NMR Biomed.2006;19: 527–534
Published online 5 April 2006 in Wiley InterScience (www.interscience.wiley.com). DOI:10.1002/nbm.1014
Single-scan quantitative T
2* methods with susceptibility artifact reduction
Florence Franconi,1* Pierre Mowat,2Laurent Lemaire,2Pascal Richomme1and Jean-Jacques Le Jeune2
1Service Commun d’Analyses Spectroscopiques, UFR Sciences, 2 boulevard Lavoisier, 49045 Angers cedex, France
2INSERM U646, Universite´ d’Angers, F-49100 Angers, France
Received 11 July 2005; Revised 18 October 2005; Accepted 21 November 2005
ABSTRACT: Two imaging methods, MSSAVE (Multiple echo SubSlice AVEraging imaging), based on sub-slice averaging and MGESEPI (Multiple echo Gradient-Echo Slice-Excitation Profile Imaging), based on over-sampling in the slice direction, are proposed for single-scan quantitativeT2* evaluation with susceptibility artifact compensation. Their potentials in terms of sensitivity, minimum performance time, susceptibility artifact reduction and T2* quantitation quality, were compared with existing single-scan methods such as classical FLASH two- or three-dimensional orz-shimmed methods bothin vitroandin vivoin normal rat brain. MGESEPI offered good qualityT2* maps nearly free of artifacts but required a long acquisition time. MSSAVE was faster, but at the expense of reduced artifact compensation and the achievable T2* quantitation quality. Copyright#2006 John Wiley & Sons, Ltd.
KEYWORDS: susceptibility artifact compensation; single-scan quantitativeT2* magnetic resonance imaging; high-field magnetic resonance imaging
INTRODUCTION
T2*-weighted imaging, owing to its sensitivity to mag- netic susceptibility, is widely used in functional MRI (1), iron deposition pathologies (2), super-paramagnetic con- trast-enhanced imaging (3), cellular imaging (4) and biological microstructure investigation (5).
However, T2*-weighted sequences, based on gradient- echo acquisition, are sensitive not only to microscopic magnetic field inhomogeneities due to the inherent differ- ences between tissues or differences induced by contrast agents, but also to macroscopic field inhomogeneities which arise from magnet imperfections, imperfect shim- ming and air–tissue interfaces. Both global and local field inhomogeneities lead to signal decay. Macroscopic field inhomogeneities that appear as artifacts on the image mask the desired T2* information content of microscopic field inhomogeneities. Enhancing the contrast while minimizing artifacts is difficult since both have the same common physical mechanism. In the past, efforts have been made to remove macroscopic susceptibility artifacts while preser- ving sensitivity toT2* contrast. Ro and Cho proposed the
use of tailored radiofrequency (RF) pulses to map the susceptibility effect (6). Assuming that the greatest effect is related to the largest dimensions, i.e. the slice selection direction, then thin slice imaging (7), sub-slice averaging imaging (SSAVE) (8) and three-dimensional (3D) gradient echo sequence (9) were proposed to limit the susceptibility artifact. Thez-shim methods, originally proposed by Frahm et al. (10) and later extended by Ordidge et al. (2) and Constable (11), compensate susceptibility artifacts using a series of compensation gradients applied along the slice-selection direction. Alternatively, an over-sampled 3D method in the slice direction was introduced by Yang et al. (12) and named GESEPI (Gradient-Echo Slice Excitation Profile Imaging) whereby images are acquired with an incremental slice refocusing gradient offset and integrated with a Fourier transform.T2*-weighted imaging susceptibility artifact compensation methods were com- pared by Wadghiriet al.(8) and Merboldtet al.(13).
