to account for the bright signal observed in FSE images for both corn oil and fat. MAGNETIZATION TRANSFER AND DIRECT SATURATION EFFECTS
Theory Relaxation in protein solutions and tissues largely reflects magnetic cross-relaxation between solvent water molecules and protons at the surfaces of macromolecules. It is now well established that such crossrelaxation can be affected by selective saturation or excitation of the broadened resonances in the macromolecule, and the consequences of pre-irradiation of tissues by saturating pulses applied off-resonance from water has been exploited in magnetization transfer contrast imaging. I1912Tissues in which water interacts
Contrast in fast spin-echo
MRI
with protonated surfaces undergo magnetization transfer (MT) and show a decrease in the longitudinal magnetization, Mz, following pre-saturation off-resonance produced either by a long low amplitude pulse,“~12 or a series of short RF pu1ses,‘3,28*29applied prior to an imaging experiment. If the linewidth of the solid component in the tissue of interest is very broad, then MT may take place even when the pre-pulses are relatively far off-resonance. In conventional single echo SE imaging, multiple slices may be excited within a given TR interval. For each slice, a 90” and a 180” RF pulse are applied in the presence of a slice selective gradient. These RF pulses excite not only the on-resonance slice of interest, but adjacent slices are also irradiated at some off-resonance frequency. The amount by which the pulses are off-resonance on adjacent regions depends both on the distance from the excited slice and the strength of the slice select gradient. The precise influence of these pulses is dependent on the total power applied off-resonance the line-width and shape of the broad proton component in the tissue of interest, and the time between the application of the off-resonance pulses and the imaging of the slice of interest. Dixon et al. l3 have shown that the off-resonance excitation of adjacent slices in a multi-slice conventional SE imaging experiment can decrease tissue signal as a result of MT. In multi-slice FSE imaging, a large number of 180” RF pulses are applied at a given off-resonance frequency in rapid succession and hence it is expected that a much greater MT effect should be observed. Indeed, as the results below indicate, MT plays a significant role in determining contrast in FSE imaging, and is the dominant cause of signal loss compared to SE imaging in brain tissue. It is important to distinguish MT effects which arise from off-resonance irradiation in multi-slice FSE imaging from the direct saturation of overlapping slice profiles. Such slice interference arises when small inter-slice gaps are used and the imperfect slice profiles from adjacent slices overlap. The spins in the overlap region between two adjacent slices are then partially saturated as they experience the RF pulses for both slices. This leads to a loss of signal from all tissues in the overlap region. Direct partial saturation of the tissue water resonance can also be produced by the off-resonance RF pulses (on-resonance for the slice selected but off-resonance for all other slices) because even though they are offset in frequency, they may still cause some effect on the spins at the appropriate onresonance frequency. Thus both MT and direct saturation mechanisms can lead to signal loss and tend to mimic each other. Furthermore, each of these mechanisms is more pronounced as the slice gap is narrowed.
0 R.T.
CONSTABLEET AL.
503
Narrowing the slice gap moves the off-resonance RF pulses closer to the resonance frequency of the slice of interest and increases the slice overlap and direct saturation effects. The focus of this section is on separating these direct saturation effects from MT effects and to quantify the severity of the MT effects, as it has the potential for producing the altered contrast of FSE imaging. The effects of direct saturation and MT may be separated by observing changes in magnetization, and hence signal intensity, as a function of slice gap for materials which experience different amounts of MT. Materials that demonstrate no MT effects, such as GdDTPA doped water samples, will demonstrate signal changes only as a function of the direct saturation effect. Material such as cross-linked albumin and agarose gels have relaxation times that depend strongly on crossrelaxation, and thus will experience both mechanisms of signal loss. Separating direct saturation effects, which are identical for materials with the same value of T,/T2, from MT effects allows residual signal changes to be attributed to MT effects. Experiments designed to address these issues are described below. Figure 4 shows the results of computer calculations of the change in MZ predicted to result from direct irradiation by a train of 16, 180” RF pulses as a func-
.9. .8. .7. :
.6.
z A i!
.5.
g
.4.
a TVl’2=2 o Tlfl’2=5 n TU’l%lO D TlA’2=20
.3* .2. .l. 0 -200 1
500
1000 1500 2000 2500 3000 3500 4000 4 10
Offset (Hz) Fig. 4. Computer
simulation demonstrating the direct saturation effects followed a train of off-resonance RF pulses. The pre-pulses applied off-resonance consisted of 16 180’ RF pulses of 3.2 msec duration separated by 16.8 msec delays. Note that at high r, /T, ratios much more of a direct saturation effect is observed.
504
Magnetic Resonance Imaging 0 Volume 10, Number 4, 1992
tion of the off-resonance frequency for the set of 16 pulses. Each 180” RF pulse was assumed to have a duration of 3.2 msec (five lobes) and an interpulse delay of 16.8 msec. The longitudinal magnetization, Mz, was calculated by solving the Bloch equations, with appropriate Tr and T2 values, for the whole time evolution of the sequence and the final Mz was plotted as a function of offset frequency. T, was 1000 msec and Tz varied between 50 and 500 msec to yield T,/T, = 2, 5, 10, and 20. At small offset frequencies, there are some oscillations in the amplitude of the net magnetization, Mz, since the 180” pulses flip the spins back and forth. The figure shows that at large off-resonance frequencies, the direct saturation effect is strongest for materials with a large T,/T, ratio, but that the direct effect is ~30% for materials such as brain for pulses that are offset by 21500 Hz, which is equivalent to one slice thickness under typical imaging conditions. The results for offsets of several hundred Hz or larger and for T2s of 50 msec or more are a function of T,/T, only, instead of Tl or T2, when a steady state is reached after the pulse train. This is true for both simple square pulses and the sine pulse train we used, and was confirmed by using different Tl and T2 values but the same T,/T,. Experiments To demonstrate that both MT and direct saturation effects are present in multi-slice FSE imaging, twoslice FSE imaging was performed with slice gaps ranging from 1 to 10 mm, and the slice thickness fixed at 4 mm. Modifying the slice gap shifts the off-resonance frequency of the adjacent slice RF pulses (from 905 to 2534 Hz off-resonance corresponding to the above slice gap limits) and also alters the degree of direct saturation effects. A phantom was constructed containing a number of test tubes filled with one of the following samples; Gd-DTPA doped water, albumin (20%) or agarose (5%) gels (doped with Gd-DTPA and MnCl,). These samples provided a range of Tl /T2 ratios in addition to the aforementioned differences in MT properties. The FSE imaging parameters were; ETL = 16, r= 18 msec, TE=36msec, TR =2500msec. A second phantom composed of samples with matched T,s and T2s was also constructed to confirm that MT effects were responsible for the differential effects observed in the Gd-DTPA doped samples versus the gels and that these effects were not due to the differences in T,/T2 ratios of the samples. This second phantom was used to measure the MT and direct saturation effects as a function of slice gap (resonance offset) and as a function of the number of slices selected. These samples were used to estimate the influ-
ence of both the total RF power applied off-resonance and the time between the application of the off-resonance pulses and imaging. The number of slices was changed from one to seven with the time between slices changing from 46 to 1186 msec while maintaining a fixed TR. Since the different slices are interleaved in TR (i.e., all slices are excited within the TR of the first slice), and since the echo trains for each slice are evenly spaced in the TR interval, as the number of slices increases, the time between respective echo trains must decrease. The FSE imaging parameters were; ETL = 16, r = 15 msec, TE = 15 msec, TR = 2000 msec. In one further experiment, the number of slices was held constant at two and only the time between consecutive slices was changed. The results, described below, confirmed that MT can be significant in FSE imaging. To illustrate directly the MT effect observed in FSE imaging and validate our computer predictions, a “FAKE” sequence was developed which applies a series of 16, 180” frequency selective RF pulses, at a given off-resonance frequency, prior to collecting a single phase-encoded spin echo for imaging. Using this sequence it is possible to mimic the off-resonance effects of the RF pulses in multi-slice FSE imaging while collecting the data in a conventional SE format. The sequence differs from the FSE sequence in two ways. First, the off-resonance pulses are the same as the slice selective refocusing pulses that would be applied to adjacent slices, but are applied without a slice select gradient. In addition, the actual imaging sequence is a single echo SE sequence and hence any changes in contrast with the application of the train of pre-pulses will be due solely to off-resonance effects. The use of a single echo, SE sequence avoids any contrast changes which may arise from either PSF effects or from the use of multiple echoes that have been influenced by the application of multiple RF pulses in the CPMG echo train used in FSE sequences. This sequence should therefore produce identical contrast to the FSE sequence in Gd-DTPA doped water samples. Results and Discussion The change in signal intensity as a function of slicegap in multi-slice FSE imaging is shown in Fig. 5. The results show both a direct saturation effect, as indicated by the increase in signal intensity of the GDDTPA doped water samples (dashed lines) as the slice gap is increased, and a MT effect, as indicated by the continued increase in signal from the albumin and agarose gels (solid lines), after the Gd-DTPA signal intensities have reached maximum values as slice gap increased. The slice interference effects are negligible beyond a 100% skip as indicated by the lack of change
Contrast in fast spin-echo MRI 0 R.T.
