The Impact of Navigator Timing Parameters and Navigator Spatial Resolution on 3D Coronary Magnetic Resonance Angiography

JOURNAL OF MAGNETIC RESONANCE IMAGING 14:311 318 (2001) Technical Note The Impact of Navigator Timing Parameters and Navigator Spatial Resolution on 3D Coronary Magnetic Resonance Angiography Elmar Spuentrup, MD, 1,2 * Matthias Stuber, PhD, 1,3 René M. Botnar, PhD, 1,3 and Warren J. Manning, MD 1,4 The impact of navigator spatial resolution and navigator evaluation time on image quality in free-breathing navigator-gated 3D coronary magnetic resonance angiography (MRA), including real-time motion correction, was investigated in a moving phantom. Objective image quality parameters signal-to-noise ratio (SNR) and vessel sharpness were compared. It was found that for improved image quality a short navigator evaluation time is of crucial importance. Navigator spatial resolution showed minimal influence on image quality. J. Magn. Reson. Imaging 2001;14: 311 318. 2001 Wiley-Liss, Inc. Index terms: magnetic resonance (MR); coronary vessel; motion study; coronary angiography; coronary artery disease SUBMILLIMETER FREE-BREATHING navigator-gated and corrected 3D coronary magnetic resonance angiography (MRA) has been successfully implemented in patients for visualization of the proximal and mid portion of the coronary vessels (1,2). Using real-time navigator technology, data are accepted if the navigator interface (commonly the lung-liver interface) is within a small (3 5 mm) end-expiratory gating window (3 5). In addition to navigator gating, a prospective real-time adaptive motion correction of the imaged volume (tracking) has been described (5,6). While studies of navigator localization and variable gating window width have been reported (5,7,8), the impact of navigator spatial resolution and navigator timing (navigator time delay) on image quality remain to be defined. We sought to study this relationship using a previously implemented 1 Department of Medicine (Cardiovascular Division), Beth Israel Deaconess Medical Center and Harvard Medical School, Boston, Massachusetts. 2 Department of Diagnostic Radiology, Technical University of Aachen, Aachen, Germany. 3 Philips Medical Systems, Best, The Netherlands. 4 Department of Radiology, Beth Israel Deaconess Medical Center and Harvard Medical School, Boston, Massachusetts. Contract grant sponsor: German Research Council; Contract grant sponsor: Established Investigator Grant of the American Heart Association; Contract grant number: 9740003N. *Address reprints requests to: E.S., Department of Diagnostic Radiology Technical University of Aachen, Pauwelsstrasse 30, 52057 Aachen, Germany. E-mail: spuenti@rad.rwth-aachen.de Received October 25, 2000; Accepted May 15, 2001. ex vivo periodically moving phantom (9), which allows for methodical investigation of a defined model motion. MATERIALS AND METHODS The impact of real-time navigator performance on image quality was investigated utilizing an MR-compatible, periodically moving phantom (9) designed to mimic physiologic respiratory motion. Image data were acquired using a previously described navigator-based coronary MRA technique (1). Objective image quality parameters, including signal-to-noise ratio (SNR) (3) and vessel sharpness (1), of a small diameter tube (2.8 mm) attached to the moving phantom were subsequently analyzed. MRI MR imaging (MRI) was performed on a commercial 1.5 T Gyroscan ACS-NT whole-body MR system (Philips Medical Systems, Best, The Netherlands) equipped with cardiac research software (INCA 2) and a commercial PowerTrak 6000 gradient system (23mT/m, 219- sec rise time). For signal acquisition, a circular surface coil (C1; diameter, 17 cm) was positioned anterior to the moving phantom. Phantom Configuration An MR-compatible moving phantom (9) with a periodically moving head (Fig. 1) was used. This phantom is operated by a DC motor positioned 250 cm from the magnet iso-center and generated a sinusoidal translational motion in the foot-head direction with an amplitude