Although T2* contrast is often sufficient to localize a contrast agent tagging area or to identify functional changes, quantitative T2* information significantly en- hances the sensitivity of the method, the variation inT2* being more stable and significant than the signal variation measured at one particular echo time. Furthermore, absolute determination of T2* allows the detection of slight variations, comparison from scan to scan or from MRI facility to MRI facility and is widely used for vessel size imaging (14), biological microstructure exploration (5) and iron overload/defect (15) evaluation, for which besides the necessity for quantitative T2* imaging and more than temporal resolution, susceptibility artifact-free
*Correspondence to:F. Franconi, Service Commun d’Analyses Spec- troscopiques, UFR Sciences, 2 boulevard Lavoisier, 49045 Angers cedex, France. E-mail: fl[email protected]
Abbreviations used: 2D, two-dimensional; 3D, three-dimensional;
FLASH, fast low-angle shot; FOV, field-of-view; FT, Fourier transfor- mation; GESEPI, gradient-echo slice excitation profile imaging;IE, inter-echo time; MGESEPI, multiple-echo gradient-echo slice excitation profile imaging; MSSAVE, multiple-echo sub-slice averaging imaging;
Nex, number of experiments; RF, radiofrequency; SSAVE, sub-slice averaging imaging;Tim, imaging time; WILD, Wildz-shimmed method.
methods are needed. Indeed, susceptibility artifacts are exacerbated with the high magnetic field that is required to perform high-resolution scans. In these conditions, high-quality, rapid imaging such as echo planar acquisi- tion could be difficult to achieve (16). Therefore, single- scan multiple-gradient-echo approaches, providing in one shot the compensated data points for an accurate estimate ofT2*, are of great interest. Wildet al.proposed the first implementation of a single-scan method ofT2* mapping with susceptibility-gradient compensation (17).
The method extends Ordidgeet al.’s sequence (2) in order to perform compensation and signal decay envelopes in a single scan. Posseet al. (18) also proposed a single-shot T2* mapping method based on volume-of-interest-specific compensation gradients integrated into a multi-echo echo-planar sequence. However, this method, designed for functional MRI, trades off the spatial resolution and T2* determination accuracy against temporal resolution.
Here, we propose two single-scan methods ofT2* map- ping with susceptibility-gradient compensation, one being an extension of the GESEPI sequence to a multi- echo method referred to as Multi-Gradient-Echo Slice Excitation Profile Imaging (MGESEPI). Similarly, a multiple echo SSAVE extended from the method pro- posed by Wadghiri et al. (8), called MSSAVE, is also proposed. Potentials of the different dephasing artifact reduction single-scan quantitative T2* methods were evaluated and compared in vitro and in vivo with the standard multi-echo 2D and 3D Fast Low-Angle SHot sequence (FLASH) (19) and thez-shim sequence devel- oped by Wildet al.(17).
THEORY
All the methods presented allowed single-scan quantita- tive evaluation of T2* by the use of multiple-gradient- echo sequences. The inherent sensitivity of gradient-echo to magnetic field inhomogeneities results in susceptibility artifacts, most markedly in the largest voxel dimension, usually in the slice-encoding direction. Local suscept- ibility gradients G are responsible for distortions and spin dephasing that induce signal losses. At an echo time TEand for a positionzin the slice encoding direction, the magnetization distribution in the presence of a local susceptibility gradientGis
Mðz;TEÞ ¼M0ðzÞexpðiGTEzÞ ð1Þ where is the gyromagnetic ratio and M0(z) is the magnetization density including all others factors such as spin density and relaxation effects. A corresponding NMR signal is provided by using the Fourier transforma- tion (FT) of eqn (1):
Sðkz;TEÞ ¼FTfM0ðzÞ:expðiGTEzÞg ð2Þ In case of linearG, the maximum signal in the Fourier space is shifted from kz¼0 to kz0¼GTE. For 2D
imaging, the slice is defined by selective excitation. If the slice profile is assumed to be an ideal rectangle, the dephasing signals in voxels therefore correspond to a sinc function profile:
Sðkz;TEÞ ¼mz0sinc½ðkzkz0Þz0 ð3Þ where m represents magnetization and z0 the slice thickness.