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CONSTABLE ET AL.
_
Tl/T2=22.14(Albumin)
_
Tl!l2=23.52(Albumin)
+
Tl/T2=44.14(Agar)
-
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..-..+..
r.,
0
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8 Separation
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I.,
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16
(mm)
Fig. 5. Plot of signal intensity as a function of slice separation in a multi-slice imaging experiment for Gd-DTPA doped water and cross-linked albumin and agarose gels. The doped water solutions, which do not experience MT, show no further recovery of magnetization beyond slice gaps of approximately 4 mm, indicating that direct saturation effects are negligible beyond this point. In contrast, the signal intensity for the gels, which are known to experience MT, continues to increase as the slice gap increases beyond a 100% skip. This further increase in the net magnetization is attributed to a decrease in MT as the adjacent slices are moved further off-resonance.
of the signal intensity
of the Gd-DTPA doped water samples as the slice gap increased further. However, the singles from both the albumin and agarose gels, in
which cross relaxation is known to be significant,30 continue to increase with slice gaps beyond 4 mm, or beyond the regime of direct saturation effects. This further increase in signal for the gels provides strong evidence for the occurrence of the MT effect in multislice FSE imaging. The Gd-DTPA and gel samples with matched Z”, and T,showed similar trends to the above experiment and confirmed that MT an have a significant effect on image contrast in multi-slice FSE imaging. These samples were also used to quantitate the MT effect as a function of the number of slices excited within a TR interval and as a function of the time between slice excitations. The results, shown in Fig. 6, indicate that increasing the number of slices selected increases the significance of the MT effect. Figure 7 indicates that as the time interval between exciting adjacent slices is decreased the MT effect increases. The results show that not only do the RF pulses for the slices immediately adjacent to the slice of interest produce an MT
effect, but that the RF pulses for slices much further away also play a role. This is evident in Fig. 7 as the change in Mz for the double slice experiment is always smaller than the change observed in the seven slice experiment for the same inter-slice time delay. This is reasonable because the macromolecular linewidth in the gels extends over a frequency range much larger than the offset of each slice. To summarize, doubling the time interval between exciting adjacent slices can lead to a 10% decrease in the MT effect and hence an increase in signal intensity for those tissues experiencing MT. Furthermore, if this increased inter-slice time interval is obtained through an increase in TR then a synergistic effect results which provides a gain in SNR greater than that found when increasing TR alone. The “FARE” data confirm that the 180’ refocusing RF pulses used in multi-slice FSE imaging are capable of inducing significant MT effects. The results shown in Fig. 8 demonstrate that in a cross-linked gel, signal loss from these off-resonance pulses can be greater than 10% at typical adjacent slice offsets of 1500 Hz, and this result is for only a two-slice experiment. As the slice gap is narrowed and the off-resonance pulses
Magnetic Resonance Imaging 0 Volume 10, Number 4, 1992 1.051
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Intra-slice
6. Magnetization transfer versus T, /T, samples. The results indicate MT is takof signal with decreasing is not a r1 effect.
approach MT effects increase,
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J-COUPLING
Theory The high resolution spectra of lipids and fat contain multiple resonances with different chemical shifts and often with multiplet resonances from spin-spin couplings. Such J-couplings occur largely from homonuclear interactions between neighboring protons in hydrocarbon chains and may complicate the interpretation of relaxation time measurements, especially estimates of T2 obtainable by CPMG or similar pulse sequences. In homonuclear-coupled systems, the phase evolution of coupled spins is not reversed by a perfect 180” pulse, whereas all chemical shift and field inhomogeneity effects are reversed. The frequency differences between homonuclear coupled nuclei arising from spin-spin effects therefore causes a continual phase dispersal through out a spin echo or CPMG sequence, and the evolution of this spread in phase causes the NMR signal to be modulated in an oscillatory fashion in addition to its intrinsic T2 decay. Even in simple spectra, the echo amplitude at short TE may appear smaller than would be indicated by T2 alone because of such modulation effects, although the amplitude may actually increase again at longer times as the phase difference between coupled spins undergoes cyclic variations. For a doublet, the modulation is of
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the form cos( rnJ7); where, hertz, IZis the echo number For example, for a doublet crements in TE in a single
J = coupling constant in and 7 is the echo spacing. separated by J, small inecho SE experiment will
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Fig. 8. To confirm that MT effects can be induced by pulsed irradiation, a “FAKE” imaging sequence was developed that consists of a train of 16 180” frequency selective RF pulses followed by a 90”, 180” spin-echo imaging sequence. The FAKE data confirm that pulsed irradiation with the same equivalent off-resonance power as that used in a typical double slice FSE imaging experiment can induce MT effects.
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yield an oscillating response. The signal oscillates as the peaks are shifted from being in phase (maximum signal at TE = l/J) to each peak being shifted by 180” with respect to its coupled neighbor, (minimum signal at TE = l/25), producing complete cancellation of the signal in this example. For complex spectra, the summation of several different J-modulated decays may appear more like a monotonically decreasing signal whose apparent T, is dominated by the number and strength of the couplings present, rather than the intrinsic relaxation rates of each resonance. The measured T2 in such circumstances is dependent on the precise details of the measurement. Apparent relaxation times in complex coupled systems may depend strongly on the-number and rate of the refocusing pulses in a multiple echo sequence. Allerhand’ derived a relationship between the number of homonuclear coupled spins, their relative chemical shifts, and the echo spacing in a CPMG experiment. The increase in signal observed in FSE imaging for a coupled system as the echo spacing is decreased and the number of echoes increased is properly described using quantum mechanics. It can, however, be understood by the following classical argument. As the pulse rate becomes large relative to the chemical shifts, phase deviations due to the shifts approach zero, and hence can be neglected. Although the 180” pulses do not refocus the J-coupling evolution, under these conditions the cumulative motion of each spin is a precession around the field due to the spins to which it is coupled. The resulting motion is complicated, but preserves the net spin vector, and hence the signal, at the echo peaks. This result may also be understood in terms of the well-known theorem that J-coupling between equivalent spins is unobservable. 32 Rapid pulsing in a CPMG sequence removes the effects of chemical shift, rendering all spins “equivalent,” and results in a signal unmodulated by the coupling. Experiments and Results All imaging to study J-coupling effects was performed on single slices to avoid direct saturation and MT effects. We chose to study the behavior of corn oil as well as fat because they behave similarly when imaged using FSE and SE sequences. Thus many of the experiments described below compare corn oil and water signals. Conventional SE imaging was performed (256 x 128 x 2 nex, TR = 2000 msec) with small increments in TE to examine if modulation of the T2 decay curve from J-coupling effects could be observed using a phantom containing core oil and Gd-DTPA doped water solutions. The TE was varied from 15 to 100 msec in 5-msec increments. It was found that small increments in TE in a conventional SE sequence do not
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demonstrate modulation of signal in either sample. The lack of modulation implies that a number of different couplings may be present in corn oil and that while some coupled spins are moving out of phase, others may be moving into phase with the net result an essentially monotonic decay. The decay rate largely reflects the number and separation of component frequencies rather than the intrinsic linewidth of any one resonance. Multiple J-coupled peaks are indeed present in corn oil as confirmed by high resolution NMR spectra (not shown). Similar results have been obtained in fat. A J-coupled system should demonstrate an increase in signal as the CPMG echo train effectively decouples the spin system. A series of imaging experiments were performed using the FSE sequence with a range of values of echo spacing 7, and ETL. These experiments were designed to demonstrate the decoupling effect of the CPMG sequence as r is shortened and ETL increased. Signal intensities obtained using the FSE sequence and the conventional SE sequence were compared using corn oil and water samples. The FSE imaging was carried out using a constant effective TE but with different ETLs and 7 values (ETL = 8, r = 15 msec; ETL = 4, 7 = 30 msec; ETL = 2, r = 60 msec; with TE = 120 msec in each case). The use of variable echo spacing (7) and ETL with fixed effective TE, changes the degree of decoupling and hence signal intensity in J-coupled systems. Comparing the FSE sequence with the conventional signal slice, single echo SE imaging sequence with these changes in ETL and r, demonstrates the effects of the train of 180” RF pulses. Figure 9 shows the preferential enhancement of signal from a corn oil phantom compared with that from water as the echo train length in FSE imaging is increased. It also shows that the signal intensity for corn oil is much brighter in images obtained using the FSE sequence than in those obtained using the conventional SE sequence. This preferential signal enhancement is a characteristic feature of the FSE imaging sequences. Since the bright fat signal is essentially a consequence of a decoupling phenomenon, it is useful to consider alternate refocusing schemes that produce spin echoes, but which are less effective at decoupling. Shaka and Keeler33 differentiate the relative decoupling performances of the two sequences, [7/2 - 180”(x) - T - 180”(x) - 7/2]” and [7/2 - 180” (x) - 7 - 180” (-x)
