- Email: [email protected]
Determination of saturation transfer parameters of human tissues in vivo
Magnetic
Imaging, Vol. 14, No. 4, pp. 413-417, 1996 Copyright 0 1996 Elsevier Science Inc. Printed in the USA. All rights reserved 0730-7251096 $15.00 + .OO
0730-725X( 96) 00029-S
ELSEVIER
l
Resonance
Original Contribution DETERMINATION
OF SATURATION TRANSFER OF HUMAN TISSUES IN VIVO
SAMI KAJANDER,
MARKKU
KOMU,
F%KKA NIEMI,
PARAMETERS
AND MARTTI
KOFCMANO
Diagnostic Imaging Center, Turku University Hospital, Turku, Finland In order to study the applicability of magnetization transfer contrast (MTC) to tissue differentiation, the determination of the magnetization transfer (MT) parameters of normal tissues is necessary for the evaluation of pathological conditions. The time-dependent saturation transfer technique was used to investigate the observed magnetization transfer parameters in several human tissues in vivo at 0.1 T. The length of the off-resonance saturation pulse varied from 0 to 750 ms. The magnetization transfer contrast (MTC) was 0.71 in striated muscle, 0.49 in liver, 0.49 in renal cortex, and 0.50 in spleen. The observed magnetization transfer rates (R,) were 5.5 s-’ for muscle, 3.1 s-’ for liver, and 1.5 s-’ for both renal cortex and spleen. Our results indicate that measuring R,, and possibly other relaxation parameters could be useful in tissue differentiation. Keywords:
MRI; Low-strength
imaging; Magnetization
INTRODUCTION
5123195;
ACCEPTED
contrast; Tissue characterization.
the water proton RF excitation pulse and water proton signal can be collected using a routine imaging sequence. The magnitude of water proton signal intensity reduction will be due to the magnetization transfer between the macromolecular and water pools. This signal reduction depends on several factors, among them Rwm, water and macromolecular relaxation times T, and TZ, and the properties of the saturating field (power, frequency, and duration). The saturation pulse can be combined with both T,- and T,-weighted sequences and therefore the MTC value seems to give information that is additional, not alternative, to that of conventional methods. 4*7-9 This may aid in tissue characterization, which is one of the two main applications for clinical MT imaging (contrast augmentation being the other) and which has been hampered by the significant overlapping between T,s (and Tzs) of different types of tissues.” Indeed, several in vivo and in vitro studies have attempted to relate MTC to tissue composition, 12-‘* with varying results. There are few reports in which the emphasis has been on analyzing MT parameters other than just MTC. Eng et al.,’ however, constructed magnetization transfer parameter maps from a rabbit kidney in vivo at 4.7
Magnetization transfer (MT) is a consequence of magnetic dipolar coupling between restricted, macromolecular protons and protons of free water. Due to this coupling, relaxation of one of the populations will influence the relaxation of the other (cross-relaxation) .lm4 Although the MT phenomenon has been studied for years (even before modem MR imaging technology)* 1-3 greater understanding of these rather complex processes and the capability of MT imaging have been acquired only recently.4-6 The current models consider the water and macromolecular protons as two systems with equilibrium established within the macromolecular pool through spin diffusion and within the water pool through molecular exchange and self-diffusion. MT from free water to macromolecular proton is detected in the bulk water phase.4 The magnetization transfer contrast (MTC) can be generated by selectively saturating the macromolecular proton pool with a radiofrequency (RF) pulse offset from the Larmor resonance frequency (saturation transfer technique) .4*6This pulse is applied just before RECEIVED
transfer
tic Imaging Center, Turku University Hospital, Kiinamyllynkatu 4-8, FIN 20520 Turku, Finland.
212196.