of 25 mm and a frequency of 18 cycles/minute. To provide an interface for the navigator, a cup was filled with ULTRA PHONIC conductivity gel (Pharmaceutical Innovations, Inc., Newark, NJ) and mounted on the moving phantom head (Fig. 1). An additional phantom bottle filled with gelled water (JELL-O, Kraft Foods, NY) was connected with a 2.8-mm inner diameter tube (Interlink System, Baxter, Healthcare Corporation, Deerfield, IL), filled with baby oil (Johnson & Johnson Consumer Products, Inc., Skillman, NJ) to model a coronary artery. This coronary had a curved loop (white 2001 Wiley-Liss, Inc. 311

312 Spuentrup et al. 2D Selective Navigator and Navigator Position Determination A 2D selective pencil excitation pulse with a subsequent flow-compensated readout of 256 data points in the longitudinal direction was used as the navigator (5 7,10). The navigator interface position was calculated by cross-correlation (10) of the actual navigator signal and an end-expiratory reference navigator profile. Hereby, the shifts of all data points on the signal slope in the kernel (Fig. 2) are used. This allowed for subpixel resolution of the calculated interface position. If the navigator-detected interface position falls into a userspecified range (gating window), the data are accepted for image reconstruction; otherwise the data are dis- Figure 1. Phantom head with a paper cup filled with gelled water for navigator interface detection (a) and a parallel phantom bottle with a model coronary vessel (b). This coronary had a curved loop (bright arrow), which was perpendicular-orientated to the moving direction, and a straight portion, which was orientated along the motion direction (black arrowhead). The amplitude of the movement is marked with a black double arrow frontally. Bo direction of main magnetic field. The coronary was blackened for this picture. arrow in Fig. 1), which was orientated perpendicular to the primary direction of phantom motion, and a linear/ straight portion (black arrowhead in Fig. 1) orientated along the direction of phantom motion. Both the navigator interface and the bottle with the attached coronary moved synchronously. Anterior to the phantom head, a nonmoving cooling gel-container (NEXCARE Cold Pack, 3M Health Care, MN) was positioned in order to have a static structure in the MR images. MRI Sequence A previously described 3D segmented-k-space gradient-echo imaging sequence (1) (without T2 preparation or fat saturation) was used as the MRA sequence. The repetition time (TR) was 9.9 msec, and the echo time (TE) was 2.9 msec. The sequence was ECG triggered using an artificial ECG with a frequency of 80 beats/ minute (effective RR interval of 750 msec). Eight RF excitations were applied during each RR interval resulting in an acquisition window of 79 msec. Data were acquired using centric k-space ordering with priority for the central k-space profiles. The acquired 3D volume included five 3-mm-thick slices interpolated (zero filling) to 10 contiguous 1.5-mm-thick slices. The 3D volume was positioned sagittally with the coronary orientated in-plane and in the center of the slab. The field of view (FOV) was 256 mm, with a matrix of 512 512, resulting in an in-plane resolution of 0.5 0.5 mm 2. The phase encoding direction was anterior-posterior. Figure 2. Signal intensity curves and navigator display of 24 navigator beams (voxel size, 1mm) under nonmoving conditions. A high reproducibility of the curves is seen. The SD of the calculated interface position for this navigator spatial resolution was 0.0091 mm (mean for all investigated navigator spatial resolutions, 0.011 0.003 mm). The interface offset of the navigator beam compared to a reference navigator beam is calculated by cross-correlation, which uses the shift of all signal data of the signal slope in the kernel (marked with long vertical lines). Therefore, a subpixel resolution of the detected interface position can be obtained. ROIsmax and ROInoise are the user-specified ROIs for SNR calculation. a.u. arbitrary units.