- 7/2]“,
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of the brain demonstrate similar contrast in all of the images, but notice that the fat signal increases with increasing ETL. This is a direct result of the removal of J-modulation effects for the coupled spins in fat as the number of 180” RF pulses increases and the spacing between the pulses decreases, as described above. Measurements were also made to quantify the effect of pulse rate on echo intensity from J-coupled nuclei using a spectrometer operating at 20 MHz and the CPMG sequence. We studied the echo intensity at a fixed TE (100 msec) for variable numbers and pulse rates of 180” pulses, for water and corn oil as well as hexane and 1-decanol. Hexane contains no polar group or exchangeable protons. The amplitude of the echo occurring at 100 msec was measured and the number of echoes preceding that echo was varied from 1 to 50, with 7 ranging from 50 to 2 msec and ETL from 2 to 50, respectively, Figure 11 shows the results of the spectrometer measurements for water doped with MnC& (same T2 as corn oil), corn oil, hexane, and l-decanol. The increase in signal intensity at short pulse intervals occurs at intervals below 20 msec, consistent with the chemical shift range of 2.5 ppm for coupled nuclei. This range corresponds to the range of shift visible in the high resolution spectrum. The increase in signal intensity, shown in the above examples for coupled systems, as the ETL is increased and the echo spacing is decreased, agrees with the results predicted by Allerhand for J-coupled systems. CONCLUSIONS
wherein the latter sequence uses refocusing pulses with
alternating phases. These sequences behave quite differently when employed as potential decoupling schemes for heterogeneous decoupling. In the limit as r + 0, the constant phase sequence becomes a conventional decoupling scheme, whereas the sequence with alternating phase does not decouple at all. By analogy therefore, it may be predicted that alternating the phases of the refocusing pulses in FSE should be less effective at removing J-modulation effects and will cause fat to be less bright. Evidence for this is provided experimentally in Fig. 9, where the results for the alternating phase sequence fall between that of the conventional FSE and SE imaging sequences. Figure 10 shows a conventional SE image of a normal brain a) with ETL = 1, and FSE images, b) ETL = 4, r = 28 msec, and c) ETL = 16, r = 14 msec. In each image, the TE = 112 msec and TR = 2000 msec. Note that the imaging time increase is inversely proportional to the echo train length. The soft tissues
The origins of the contrast differences between FSE and conventional SE images have been accounted and quantified. PSF effects are important only for long ETL and r values and short T2 materials. In most cases when using an ETL I 8, the PSF does not markedly alter the contrast of Tz-weighted images. Stimulated echoes, while contributing to the overall signal in an FSE experiment, do not significantly alter the T2 contrast. Gains of approximately 10% in signal intensity for soft tissues may be realized in single slice FSE imaging, from stimulated echo effects, but these are not realized in multi-slice FSE as the gain is offset by MT losses. The FSE sequence can regain some of the signal lost through diffusion processes (<2%), but this gain in signal intensity for soft tissues such as brain tissues, is also outweighed by the MT losses which occur in multi-slice FSE. Magnetization transfer does take place in multi-slice FSE imaging and this serves to darken preferentially the gray and white matter in the brain, but does not affect water (CSF) or fat significantly. In body imaging we expect MT to be important
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Fig. 10. Single slice SE image (A) ETL = 1, and FSE images (B) ETL = 4, T = 28 msec, and (C) ETL = 16, r = 14 msec. For all images TE = 112 msec and TR = 2000 msec. Note the soft tissues of the brain have similar signal intensity in each image while the subcutaneous fat layer surrounding the brain increases in intensity as the ETL increases and the echo spacing, 7, decreases.
in muscle, liver, and kidney. The MT observed, increases with the number of slices excited within a TR interval and as the time interval between the excitation of these slices decreases. With small slice gaps, direct saturation effects can decrease tissue signal through imperfect excitation profiles overlapping adjacent slices, just as in conventional SE imaging, and these affects may be more pronounced in FSE because of the application of a train of off-resonance RF pulses.
It should be noted that when FSE imaging is performed in interleaved mode, collecting the odd slices first and then collecting the even slices, both the direct saturation and MT effects are reduced but not eliminated. The attenuation of signals in J-coupled systems may be reduced with the use of a CPMG echo train, and it is the reduction of this dephasing in FSE imaging which is partly responsible for the bright fat signal observed. The fat also appears relatively brighter in
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Fig. 11. Signal intensity at TE = 100 msec for CPMG sequences with different 7 and echo numbers at 20 MHz. The signal, for the corn oil, 1-decanol, and hexane, increases with decreasing echo spacing and increasing ETL while the doped water signal remains constant. The I-decanol and hexane samples are included to rule out exchange effects.
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Acknowledgments- We gratefully acknowledge the support of grant CA40675 awarded by the National Cancer Institute and the support of G.E. Medical Research Systems, Milwaukee, WI. A.W.A. acknowledges support from an NIH Training Grant NC1 CA-09549. Drs. R. Mark Henkelman, Richard Kennan, Gordon Sze, and Robert Smith are also thanked for useful discussions. REFERENCES 1.
2.
3.
4.
5.
6.
Melki, P.S.; Mulkern, R.V.; Panych, L.P.; Jolesz, EA. Comparing the FAISE method with conventional dualecho sequences. J. Magn. Resort. Imag. 1:319-326; 1991. Constable, R.T.; Smith, R.; Gore, J.C. Signal to noise and contrast in fast spin echo (FSE) imaging and inversion-recovery FSE. J. Comput. Assist. Tomog. 16(l): 41-47; 1992. Meiboom, S.; Gill, D. Modified spin-echo method for measuring nuclear relaxation times. Rev. Sci. Z&rum. 29(8): 688-691; 1958. Hennig, J.; Naureth, A.; Friedburg, H. RARE imaging: A fast imaging method for clinical MR. Magn. Reson. Med. 3:823-833; 1986. Hennig, J.; Friedburg, H. Clinical applications and methodological developments of the RARE technique. Magn. Reson. Imaging 6:391-395; 1988. Constable, R.T.; Gore, J.C. The loss of small object contrast in variable TE imaging: Implications for FSE, RARE and EPI. Magn. Reson. Med. (in press).
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Bovey, EA. Nuclear Magnetic Resonance Spectroscopy. San Diego, CA: Academic Press; 1988. 8. Carver, J.P.; Richards, R.E. A general two-site solution for the chemical exchange produced dependence of Tz upon the Carr-Purcell pulse separation. J. Magn. Reson. 6:89-105; 1972. 9. Allerhand, A. Analysis of Carr-Purcell spin-echo NMR experiments on multiple-spin systems. I. The effect of homonuclear coupling. J. Chem. Phys. 44(1):1-9; 1966. 10a. Yeung, H.N. Tissue relaxation under periodic pulsed saturation. San Francisco: Proceedings of the 10th Annual SMRM; 1991: abstract #174. lob. Yeung, H.N.; Aisen, A.M. Magnetization transfer contrast with periodic pulsed saturation. Radiology 183:209-214; 1992. 11. Eng, J.; Ceckler, T.L.; Balaban, R. Quantitative ‘H magnetization transfer in vivo. Magn. Reson. Med. 17: 304-314; 1991. 12. Wolff, SD.; Balaban, R.S. Magnetization transfer contrast (MTC) and tissue water proton relaxation in vivo. Magn. Reson. Med. 10:135-144; 1989. 13. Dixon, W.T.; Engels, H.; Castillo, M.; Sardashti, M. Incidental magnetization transfer contrast in standard multislice imaging. Magn. Reson. Imaging 8:417-422; 1990. R.M. A stimulated echo 14. Crawley, A.P.; Henkelman, artifact from slice interference in magnetic resonance imaging. Med. Phys. 14(5):842-848; 1987. S.; Orphanoudakis, S.C.; Gmitro, A.; 15. Majumdar, d’Donne1, M.; Gore, J.C. Errors in the measurements of T2 using multiple-echo MRI techniques, I. Effects of radiofrequency pulse imperfections. Magn. Reson. Med. 3:397-417; 1986. 16. zarr, H.Y.; Purcell, E.M. Effects of diffusion on free Irecision in nuclear magnetic resonance experiments. Dhys. Rev., 94(3):630-638; 1954. 17. Itejskal, E.O.; Tanner, J.E. Spin diffusion measurenents: Spin echoes in the presence of a time dependent ‘ield gradient. J. Chem. Phys. 42(1):288-292; 1965. 18.
