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JOURNAL OF MAGNETIC RESONANCE IMAGING 28:1273–1279 (2008) Technical Note Inflow Quantification in Three-Dimensional Cardiovascular MR Imaging Reza Nezafat, PhD,1* Daniel Herzka, MD,3* Christian Stehning, PhD,4* Dana C. Peters, PhD,1 Kay Nehrke, MD,4 and Warren J. Manning, MD1,2 Purpose: To investigate blood inflow enhancement (or lack thereof) in three-dimensional (3D) cardiovascular MR for both single phase whole-heart and cine biventricular functions. Materials and Methods: A 3D imaging sequence is proposed in which radiofrequency excitation gradient is changed without modifying image acquisition or phase/ slice encoding. This imaging sequence enables direct inflow measurement while retaining static voxel signal-to-noise ratio. Inflow measurements were performed for both spoiled gradient-echo (GRE) imaging and balanced steadystate free precession (SSFP) in 18 healthy subjects. Results: For single phase imaging, increasing slab thickness from 3 to 10 cm lead to 73% and 59% reductions in contrast-to-noise ratio (CNR) with GRE and SSFP, respectively. For cine acquisitions, systolic CNR was reduced by 85% and 50% for the GRE and SSFP acquisitions, respectively, while diastolic CNR was reduced by 64% and 42%. Conclusion: There is significant loss of CNR between blood and myocardium when using larger 3D slabs due to saturation of inflowing spins. The loss of contrast is less pronounced for SSFP than for GRE, though both acquisition techniques suffer. Key Words: 3D cardiac imaging, 3D whole heart imaging, inflow quantification J. Magn. Reson. Imaging 2008;28:1273–1279. © 2008 Wiley-Liss, Inc. IN TWO-DIMENSIONAL (2D) balanced steady-state free precession (SSFP) and spoiled gradient echo (GRE) acquisitions, fresh and unsaturated spins are continuously flowing into the imaging slice. These inflow spins 1 Departments of Medicine (Cardiovascular Division) and 2Radiology, Beth Israel Deaconess Medical Center and Harvard Medical School, Beth Israel Deaconess Medical Center, Boston, Massachusetts. 3 Clinical Sites Research Program, Philips Research North America, Briarcliff Manor, NY. 4 Tomographic Imaging, Philips Research Europe, Hamburg, Germany. The first three authors contributed equally to this work. *Address reprint requests to: R.N., Beth Israel Deaconess Medical Center, 330 Brookline Avenue, Boston, MA, 02215. E-mail: [email protected] Received February 19, 2008; Accepted June 6, 2008. Grant sponsor: American Heart Association; Grant number AHA SDG0730339N DOI 10.1002/jmri.21493 Published online in Wiley InterScience (www.interscience.wiley.com). © 2008 Wiley-Liss, Inc. tend to improve the blood–myocardium contrast-tonoise ratio (CNR) as they yield higher blood signal to noise ratio (SNR), particularly for SSFP sequences that use high imaging flip angles. The 3D cardiovascular MR (CMR) imaging is appealing due to its potentially higher SNR compared with 2D imaging (1). The 3D CMR imaging has recently gained importance due to the availability of phased arrays with a higher number of coil elements that enable the use of 2D parallel imaging to reduce total acquisition time (2– 6). Whole heart 3D coronary imaging, which applies a single large slab over the entire heart in an axial orientation obviates the use of multiple thin slab 3D acquisitions targeted to the right and left coronary arteries (6 – 8). Similarly, single breath-hold (9) or self-navigated 3D cine acquisitions are now becoming feasible. In addition to SNR gain in 3D cine, slice misregistration is typically improved, permitting more accurate evaluation of cardiac function. Both SSFP and GRE sequences are used in 3D CMR, with SSFP imaging more commonly used at 1.5 Tesla (T) due to its superior SNR and CNR (10,11). GRE imaging is more robust at higher magnetic fields (e.g., 3.0T), which display increased field inhomogeneity. Imaging of coronary vein anatomy for evaluation of patients undergoing cardiac resynchronization therapy has also been shown to be more robust with GRE than SSFP even