Navigator Timing on 3D Coronary MRA 313 carded and remeasured until the gating window is satisfied. To analyze the impact of navigator spatial resolution, the navigator FOV was adjusted from 64 to 512 mm, corresponding to a navigator spatial resolution of 0.25 2.00 mm. The navigator excitation flip angle was 60, with a navigator diameter of 25 mm and 12 cycles in k-space using a jinc-shaped RF pulse (11). For all studies, a previously implemented gating window of 5 mm was used. In addition to navigator gating, prospective adaptive real-time motion correction (tracking) of the imaged 3D volume position in all three spatial coordinates was performed. This was accomplished by the prospective adaptation of the frequency of the RF excitation and by the adaptation of acquisition phase and frequency (5). The correction factor describing the relative navigator interface and coronary displacement was set to 1.00 (rigid motion). Description of Experiments and Analysis of Data Static Phantom: Navigator Signal and Navigator Interface Definition For the investigation of the impact of the navigator signal for interface definition, 24 consecutive navigator beams were sampled with a repetition time of 750 msec. The navigator FOV was incrementally increased from 64 to 512 mm in 29 steps of 16 mm each, resulting in a voxel size variation of 0.250 2.00 mm. The signal profile of the navigator beam and the position of the detected interface positions were automatically written to a log file on the system. From a user-specified region of maximal signal intensity (S(ROImax), Fig. 2) and the SD of the background noise (SD(ROInoise), Fig. 2), the SNR of the navigator signal was calculated according to SNR S(ROImax)/SD(ROInoise). The sharpness of the detected navigator interface was analyzed by using the pixel-by-pixel first-order derivative in readout direction (perpendicular to the interface), corrected for the spatial resolution of the navigator voxel size. Because reliable measurements of the navigator interface position inside the magnet using an external device are difficult, the reproducibility of the navigator interface position detection was determined by calculating the SD of the detected interface position of 24 subsequent navigator beams for each navigator spatial resolution. Moving Phantom: Coronary MRA Image Quality Initially and as a reference, MR images were acquired without superimposed motion to the phantom. Then the impact of navigator processing time (navigator time delay) and navigator spatial resolution on MR image quality was studied as follows. First, the time delay between navigator and imaging part of the sequence (navigator time delay) was increased from 20 to 160 msec in 20-msec increments. Two sets of measurements were acquired: one data set with gating alone and a second data set with combined gating and prospective adaptive motion correction (tracking) (6). Second, the impact of spatial navigator resolution on image quality was investigated by increasing the navigator FOV from 64 to 512 mm in steps of 32 mm each, resulting in a voxel size of 0.250 2.00 mm. For this navigator spatial resolution study, a minimal navigator time delay of 20 msec and combined gating with tracking were used. For objective assessment of image quality, the MR data were transferred to a personal computer. The signal intensity was measured in a user-specified region of interest (ROI) from the curved portion of the model coronary loop (orientated perpendicular to the moving direction, Fig. 1, white arrow) and in the straight portion ventral to the bottle (orientated parallel to the primary direction of motion, Fig. 1, black arrowhead). A further ROI was placed in the background noise at the level of the curved and straight portion of the model coronary. The SNRs of the curved and straight coronary were calculated as previously described for the static phantom study. For objective quantification of vessel definition, we applied a previously described edge detection tool (1). This algorithm is primarily based on the first-order derivative of the vessel edge signal and was applied for measurements of vessel sharpness (1). Statistics Six repetitions for gating alone and combined gating with tracking for each navigator time delay were performed. Comparisons of the vessel sharpness and SNR as a function of navigator time delay utilizing gating alone and gating with tracking were made by calculation of the slope in a linear regression model. A P value of 0.01 was considered significant. A further linear regression analysis was made for the vessel sharpness and the navigator spatial resolution. For comparison of the image quality for each single navigator time delay, a paired, bidirectional Student s t-test was calculated. Hereby, a