As the signal is collected only at the center ofk-space, substantial signal loss occurs asGincreases:
Sðkz ¼0;TEÞ ¼mz0sinc½ð=2ÞGTEz0 ð4Þ Conversely, for 3D imaging, the slice is defined by phase encoding. Kz-space is sampled from kzmax to þkzmaxwithkzmaxequal toNp=z0,Npbeing the number of phase encoding steps. Signal loss occurs forkzmax<kz0
since the maximum of the signal is shifted out of the sampling window. Figure 1 shows the signal response to local susceptibility gradients calculated from this model.
Two main methods exist to reduce susceptibility artifacts. On the one hand, artifact outcome can be limited by either reducing the slice thickness or phase encoding in the slice direction. On the other hand, artifacts can be compensated with z-shimmed-based methods where a compensation gradient lobe offset, Gc, is applied for the duration tc, rewinding a specific amount of phase dispersion created by G such as Gctc¼GTE. The acquisition point is then shifted from the center ofk-space tokz0. However, as the phase dispersion may vary from one voxel to another, a series of compensation gradients has to be used in order to compensate for each particular level of susceptibility- induced gradient.
Four artifact-compensated methods are described and compared with the reference, the multi-echo FLASH-2D sequence.
Figure 1. Signal response to local susceptibility gradients Gfor FLASH-2D (–), MSSAVE (–), WILD (. . .) and 3D (- -).
Simulation assuming a slice thickness of 3 mm, a correction factor N¼3, an echo time TE¼10 ms and G varying between 2.3 mT/m
Multi-echo FLASH-3D
A phase-encoding gradient is applied in the slice direc- tion to separate the slice into thin sub-slices. A final set of images of a multi-echo thick slice is obtained by adding the sub-slices using an echo-by-echo basis.
MSSAVE
Several multi-echo FLASH 2D thin sub-slices are ac- quired within the same repetition time. The final set of images of a multi-echo thick slice is formed by averaging sub-slices using an echo-by-echo basis.
MGESEPI
An incremental slice refocusing gradient offset is applied in the slice direction. The maximum compensation gradient offset needed,Gcmax, is determined by the maximum value ofGencountered in the image plane,Gmax, and by the echo timing. Every dispersion induced, even on the last echo of the echo train, must be refocused or, in other words, Gcmaxtcmust be higher thanGmaxTE(last echo)in order to keep the echo peak within the acquisition window. The acquired data set is then reconstructed with a 3D complex Fourier transformation. The final image is obtained by adding, on an echo-by-echo basis, the magnitude images containing appreciable signals within the excited slice.
WILD
A succession of triple gradient-echo cycles throughout the transverse decay envelope is acquired with successive slice-refocusing gradient increments. Effective back- ground gradient compensation then requires a combina- tion of images as the square root of the sum of the squares of the pixel magnitudes.
EXPERIMENTAL MR imaging
MR imaging experiments were performed on a Bruker Avance DRX 300 system (Bruker Biospin, Wissembourg, France) equipped with a 150 mm vertical super-wide- bore magnet operating at 7 T, an 84 mm i.d. shielded gradient set capable of 144 mT/m maximum gradient strength and a standard 64 mm diameter birdcage reso- nator, operating on a Paravision (version 2.1.1) software platform (Bruker Biospin).
In vitroexperiments
In order to compare performances, T2* mapping was performed with the same imaging time of 6 min 55 s, the
same spanning of the T2* decay curve and the same correction factor on a 26 mm i.d. phantom filled with 0.1 mM NiCl20.1 mM MnCl2-doped solution (T1¼ 840 ms, T2¼53 ms). The imaging parameters were an effectiveTRof 540 ms, a 4 ms three-lobe sinc RF pulse, a flip angle of 56, a field-of-view (FOV) of 30 mm and a matrix size of 128. The T2* decay curve was sampled with a first echo time at 6.2 ms, an effective inter-echo time, IE, of 15 ms and six data samples, except for the WILD sequence, where 18 echoes were needed (TE¼6.2 ms andIE¼5 ms) as each triple-echo compen- sation cycle was used to calculate each compensated image. For 3D methods, six phase-encoding steps were applied, resulting in an over-sampling of 2. The same correction factor of N¼3 was used for all methods, calculated as the ratio of the final slice thickness (3 mm) to the sub-slice thickness for MSSAVE, FLASH 3D and MGESEPI and corresponding to the amount of rephasing gradient amplitude for WILD. In the last sequence, the refocusing factor,", was optimized experi- mentally to its best value of 0.2. These parameters were taken as they represented one of the best compromises achievable for all sequences. The number of experiments (Nex) was chosen to equalize the imaging time, 6 for 2D methods and 1 for 3D methods.