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23. Holsinger, A.E.; Riederer, S.J. The importance of phase-encoding order in ultra-short TR snapshot MR imaging. Magn. Reson. Med. 16:481-488; 1990. 24. Chien, D.; Atkinson, D.J.; Edelman, R.R. Strategies to improve contrast in TurboFlash imaging: Reordered phase encoding and k-Space segmentation. J. Magn. Reson. Imag. 1:63-70; 1991. 25. Oshio, K.; Singh, M. A computer simulation of T2 decay effects in echo planar imaging. Magn. Reson. Med. 11:389-397; 1989. 26. Constable, T.R.; Gore, J.C. The loss of small object contrast in fast spin echo imaging. San Francisco: Proceedings of the 10th Annual SMRM; 1991: abstract # 1238. 27. Bishop, J.E.; Plewes, D.B. Stimulated echo effects in fast spin echo images. San Francisco: Proceedings of the 10th Annual SMRM; 1991: abstract # 1025. 28. Hu, B.; Nishimura, D.; Macovski, A. Pulsed magneti-
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32. 33.
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zation transfer contrast. New York: Proceedings of the 9th Annual SMRM; 1990: abstract # 352. Melki, P.S.; Mulkern, R.V. Magnetization transfer effects in multi-slice RARE sequences. Magn. Reson. Med. 24:189-195; 1992. Mendelson, D.A.; Heinsbergen, J.F.; Kennedy, SD.; Szczepaniak, L.S.; Lester, C.C.; Bryant, R.G. Comparison of agarose and cross-linked protein gels as magnetic resonance imaging phantoms. Magn. Reson. Imaging 9: 975-978; 1991. Freeman, R.; Hill, H.D.W. Phase and intensity anomalies in Fourier transform NMR. J. Magn. Reson. 4: 366-383; 1971. Abragam, A. Principles of Nuclear Magnetism. New York: Oxford University Press; 1985: p. 481. Shaka, A.; Keeler, J. Broadband spin decoupling in isotropic liquids. Progr. NW/. Magn. Reson. Spect. 19:47129; 1987.
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l Original Contribution
FACTORS
INFLUENCING
CONTRAST
IN FAST SPIN-ECHO
MR IMAGING
R.T. CONSTABLE, A.W. ANDERSON, J. ZHONG, AND J.C. GORE Department of Diagnostic Radiology, Yale University School of Medicine, New Haven, CT 06510, USA Multi-echo pulse sequences for producing T.-weighted images in much reduced imaging times have recently been developed for romine dhricaf use. A number of recent articles have described the contrast obtained with fast spinecho (FSE) sequences and have generally indicated that they depict tissues very similarly to conventional spin-echo (SE) imaging. There are, however, some important differences in contrast between some tissues in FSE images. This work presents a detailed study of the contrast obtained with FSE imaging sequences and examines the image sequence and tissue parameters which influence contrast. The use of multiple refocusing pulses produces several subtle effects not seen in conventional SE imaging sequences, and in this study the precise nature and extent of such effects are described. The relative contributions to image contrast of magnetization transfer, the decoupling of J-modulation effects, the production of stimulated echoes and direct saturation effects, of diffusion and of the effects of the differential attenuation of different spatial frequencies, are each quantified. The mechanisms responsible for the brighter fat signal seen in FSE images, as well as the loss of signal from some other tissues, are explained. Computer simulations, phantom experiments, and clinical images are all used to support the conclusions.
Keywords: Magnetic resonance; Fast imaging; Contrast; NMR.
than in single or multi-slice conventional SE imaging. Figure 1 demonstrates these features clearly, showing single and multi-slice conventional SE and FSE images of the brain of a normal volunteer. Table 1 and Fig. 1 demonstrate that there are effects which alter contrast in FSE imaging that are either absent or of greatly reduced influence in conventional SE imaging. The investigation presented here examines the effects on
The image contrast (namely, the variations in NMR signal between different tissues) is not the same in fast spin-echo (FSE) imaging as in conventional spin echo (SE) imaging, even when the effective-TE values are the same. Although FSE images show contrast very similar to conventional SE images,‘,’ there are several subtle but important differences which indicate that different contrast mechanisms may be influential. A proper understanding of these differences may be important in optimizing the design of fast sequences. Some typical absolute signal intensities measured in single and multi-slice SE and FSE magnitude images of the brain of a normal volunteer are summarized in Table 1, where several differences in tissue signals, comparing the two imaging sequences, are apparent. (Imaging was performed on a 1.5-T GE Signa with the transmit and receive gains the same in all cases.) The primary features are that the absolute signal intensity for brain tissue is reduced (by 25-30’70) in multi-slice FSE imaging compared with conventional multi-slice SE imaging, although it is increased (by 15-20’70) in single slice versions, and that in both single and multislice FSE imaging the fat is typically much brighter
RECENED
Address
2/S/92; ACCEPTED 5/23/92. correspondence to Dr. R. Todd’Constable,
Table 1. Absolute
signal intensities Sequence
Gray matter White matter CSF Fat
SE
SE
(1 slice)
(7 slices)
FSE (1 slice)
FSE (7 slices)
324 226 603 192
311 218 581 201
377 262 623 464
240 148 568 435
imaging parameters: TE = 112msec, TR = 2500 msec, FSE: T = 14 msec, ETL = 16.
partment of Diagnostic Radiology, Yale University School of Medicine, 333 Cedar Street, New Haven, CT 06510.
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Fig. 1. Axial brain images of a normal volunteer. TE = 112 msec, TR = 2500 msec. (A) and (C) are single echo spin-echo images with (A) obtained in a seven slice acquisition and (C) obtained in single slice mode. (B) and (D) are FSE images (ETL = 16, 7 = 14 msec), obtained in (B) seven slice and (D) single slice acquisitions. Note that the subcutaneous fat surrounding the brain is very bright in the FSE images (B,D) but dark in conventional SE imaging (A,C), and that the brain tissues are much darker in the multi-slice FSE image (B).
tissue contrast of using multiple refocusing pulses3 in NMR imaging. FSE imaging sequences, based on the rapid acquisition with relaxation enhancement (RARE) technique first introduced by Hennig et a1.,4v5produce multiple, closely spaced echoes, each separately phase encoded to represent a different line in the spatial frequency domain or k-space, for a given slice within a single TR interval. Such multiple echo trains have different sensitivities to some effects than single echo sequences, and these are a primary source of contrast differences between FSE and conventional SE imaging. However, the precise roles of individual mechanisms have not been clearly separated before. The mechanisms for altered contrast in FSE imaging may include an altered effective spatial point spread function (PSF) for each tissue type6 which arises because different phase-encoding steps are collected at differ-
ent echo times; the effect of multiple refocusing pulses on removing the modulation due to J-Coupling7-9 in tissue; multi-slice effects including magnetization transfer from off-resonance saturation as well as direct saturation from the overlap of slice profiles;‘O-l3 the influence of stimulated echoes’4*‘5 produced by imperfect pulses; and the effects of diffusion through imaging gradients. w’* Each of these mechanisms is examined both theoretically and experimentally. Phantom and clinical images are used to support the conclusions. The different mechanisms are treated in the sections which follow with each section examining the role of a separate contrast mechanism. Two mechanisms which are not examined and which are not felt to impact significantly on image contrast are chemical exchange effects and the transient approach to equilibrium within an extended echo
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train. Chemical exchange, as well as diffusion between adjacent areas of different susceptibility, are not believed to be of significant influence in affecting transverse relaxation at fields of 1.5 T or below,” even though T, is sensitive in exchanging systems to the number and spacing of 180” radiofrequency (RF) pulses in a Carr-Purcell-Meiboom-Gill (CPMG) sequence.20 The approach to equilibrium affects the amount of signal recorded at the first few echoes in an echo train,21,22 which can affect the PSF but should not significantly alter tissue contrast, particularly if long effective TEs are used. An analysis of motion/flow effects is beyond the scope of this study although in the spine, cerebrospinal fluid (CSF) flow may have a significant impact on contrast. While emphasis is given to explaining the differential enhancement of the fat signal in FSE images when compared to conventional SE images, the contrast behavior of other tissues, such as gray and white matter and CSF, is also examined. The discrepancies in tissue contrast between FSE and SE imaging are explained in terms of the above effects. Each of these mechanisms impacts on all tissues, but the magnitude of each particular effect is dependent, in part, on tissue properties such as the number of resonance peaks and the coupling between them, (in the case of J-Coupling), the relaxation times, Tl and T2, of the individual tissues, (as in the PSF), and the role of cross-relaxation (as in magnetization transfer effects). The magnitude of each effect is quantified for a number of specific tissues. THE EFFECTIVE