Address correspondence to Dr. Sami Kajander, Diagnos413
Magnetic Resonance Imaging 0 Volume 14, Number 4, 1996
414
T. They found that R,, maps had significantly greater contrast than those with other parameters. Niemi et a1.9 found out that most abdominal organs consist of tissues that have strong MTC values and increased values of RWItI.Most tissues also showed typical and quite specific patterns of signal decline in vitro. The purpose of this article is to study the applicability of these results to many human tissues in vivo and examine the role of R,, measurements in MT protocols. Specifically, the emphasis will be on using sequences and protocols that can be utilized in a clinical setting. MATERIALS
AND METHODS
Eight healthy volunteers, four men and four women, between the ages of 23 and 30 yr, were imaged on a permanent-magnet 0.1-T clinical MRI device (Merit, Picker Nordstar, Helsinki, Finland). A body coil was used with a tensive belt around the abdomen to decrease respiratory motion. Slice thickness was 10 mm and the field-of-view was 645 X 484 mm with a pixel size of 2.5 X 2.5 mm. After a coronal localizer scan, a single transaxial slice was acquired. The slice included the organs of interest which were liver, spleen, paraspinal muscle, and kidney. The T1 value of tissue was determined with five standard inversion-recovery (IR) sequences (TR/TE/TI = 1000/20/30, 100, 300, 400, 500 ms). From signal intensities, operator-defined region-of-interest (ROI) tracings were generated on magnified images with Picker software that allowed them to be stored and transferred to other images while maintaining the corresponding position. Using this method, ROIs were always defined from the image that had the best contrast. The pulse sequences used to measure MT consisted of two sequences that differed only by the presence or .absence of an off-resonance saturation RF pulse. This Gaussian pulse with a 20-PT RF field and a 4-kHz offset frequency was launched before the 90” excitation pulse and the signal was collected as a gradient echo ( lOOO/ 20 ms). The long TR and short TE diminished undesirable T1 or T2 weighting. The durations of the saturation pulses were 50, 100, 200, 300, 500, and 750 ms. This provided enough data for MT calculations even though the parameters measured differed considerably from one type of tissue to another. A detailed description of the method has been published elsewhere.’ To record the signal intensities, a ROI tracing method similar to the one used in T1 measurements was applied. A single transaxial slice from each volunteer, as in T, measurements, was studied: the employment of just one preselected slice prevented crossover
of MT signal,” known to plague several multislice protocols, and decreased imaging time, which may be a limiting factor in MT imaging. In the time-dependent saturation transfer method, the exponential decay of longitudinal magnetization in the tissues was determined by fitting, with the leastsquares-sum method, the measured intensity values as a function of saturation pulse length to the following equation: M(t)
= Mae-“T’s, + MS
(1)
where M(t) is the proton magnetization along the zaxis after a saturation pulse, t is the length of this saturation pulse, M, is the magnitude of the reduction of magnetization, MS is the residual magnetization at the steady-state level toward which the magnetization decays, and Tls, is the longitudinal relaxation time in the presence of irradiation of the restricted proton pool. The equilibrium magnetization in the absence of the saturation pulse, MO, is: M, = Ma + M,
(2)
and MTC is simply: MTC = 1 - MJM, The observed magnetization then be expressed as:
= Ma/MO
transfer rate, R,,,
Rwm= MTCITrsal
(3) can
(4)
From each volunteer, MTC and R,, were calculated for each type of tissue. We then compared the differences among tissues in both these parameters using an unpaired nonparametric t-test. The average whole body specific absorption rates (SARs) were calculated, and were below the FDA limits (0.4 W/kg for body and extremities) with a maximum SAR of 0.17 W/kg. RESULTS The main results are summarized in Table 1. As expected from previous studies,4*‘1,20 striated muscle showed both the greatest MTC (0.7 1) and the highest value of R,, (5.5 s-l). The MTC values of the other tissues, however, did not differ from each other significantly. The value of the MTC was 0.49 in both liver and kidney, and 0.50 in spleen. On the other hand, R,, of the liver ( 3.1 s -’ ) was significantly greater than that of either kidney cortex ( 1.5 s-r) or spleen ( 1.5
Saturation transfer parameters
0 S. IWANDER
41.5
m AL.
Table 1. Measured T, values and magnetization transfer parameters in human tissues in vivo at 0.1 T Tissue
Tlobs(ms)
Tl,, b-d
Muscle Liver Kidney cortex Spleen
258 212 470 420
130 + 164 t 342 + 332 ?
+ 12 t 12 k 71 -c 28
14 29 65 37
MTC 0.71 0.49 0.49 0.50
+ 0.01 ?I 0.06 c 0.04 2 0.07
Rwm(s-l) 5.5 3.1 1.5 1.5
” t ‘+
0.6 0.8 0.3 0.1
Note: All data expressed as mean + standard deviation. Muscle refers to striated muscle, kidney to renal cortex, T,,, = measured T, of tissue; T,_, = T, in the presence of saturation pulse; MTC = magnetization transfer contrast; R,, = observed magnetization transfer rate.