The influence of the T2* decay curve data sampling was then explored by acquiring the same data sets but with 12 echo images and an effective inter-echo time of 7.5 ms. The influence of over-sampling was also explored for 3D methods by reducing the over-sampling factor n from 2 to 1 while doublingNexto preserve imaging time.
Finally, the influence of the correction factor was ex- plored in a separate experiment (i.e. a different inhomo- geneity pattern) by acquiring the same data sets with N¼3 and 6. The imaging time was then doubled to 13 min 50 s to accommodate 3D scans,Nexbeing either 1 (N¼6) or 2 (N¼3) for 3D imaging and 12 for 2D imaging.
Theoretical sensitivities,, and minimum performance times,Tp, were calculated as expressed by Wadghiriet al.
(8) based on the definition of Brunner and Ernst (20).
Experimental sensitivities were determined as the signal- to-noise ratio of the first echo image divided by the square root of the imaging time, normalized to 1 for FLASH-2D.
Susceptibility artifact reduction was evaluated by the standard deviation of theT2* map intensity on the entire phantom, expressed as a percentage. If the artifact is well compensated, the standard deviation of theT2* intensity of the homogeneous phantom should be small. T2* quantitation quality was evaluated by the standard devia- tion of theT2* fit, expressed as a percentage.
In vivoexperiments
Animal care and use were in strict accordance with the regulations of the French Ministry of Agriculture.
2
Sprague–Dawley rat axial, coronal and sagittal brainT2* maps were obtained for each of the methods with the following parameters: FOV of 40 mm, matrix size of 128, effective TE¼3.1 ms and effective inter-echo time of 6.4 ms, six echoes,TR¼250ms, a 1.4 ms three-lobe sinc RF pulse and a flip angle of 29, a final slice thickness of 3 mm and a correction factor of 3. 3D methods were performed with over-sampling:n¼2 for MGESEPI and n¼4 for the FLASH-3D method (determined by the minimum slab thickness authorized, 12 mm). The ima- ging time was 3 min 30 s (except for FLASH-3D, where the imaging time was 7 min.).Nexwas 6 for 2D methods and 1 for 3D methods. The refocusing factor for the WILD sequence,", was optimized experimentally to its best value of 0.2.
RESULTS
For each method, in vitro performance parameters are summarized in Table 1. Compared with the FLASH-2D reference method, compensation induced a loss of sensi- tivity, 2D methods offering a minimum performance time much shorter than for 3D methods.
Figure 2 shows theT2* maps for anin vitroacquisition and the corresponding measures of the percentage stan- dard deviation ofT2* map intensity on the entire phantom (representing the susceptibility artifact reduction quality) versusT2*. The symbol diameters are proportional to the T2* quantitation quality measured by the percentage standard deviation ofT2* fit. A comparison of methods is illustrated in Fig. 2(a). The FLASH-2DT2*-map image displays massive susceptibility artifacts. Artifacts of limited intensity are nearly completely removed, what- ever compensation method is used, conversely to highly artifacted areas where the signal is not completely re- covered. Artifact level improvement of compensation methods compared with FLASH 2D is demonstrated by the decrease in the percentage standard deviation ofT2* map intensity on the entire phantom and by the increase
inT2* values. The influence of increasing data sampling from 6 to 12 along the T2* decay curve is illustrated in Fig. 2(b). The difference between T2* values acquired with 6 and 12 TEs is kept low (under 1.6%). However, increasing the number of data samples increases the
‘goodness of fit’, since the average decrease in the standard deviation percentage on the T2* fit is 27% for the four methods. The influence of the over-sampling factor for the 3D scan is illustrated in Fig. 2(c). The use of a sampling factor of 1 induces large artifacts especially with the MGESEPI acquisition. The influence of the correction factor is illustrated in Fig. 2(d). When N is increased from 3 to 6, the standard deviation decreases from 7.8 to 4.2% for the MSSAVE method, from 4.9 to 4.5% for the FLASH 3D method and from 7.91 to 2.2%
for the MGESEPI method whereas the measured T2* values increase for all the methods.