PSF
Several reports have been published6,21-26 describing the effect of collecting different phase encode steps at different TE times or under non-steady state conditions. In FSE, k-space is separated into NJETL segments (NY = # phase encode steps) and each segment is assigned to a particular echo in the echo train. One k,-line per segment is collected in each TR. The lines for adjacent k-space segments are collected on adjacent echoes within the echo train. The effective TE is determined by the echo number (1 I echo # I ETL) at which the segment containing the 0th order phase encode line is located. The value of TE for the k, = 0 step uniquely determines the integrated signal for the reconstructed object. The sampling strategy for traversing k-space in FSE has been previously described in detail.1*6 The effective PSF is dependent upon the T2 weighting assigned to the different spatial frequencies according to the echo time at which each spatial frequency is measured. In most cases, the PSF leads to either slight edge enhancement, if the high spatial frequencies are collected at the early echoes, or slight
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blurring, if the high spatial frequencies are collected at late echoes. Significant effects are observed however, in proton density weighted (or short effective TE) images, for small objects with short T2, in which case the blurring can be severe and results in a loss of object peak signal intensity and contrast. In this context small means that the object spans only a few pixels in the phase encode direction since the blurring occurs only in that direction. A long thin object running perpendicular to the phase encode direction may be lost through this blurring mechanism. Alternatively, small object contrast is enhanced if the high order phaseencode steps are collected at the early echoes.6,26 As described by Constable and Gore,6 PSF effects can lead to an enhancement of the thin layer of fat surrounding the brain, but this should only enhance the signal in regions where the fat tissue is thin in the phase-encode direction. Additionally, this only occurs if the effective echo time is near the end of the echo train such that the high spatial frequency information is collected at early echoes and thus, is enhanced. In this situation, increased signal may be expected when thin sections of the subcutaneous fat tissue layer are perpendicular to the phase-encode direction, whereas in the regions where the tissue is aligned parallel to the phase-encode direction the signal should remain unchanged. In the FSE image shown in Fig. lB, the subcutaneous fat is noticeably brighter around the entire brain in comparison to the conventional SE image shown in Fig. 1A. In FSE sequences typically used for T2-weighted clinical studies with an effective TE of 80-120 msec, an inter-echo spacing (7) of 14-20 msec, and echo train lengths of 8-16 echoes, the PSF is not sufficiently altered to produce this type of edge enhancement. An edge/small object enhanced FSE image (r = 14 msec, ETL = 16, TE = 224 msec, TR = 2000 msec) is shown in Fig. 2A, where the fat is bright around the entire brain, while a blurred image (7 = 14 msec, ETL = 16, TE = 28 msec, TR = 2000 msec) is shown in Fig. 2B. The markedly different PSFs, primarily affect the sharpness of the image obtained as the extreme examples of Fig. 2 demonstrate. The uniformity of the fat signal in each image argues against PSF effects significantly altering contrast. Thus, the altered PSF produced in FSE imaging may account for slight changes in contrast for objects spanning only a few pixels in the phase-encode direction, but does not lead to the gross contrast changes seen in FSE when compared to conventional SE imaging. STIMULATED
ECHOES
Theory Stimulated echoes have also been suggested as a possible explanation for the bright fat signal seen on
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Fig. 2. Axial brain FSE images of a normal volunteer. The imaging parameters are (A) TE = 224 msec, (B) TE = 28 msec. In each image, r = 14 msec, ETL = 16, and TR = 2000 msec. Note the edge enhancement effects in (A) and the blurring due to PSF effects in (B). Edge enhancement occurs for objects which span only a few pixels in the phase encode direction. This
does not account for the bright fat observed in Figs. 1B and ID.
late echo FSE images. Stimulated echoes may comprise a significant amount of the signal present in an FSE image4 but previous studies have not made clear where these echoes alter tissue contrast. Whenever three or more RF pulses of less than 180” flip angle are applied serially, stimulated echoes will be produced. l4 Stimulated echoes arise from regions at the edges of slices where the selective RF pulses produce smaller nutations than at the center of each slice. In conventional SE imaging, stimulated echoes may cause image artifacts’4*‘5 and therefore are removed by additional dephasing gradients. In FSE imaging, rather than suppressing the stimulated echoes, the echoes are assigned the correct phase and are added to the regular spin echo.‘** In the following experiments, performed on a healthy volunteer, the contribution of stimulated echoes to the total signal was altered and contrast measurements were performed.
signals was altered by changing the amplitude of the slice selection gradient during each 180”RF pulse, thereby changing the fractional contribution of stimulated echoes to the total signal in a single slice FSE sequence. If stimulated echoes play a significant role in determining tissue contrast, then changing the contribution of the stimulated echo component in this way should significantly alter the tissue contrast. Four experiments were performed with four different amplitudes of the slice-selection gradient for the 180”RF refocusing pulses. The slice-selection gradient for the 90” RF pulse was held constant and none of the RF pulses were changed. It should be noted that changing the gradients should have no effect on regions at the center of each slice where the flip angle remains 180”) but only serves to alter the fraction of each slice that produces stimulated echoes.
Results and Discussion Experiments Stimulated echoes are produced because each refocusing pulse is significantly off-resonance towards the edge of each slice, the slice selection gradient acting to tip the effective field out of the transverse plane. In our experiments, the magnitude of the stimulated echo
In Fig. 3, FSE images are shown in which the magnitude of the slice selective gradient for the 180” RF pulses was varied. In Fig. 3A, the gradient is set to its normal value. In Figs. 3B and 3C, the gradient is reduced to l/2 and then l/3 of its normal value, respectively, thereby decreasing the contribution of stimu-
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Fig. 3. FSE images (TE,rr = 100 msec, TR = 2800 msec, ETL = 16, and r = 20 msec) are shown in which the contribution of stimulated echoes to the overall signal is altered by changing the magnitude of the gradient applied with the 180” refocusing RF pulses. In (A), the gradient is set to its normal value. In (B) and (C), the gradient is reduced to l/2 and then l/3 of its normal value, respectively, decreasing the contribution of stimulated echoes. The contrast is virtually identical in (A-C). In (D), the gradient is increased to twice the normal value which leads to sharp signal loss because of the much thinner slice refocused
lated echoes. It may be noted that the contrast is virtually identical in Figs. 3A-C. In Fig. 3D, the gradient is increased to twice the normal value, which leads to sharp but uniform signal loss because the slice that is refocused is now much thinner. Table 2 summarizes the intensity measurements from the images of Fig. 3 for gray and white matter, fat, and CSF. It may be concluded from these results that stimulated echoes are not responsible for the differences in contrast between FSE and SE sequences, as they have a negligible impact on contrast. This lack of change in tissue contrast confirms an earlier study2’ indicating that stimulated echoes, while contributing to the overall signal, do not strongly influence T2-weighted tissue contrast. It should be borne in mind however, that stimulated echo effects may account for the 10% increase in signal seen in FSE images over SE images in single slice
imaging. Furthermore, the slice profiles in FSE can be different from those produced in SE imaging because of the cumulative effect of multiple selective refocusing pulses, and thus absolute signal values are not ex-
Table 2. Absolute
signal intensities
Relative Z-gradient
Gray matter White matter CSF Fat Imaging parameters: ETL = 16.