s-r). Between spleen and kidney cortex, no difference in R,, was noted. Other abdominal organs were not included in the study due to difficulties in reliably defining the rather small organs (e.g., pancreas) in a single slice with enough tissue material to avoid the partial volume effect. In kidneys, differentiation of cortex and medulla was impossible during some MT pulse lengths, and patient movement prevented the use of transferable ROIs (described previously). Therefore, kidney measurements consisted of renal cortex only. Fat was excluded because these young volunteers had very little fat and fat is known from the literature to have almost no MTC. The image quality varied between individuals and respiratory motion did cause some artifacts despite the precautions. In general, however, images were of satisfactory quality (Fig. 1) and in no individual they were unacceptable. DISCUSSION The significant, apparently tissue-dependent differences of signal intensities acquired with saturation pulses seem to reflect the MT phenomenon itself, rather than changes in T, or T2. Tissues that have similar appearances in conventional spin-echo or gradient-echo images may show contrast in MT, and this contrast is different from that in conventional imaging. The clinical application of MTC for tissue differentiation purposes has, however, been hampered by some limitations and has showed only modest success.” First, it seems that, between many neighboring tissues (e.g., kidney, spleen, and liver), differences in signal intensities using long saturation pulses (and thus achieving maximal h4TC) are practically nonexistent, as the MTC curves approach each other near the plateau. Second, it is often difficult to separate intensity changes caused by crossrelaxation and MT itself from changes induced by differences in T1 or T2 onto MTC.49” Both theoretically and in our experience the measurements of the rate of magnetization hold promise in tissue differentiation. Although the observed MT rate parameter
R,, in this study contains some direct saturation effects4
we believe that this observed parameter reflects the rate of magnetization transfer reasonably well. In our series, only striated (paraspinal) muscle had both R,, and MTC that were significantly greater than those of other tissues analyzed. Liver, kidney cortex, and spleen each achieved maximal signal reduction (MTC) of about 50% with no significant differences between them. On the other hand, the R, of liver was markedly higher than those of kidney and spleen. Although the optimal RF power and the magnitude of the resonance offset of the MT pulse can be calculated,798 the effect of field strength on MT parameters is, in our opinion, still a largely unsettled issue. It is usually believed that MTC becomes greater as the field strength increases.“~24 However, some investigators have concluded that the field strength of the MRI apparatus does not contribute to the magnitude of the MTC.22 On the other hand, it seems that R,, may be larger in lower fields 7.9,19but the exact nature of this relationship remains, to our knowledge, unknown. The rate of magnetization transfer in kidney cortex, for example, has been reported as 1 s-’ at 4.7 T,’ 1.4 s-l at 1.5 T,” and near 4 SC’ at 0.1 T.9 In our series, we obtained values of around 1.5 s-l, considerably lower than obtained at 0.1 T in vitro but still higher than those acquired at high fields. A comprehensive study of the impact of field strength on the MT parameters remains needed both in vitro and in vivo. Liver tissue showed considerable variation of MT parameters among the volunteers, with R,, ranging from 2.3 to 4.8 SK’ and MTC from 0.44 to 0.62. The reasons for this variance are still to be studied; however, it is possible that iron deposits of liver may cause the diversity of MTC. Another possible explanation is the role of diet, which may have an effect on MTC in liver.25 In this study, neither the amount of fat and connective tissue within liver was known in the volunteers (although they were all in the same age group, none of them were grossly overweight and their alcohol consumption was moderate) nor was their dietary intake prior to MR measurement controlled in any way.
Magnetic Resonance Imaging 0 Volume 14, Number 4, 1996
416
(4
(B)
m
(D)
Fig. 1. The effect of magnetizationtransferon MRI imagesof severalhumantissuesat 0.1 T in vivo. Gradient echo ( lOOO/ 20 ms) imageswith no MT pulse(A), MT pulseof 100 ms (B), 300 ms (C), and 7.50ms (D) .
The in vivo determination of MTC parameters of normal tissue is necessary for future studies of pathological conditions. We also believe that analysis of different age groups, nurturing conditions, and other determinants will be needed to further evaluate the MT effect in humans. In addition to a change in water content, many tissues may show signs of aging or other physiological processes having MRI appearances that are still unknown. In the abdominal area, despite some studies on liver pathology, 2223reports of MT imaging on parenchymal diseases of other organs have been scarce. However, some exciting possibilities in tissue characterization may be seen. Clinically important, for example, would be the evaluation of diseases, such as amyloidosis, that may diffusely infiltrate several organs without macroscopic tumor formation. The MRI appearance and relaxometry of such diseases is usually unspecific and difficult to interpret. Tissue-specific MT analysis could be of assistance and estimates of the amount of pathological tissue
could be made. Parametric images based on R,, or other MT parameters can also be created, and our preliminaty results have been encouraging. Despite its rather modest spatial resolution, lowfield imaging may offer some advantages over higher field strengths for MTC imaging. Because the longitudinal relaxation rates are shorter, shorter saturation pulses are needed for steady-state magnetization. Magnetization transfer rates may be, as mentioned above, greater, and power requirements lesser. Also, FDA guidelines may be broken when long off-resonance pulses are applied in high fields. Therefore, we think that low-field MRI also has its place in MT measurements. CONCLUSIONS 1. The results of the MT measurements reflect the properties of tissues and are measurable in the ab-
Saturationtransfer parameters0 S. KAJANDER
dominal and paraspinal area in vivo. Knowing the normal values of these parameters aids in future attempts in tissue differentiation. 2. The observed magnetization transfer rate, R,,, seems to be a better indicator of the MT phenomenon than MTC. REFERENCES
2. Edzes, H.T.; Samulski, E.T. The measurement of cross-
13. 14. 15.
16.
relaxation effects in proton NMR spin-lattice relaxation 3. 4.
5.
6. 7.
8.
9.
10.