P value of 0.01 was considered significant. RESULTS Representative MR images from the moving phantom studies are shown in Figure 3. Data regarding vessel sharpness and SNR for the curved and straight portion of the coronary are presented in Figures 4 and 5, respectively. The data were normalized to static/nonmoving conditions (1.00). Gating Without Tracking MR Images: Visual Comparison In the static (nonmoving) phantom (Fig. 3a), the curved and straight portion of the coronary are shown with high contrast and sharp edges. Under moving conditions, with navigator gating alone, the curved portion of the coronary loop showed enhanced blurring and motion artifacts, especially at the inner boundary of the loop, independent of the navigator time delay (Fig. 3b, d, f, and h). Objective Image Quality Parameters The vessel sharpness of the curved coronary utilizing gating alone was reduced to 0.8 of the value in static

314 Spuentrup et al. Figure 3. 3D coronary MRA (slice thickness, 1.5 mm; in plane resolution, 0.5 0.5 mm 2 ) static (a) and with moving phantom (b i) for different navigator time delays. b, d, f, and h: Gating alone. c, e, g, and i: Combined gating with tracking. a: Labeled portions of the curved coronary and straight coronary for data evaluation. Images in the upper row show a 240% magnification of the curved loop for enhanced reader visibility. Utilizing gating only (b, d, f, and h), a veil at the inner boundary of the curved coronary loop is seen. c: Superior image quality was found using gating with tracking and a minimal navigator time delay of 20 msec. For prolonged navigator time delays, the curved coronary shows enhanced blurring at both edges (arrows in g). For the longest (160 msec) presently investigated navigator time delay, a double contour at the curved coronary is visible (i, arrow). Utilizing tracking (c, e, g, and i), enhanced phase errors at the level of the static tissue were found. conditions (P 0.01) (Fig. 4) and was independent of navigator time delay (P 0.5). For the straight coronary, no change in vessel sharpness as a function of the navigator time delay was found (P 0.5) (Fig. 4). The SNR for the curved coronary was significantly reduced with 0.26 0.39 of the static condition (P 0.01) (Fig. 5). This was also independent of the navigator time delay (P 0.5).

Navigator Timing on 3D Coronary MRA 315 Figure 4. Vessel sharpness of the curved (rhombs) and straight (circles) model coronary with varying navigator time delay utilizing gating only (open) and gating combined with tracking (filled). All values are normalized to nonmoving conditions (1.00). The best vessel sharpness of the curved coronary was found utilizing gating with tracking in settings of a short (20 msec) navigator time delay (P 0.01). Prolonged navigator time delays resulted in reduced vessel sharpness using gating with tracking (negative slope in the linear regression, P 0.01), and the vessel sharpness utilizing tracking was reduced compared to gating alone for 80 msec. The vessel sharpness of the straight coronary demonstrated no difference compared to static conditions (P 0.5). Statistically significant values of vessel sharpness between gating alone and gating with tracking for the curved coronary are marked with a star. coronary showed a blurring (Fig. 3e, g, and i, arrows in Fig. 3g), and increased motion artifacts were seen. For the longest navigator time delays investigated in the present study (160 msec), a double contour at the inner boundary of the artificial coronary was visible while utilizing gating with tracking (arrow in Fig. 3i). Since such a double contour may influence the vessel sharpness calculation, the data points 140 and 160 msec navigator time delays were excluded from regression analyzes. Objective Image Quality Parameters By using navigator gating with tracking, superior vessel sharpness of the curved coronary was seen utilizing the Figure 3. (Continued) Gating With Real-Time Tracking MR Images: Visual Comparison By using gating with tracking, superior image quality was seen by utilizing a minimal navigator time delay of 20 msec (Fig. 3c). No or minor motion artifacts were visible and the curved coronary showed a sharp contour. However, by using longer navigator time delays ( 40 msec), the image quality was reduced. The curved Figure 5. SNR of the curved coronary utilizing gating only and combined gating with tracking. All data are normalized to static conditions (1.00). The best SNR was found for gating with tracking and a short (20 msec) navigator time delay (P 0.01). Long navigator time delays demonstrated reduced SNR (linear regression model, P 0.01). Statistically significant (P 0.01) values of SNR between gating alone and gating with tracking are marked with a star.