strength
l/4
l/2
1
2
516 332 817 535
519 340 806 565
527 322 812 551
338 222 583 357
TE = 102
msec,
TR = 2800
msec, 7 =
17
msec,
MagneticResonanceImaging0 Volume10, Number4, 1992
502
petted to be identical. Such slice profile changes will be reported separately. DIFFUSION
EFFECTS
We have also examined diffusion effects in FSE imaging as a possible mechanism for contrast changes. It is well understood’6-‘9 that the diffusion of water molecules through a nonuniform magnetic field can result in signal loss through intravoxel dephasing of the spins. It is also known that a train of multiple 180” RF pulses can serve to partially refocus diffusing spins if the echo spacing is short relative to the diffusion time. The loss of signal induced by the random motion of diffusing nuclei in a linear gradient led to the development of pulsed gradient spin-echo sequences (PGSE) for measuring diffusion coefficients. ” The attenuation of a spin echo in a PGSE experiment is given by, A(G)/A(O)
= exp[-y2G2h2(A
- a/3)0]
,
(1)
where A(G) and A (0) are the spin-echo amplitudes with and without a pulsed magnetic field gradient present, respectively, y is the gyromagnetic ratio, 6 is the duration of each segment of the pulsed field gradient, A is the separation between the centers of the pulsed field gradient lobes, and D is the diffusion coefficient of the material studied. ” Using Eq. (l), the effect of diffusion in both a conventional SE imaging sequence and a FSE sequence may be modelled. The FSE sequence may be modelled as a series of n separate SE imaging sequences, where 12= ETL. While the results of these calculations are dependent on the actual gradient moments in the two sequences, it is possible to derive some conclusions from the model. Consider, for example, the diffusion effects in the direction of the read gradient in the two sequences. If the initial dephasing lobe of the read gradient in the conventional SE sequence is applied following the 90” pulse, but before the 180’ pulse, then the model indicates that as the echo time increases the diffusion effects are slightly greater in conventional SE imaging than in FSE imaging. The amount by which the signal loss is greater in conventional SE imaging depends, of course, on the duration of the applied gradients and the diffusion coefficient of the material imaged. However, if the dephase lobe of the read gradient is moved to a position immediately prior to the readout gradient in the conventional SE imaging sequence, then the diffusion effects in this direction will be minimized and the FSE sequence will suffer the greater signal loss through diffusion effects. Similar results are found for the diffusion effects in the slice select and phase encode directions. Comparing the conventional SE se-
quence with the FSE sequence, diffusion effects for any tissue contribute to a differential change in signal intensity of less than 2% for that tissue, with the FSE sequence providing this slight signal gain through the refocusing of some diffusive losses by the multiple 180” refocusing pulses. That the signal gained through the refocusing of diffusion effects in FSE imaging of brain tissues is small, is reasonable, as the refocusing action of the multiple 180” pulses is partially offset by the addition of more gradients, which induce diffusion losses between the 90” pulse and the final readout gradient located at the effective TE. These “diffusion weighting” gradients are in fact the read gradients for the early phase-encoded echoes collected in a FSE imaging experiment. The relevant echoes considered here for diffusion effects are those echoes preceding the echo at which the zero phase encode step is located; thus preceding the echo which defines the effective TE. Before dismissing diffusion losses completely, it is essential to have an idea of D for fat protons. As corn oil mimics the behavior of fat in both conventional SE and FSE imaging (see below) we compared the diffusion coefficients of corn oil and water at room temperature. In diffusion experiments performed at 2.0 T, the water signal experienced a much greater decrease in signal intensity compared to oil as the gradient factor b, where b = y2G26’(A - 6/3) was increased. The diffusion coefficient for water was determined to be 2.9 x 10m3mm2/sec, well within the range of earlier published values, while the diffusion coefficient for corn oil was found to be 15.5 x 1O-5 mm2/sec; a factor more than 18 times lower than that of water. Based on these measurements and the above analysis, it may be concluded that diffusion effects play a minor role in altering contrast in FSE imaging and that these effects are not great enough to account for the bright signal observed in FSE images for both corn oil and fat. MAGNETIZATION TRANSFER AND DIRECT SATURATION EFFECTS
Theory Relaxation in protein solutions and tissues largely reflects magnetic cross-relaxation between solvent water molecules and protons at the surfaces of macromolecules. It is now well established that such crossrelaxation can be affected by selective saturation or excitation of the broadened resonances in the macromolecule, and the consequences of pre-irradiation of tissues by saturating pulses applied off-resonance from water has been exploited in magnetization transfer contrast imaging. I1912Tissues in which water interacts
Contrast in fast spin-echo
MRI
with protonated surfaces undergo magnetization transfer (MT) and show a decrease in the longitudinal magnetization, Mz, following pre-saturation off-resonance produced either by a long low amplitude pulse,“~12 or a series of short RF pu1ses,‘3,28*29applied prior to an imaging experiment. If the linewidth of the solid component in the tissue of interest is very broad, then MT may take place even when the pre-pulses are relatively far off-resonance. In conventional single echo SE imaging, multiple slices may be excited within a given TR interval. For each slice, a 90” and a 180” RF pulse are applied in the presence of a slice selective gradient. These RF pulses excite not only the on-resonance slice of interest, but adjacent slices are also irradiated at some off-resonance frequency. The amount by which the pulses are off-resonance on adjacent regions depends both on the distance from the excited slice and the strength of the slice select gradient. The precise influence of these pulses is dependent on the total power applied off-resonance the line-width and shape of the broad proton component in the tissue of interest, and the time between the application of the off-resonance pulses and the imaging of the slice of interest. Dixon et al. l3 have shown that the off-resonance excitation of adjacent slices in a multi-slice conventional SE imaging experiment can decrease tissue signal as a result of MT. In multi-slice FSE imaging, a large number of 180” RF pulses are applied at a given off-resonance frequency in rapid succession and hence it is expected that a much greater MT effect should be observed. Indeed, as the results below indicate, MT plays a significant role in determining contrast in FSE imaging, and is the dominant cause of signal loss compared to SE imaging in brain tissue. It is important to distinguish MT effects which arise from off-resonance irradiation in multi-slice FSE imaging from the direct saturation of overlapping slice profiles. Such slice interference arises when small inter-slice gaps are used and the imperfect slice profiles from adjacent slices overlap. The spins in the overlap region between two adjacent slices are then partially saturated as they experience the RF pulses for both slices. This leads to a loss of signal from all tissues in the overlap region. Direct partial saturation of the tissue water resonance can also be produced by the off-resonance RF pulses (on-resonance for the slice selected but off-resonance for all other slices) because even though they are offset in frequency, they may still cause some effect on the spins at the appropriate onresonance frequency. Thus both MT and direct saturation mechanisms can lead to signal loss and tend to mimic each other. Furthermore, each of these mechanisms is more pronounced as the slice gap is narrowed.
0 R.T.
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Narrowing the slice gap moves the off-resonance RF pulses closer to the resonance frequency of the slice of interest and increases the slice overlap and direct saturation effects. The focus of this section is on separating these direct saturation effects from MT effects and to quantify the severity of the MT effects, as it has the potential for producing the altered contrast of FSE imaging. The effects of direct saturation and MT may be separated by observing changes in magnetization, and hence signal intensity, as a function of slice gap for materials which experience different amounts of MT. Materials that demonstrate no MT effects, such as GdDTPA doped water samples, will demonstrate signal changes only as a function of the direct saturation effect. Material such as cross-linked albumin and agarose gels have relaxation times that depend strongly on crossrelaxation, and thus will experience both mechanisms of signal loss. Separating direct saturation effects, which are identical for materials with the same value of T,/T2, from MT effects allows residual signal changes to be attributed to MT effects. Experiments designed to address these issues are described below. Figure 4 shows the results of computer calculations of the change in MZ predicted to result from direct irradiation by a train of 16, 180” RF pulses as a func-
.9. .8. .7. :
.6.
z A i!
.5.
g
.4.
a TVl’2=2 o Tlfl’2=5 n TU’l%lO D TlA’2=20
.3* .2. .l. 0 -200 1
500
1000 1500 2000 2500 3000 3500 4000 4 10
Offset (Hz) Fig. 4. Computer
simulation demonstrating the direct saturation effects followed a train of off-resonance RF pulses. The pre-pulses applied off-resonance consisted of 16 180’ RF pulses of 3.2 msec duration separated by 16.8 msec delays. Note that at high r, /T, ratios much more of a direct saturation effect is observed.
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Magnetic Resonance Imaging 0 Volume 10, Number 4, 1992
tion of the off-resonance frequency for the set of 16 pulses. Each 180” RF pulse was assumed to have a duration of 3.2 msec (five lobes) and an interpulse delay of 16.8 msec. The longitudinal magnetization, Mz, was calculated by solving the Bloch equations, with appropriate Tr and T2 values, for the whole time evolution of the sequence and the final Mz was plotted as a function of offset frequency. T, was 1000 msec and Tz varied between 50 and 500 msec to yield T,/T, = 2, 5, 10, and 20. At small offset frequencies, there are some oscillations in the amplitude of the net magnetization, Mz, since the 180” pulses flip the spins back and forth. The figure shows that at large off-resonance frequencies, the direct saturation effect is strongest for materials with a large T,/T, ratio, but that the direct effect is ~30% for materials such as brain for pulses that are offset by 21500 Hz, which is equivalent to one slice thickness under typical imaging conditions. The results for offsets of several hundred Hz or larger and for T2s of 50 msec or more are a function of T,/T, only, instead of Tl or T2, when a steady state is reached after the pulse train. This is true for both simple square pulses and the sine pulse train we used, and was confirmed by using different Tl and T2 values but the same T,/T,. Experiments To demonstrate that both MT and direct saturation effects are present in multi-slice FSE imaging, twoslice FSE imaging was performed with slice gaps ranging from 1 to 10 mm, and the slice thickness fixed at 4 mm. Modifying the slice gap shifts the off-resonance frequency of the adjacent slice RF pulses (from 905 to 2534 Hz off-resonance corresponding to the above slice gap limits) and also alters the degree of direct saturation effects. A phantom was constructed containing a number of test tubes filled with one of the following samples; Gd-DTPA doped water, albumin (20%) or agarose (5%) gels (doped with Gd-DTPA and MnCl,). These samples provided a range of Tl /T2 ratios in addition to the aforementioned differences in MT properties. The FSE imaging parameters were; ETL = 16, r= 18 msec, TE=36msec, TR =2500msec. A second phantom composed of samples with matched T,s and T2s was also constructed to confirm that MT effects were responsible for the differential effects observed in the Gd-DTPA doped samples versus the gels and that these effects were not due to the differences in T,/T2 ratios of the samples. This second phantom was used to measure the MT and direct saturation effects as a function of slice gap (resonance offset) and as a function of the number of slices selected. These samples were used to estimate the influ-
ence of both the total RF power applied off-resonance and the time between the application of the off-resonance pulses and imaging. The number of slices was changed from one to seven with the time between slices changing from 46 to 1186 msec while maintaining a fixed TR. Since the different slices are interleaved in TR (i.e., all slices are excited within the TR of the first slice), and since the echo trains for each slice are evenly spaced in the TR interval, as the number of slices increases, the time between respective echo trains must decrease. The FSE imaging parameters were; ETL = 16, r = 15 msec, TE = 15 msec, TR = 2000 msec. In one further experiment, the number of slices was held constant at two and only the time between consecutive slices was changed. The results, described below, confirmed that MT can be significant in FSE imaging. To illustrate directly the MT effect observed in FSE imaging and validate our computer predictions, a “FAKE” sequence was developed which applies a series of 16, 180” frequency selective RF pulses, at a given off-resonance frequency, prior to collecting a single phase-encoded spin echo for imaging. Using this sequence it is possible to mimic the off-resonance effects of the RF pulses in multi-slice FSE imaging while collecting the data in a conventional SE format. The sequence differs from the FSE sequence in two ways. First, the off-resonance pulses are the same as the slice selective refocusing pulses that would be applied to adjacent slices, but are applied without a slice select gradient. In addition, the actual imaging sequence is a single echo SE sequence and hence any changes in contrast with the application of the train of pre-pulses will be due solely to off-resonance effects. The use of a single echo, SE sequence avoids any contrast changes which may arise from either PSF effects or from the use of multiple echoes that have been influenced by the application of multiple RF pulses in the CPMG echo train used in FSE sequences. This sequence should therefore produce identical contrast to the FSE sequence in Gd-DTPA doped water samples. Results and Discussion The change in signal intensity as a function of slicegap in multi-slice FSE imaging is shown in Fig. 5. The results show both a direct saturation effect, as indicated by the increase in signal intensity of the GDDTPA doped water samples (dashed lines) as the slice gap is increased, and a MT effect, as indicated by the continued increase in signal from the albumin and agarose gels (solid lines), after the Gd-DTPA signal intensities have reached maximum values as slice gap increased. The slice interference effects are negligible beyond a 100% skip as indicated by the lack of change
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CONSTABLE ET AL.