11.
of water in biological systems: Hydrated collagen and muscle. J. Magn. Reson. 31:207-229; 1978. Edzes, H.T.; Samulski, E.T. Cross-relaxation and spin diffusion in the proton NMR of hydrated collagen. Nature (Lond.) 265521-523; 1977. Balaban, R.S.; Ceckler, T.L. Magnetization transfer contrast in magnetic resonance imaging. Magn. Reson. Q. 8:116-137; 1992. Koenig, S.H.; Brown, R.D. III. A molecular theory of relaxation and magnetization transfer: Application to cross-linked BSA, a model for tissue. Magn. Reson. Med. 6:685-695; 1993. Wolff, S.D.; Balaban, R.S. Magnetization transfer contrast (MTC) and tissue water proton relaxation in vivo. Magn. Reson. Med. 10:135-144; 1989. Eng, J.; Ceckler, T.L.; Balaban, R.S. Quantitative 1H magnetization transfer imaging in vivo. J. Magn. Reson. 17:304-314; 1991. Komu, M.E.S. Analysis of longitudinal relaxation rate constants from magnetization transfer MR images of human tissues at 0.1 T. Magn. Reson. Imag. 10:35-40; 1992. Niemi, P.T.; Komu, M.E.S.; Koskinen, S.K. Tissue specificity of low-field-strength magnetization transfer contrast imaging. Magn. Reson. Imag. 2:197-201; 1992. Bottomley, P.A.; Hardy, C.J.; Argensinger, R.E.; AllenMoore, G. A review of 1H NMR relaxation in pathology: Are Tl and T2 diagnostic? Med. Phys. 14:1-37; 1987. Wolff, S.D.; Balaban, R.S. Magnetization transfer imaging : Practical aspects and clinical applications. Radiology 192:593-599; 1994.
417
12. Wolff, S.D.; Balaban, R.S. Magnetization transfer con-
1. Koenig, S.H.; Bryant, R.G.; Hallenga, K.; Jacob, G.S.
Magnetic cross relaxation among protons in protein solutions. Biochemistry 17:4348-4358; 1978.
ETAL.
17.
18. 19.
20.
21.
22.
23.
24. 25.
trast: MR imaging of the knee. Radiology 179:133-137; 1991. Koskinen, S.K.; Komu, M.E. Low-field srength magnetization transfer contrast imaging of the pate&r cartilage. Acta Radiol. 34:124-126; 1993. Lundbom, N. Determination of magnetization transfer contrast in tissue: An MR imaging study of brain tumors. AJR 159:1279-1285; 1992. Ordridge, R.J.; Helpem, J.A.; Knight, R.A.; Qing, Z.X.; Welch, K.M. Investigation of cerebral ischemia using magnetization transfer contrast (MTC) MR imaging. Magn. Reson. Imag. 9:895-902; 1991. Balaban, R.S.; Chesnick, S.; Hedges, K.; Samaha, F.; Heineman, F.W. Magnetization transfer contrast in MR imaging of the heart. Radiology 180:671-675; 1991. Mittl, R.L., Jr.; Gomori, J.M.; Schnall, M.D.; Holland, G.A.; Grossman, R.I.; Atlas, S.W. Magnetization transfer effects in MR imaging of in vivo intracranial hemorrhage. AJNR 14:881-891; 1993. Gomori, J.M.; Grossman, R.I.; Asakura, T.; et al. An in vitro study of magnetization transfer and relaxation rates of hematoma. AJNR 14:871-880; 1993. Dixon, W.T.; Engels, H.; Castillo, M.; Sardashti, M. Incidental magnetization transfer contrast in standard multislice imaging. Magn. Reson. Imag. 8:417-422, 1990. Hajnal, J.V.; Baudouin, C.J.; Oatridge, A.; Young, I.R.; Bydder, G.M. Design and implementation of magnetization transfer pulse sequences for clinical use. J. Comput. Assist. Tomogr. 16:7-18; 1992. Grad, J.; Mendelson, D.; Hyder, F.; Bryant, R.G. Direct measurements of longitudinal relaxation and magnetization transfer in heterogenous systems. J. Magn. Reson. 86:416-419; 1990. Outwater, E.; Schnall, M.D.; Braitman, L.E.; Dinsmore, B.J.; Kressel, H.Y. Magnetization transfer of hepatic lesions: Evaluation of a novel contrast technique in the abdomen. Radiology 182:535-540; 1992. Kahn, C.E., Jr.; Perera, S.D.; Sepponen, R.E.; Tanttu, J.I.; Tierala, E.K.; Lipton, M.J. Magnetization transfer imaging of the abdomen at O.lT: Detection of hepatic neoplasms. Magn. Reson. Imag. 11:67-71; 1993. Ceckler, T.L.; Balaban, R.S. Field dispersion in watermacromolecular proton magnetization transfer. J. Magn. Reson. Series B 105:242-248; 1994. Salo, S.; Alanen, A.; Komu, M.; Kallio, T.; Bondestam, S. Effect of fasting and food intake on magnetization transfer of human liver tissue. (submitted).
