Origin and Minimization of Residual Motion-Related Artifacts in Navigator-Corrected Segmented Diffusion-Weighted EPI of the Human Brain

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1 Magnetic Resonance in Medicine 47: (2002) DOI /mrm Origin and Minimization of Residual Motion-Related Artifacts in Navigator-Corrected Segmented Diffusion-Weighted EPI of the Human Brain Hangyi Jiang, 1,2 Xavier Golay, 1,2 Peter C.M. van Zijl, 1,2 and Susumu Mori 1,2 * Motion sensitivity in diffusion-weighted imaging (DWI) can be effectively suppressed using single-shot echo-planar imaging (EPI). However, segmented (multishot) EPI is often used to increase resolution and reduce spatial distortions, which in turn increases susceptibility to brain motion. The sources of these residual motion artifacts in navigator-echo-corrected segmented EPI images of the brain were investigated. The results indicate that the dominant source of these artifacts is cardiac pulsation with occasional involuntary movement of the subject. The relationship between the cardiac cycle and motion artifacts shows that optimum timing for the data acquisition is possible. In addition it is shown that the effects of involuntary motion can be removed by swapping k-space data between redundant datasets. Magn Reson Med 47: , Wiley-Liss, Inc. Key words: diffusion-weighted imaging; navigator correction; brain motion; cardiac pulsation; artifacts Diffusion-weighted imaging (DWI) is a unique technique that provides images sensitive to molecular motion (diffusion) of water molecules (1,2). The recent advance of diffusion tensor imaging (DTI) (3) offers the exciting possibility of mapping dominant fiber orientations in the brain by using anisotropy-based color coding (4) or actual fiber tracking (5,6). However, the large gradients required for diffusion weighting make this approach very sensitive to motion, which may induce large phase changes in the MR signals and result in serious image ghosting artifacts. Several methods are now available to reduce these problems (7). The advance of these techniques has been stimulated by the availability of single-shot echo-planar imaging (EPI). However, although this approach can generally avoid motion artifacts (2,8), the resulting images have limited resolution and often suffer from severe distortions due to the presence of susceptibility artifacts. The consequent anatomical uncertainties hamper the combination of DTI approaches with conventional image approaches. Navigator-corrected multishot EPI (9 11), on the other hand, can be used to improve image resolution and reduce susceptibility artifacts, but patient motion may lead to residual small, incoherent phase changes between segments, again 1 Department of Radiology, Division of MRI Research, Johns Hopkins University School of Medicine, Baltimore, Maryland. 2 F.M. Kirby Research Center for Functional Brain Imaging, Kennedy Krieger Institute, Baltimore, Maryland. Grant sponsor: Whitaker Foundation; Grant sponsor: National Institutes of Health; Grant number: RO1 HD37931; Grant sponsor: NCRR; Grant number: P41 RR *Correspondence to: Susumu Mori, Ph.D., Department of Radiology, 217 Traylor Bldg., 720 Rutland Ave., Johns Hopkins University School of Medicine, Baltimore, MD susumu@mri.jhu.edu Received 25 September 2001; revised 14 November 2001; accepted 18 November Wiley-Liss, Inc. 818 resulting in severe image artifacts. Even the use of electrocardiographic (ECG) triggering or cardiac gating to synchronize the data acquisitions in combination with the simple readout (or 1D) navigator-echo approach cannot suppress all motion-related phase errors, and residual artifacts are often found in the corrected images. The overall aim of this study was to understand the dominant source of these residual artifacts in cardiactriggered navigator-corrected diffusion-weighted EPI (DW- EPI), and to design data acquisition and postprocessing schemes that minimize the motion-related artifacts when using segmented EPI readout techniques. METHODS MRI Data Acquisition Experiments were performed using a 1.5 T Philips Gyroscan NT system (Philips Medical Systems, The Netherlands) with standard body-coil excitation and birdcage coil reception. DWI was accomplished using multislice segmented EPI sequences, with cardiac triggering and navigator-echo phase monitoring. Images were acquired coronally (four segmented shots) or axially (three segmented shots) with thicknesses of 3 mm and in-plane resolution of 2 mm for both orientations, respectively. Seventeen echoes were