316 Spuentrup et al. minimal navigator time delay (20 msec) (0.95 of the static conditions, Fig. 4). However, increasing navigator time delays were associated with continuously reduced vessel sharpness (Fig. 4, significantly negative slope in the regression analyzes, P 0.01). For the straight coronary, no different values for vessel sharpness compared to static conditions were found (Fig. 4, P 0.5). Similarly, the SNR of the curved coronary was 0.96 of the static conditions using the minimal navigator time delay (20 msec) (Fig. 5). SNR decreases as a function of navigator time delay (negative slope in the linear regression model, P 0.01). Comparing both, combined gating with tracking and gating alone, the objective image quality parameters vessel sharpness and SNR were significantly improved for gating with tracking while using short ( 40 msec for vessel sharpness, 80 msec for SNR) navigator time delays (P 0.01, Figs. 4 and 5). The best image quality was achieved using combined navigator gating with tracking and a 20-msec navigator time delay. For navigator time delays of 80 msec, opposite proportions were found with significantly inferior vessel sharpness for gating with tracking compared to gating alone (P 0.01, Fig. 4). Only for combined gating with tracking was a significant decrease of vessel sharpness as a function of navigator time delay found. Spatial Navigator Resolution Reproducibility of Navigator Interface Position Detection The detected navigator interface position in static conditions (Fig. 2) for the different navigator spatial resolutions (24 navigator beams for each spatial navigator spatial resolution) showed a mean SD of 0.011 0.003 mm. Objective Image Quality Parameters The vessel sharpness of the curved coronary found in the moving phantom was independent of the navigator voxel sizes in the investigated range of 0.250 2.00 mm (P 0.5, Fig. 6). For navigator SNR (static phantom), an increase was seen as a function of the incrementally reduced navigator spatial resolution. For the navigator interface sharpness, a continuous decrease was found. DISCUSSION 3D free-breathing coronary MRA has been successfully acquired using navigators for suppression of respiratory artifacts (1,2,5,8). Although these techniques have already shown to be potentially useful in small patient studies (1,2), coronary MRA spatial resolution still remains to be improved and may depend on the accuracy of the navigator interface detection. We and others have previously utilized navigators with 1.00 mm spatial resolution (5 7,12,13). In this study, we found that a variation of navigator spatial resolution yielded no significant impact on image quality (Fig. 6). Although higher navigator interface sharpness was seen as a function of navigator spatial resolution, no improved vessel sharpness could be found. Figure 6. Model vessel sharpness (filled), navigator beam SNR, and navigator interface sharpness (open) for different navigator spatial resolutions. All values are normalized to a 1.00-mm voxel size. Although the navigator interface sharpness increases as a function of reduced navigator voxel size (on cost of SNR), no impact on vessel sharpness is seen (P 0.5). The navigator beam signal and the navigator interface position detection were highly reproducible and in the subpixel range (mean SD of the detected position, 0.011 mm), which can be explained by the used crosscorrelation algorithm (10). Using this algorithm, all the navigator signal data points on the curved slope in the kernel (Fig. 2) are used for interface position calculation, allowing for subpixel resolution position calculation. Such a navigator position detection algorithm may be robust even in a less sharply defined interface configuration. The presented results suggest that the cross-correlation algorithm for navigator interface position detection (10) is a stalwart algorithm in the investigated range of lower spatial navigator resolutions (down to 2.0 mm) as well. A major impact on image quality was seen using variable navigator time delays and utilizing gating alone or gating with tracking. Hereby, we found a good agreement between visual MR image quality and objectively determined parameters, such as vessel sharpness (1) and SNR (3). Superior image quality (close to static conditions) was found for gating with tracking in settings of a minimal (20 msec) navigator time delay. A major negative impact on image quality for prolonged navigator time delays (Figs. 3 and 4) was observed. This suggests that the improved vessel sharpness by using gating with tracking is dependent on a short navigator time delay. A possible mechanism of these effects is shown in Figure 7. During tracking, the