Imaging, Vol. 14, No. 4, pp. 413-417, 1996 Copyright 0 1996 Elsevier Science Inc. Printed in the USA. All rights reserved 0730-7251096 $15.00 + .OO
0730-725X( 96) 00029-S
ELSEVIER
l
Resonance
Original Contribution DETERMINATION
OF SATURATION TRANSFER OF HUMAN TISSUES IN VIVO
SAMI KAJANDER,
MARKKU
KOMU,
F%KKA NIEMI,
PARAMETERS
AND MARTTI
KOFCMANO
Diagnostic Imaging Center, Turku University Hospital, Turku, Finland In order to study the applicability of magnetization transfer contrast (MTC) to tissue differentiation, the determination of the magnetization transfer (MT) parameters of normal tissues is necessary for the evaluation of pathological conditions. The time-dependent saturation transfer technique was used to investigate the observed magnetization transfer parameters in several human tissues in vivo at 0.1 T. The length of the off-resonance saturation pulse varied from 0 to 750 ms. The magnetization transfer contrast (MTC) was 0.71 in striated muscle, 0.49 in liver, 0.49 in renal cortex, and 0.50 in spleen. The observed magnetization transfer rates (R,) were 5.5 s-’ for muscle, 3.1 s-’ for liver, and 1.5 s-’ for both renal cortex and spleen. Our results indicate that measuring R,, and possibly other relaxation parameters could be useful in tissue differentiation. Keywords:
MRI; Low-strength
imaging; Magnetization
INTRODUCTION
5123195;
ACCEPTED
contrast; Tissue characterization.
the water proton RF excitation pulse and water proton signal can be collected using a routine imaging sequence. The magnitude of water proton signal intensity reduction will be due to the magnetization transfer between the macromolecular and water pools. This signal reduction depends on several factors, among them Rwm, water and macromolecular relaxation times T, and TZ, and the properties of the saturating field (power, frequency, and duration). The saturation pulse can be combined with both T,- and T,-weighted sequences and therefore the MTC value seems to give information that is additional, not alternative, to that of conventional methods. 4*7-9 This may aid in tissue characterization, which is one of the two main applications for clinical MT imaging (contrast augmentation being the other) and which has been hampered by the significant overlapping between T,s (and Tzs) of different types of tissues.” Indeed, several in vivo and in vitro studies have attempted to relate MTC to tissue composition, 12-‘* with varying results. There are few reports in which the emphasis has been on analyzing MT parameters other than just MTC. Eng et al.,’ however, constructed magnetization transfer parameter maps from a rabbit kidney in vivo at 4.7
Magnetization transfer (MT) is a consequence of magnetic dipolar coupling between restricted, macromolecular protons and protons of free water. Due to this coupling, relaxation of one of the populations will influence the relaxation of the other (cross-relaxation) .lm4 Although the MT phenomenon has been studied for years (even before modem MR imaging technology)* 1-3 greater understanding of these rather complex processes and the capability of MT imaging have been acquired only recently.4-6 The current models consider the water and macromolecular protons as two systems with equilibrium established within the macromolecular pool through spin diffusion and within the water pool through molecular exchange and self-diffusion. MT from free water to macromolecular proton is detected in the bulk water phase.4 The magnetization transfer contrast (MTC) can be generated by selectively saturating the macromolecular proton pool with a radiofrequency (RF) pulse offset from the Larmor resonance frequency (saturation transfer technique) .4*6This pulse is applied just before RECEIVED
transfer
tic Imaging Center, Turku University Hospital, Kiinamyllynkatu 4-8, FIN 20520 Turku, Finland.
212196.