acquired per excitation. A diffusion-encoding gradient pair was applied on both sides of a 180 radiofrequency (RF) pulse, directly followed by the acquisition of a navigator echo to monitor motion-induced phase errors in the segmented EPI readout. Subsequently, a second 180 pulse was applied, followed by the segmented EPI readout. Diffusion encoding (b 600 s/mm 2 ) was applied along six independent axes, including X (readout encoding), Y (phase encoding), Z (slice selection), and XY, XZ, and YZ. Additionally, a reference-image without diffusion weighting was also recorded. This data acquisition was repeated two times (or more if necessary) to produce higher signalto-noise ratios for high-resolution fiber tracking and redundant data sets for k-space data swapping in postprocessing. The imaging sequence was synchronized using cardiac trigging. To study the relationship between the cardiac cycle and phase errors, a series of images was collected with different cardiac delays within the R-R interval of the subject s cardiac cycle. A 1D Fourier transform (FT) was performed on the navigator echoes of all separated shots and all images. The magnitude of the navigator projection was calculated and the area under the magnitude curve was taken as the measurement of the image quality. The studies were performed for two healthy subjects (one 41- year-old male and one 29-year-old female).

2 Residual Motion Artifacts in DW-EPI 819 FIG. 1. Reconstructed diffusion-weighted images and their navigator projections for different cardiac delays ( ms). The diffusion gradient was applied in the phase-encoding (right/left) direction. The navigator curves were normalized using their minimum and maximum values. Image b, acquired at a trigger delay of 100 ms, represents the worst image quality as compared with the other images, and its corresponding navigator projection has the lowest profile integral. The navigator curves were reproducible for all four image segments. MRI Data Correction The phase difference between navigator echos of different shots after 1D-FT (10,11) was calculated and used to correct motion-induced phase error in the multishot imaging echoes. This was done in multiple steps. First, the navigator echo NAV(kx) was masked with a rectangle window around its center kxc so that only those samples that have significant signal remained (9), i.e., 再 NAV共kx兲 NAV共kx兲 0 ifpkx kxc P otherwise [1] FIG. 2. The area under the navigator profile as a function of cardiac delay. The area values are normalized using their maximum. Diffusion gradients were applied in the (a) readout-, (b) phase-, and (c) slice-encoding directions. The combined data of four image segments are shown. Curve d was obtained without diffusion weighting, indicating minimal sensitivity to cardiac pulsations in this case.

3 820 Jiang et al. in which kx is the index number in k-space and the half of the window width. This step leads to a smoother navigator projection after its FT, because masking the signal with a rectangular window in the k-space domain is in fact equivalent to convolution with a low-pass sinc-shaped filter in the spatial domain. The adjustable mask window size was chosen as 10% of the echo length as a default selection in our experiments. Then, the navigator echo with the largest amplitude area after FT was chosen as the reference navigator. Because motion artifacts tend to reduce image intensity and change phase, the navigator profile with the largest area is likely to have less motion effects (12). Subsequently, the phase difference functions, Diff (x), between the reference and each of the navigator projections were calculated from Diffi x NAVi x NAV0 x i 1..m, where NAVi x indicates the navigator phase response of the ith imaging shot after its 1D-FT, m the total number of shots, and NAV0 x the reference navigator phase response. Subsequently, a point-by-point phase adjustment was made between Fourier-transformed image echoes and corresponding navigator phase difference for every shot: collected at various time points over the cardiac cycle. Figure 1 shows some of the selected images (after navigator correction) obtained at different cardiac delays for the particular coronal slice in which motion artifacts have been observed most frequently in our experience. The heart rate of this particular healthy volunteer was approximately 72 beats per min (833 ms per cycle). The imaging sequences were synchronized using cardiac delays from 50 ms to 650 ms, with 50-ms increments. This yielded a total of 13 images with