at 1.5T, mainly due to increased B0 field inhomogeneity surrounding the veins (12). The transition from targeted thin 3D slab acquisitions to thick 3D slab whole heart should be advantageous with both GRE and SSFP due to higher spatial coverage and potential SNR increase of 公Nz, where Nz is the number of slice partitions or encoding steps (1). The effects of slab size on inflow-based image contrast are yet to be explored in a systematic manner. In 2D and small slab 3D imaging, fresh inflow results in blood signal enhancement thus blood–myocardium CNR improvement. With large 3D slab acquisitions, inflowing spins experience a higher number of radiofrequency (RF) pulses within the imaging volume, driving the magnetization closer to its steady-state magnitude (13). Therefore, the effect of inflowing spins on image contrast is expected to be greatly reduced (13). In the absence of inflow enhancement, GRE produces a T1 weighting, leading to a reduction of blood–myocardium CNR because both tissues have very similar T1s (1). In 1273 1274 Nezafat et al. Figure 1. Image acquisition set-up: a large 3D axial 10-cm slab (100 slices of 1 mm thickness) is prescribed to cover the entire heart. As usual, frequency (x), phase (y), and slice (z) encoding are performed. In the reference experiment, the imaging RF pulse excites the entire 10 cm volume, as shown, with an extra 20% slice encoding to prevent fold-over from RF imperfections. In subsequent experiments, b, c, and d, the excitation volume is reduced by changing the magnitude of the RF encoding gradient, reducing the effective slab width to 6.0, 4.0, and 3.0 cm, respectively. The frequency, slice and phase encoding gradients are left unchanged and cover the entire volume. This experiment maintains voxel size and scan time constant, and, therefore, changes in SNR and CNR can be attributed to inflowing spins. contrast, SSFP yields images with T2/T1 contrast, which in the absence of any inflow enhancement, might be more advantageous in terms of blood–myocardium CNR (1). Also, because the innate contrast produced by these sequences may not be as good as that observed with thin slab acquisitions, prepulses such as T2 magnetization preparation (14,15) or magnetization transfer (MT) may be used to create additional contrast (12,16). However, during the long transit time of the blood through a thick 3D-slab, considerable accumulative phase errors may evolve due to flow encoding of the phase encoding gradients (17). In this study, we sought to investigate inflow saturation in 3D single phase and cine CMR imaging by quantifying the contribution of inflowing spins to SNR and CNR with varying slab thickness. An imaging sequence is designed such that RF excitation slab thickness is varied in different acquisitions, thereby enabling direct inflow measurement while retaining static voxel SNR. MATERIALS AND METHODS To study the effect of inflowing blood in a large slab prescription, a series of imaging studies were performed in each subject. The RF excitation to image encoding ratio, that is, excitation slab thickness relative to image slab size, was changed between different image acquisitions by changing the amplitude of the slab selection gradient of Gz (i.e., reducing the excitation size) between different acquisitions, as shown in Figure 1. A slice selective 3D acquisition is used with excited RF profile covering the entire imaging volume. The excitation slab thickness was reduced manually in consecutive acquisitions without modifying the image acquisition volume, that is, frequency and phase encoding steps. By keeping the imaging parameters of slice, phase and frequency encoding the same between all acquisitions, the SNR and CNR changes between different acquisitions can be associated with the saturation of inflowing spins. The flowing spins into the imaging plane will experience different RF histories based on the excitation profile. MR imaging studies were performed on two 1.5T Philips Achieva (Philips Medical Systems, Best, NL). One system was equipped with a 