slice position is prospectively adapted. For the positions with movement between navigator position detection (points A1 and A2) and imaging, the two positions will move superior (A1) or inferior (A2), resulting in a bidirectional offset error. For moving structures, an offset always exists. Therefore, during tracking some blurring of the moving coronary voxels occurs, resulting in a reduced vessel sharpness. This effect is strongly dependent on the navigator time delay. For long (160 msec) navigator time delays (Fig. 3i) an increased spreading of the coronary voxels occurs, which may explain the double contour observed in Fig. 3i (arrow). In contrast to this, the artifacts by using gating alone and without tracking may be caused by two different

Navigator Timing on 3D Coronary MRA 317 Figure 7. Schematic of moving structures in a gating window during free breathing to explain the impact of navigator time delay on image quality. mechanisms. First, because no adaptation of the imaging volume is performed, measurement points with a former offset may move during the navigator time delay to central positions and will then be measured without any offset error. Second, for points with higher velocity (at the inner boundary of the gating window in Fig. 7) a larger offset, as with tracking, exists. This may explain the larger, but lower dense veil at the inner boundary of the curved coronary while using gating alone (Fig. 3b, d, f, and h). Both possible mechanisms are only minimally dependent on the navigator time delay: prolonged navigator time delays using gating alone result only in a slightly increased gating window interval due to the points moving out of the given gating window, but without additional offset artifacts. This may explain why no impact on navigator time delay was found in the investigated range (20 160 msec) if gating alone was utilized. Therefore, gating alone is potentially less susceptible to prolonged navigator time delays than using combined gating and tracking. The potentially different reasons for artifacts utilizing gating alone and combined gating with tracking may explain the superior vessel sharpness for combined gating and tracking with a short ( 40 msec) navigator time delay, whereas for prolonged navigator time delays ( 80 msec), gating alone was superior. Overall, the best image quality was obtained using gating with tracking in settings of a minimal short navigator time delay (20 msec). Combined gating and tracking only results in a superior image quality, if short navigator time delays ( 40 msec) can be achieved. These results may be of high importance for the design of imaging sequences and prepulses. Since prepulses (T2 preparation, inversion, etc.) are independent of the imaged volume position, the navigator beam should be positioned as close as possible to the imaging sequence without any prepulses interspersed between the navigator and the imaging part of the sequence. In addition, the central k-space lines of the imaging sequences, which mainly induce the motion errors in coronary MRA (14), should be acquired as early as possible. Since the image quality was independent of the spatial navigator resolution (Fig. 6), a shorter navigator time delay can be traded for a reduced navigator resolution in order to increase the accuracy of the navigator interface detection and therefore to obtain higher image quality, which might be important for increased image spatial resolution. However, the potential of such a trade-off might depend on scanner-specific navigator processing capabilities. In addition to the curved coronary loop, which showed a major impact of navigator time delay on vessel sharpness, a portion of the straight coronary was investigated. Because this portion was orientated parallel with the moving direction (Fig. 1), the motion error offsets are in parallel with this portion of the coronary. As a consequence, no locally reduced vessel sharpness was observed under moving conditions (Figs. 3 and 4). Therefore, in vivo the vessel sharpness of different portions of the coronary tree may be differently influenced by respiratory motion, depending on the orientation to the respiratory motion direction. During our experiments, we also noted enhanced artifacts at the level of static tissue, if tracking was applied (Fig. 3c, e, g, and i). In contrast with gating alone these artifacts were minimized (Fig. 3b, d, f, and h). By using tracking, static tissue appears with a relative different velocity, which results in enhanced artifacts in a phase encoding direction. As a consequence, suppression of the signal from static tissues, like the chest wall, in vivo may serve to minimize these artifacts. In our study, we investigated the impact of navigator timing in a periodically moving phantom. With such a model, the impact of different timing parameters can be methodically investigated. Furthermore, a static (nonmoving) condition can be used as a reference scan. A limitation with such a model may be a different condition in vivo. In vivo, typically a 0.6 correction factor (15) is used (1,2,5), resulting in a smaller tracking interval when compared to our in vitro conditions (only one