_
Tl/T2=22.14(Albumin)
_
Tl!l2=23.52(Albumin)
+
Tl/T2=44.14(Agar)
-
TlIT2=2293(Albumin)
___... . ...-.
T,,T2
..-.........-.
T,m2=,
6.19
. . ..-..+.--..
T,n&,
2.84
---k---
T, ,T24.33
=,,,
65
!
Gd-DTPA
..-..+..
r.,
0
I.
4 Slice
8 Separation
_...
T,m2&,58
+ MI-CL2
doped water
I.,
12
16
(mm)
Fig. 5. Plot of signal intensity as a function of slice separation in a multi-slice imaging experiment for Gd-DTPA doped water and cross-linked albumin and agarose gels. The doped water solutions, which do not experience MT, show no further recovery of magnetization beyond slice gaps of approximately 4 mm, indicating that direct saturation effects are negligible beyond this point. In contrast, the signal intensity for the gels, which are known to experience MT, continues to increase as the slice gap increases beyond a 100% skip. This further increase in the net magnetization is attributed to a decrease in MT as the adjacent slices are moved further off-resonance.
of the signal intensity
of the Gd-DTPA doped water samples as the slice gap increased further. However, the singles from both the albumin and agarose gels, in
which cross relaxation is known to be significant,30 continue to increase with slice gaps beyond 4 mm, or beyond the regime of direct saturation effects. This further increase in signal for the gels provides strong evidence for the occurrence of the MT effect in multislice FSE imaging. The Gd-DTPA and gel samples with matched Z”, and T,showed similar trends to the above experiment and confirmed that MT an have a significant effect on image contrast in multi-slice FSE imaging. These samples were also used to quantitate the MT effect as a function of the number of slices excited within a TR interval and as a function of the time between slice excitations. The results, shown in Fig. 6, indicate that increasing the number of slices selected increases the significance of the MT effect. Figure 7 indicates that as the time interval between exciting adjacent slices is decreased the MT effect increases. The results show that not only do the RF pulses for the slices immediately adjacent to the slice of interest produce an MT
effect, but that the RF pulses for slices much further away also play a role. This is evident in Fig. 7 as the change in Mz for the double slice experiment is always smaller than the change observed in the seven slice experiment for the same inter-slice time delay. This is reasonable because the macromolecular linewidth in the gels extends over a frequency range much larger than the offset of each slice. To summarize, doubling the time interval between exciting adjacent slices can lead to a 10% decrease in the MT effect and hence an increase in signal intensity for those tissues experiencing MT. Furthermore, if this increased inter-slice time interval is obtained through an increase in TR then a synergistic effect results which provides a gain in SNR greater than that found when increasing TR alone. The “FARE” data confirm that the 180’ refocusing RF pulses used in multi-slice FSE imaging are capable of inducing significant MT effects. The results shown in Fig. 8 demonstrate that in a cross-linked gel, signal loss from these off-resonance pulses can be greater than 10% at typical adjacent slice offsets of 1500 Hz, and this result is for only a two-slice experiment. As the slice gap is narrowed and the off-resonance pulses
Magnetic Resonance Imaging 0 Volume 10, Number 4, 1992 1.051
.95.
-2 i
.a.
+ Tl/T2-matched solution
.6.
White matter v Gray matter n
‘? .a5
C Z s % ;
.75.
$z
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.a. 0 Solution (7 slices) Gel (7 slices) v Gel (2 slices) +
.65. .6.
.41!
0
1 10
30
40
6.0
70
,551 0
200
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600
Intra-slice
6. Magnetization transfer versus T, /T, samples. The results indicate MT is takof signal with decreasing is not a r1 effect.
approach MT effects increase,
400
of the slice as one would
inter-
J-COUPLING
Theory The high resolution spectra of lipids and fat contain multiple resonances with different chemical shifts and often with multiplet resonances from spin-spin couplings. Such J-couplings occur largely from homonuclear interactions between neighboring protons in hydrocarbon chains and may complicate the interpretation of relaxation time measurements, especially estimates of T2 obtainable by CPMG or similar pulse sequences. In homonuclear-coupled systems, the phase evolution of coupled spins is not reversed by a perfect 180” pulse, whereas all chemical shift and field inhomogeneity effects are reversed. The frequency differences between homonuclear coupled nuclei arising from spin-spin effects therefore causes a continual phase dispersal through out a spin echo or CPMG sequence, and the evolution of this spread in phase causes the NMR signal to be modulated in an oscillatory fashion in addition to its intrinsic T2 decay. Even in simple spectra, the echo amplitude at short TE may appear smaller than would be indicated by T2 alone because of such modulation effects, although the amplitude may actually increase again at longer times as the phase difference between coupled spins undergoes cyclic variations. For a doublet, the modulation is of
a00
1000
1200
ld
time delay (ms)
Fig. 7. Ratio of multi-slice to single-slice signal as a function of intra-slice time deIay. Increasing the number of slices while maintaining TR fixed leads to an increase in signal loss through MT. Alternatively, maintaining a fixed number of slices while decreasing the time interval between excitation of the slices leads to an increase in MT effects. This example demonstrates both effects.
the form cos( rnJ7); where, hertz, IZis the echo number For example, for a doublet crements in TE in a single
J = coupling constant in and 7 is the echo spacing. separated by J, small inecho SE experiment will
+
Gel
WSolution
-5
0
5
10 15 20 25 30 Frequency offset &Hz)
35
40
. 45
Fig. 8. To confirm that MT effects can be induced by pulsed irradiation, a “FAKE” imaging sequence was developed that consists of a train of 16 180” frequency selective RF pulses followed by a 90”, 180” spin-echo imaging sequence. The FAKE data confirm that pulsed irradiation with the same equivalent off-resonance power as that used in a typical double slice FSE imaging experiment can induce MT effects.