Address correspondence to Dr. Sami Kajander, Diagnos413
Magnetic Resonance Imaging 0 Volume 14, Number 4, 1996
414
T. They found that R,, maps had significantly greater contrast than those with other parameters. Niemi et a1.9 found out that most abdominal organs consist of tissues that have strong MTC values and increased values of RWItI.Most tissues also showed typical and quite specific patterns of signal decline in vitro. The purpose of this article is to study the applicability of these results to many human tissues in vivo and examine the role of R,, measurements in MT protocols. Specifically, the emphasis will be on using sequences and protocols that can be utilized in a clinical setting. MATERIALS
AND METHODS
Eight healthy volunteers, four men and four women, between the ages of 23 and 30 yr, were imaged on a permanent-magnet 0.1-T clinical MRI device (Merit, Picker Nordstar, Helsinki, Finland). A body coil was used with a tensive belt around the abdomen to decrease respiratory motion. Slice thickness was 10 mm and the field-of-view was 645 X 484 mm with a pixel size of 2.5 X 2.5 mm. After a coronal localizer scan, a single transaxial slice was acquired. The slice included the organs of interest which were liver, spleen, paraspinal muscle, and kidney. The T1 value of tissue was determined with five standard inversion-recovery (IR) sequences (TR/TE/TI = 1000/20/30, 100, 300, 400, 500 ms). From signal intensities, operator-defined region-of-interest (ROI) tracings were generated on magnified images with Picker software that allowed them to be stored and transferred to other images while maintaining the corresponding position. Using this method, ROIs were always defined from the image that had the best contrast. The pulse sequences used to measure MT consisted of two sequences that differed only by the presence or .absence of an off-resonance saturation RF pulse. This Gaussian pulse with a 20-PT RF field and a 4-kHz offset frequency was launched before the 90” excitation pulse and the signal was collected as a gradient echo ( lOOO/ 20 ms). The long TR and short TE diminished undesirable T1 or T2 weighting. The durations of the saturation pulses were 50, 100, 200, 300, 500, and 750 ms. This provided enough data for MT calculations even though the parameters measured differed considerably from one type of tissue to another. A detailed description of the method has been published elsewhere.’ To record the signal intensities, a ROI tracing method similar to the one used in T1 measurements was applied. A single transaxial slice from each volunteer, as in T, measurements, was studied: the employment of just one preselected slice prevented crossover
of MT signal,” known to plague several multislice protocols, and decreased imaging time, which may be a limiting factor in MT imaging. In the time-dependent saturation transfer method, the exponential decay of longitudinal magnetization in the tissues was determined by fitting, with the leastsquares-sum method, the measured intensity values as a function of saturation pulse length to the following equation: M(t)
= Mae-“T’s, + MS
(1)
where M(t) is the proton magnetization along the zaxis after a saturation pulse, t is the length of this saturation pulse, M, is the magnitude of the reduction of magnetization, MS is the residual magnetization at the steady-state level toward which the magnetization decays, and Tls, is the longitudinal relaxation time in the presence of irradiation of the restricted proton pool. The equilibrium magnetization in the absence of the saturation pulse, MO, is: M, = Ma + M,
(2)
and MTC is simply: MTC = 1 - MJM, The observed magnetization then be expressed as:
= Ma/MO
transfer rate, R,,,
Rwm= MTCITrsal
(3) can
(4)
From each volunteer, MTC and R,, were calculated for each type of tissue. We then compared the differences among tissues in both these parameters using an unpaired nonparametric t-test. The average whole body specific absorption rates (SARs) were calculated, and were below the FDA limits (0.4 W/kg for body and extremities) with a maximum SAR of 0.17 W/kg. RESULTS The main results are summarized in Table 1. As expected from previous studies,4*‘1,20 striated muscle showed both the greatest MTC (0.7 1) and the highest value of R,, (5.5 s-l). The MTC values of the other tissues, however, did not differ from each other significantly. The value of the MTC was 0.49 in both liver and kidney, and 0.50 in spleen. On the other hand, R,, of the liver ( 3.1 s -’ ) was significantly greater than that of either kidney cortex ( 1.5 s-r) or spleen ( 1.5
Saturation transfer parameters
0 S. IWANDER
41.5
m AL.
Table 1. Measured T, values and magnetization transfer parameters in human tissues in vivo at 0.1 T Tissue
Tlobs(ms)
Tl,, b-d
Muscle Liver Kidney cortex Spleen
258 212 470 420
130 + 164 t 342 + 332 ?
+ 12 t 12 k 71 -c 28
14 29 65 37
MTC 0.71 0.49 0.49 0.50
+ 0.01 ?I 0.06 c 0.04 2 0.07
Rwm(s-l) 5.5 3.1 1.5 1.5
” t ‘+
0.6 0.8 0.3 0.1
Note: All data expressed as mean + standard deviation. Muscle refers to striated muscle, kidney to renal cortex, T,,, = measured T, of tissue; T,_, = T, in the presence of saturation pulse; MTC = magnetization transfer contrast; R,, = observed magnetization transfer rate.