A comparison ofin vivorat axial, coronal and sagittal brain T2* maps is provided in Fig. 3. The percentage standard deviation ofT2* map intensity versusT2* in the whole brain is also presented. The symbol diameters are proportional to the percentage standard deviation of T2* fit. Errors bars represent the standard deviation on the three values measured axially, coronally and sagittally in the whole brain. The FLASH-2D images showed regions of massive susceptibility artifact in the brain. Even with a small correction factor of 3, all the compensation tech- niques significantly reduce the artifact level. All com- pensation methods allow the correction of artifacts of limited intensity such as the close skull artifacts, but present poor compensation for highly artifacted areas such as the areas close to the ear canals. The compensa- tion effectiveness of MSSAVE and MGESEPI generate good-qualityT2* maps.
DISCUSSION
We have presented two methods, MSSAVE, based on sub-slice averaging, and MGESEPI, based on 3D Table 1. In vitroperformance parameters for single-scanT2* quantitative methods: correction factorN, theoretical (g¼1=p
N) and experimental (g¼SNR=p
Tim) sensitivities and minimum performance time are given for the five single-scanT2* mapping methods
Method In vitroperformance parameter
Theoretical Experimental sensitivitya,c Correction factor Minimum performance timea Sensitivitya
FLASH-2D 1 1 1.00 1.00
FLASH-3D 3 nN¼6b 0.58 0.39
MSSAVE 3 1 0.58 0.29
MGESEPI 3 nN¼6b 0.58 0.28
WILD 3 1 0.58 0.50
aMinimum performance time and sensitivity were normalized to 1 for the FLASH-2D acquisition.
bThe MGESEPI and FLASH-3D over-sampling factor,n, was 2.
cExperimental sensitivity was calculated as¼SNR=p
Tim. The imaging time (Tim) was 415 s for all sequences. The signal-to-noise ratio (SNR) was measured on the first echo by the ratio of the signal mean value in a region of interest located in the phantom to the standard deviation of the noise.
Figure 2. In vitro comparison of single-scan T2* quantitative methods: T2* maps and the corresponding measures of the percentage of standard deviation of T2* map intensity on the entire phantom (representing the susceptibility artifact reduction quality) versusT2*. The better the artifact is compensated, the smaller is the standard deviation of theT2* intensity of the phantom. The symbol diameters are proportional to the percentage standard deviation ofT2* fit, representing theT2* quantitation quality on a 0.1 mMNiCl20.1 mMMnCl2solution. Acquisition parameters were TR¼540 ms, 4 ms three-lobe sinc RF pulse (56), FOV¼30 mm, matrix¼128, slice thick- ness¼3 mm, "¼0.2. (a)Comparison of methods: six echo times (TE¼6.2 ms, IE¼15 ms), N¼3, n¼2, Nex¼6 (2D) or 1 (3D).Tim¼6 min 55 s. (b)Influence of data sampling(number ofTEs):IE¼15 ms (6TEs) or 7.5 ms (12TEs).