increasing ECG triggering delays. TR was set to four times the R-R interval, i.e., TR 3333 ms. The diffusion gradient was in the phase-encoding (Y) direction. From these images, it can be seen that images at cardiac delays of ms were corrupted and the corresponding navigator projections had significantly lower intensities and different shapes. The similar relationship between the image quality and the cardiac cycle was also found in different slice locations. S i x S i x e j i x Diffi x i 1..m [2] in which S i (x) S i (x) e j i(x) denotes the image data of the ith imaging shot after 1D-FT. Once all image data were corrected (i.e., all echoes having the same phase shift as the reference navigator), the resultant matrix was Fourier transformed in the phase-encoding direction to get a complex image. Finally, the magnitude image was calculated and displayed. Measurement of Blood Velocity To verify the cardiac-cycle dependency of the navigator projection intensity, velocity profiles were acquired in a plane perpendicular to the right common carotid artery of the same volunteer using a 10-cm circular surface coil. For this purpose a gradient-echo sequence was implemented using the following parameters: TE 9.1 ms, TR 20 ms, flip angle 15, FOV 15 cm, matrix , rectangular FOV 70%, slice thickness 5 mm. Bipolar flow-encoding gradients were set in the slice-selection direction, with a maximum velocity encoding of 90 cm/s in order to avoid any phase wrap. Images were reconstructed using a retrospective cardiac gating scheme, in which the phases are reconstructed on the basis of the RR-intervals recording during a nontriggered scan. It can be regarded as a cardiac-triggered multiphase technique, which will therefore increase the image quality in a short-tr gradientecho sequence, such as the one used in absolute blood velocity imaging (13 15). Twenty phases were recorded 30 ms apart during the 700 ms of the cardiac cycle. RESULTS Relationship Between Corrupted Scans and Cardiac Cycle To investigate the effects of cardiac pulsation on the uncorrectable motions, images and navigator echoes were FIG. 3. a: Time profile of the integrated blood velocity in the right common carotid artery. b: Correlation plots of the blood velocity to the area of navigator projections. The obtained regression lines are given by: y x, r 0.78, gradient X, solid line; y x, r 0.72, gradient Y, dashed line; and y x, r 0.84, gradient Z, dotted line.

4 Residual Motion Artifacts in DW-EPI 821 FIG. 4. Identification of motion-affected scans and their correction by k-space swapping. a: Phase-corrected images (left column) and the associated three navigator echoes of two redundant datasets. b: Images after the k-space swapping. The third segmented scan of the first image (numbered 00:2) was discarded and replaced with a corresponding segment in the second dataset (numbered 01:2). The segment numbered 01:1 was substituted with 00:1. Images were copied from the display of our inhouse-developed software that allows us to choose the desired combination of k-space segments automatically or manually for reconstruction. The normalized total areas (over all four shots) under the navigator profiles are plotted as a function of cardiac delay in Fig. 2. This figure represents three data sets with different orientations of diffusion weighting. The relationship between the cardiac delay and navigator projections is very reproducible in that low intensities of navigator projections were observed for cardiac delays of ms (Fig. 2a, X orientation) in all four image segments. A similar relationship was observed for different gradient orientations (Y (Fig. 2b)) and Z (Fig. 2c)), although there were slight differences in the time length of the low-intensity period, which extended as long as 250 ms for the Z gradient orientation (the anterior-posterior direction). For comparison, the navigator area as a function of cardiac delay is also given for the images without diffusion weighting (Fig. 2d), showing almost negligible area fluctuation. Similar results were observed for the other healthy volunteer. Correlation With Blood Pulsatile Flow Through the Right Common Carotid Artery To verify that the time-profile of the navigator projection intensity is caused by brain pulsation, the time-profile of blood flow in the right common carotid artery was mea- sured (Fig. 3a). It can be seen that the time period over which the navigator projection profiles change drastically coincides with the time when blood flows into the brain most rapidly. The linear correlation