s-r). Between spleen and kidney cortex, no difference in R,, was noted. Other abdominal organs were not included in the study due to difficulties in reliably defining the rather small organs (e.g., pancreas) in a single slice with enough tissue material to avoid the partial volume effect. In kidneys, differentiation of cortex and medulla was impossible during some MT pulse lengths, and patient movement prevented the use of transferable ROIs (described previously). Therefore, kidney measurements consisted of renal cortex only. Fat was excluded because these young volunteers had very little fat and fat is known from the literature to have almost no MTC. The image quality varied between individuals and respiratory motion did cause some artifacts despite the precautions. In general, however, images were of satisfactory quality (Fig. 1) and in no individual they were unacceptable. DISCUSSION The significant, apparently tissue-dependent differences of signal intensities acquired with saturation pulses seem to reflect the MT phenomenon itself, rather than changes in T, or T2. Tissues that have similar appearances in conventional spin-echo or gradient-echo images may show contrast in MT, and this contrast is different from that in conventional imaging. The clinical application of MTC for tissue differentiation purposes has, however, been hampered by some limitations and has showed only modest success.” First, it seems that, between many neighboring tissues (e.g., kidney, spleen, and liver), differences in signal intensities using long saturation pulses (and thus achieving maximal h4TC) are practically nonexistent, as the MTC curves approach each other near the plateau. Second, it is often difficult to separate intensity changes caused by crossrelaxation and MT itself from changes induced by differences in T1 or T2 onto MTC.49” Both theoretically and in our experience the measurements of the rate of magnetization hold promise in tissue differentiation. Although the observed MT rate parameter
R,, in this study contains some direct saturation effects4
we believe that this observed parameter reflects the rate of magnetization transfer reasonably well. In our series, only striated (paraspinal) muscle had both R,, and MTC that were significantly greater than those of other tissues analyzed. Liver, kidney cortex, and spleen each achieved maximal signal reduction (MTC) of about 50% with no significant differences between them. On the other hand, the R, of liver was markedly higher than those of kidney and spleen. Although the optimal RF power and the magnitude of the resonance offset of the MT pulse can be calculated,798 the effect of field strength on MT parameters is, in our opinion, still a largely unsettled issue. It is usually believed that MTC becomes greater as the field strength increases.“~24 However, some investigators have concluded that the field strength of the MRI apparatus does not contribute to the magnitude of the MTC.22 On the other hand, it seems that R,, may be larger in lower fields 7.9,19but the exact nature of this relationship remains, to our knowledge, unknown. The rate of magnetization transfer in kidney cortex, for example, has been reported as 1 s-’ at 4.7 T,’ 1.4 s-l at 1.5 T,” and near 4 SC’ at 0.1 T.9 In our series, we obtained values of around 1.5 s-l, considerably lower than obtained at 0.1 T in vitro but still higher than those acquired at high fields. A comprehensive study of the impact of field strength on the MT parameters remains needed both in vitro and in vivo. Liver tissue showed considerable variation of MT parameters among the volunteers, with R,, ranging from 2.3 to 4.8 SK’ and MTC from 0.44 to 0.62. The reasons for this variance are still to be studied; however, it is possible that iron deposits of liver may cause the diversity of MTC. Another possible explanation is the role of diet, which may have an effect on MTC in liver.25 In this study, neither the amount of fat and connective tissue within liver was known in the volunteers (although they were all in the same age group, none of them were grossly overweight and their alcohol consumption was moderate) nor was their dietary intake prior to MR measurement controlled in any way.
Magnetic Resonance Imaging 0 Volume 14, Number 4, 1996
416
(4
(B)
m
(D)
Fig. 1. The effect of magnetizationtransferon MRI imagesof severalhumantissuesat 0.1 T in vivo. Gradient echo ( lOOO/ 20 ms) imageswith no MT pulse(A), MT pulseof 100 ms (B), 300 ms (C), and 7.50ms (D) .
The in vivo determination of MTC parameters of normal tissue is necessary for future studies of pathological conditions. We also believe that analysis of different age groups, nurturing conditions, and other determinants will be needed to further evaluate the MT effect in humans. In addition to a change in water content, many tissues may show signs of aging or other physiological processes having MRI appearances that are still unknown. In the abdominal area, despite some studies on liver pathology, 2223reports of MT imaging on parenchymal diseases of other organs have been scarce. However, some exciting possibilities in tissue characterization may be seen. Clinically important, for example, would be the evaluation of diseases, such as amyloidosis, that may diffusely infiltrate several organs without macroscopic tumor formation. The MRI appearance and relaxometry of such diseases is usually unspecific and difficult to interpret. Tissue-specific MT analysis could be of assistance and estimates of the amount of pathological tissue
could be made. Parametric images based on R,, or other MT parameters can also be created, and our preliminaty results have been encouraging. Despite its rather modest spatial resolution, lowfield imaging may offer some advantages over higher field strengths for MTC imaging. Because the longitudinal relaxation rates are shorter, shorter saturation pulses are needed for steady-state magnetization. Magnetization transfer rates may be, as mentioned above, greater, and power requirements lesser. Also, FDA guidelines may be broken when long off-resonance pulses are applied in high fields. Therefore, we think that low-field MRI also has its place in MT measurements. CONCLUSIONS 1. The results of the MT measurements reflect the properties of tissues and are measurable in the ab-
Saturationtransfer parameters0 S. KAJANDER
dominal and paraspinal area in vivo. Knowing the normal values of these parameters aids in future attempts in tissue differentiation. 2. The observed magnetization transfer rate, R,,, seems to be a better indicator of the MT phenomenon than MTC. REFERENCES
2. Edzes, H.T.; Samulski, E.T. The measurement of cross-
13. 14. 15.
16.
relaxation effects in proton NMR spin-lattice relaxation 3. 4.
5.
6. 7.
8.
9.
10.