Tim¼6 min 55 s. (c)Influence of the over-sampling factor for 3D scan,n:Nex¼1 (n¼2) or 2 (n¼1).Tim¼6 min 55 s. Spreading of the values required the use of a vertical logarithmic scale and does not allow representation of the percentage standard deviation of theT2* fit as symbol diameter (FLASH 3D:n¼1, 2.5% andn¼2, 4.6%; MGESEPI:
n¼1, 0.9% and n¼2, 187.6%). The corresponding FLASH 2D T2* map was also displayed to visualize the inhomogeneity pattern. (d)Influence of the correction factor,N:Nex¼1 (N¼6) or 2 (N¼3) for 3D imaging and 12 for 2D imaging. Tim¼13 min 50 s. The corresponding FLASH 2D T2* map was also displayed to visualize the inhomogeneity pattern
2
over-sampling in the slice direction, for single-scan quantitative T2* evaluation with susceptibility artifact compensation. Their potentials were compared with those of existing methods such as the classical multi- echo FLASH 2D or 3D method and the z-shimmed method described by Wildet al.(17).
In 2002, Wildet al.(17) proposed the first implemen- tation of a single-scan method of T2* mapping with susceptibility compensation. The sequence was based on the acquisition compensation and signal decay envel- ope acquired in a single scan. However, although MGE- SEPI and MSSAVE methods also allow quantitative assessment ofT2* with susceptibility artifact in a single
scan, all the echoes from the echo train may be used to sample the relaxivity curve better since the compensation scheme is performed by phase-encoding gradients in the slice direction and thin-slice averaging, respectively.
Sensitivity
Standard FLASH-2D offers the best sensitivity for single- scan T2* mapping but can be spoiled by susceptibility artifacts. All the compensation methods described in this paper reduce the artifact level to the detriment of sensi- tivity. As pointed out by Wadghiri et al. (8), in the absence of inhomogeneity, 2D multi-slicing methods Figure 3. In vivocomparison of single-scanT2* quantitative methods: axial, coronal and sagittal rat brainT2* maps and the corresponding measures of the percentage standard deviation of T2* map intensity (representing the susceptibility artifact reduction quality) versus T2* in the whole brain. The better the artifact is compensated, the smaller is the standard deviation of theT2* intensity. The symbol diameters are proportional to the percentage standard deviation ofT2* fit representing theT2* quantitation quality
such as the MSSAVE did not shown much reduced sensitivity compared with 3D methods. Differences ob- served experimentally between sensitivities were mainly due to processing choices. All the compensation methods required a combination of images to form composite compensated images. Combination by summation may reduce sensitivity, since noise from the images is added in the composite image such as in the MSSAVE method, whereas summation of squares magnitude, as in the WILD method, may increase sensitivity, since the main contribution to the composite image comes from images with substantial signals.
Efficacy in reducing susceptibility artifacts
The signal response to susceptibility gradients is different between 2D methods, where the slice is defined by selective excitation, and 3D methods, where the slice is defined by phase encoding. With 3D methods, no signal is lost until a cut-off gradient value is reached; above this value, all signals are lost, the transition occurring abruptly. The difference between FLASH-3D and MGE- SEPI is that the cut-off gradient value in MGESEPI is optimized to compensate for intravoxel phase dispersion, whereas in FLASH-3D, the value is determined by the chosen slice thickness. Conversely, with 2D methods, the signal decreases slowly with increasing inhomogeneity (sinc function). For WILD, each of the three echoes presents a signal response profile as a sinc function and is aimed to compensate for some specific values of susceptibility gradients and less for others. Well-opti- mized 3D techniques may offer better compensation than 2D methods. However, 2D methods are less prone to misadjustment, since signal reduction evolves only slowly with a gradient increase in inhomogeneity.
Minimum performance time
3D techniques require a long acquisition time since each phase-encoding step has to be repeatedNntimes. The samplingNmust be fine enough to define well the shifted echo peak. Accurate sampling determines the compo- site image quality. An over-sampling factor of at least 2 prevents, in most cases, potential aliasing into the image of the frequency shifts exceeding the slice band- width. However, in some cases, distribution of suscept- ibility gradients within the slice would require the use of a larger over-sampling factor. Over-sampling also bene- fits FLASH-3D sequences by limiting slab boundary artifacts since only central slices are used to form the composite image. These slab boundary artifacts are caused by imperfect RF profiles or small geometric discrepancies between the selected slab and the encoded slab, caused by local field inhomogeneity. 2D techniques, as opposed to 3D methods, offer the advantage of speed.