coefficient between the time-profile of blood velocity and navigator projection was given by 0.78 (X), 0.72 (Y), and 0.84 (Z), respectively (Fig. 3b). Negative value here indicates that blood flow and the navigator projection are inversely correlated, i.e., the area of navigator profile reduces at higher blood velocity. DISCUSSION At present, most DTI studies are performed using singleshot EPI because of its insensitivity to subject motion and subsequent image ghosting. However, single-shot EPI has severe limitations in image quality due to its intrinsic low resolution and related image distortions. This hampers correct anatomical interpretation of color maps and fiber track locations. Therefore, a method to increase the success rate of navigator-corrected segmented EPI for DTI studies is needed. In this study, we tried to understand the dominant source of residual motion artifacts, which could

5 822 Jiang et al. not be corrected using the conventional navigator-echo based phase correction. The results in Fig. 1 show that severe motion effects judged from the navigator projection intensity are highly reproducible and synchronized with the cardiac cycle. The low echo intensity period observed at ms after the triggering coincided with the time of peak blood velocity in the right common carotid artery, suggesting that brain pulsations are the dominant source of these observed motion-related artifacts. The results indicate that it should be possible to minimize the pulsation effects by avoiding acquiring data during the low echo intensity period ( ms). Therefore, the recommended timing for data acquisition with respect to cardiac cycle is within 50 ms, or at times longer than 250 ms, after R-wave detection. In addition to the reproducible motion effects due to brain pulsation, there are also occasional involuntary motions that are too severe to be corrected by the navigator echoes. These are especially a problem when studying young children and more difficult populations, such as individuals with learning disabilities or certain psychiatric disorders. An example of such a case is shown in Fig. 4a, in which only one of three segmented scans was affected. For this type of motion effects, it is very effective to acquire redundant datasets and substitute the corrupted k-space with noncorrupted data from the redundant data sources. We developed software allowing us to inspect each segmented scan for the existence of data corruption and to substitute the k-space. Figure 4b demonstrates that the software effectively suppressed the occasional motion artifacts, allowing successful image analysis in this otherwise useless data set. Recently, phase corrections on various navigator-echobased or non-navigator-echo-based approaches have been proposed to overcome the limitation of the 1D navigatorecho phase correction scheme (7,11,16 19). The present results could be combined with these new types of motion suppression schemes to further improve the quality of diffusion-weighted images. In conclusion, we demonstrated that the dominant source of residual motion artifacts in navigator-corrected segmented DW-EPI is caused by cardiac pulsation of the brain. Based on the relationship between the phase error and the cardiac cycle, it is recommended that data be acquired within 50 ms or more than ms after cardiac triggering to minimize motion-related artifacts. For the suppression of nonreproducible voluntary motions, swapping k-space data between redundant data sets was effective. ACKNOWLEDGMENTS We thank Drs. Carlo Pierpaoli, Alan Barnett, and Peter Basser for stimulating discussions regarding motion artifacts and k-space swapping. This study was supported by the Whitaker Foundation, National Institutes of Health grant RO1 HD37931 (to S.M.), and NCRR resource grant P41 RR15241 (to S.M, P.V.Z). We acknowledge Ms. Terry Brawner (F.M. Kirby Research Center at Kennedy Krieger Institute), and Drs. Paul Folkers, Frank Hoogenraad, and Arianne van Muiswinkel (Philips Medical Systems) for technical assistance. REFERENCES 1. Le Bihan D, Turner R, MacFall JR. Effects of intravoxel incoherent motions (IVIM) in steady-state free precession (SSFP) imaging: application to molecular diffusion imaging. Magn Reson Med 1989;10: Turner R, Le Bihan D, Maier J, Vavrek R, Hedges LK, Pekar J. Echoplanar imaging of intravoxel incoherent motion. Radiology 1990;177: Basser PJ, Mattiello J, Le Bihan D. Estimation of the effective selfdiffusion tensor from the NMR spin echo. 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