318 Spuentrup et al. gating window was studied). However, larger tracking intervals, as used in our phantom study, are needed to enhance navigator efficiency (8). The dome of the right hemidiaphragm, which is recommended for navigator position (5), may have a curved surface and may have different signal properties than the navigator interface model in our present phantom study. Although in vivo only a feet-head motion direction (as in our phantom study) has been corrected for coronary MRA (1,2,5,13), more complex heart motion as a consequence of respiration has been described (15), and individually dependent diaphragmatic velocities (12) have to be taken into account. Finally, sinusoidal motion does not fully reflect the normal breathing pattern. CONCLUSIONS Using a moving phantom, gating with tracking was found to be superior to gating alone only if short ( 40 msec) navigator time delays were used. If tracking is used, a minimized navigator time delay (20 msec) was found to be of crucial importance for a further improvement of image quality. In contrast, navigator spatial resolution showed minimal influence on image quality. Therefore, reduced navigator resolution can be traded for a shortened navigator time delay. A signal from static tissue may introduce image artifacts using tracking. ACKNOWLEDGMENTS Dr. Elmar Spuentrup is supported in part by the German Research Council. Dr. Warren J. Manning is supported in part by an Established Investigator Grant of the American Heart Association, Dallas, TX (9740003N). REFERENCES 1. Botnar RM, Stuber M, Danias PG, Kissinger KV, Manning WJ. Improved coronary artery definition with T2-weighted free-breathing 3D-coronary MRA. Circulation 1999;99:3139 3148. 2. Stuber M, Botnar RM, Danias PG, et al. Double oblique free-breathing high-resolution 3D coronary MRA. J Am Coll Cardiol 1999;34: 524 531. 3. Oshinski JN, Hofland L, Mukundan Jr S, Dixon WT, Parks WJ, Pettigrew RI. Two-dimensional coronary MR angiography without breath holding. Radiology 1996;201:737 743. 4. Wang Y, Rossman PJ, Grimm RC, Riederer SJ, Ehman RL. Navigator-echo-based real-time respiratory gating and triggering for reduction of respiration effects in three-dimensional coronary MR angiography. Radiology 1996;198:55 60. 5. Stuber M, Botnar RM, Danias PG, Kissinger KV, Manning WJ. Submillimeter three-dimensional coronary MR angiography with real-time navigator correction: comparison of navigator locations. Radiology 1999;212:579 587. 6. McConnell MV, Khasgiwala VC, Savord BJ, et al. Prospective adaptive navigator correction for breath-hold MR coronary angiography. Magn Reson Med 1997;37:148 152. 7. McConnell MV, Khasgiwala VC, Savord BJ, et al. Comparison of respiratory suppression methods and navigator locations for MR coronary angiography. AJR Am J Roentgenol 1997;168:1369 1375. 8. Danias PG, McConnell MV, Khasgiwala VC, Chuang ML, Edelman RR, Manning WJ. Prospective navigator correction of image position for coronary MR angiography. Radiology 1997;203:733 736. 9. Huber ME, Stuber M, Botnar RM, Boesiger P, Manning WJ. Lowcost MR-compatible moving heart phantom. J Cardiovasc Magn Reson 2000;2:181 187. 10. Ehman RL, Felmlee JP. Adaptive technique for high-definition MR imaging of moving structures. Radiology 1989;173:255 263. 11. Nehrke K, Bornert P, Groen J, Smink J, Bock JC. On the performance and accuracy of 2D navigator pulses. Magn Reson Imaging 1999;17:1173 1181. 12. Taylor AM, Jhooti P, Wiesmann F, Keegan J, Firmin DN, Pennell DJ. MR navigator-echo monitoring of temporal changes in diaphragm position: implications for MR coronary angiography. J Magn Reson Imaging 1997;7:629 636. 13. Nagel E, Bornstedt A, Schnackenburg B, Hug J, Oswald H, Fleck E. Optimization of realtime adaptive navigator correction for 3D magnetic resonance coronary angiography. Magn Reson Med 1999;42: 408 411. 14. Wang Y, Winchester PA, Yu L, et al. Breath-hold three-dimensional contrast-enhanced coronary MR angiography: motion-matched k- space sampling for reducing cardiac motion effects. Radiology 2000;215:600 607. 15. Wang Y, Riederer SJ, Ehman RL. Respiratory motion of the heart: kinematics and the implications for the spatial resolution in coronary imaging. Magn Reson Med 1995;33:713 719.