However, the minimumTRis driven by the necessity to accommodate the multi sub-slices (for MSSAVE) or multi-echo cycle (for WILD) determined by the correc- tion factor and/or the multi-echo pattern. These restric- tions in the choice of timing parameters somewhat limit the speed advantages of 2D methods.
T2* relaxivity curve spanning
The quality of the T2* decay data may be characterized by the echo time, the inter-echo time, the number of echo images spanning the decay curve and the signal-to-noise ratio. Compared with 3D techniques, 2D techniques suffer from limitations inT2* relaxivity curve spanning.
For MSSAVE, narrow slices are obtained at the price of strong slice selection gradients or long RF pulses. As gradient strength is often limited by system characteris- tics, thin slices require increased RF pulse duration and therefore increased minimum echo time. For the WILD sequence, since each three successive echo-images are combined to form a composite image, effective inter- echo spacing is limited by the switching of the gradient and useable reception bandwidth. In the experimental part of our work, the spanning of the relaxivity curve was kept low in order to fit with the WILD performance sequence. 3D method performance may easily be in- creased without increasing the imaging time, by reducing TEorIEand/or increasing the number of echo images.
In multiple-echo trains, possible distortion between odd and even echoes may occur. For all the sequences except WILD, combinations are performed on an echo- by-echo basis, so these distortions do not affect the composite images. Distortions may only be a problem when the echo train is used to calculateT2* maps (but it is always possible to make a fit on only odd or even echo series). However, with the WILD method, composite images are calculated from three echoes, odd and even echoes being mixed, hence possibly reducing the quality of resulting composite images.
Despite the limited number of quantitativeT2* studies performed on rat brain at high magnetic field values, the in vivovalues measured in our work are in the range of previously published values. Indeed, the rat brain T2* value was measured as 32 ms at 7T(21). These values increased as the magnetic field strength decreased, reach- ing 43 ms at 4.7T(21) and 45.5 ms at 2.35T(22). Even though none of these measures were performed with susceptibility artifact-compensated methods, values were obtained in the region of interest in the cortex area with fairly homogeneous T2* values (outside the outer layers of the cortex).
CONCLUSION
This study demonstrates the potential of single-scan susceptibility artifact-compensated T2* mapping methods
2
in vitroand for normal rat brain acquisitions. The choice of the appropriate sequence can be determined by the range of T2* studied, the severity of the susceptibility artifacts to correct and the available imaging time. MGESEPI may offer good T2* maps nearly free of artifacts but at the expense of a long acquisition time. WILD and MSSAVE methods have the advantage of being faster, but both artifact compensation and theT2* decay data quality are reduced.
Acknowledgement
The authors would like to thank D. Rees for editing the English text.
REFERENCES
1. Edelman RR, Mattle HP, Atkinson DJ, Hill T, Finn JP, Mayman C, Ronthal M, Hoogewoud HM, Kleefield J. Cerebral blood flow:
assessment with dynamic contrast-enhancedT2*-weighted MR imaging at 1.5 T.Radiology1990;176: 211–220.
2. Ordidge RJ, Gorell JM, Deniau JC, Knight RA, Helpern JA.
Assessment of relative brain iron concentrations using T2- weighted andT2*-weighted MRI at 3 Tesla.Magn. Reson. Med.
1994;32: 335–341.
3. Chapon C, Franconi F, Lemaire L, Marescaux L, Legras P, Saint- Andre JP, Denizot B, Le Jeune JJ. High field magnetic resonance imaging evaluation of superparamagnetic iron oxide nanoparticles in a permanent rat myocardial infarction.Invest. Radiol.2003;38: 141–146.