Contrast in fast spin-echo MRI 0 R.T.
yield an oscillating response. The signal oscillates as the peaks are shifted from being in phase (maximum signal at TE = l/J) to each peak being shifted by 180” with respect to its coupled neighbor, (minimum signal at TE = l/25), producing complete cancellation of the signal in this example. For complex spectra, the summation of several different J-modulated decays may appear more like a monotonically decreasing signal whose apparent T, is dominated by the number and strength of the couplings present, rather than the intrinsic relaxation rates of each resonance. The measured T2 in such circumstances is dependent on the precise details of the measurement. Apparent relaxation times in complex coupled systems may depend strongly on the-number and rate of the refocusing pulses in a multiple echo sequence. Allerhand’ derived a relationship between the number of homonuclear coupled spins, their relative chemical shifts, and the echo spacing in a CPMG experiment. The increase in signal observed in FSE imaging for a coupled system as the echo spacing is decreased and the number of echoes increased is properly described using quantum mechanics. It can, however, be understood by the following classical argument. As the pulse rate becomes large relative to the chemical shifts, phase deviations due to the shifts approach zero, and hence can be neglected. Although the 180” pulses do not refocus the J-coupling evolution, under these conditions the cumulative motion of each spin is a precession around the field due to the spins to which it is coupled. The resulting motion is complicated, but preserves the net spin vector, and hence the signal, at the echo peaks. This result may also be understood in terms of the well-known theorem that J-coupling between equivalent spins is unobservable. 32 Rapid pulsing in a CPMG sequence removes the effects of chemical shift, rendering all spins “equivalent,” and results in a signal unmodulated by the coupling. Experiments and Results All imaging to study J-coupling effects was performed on single slices to avoid direct saturation and MT effects. We chose to study the behavior of corn oil as well as fat because they behave similarly when imaged using FSE and SE sequences. Thus many of the experiments described below compare corn oil and water signals. Conventional SE imaging was performed (256 x 128 x 2 nex, TR = 2000 msec) with small increments in TE to examine if modulation of the T2 decay curve from J-coupling effects could be observed using a phantom containing core oil and Gd-DTPA doped water solutions. The TE was varied from 15 to 100 msec in 5-msec increments. It was found that small increments in TE in a conventional SE sequence do not
507
CONSTABLE ET AL.
demonstrate modulation of signal in either sample. The lack of modulation implies that a number of different couplings may be present in corn oil and that while some coupled spins are moving out of phase, others may be moving into phase with the net result an essentially monotonic decay. The decay rate largely reflects the number and separation of component frequencies rather than the intrinsic linewidth of any one resonance. Multiple J-coupled peaks are indeed present in corn oil as confirmed by high resolution NMR spectra (not shown). Similar results have been obtained in fat. A J-coupled system should demonstrate an increase in signal as the CPMG echo train effectively decouples the spin system. A series of imaging experiments were performed using the FSE sequence with a range of values of echo spacing 7, and ETL. These experiments were designed to demonstrate the decoupling effect of the CPMG sequence as r is shortened and ETL increased. Signal intensities obtained using the FSE sequence and the conventional SE sequence were compared using corn oil and water samples. The FSE imaging was carried out using a constant effective TE but with different ETLs and 7 values (ETL = 8, r = 15 msec; ETL = 4, 7 = 30 msec; ETL = 2, r = 60 msec; with TE = 120 msec in each case). The use of variable echo spacing (7) and ETL with fixed effective TE, changes the degree of decoupling and hence signal intensity in J-coupled systems. Comparing the FSE sequence with the conventional signal slice, single echo SE imaging sequence with these changes in ETL and r, demonstrates the effects of the train of 180” RF pulses. Figure 9 shows the preferential enhancement of signal from a corn oil phantom compared with that from water as the echo train length in FSE imaging is increased. It also shows that the signal intensity for corn oil is much brighter in images obtained using the FSE sequence than in those obtained using the conventional SE sequence. This preferential signal enhancement is a characteristic feature of the FSE imaging sequences. Since the bright fat signal is essentially a consequence of a decoupling phenomenon, it is useful to consider alternate refocusing schemes that produce spin echoes, but which are less effective at decoupling. Shaka and Keeler33 differentiate the relative decoupling performances of the two sequences, [7/2 - 180”(x) - T - 180”(x) - 7/2]” and [7/2 - 180” (x) - 7 - 180” (-x)
- 7/2]“,
Magnetic Resonance Imaging 0 Volume 10, Number 4, 1992
0
e
FSE-CornOil
w
CPMo&1mcom01l
_
FSE-Water
10
5
Number
of
15
20
Echoes
Fig. 9. Signal intensity versus ETL for FSE and SE sequences. The signal intensity of corn oil increases as the ETL is increased and the echo spacing is decreased. The signal intensity of the water (devoid of J-coupled spins) remains constant. With alternating phases on the 180” pulses, the decoupling is reduced and the signal intensity falls between that of the CPMG-FSE sequence and the conventional SE sequence.
of the brain demonstrate similar contrast in all of the images, but notice that the fat signal increases with increasing ETL. This is a direct result of the removal of J-modulation effects for the coupled spins in fat as the number of 180” RF pulses increases and the spacing between the pulses decreases, as described above. Measurements were also made to quantify the effect of pulse rate on echo intensity from J-coupled nuclei using a spectrometer operating at 20 MHz and the CPMG sequence. We studied the echo intensity at a fixed TE (100 msec) for variable numbers and pulse rates of 180” pulses, for water and corn oil as well as hexane and 1-decanol. Hexane contains no polar group or exchangeable protons. The amplitude of the echo occurring at 100 msec was measured and the number of echoes preceding that echo was varied from 1 to 50, with 7 ranging from 50 to 2 msec and ETL from 2 to 50, respectively, Figure 11 shows the results of the spectrometer measurements for water doped with MnC& (same T2 as corn oil), corn oil, hexane, and l-decanol. The increase in signal intensity at short pulse intervals occurs at intervals below 20 msec, consistent with the chemical shift range of 2.5 ppm for coupled nuclei. This range corresponds to the range of shift visible in the high resolution spectrum. The increase in signal intensity, shown in the above examples for coupled systems, as the ETL is increased and the echo spacing is decreased, agrees with the results predicted by Allerhand for J-coupled systems. CONCLUSIONS
wherein the latter sequence uses refocusing pulses with
alternating phases. These sequences behave quite differently when employed as potential decoupling schemes for heterogeneous decoupling. In the limit as r + 0, the constant phase sequence becomes a conventional decoupling scheme, whereas the sequence with alternating phase does not decouple at all. By analogy therefore, it may be predicted that alternating the phases of the refocusing pulses in FSE should be less effective at removing J-modulation effects and will cause fat to be less bright. Evidence for this is provided experimentally in Fig. 9, where the results for the alternating phase sequence fall between that of the conventional FSE and SE imaging sequences. Figure 10 shows a conventional SE image of a normal brain a) with ETL = 1, and FSE images, b) ETL = 4, r = 28 msec, and c) ETL = 16, r = 14 msec. In each image, the TE = 112 msec and TR = 2000 msec. Note that the imaging time increase is inversely proportional to the echo train length. The soft tissues
The origins of the contrast differences between FSE and conventional SE images have been accounted and quantified. PSF effects are important only for long ETL and r values and short T2 materials. In most cases when using an ETL I 8, the PSF does not markedly alter the contrast of Tz-weighted images. Stimulated echoes, while contributing to the overall signal in an FSE experiment, do not significantly alter the T2 contrast. Gains of approximately 10% in signal intensity for soft tissues may be realized in single slice FSE imaging, from stimulated echo effects, but these are not realized in multi-slice FSE as the gain is offset by MT losses. The FSE sequence can regain some of the signal lost through diffusion processes (<2%), but this gain in signal intensity for soft tissues such as brain tissues, is also outweighed by the MT losses which occur in multi-slice FSE. Magnetization transfer does take place in multi-slice FSE imaging and this serves to darken preferentially the gray and white matter in the brain, but does not affect water (CSF) or fat significantly. In body imaging we expect MT to be important
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Fig. 10. Single slice SE image (A) ETL = 1, and FSE images (B) ETL = 4, T = 28 msec, and (C) ETL = 16, r = 14 msec. For all images TE = 112 msec and TR = 2000 msec. Note the soft tissues of the brain have similar signal intensity in each image while the subcutaneous fat layer surrounding the brain increases in intensity as the ETL increases and the echo spacing, 7, decreases.
in muscle, liver, and kidney. The MT observed, increases with the number of slices excited within a TR interval and as the time interval between the excitation of these slices decreases. With small slice gaps, direct saturation effects can decrease tissue signal through imperfect excitation profiles overlapping adjacent slices, just as in conventional SE imaging, and these affects may be more pronounced in FSE because of the application of a train of off-resonance RF pulses.
It should be noted that when FSE imaging is performed in interleaved mode, collecting the odd slices first and then collecting the even slices, both the direct saturation and MT effects are reduced but not eliminated. The attenuation of signals in J-coupled systems may be reduced with the use of a CPMG echo train, and it is the reduction of this dephasing in FSE imaging which is partly responsible for the bright fat signal observed. The fat also appears relatively brighter in
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Fig. 11. Signal intensity at TE = 100 msec for CPMG sequences with different 7 and echo numbers at 20 MHz. The signal, for the corn oil, 1-decanol, and hexane, increases with decreasing echo spacing and increasing ETL while the doped water signal remains constant. The I-decanol and hexane samples are included to rule out exchange effects.
brain
imaging
ential
darkening
through
because
of the aforementioned
of the gray and white
matter
prefertissues
MT.
Acknowledgments- We gratefully acknowledge the support of grant CA40675 awarded by the National Cancer Institute and the support of G.E. Medical Research Systems, Milwaukee, WI. A.W.A. acknowledges support from an NIH Training Grant NC1 CA-09549. Drs. R. Mark Henkelman, Richard Kennan, Gordon Sze, and Robert Smith are also thanked for useful discussions. REFERENCES 1.
2.
3.
4.
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