11.
of water in biological systems: Hydrated collagen and muscle. J. Magn. Reson. 31:207-229; 1978. Edzes, H.T.; Samulski, E.T. Cross-relaxation and spin diffusion in the proton NMR of hydrated collagen. Nature (Lond.) 265521-523; 1977. Balaban, R.S.; Ceckler, T.L. Magnetization transfer contrast in magnetic resonance imaging. Magn. Reson. Q. 8:116-137; 1992. Koenig, S.H.; Brown, R.D. III. A molecular theory of relaxation and magnetization transfer: Application to cross-linked BSA, a model for tissue. Magn. Reson. Med. 6:685-695; 1993. Wolff, S.D.; Balaban, R.S. Magnetization transfer contrast (MTC) and tissue water proton relaxation in vivo. Magn. Reson. Med. 10:135-144; 1989. Eng, J.; Ceckler, T.L.; Balaban, R.S. Quantitative 1H magnetization transfer imaging in vivo. J. Magn. Reson. 17:304-314; 1991. Komu, M.E.S. Analysis of longitudinal relaxation rate constants from magnetization transfer MR images of human tissues at 0.1 T. Magn. Reson. Imag. 10:35-40; 1992. Niemi, P.T.; Komu, M.E.S.; Koskinen, S.K. Tissue specificity of low-field-strength magnetization transfer contrast imaging. Magn. Reson. Imag. 2:197-201; 1992. Bottomley, P.A.; Hardy, C.J.; Argensinger, R.E.; AllenMoore, G. A review of 1H NMR relaxation in pathology: Are Tl and T2 diagnostic? Med. Phys. 14:1-37; 1987. Wolff, S.D.; Balaban, R.S. Magnetization transfer imaging : Practical aspects and clinical applications. Radiology 192:593-599; 1994.
417
12. Wolff, S.D.; Balaban, R.S. Magnetization transfer con-
1. Koenig, S.H.; Bryant, R.G.; Hallenga, K.; Jacob, G.S.
Magnetic cross relaxation among protons in protein solutions. Biochemistry 17:4348-4358; 1978.
ETAL.
17.
18. 19.
20.
21.
22.
23.
24. 25.
trast: MR imaging of the knee. Radiology 179:133-137; 1991. Koskinen, S.K.; Komu, M.E. Low-field srength magnetization transfer contrast imaging of the pate&r cartilage. Acta Radiol. 34:124-126; 1993. Lundbom, N. Determination of magnetization transfer contrast in tissue: An MR imaging study of brain tumors. AJR 159:1279-1285; 1992. Ordridge, R.J.; Helpem, J.A.; Knight, R.A.; Qing, Z.X.; Welch, K.M. Investigation of cerebral ischemia using magnetization transfer contrast (MTC) MR imaging. Magn. Reson. Imag. 9:895-902; 1991. Balaban, R.S.; Chesnick, S.; Hedges, K.; Samaha, F.; Heineman, F.W. Magnetization transfer contrast in MR imaging of the heart. Radiology 180:671-675; 1991. Mittl, R.L., Jr.; Gomori, J.M.; Schnall, M.D.; Holland, G.A.; Grossman, R.I.; Atlas, S.W. Magnetization transfer effects in MR imaging of in vivo intracranial hemorrhage. AJNR 14:881-891; 1993. Gomori, J.M.; Grossman, R.I.; Asakura, T.; et al. An in vitro study of magnetization transfer and relaxation rates of hematoma. AJNR 14:871-880; 1993. Dixon, W.T.; Engels, H.; Castillo, M.; Sardashti, M. Incidental magnetization transfer contrast in standard multislice imaging. Magn. Reson. Imag. 8:417-422, 1990. Hajnal, J.V.; Baudouin, C.J.; Oatridge, A.; Young, I.R.; Bydder, G.M. Design and implementation of magnetization transfer pulse sequences for clinical use. J. Comput. Assist. Tomogr. 16:7-18; 1992. Grad, J.; Mendelson, D.; Hyder, F.; Bryant, R.G. Direct measurements of longitudinal relaxation and magnetization transfer in heterogenous systems. J. Magn. Reson. 86:416-419; 1990. Outwater, E.; Schnall, M.D.; Braitman, L.E.; Dinsmore, B.J.; Kressel, H.Y. Magnetization transfer of hepatic lesions: Evaluation of a novel contrast technique in the abdomen. Radiology 182:535-540; 1992. Kahn, C.E., Jr.; Perera, S.D.; Sepponen, R.E.; Tanttu, J.I.; Tierala, E.K.; Lipton, M.J. Magnetization transfer imaging of the abdomen at O.lT: Detection of hepatic neoplasms. Magn. Reson. Imag. 11:67-71; 1993. Ceckler, T.L.; Balaban, R.S. Field dispersion in watermacromolecular proton magnetization transfer. J. Magn. Reson. Series B 105:242-248; 1994. Salo, S.; Alanen, A.; Komu, M.; Kallio, T.; Bondestam, S. Effect of fasting and food intake on magnetization transfer of human liver tissue. (submitted).