4. Bulte JW, Duncan ID, Frank JA. In vivo magnetic resonance tracking of magnetically labeled cells after transplantation. J.
Cereb. Blood Flow Metab.2002;22: 899–907.
5. Kohler S, Hiller KH, Waller C, Bauer WR, Haase A, Jakob PM.
Investigation of the microstructure of the isolated rat heart: a comparison between T2* and diffusion-weighted MRI. Magn.
Reson. Med.2003;50: 1144–1150.
6. Ro YM, Cho ZH. A new frontier of blood imaging using susceptibility effect and tailored RF pulses.Magn. Reson. Med.
1992;28: 237–248.
7. Young IR, Cox IJ, Bryant DJ, Bydder GM. The benefits of increasing spatial resolution as a means of reducing artifacts due
to field inhomogeneities.Magn. Reson. Imaging1988;6: 585–
590.
8. Wadghiri YZ, Johnson G, Turnbull DH. Sensitivity and perfor- mance time in MRI dephasing artifact reduction methods.Magn.
Reson. Med.2001;45: 470–476.
9. Haacke EM, Tkach JA, Parrish TB. Reduction of T2* dephasing in gradient field-echo imaging.Radiology1989;170: 457–462.
10. Frahm J, Merboldt KD, Hanicke W. Direct FLASH MR imaging of magnetic field inhomogeneities by gradient compensation.
Magn. Reson. Med.1988;6: 474–480.
11. Constable RT. Functional MR imaging using gradient-echo echo- planar imaging in the presence of large static field inhomogene- ities.J. Magn. Reson. Imaging1995;5: 746–752.
12. Yang QX, Williams GD, Demeure RJ, Mosher TJ, Smith MB.
Removal of local field gradient artifacts in T2*-weighted images at high fields by gradient-echo slice excitation profile imaging.
Magn. Reson. Med.1998;39: 402–409.
13. Merboldt KD, Finsterbusch J, Frahm J. Reducing inhomogeneity artifacts in functional MRI of human brain activation-thin sections vs gradient compensation.J. Magn. Reson.2000;145: 184–191.
14. Tropres I, Grimault S, Vaeth A, Grillon E, Julien C, Payen JF, Lamalle L, Decorps M. Vessel size imaging.Magn. Reson. Med.
2001;45: 397–408.
15. Zywicke HA, van Gelderen P, Connor JR, Burdo JR, Garrick MD, Dolan KG, Frank JA, Bulte JW. Microscopic R2* mapping of reduced brain iron in the Belgrade rat.Ann. Neurol.2002; 52: 102–105.
16. Grieve SM, Blamire AM, Styles P. The effect of bulk suscept- ibility on murine snapshot imaging at 7.0 T: a comparison of snapshot imaging techniques.Magn. Reson. Med.2000;43: 747–
755.
17. Wild JM, Martin WR, Allen PS. Multiple gradient echo sequence optimized for rapid, single-scan mapping of R2* at high B0.Magn.
Reson. Med.2002;48: 867–876.
18. Posse S, Shen Z, Kiselev V, Kemna LJ. Single-shot T2* mapping with 3D compensation of local susceptibility gradients in multiple regions.Neuroimage2003;18: 390–400.
19. Haase A, Frahm J, Matthaei D, Hanicke W, Merboldt KD. FLASH imaging—rapid NMR imaging using low flip-angle pulses. J.
Magn. Reson.1986;67: 258–266.
20. Brunner P, Ernst RR. Sensitivity and performance time in NMR imaging. J. Magn. Reson. 1979;33: 83–106.
21. Grune M, Pillekamp F, Schwindt W, Hoehn M. Gradient echo time dependence and quantitative parameter maps for somatosensory activation in rats at 7 T.Magn. Reson. Med.1999;42: 118–126.
22. Howseman AM, Thomas DL, Pell GS, Williams SR, Ordidge RJ.
RapidT2* mapping using interleaved echo planar imaging.Magn.
Reson. Med.1999;41: 368–374.