16-channel receiver and a 5-element cardiac phased-array receiver coil for single phase whole heart coronary acquisitions. The second system used a 32-channel receiver and a 32-element cardiac phased array for 3D cine scans. In Vivo Studies Three groups of healthy adult subjects were imaged. In each group, images were acquired in 6 subjects (totaling 18 subjects, 10 females, mean age 27 years). For 3D cine imaging, 6 healthy subjects were studied using a 32-channel coil with both 3D SSFP and GRE cine. For single phase imaging, the other 12 subjects were studied using a 5-element coil using either 3D single phase GRE (in 6 subjects) or 3D single phase SSFP (in 6 other subjects). Written informed consent was obtained from all participants and the protocols were approved by both Institutional Review Boards. 3D GRE Single Phase Whole Heart CMR An electrocardiogram triggered, segmented 2-chamber SSFP cine (repetition time [TR] ⫽ 3 ms, echo time [TE] ⫽ 1.4 ms, ␣ ⫽ 60°, temporal resolution of 29 ms, spatial resolution of 1.6 ⫻ 2.0 mm2 reconstructed to 1.21 ⫻ 1.21 mm2) dataset was acquired at the level of the mid-ventricle to identify the onset of the diastolic rest period. This was followed by four whole heart CMR imaging studies that differed only in RF slab encoding gradient amplitude. A single phase sequence similar to one used for coronary artery and vein imaging was used (7,12). To improve blood–myocardium contrast, an MT pulse train consisting of eight 600° Sinc-Gaussian RF Inflow Quantification in 3D Cardiovascular MRI pulses, each with duration of 15 ms and frequency off-set of 500 Hz was used for all MT preparations (12). A spectrally selective fat saturation sequence was also used to suppress the fat signal. The navigator beam was positioned at the dome of the right hemidiaphragm with an acceptance window of 7 mm, automatic respiratory drift adaptation, and RF excitation angle of 25°. An axial volume covering the entire heart was prescribed off the two-chamber images and initial scout images. For each 3D dataset, a 10-cm-thick slab was acquired, using 50 slice encoding partitions reconstructed to 100 one-mm-thick slices. For imaging, 20 –25 RF excitations (asymmetric Gaussian-weighted-Sinc with four and two lobes and duration of 672 s) with spoiled GRE readouts: TE ⫽ 1.1 ms, TR ⫽ 3.6 ms, ␣ ⫽ 30°, BW ⫽ 383 Hz/Pixel, acquisition time of ⬃ 4:40 min for heart rate of 60 bpm. Partial echo (62.5%) was used in the acquisition. A field of view (FOV) of 270 ⫻ 270 mm2 was imaged with a scan matrix of 140 ⫻ 140 yielding a voxel size of 2 ⫻ 2 ⫻ 2 mm3 reconstructed to a 0.9 ⫻ 0.9 ⫻ 1 mm3. A total of four studies with slab thickness of 10 cm, 6 cm, 4 cm, and 3 cm were acquired. Due to the long duration of these studies, the images were prescribed with lower spatial resolution to ensure the successful completion of all 4 studies in all subjects. No parallel imaging was used to enable absolute SNR and CNR comparisons. 1275 tions and imaging artifacts with SSFP from increased field inhomogeneity. However, for consistency, in this study, 3D GRE cine images were acquired at 1.5T as well. The imaging parameters were: TR ⫽ 3.6 ms, TE ⫽ 1.3 ms, BW ⫽ 1562 Hz/Pixel, and ␣ ⫽ 20°. The scan geometry and matrix size were identical to those used for the SSFP acquisition with the same acceleration rate and coil. Both 3D GRE and SSFP cine images were acquired in one imaging session in one group of healthy subjects. 3D SSFP Single Phase Whole Heart CMR For 3D SSFP inflow quantification, a similar study to the one used for 3D GRE (previous section) was performed on six subjects. The following imaging parameters, if different from above, were used: TE ⫽ 1.8 ms, TR ⫽ 3.7 ms, ␣ ⫽ 90°, BW ⫽ 1470 Hz/Pixel. An excitation Gaussian-weighted Sinc RF pulse with three lobes and duration of 1.49 ms was used for SSFP imaging. 3D SSFP Cine CMR A 3D cardiac cine acquisition was prescribed to image the heart throughout the cardiac cycle. The imaging parameters were as follows: TR ⫽ 3.6 ms, TE ⫽ 1.8 ms, BW ⫽ 1562 Hz/Pixel, ␣⫽ 60°. A 10-cm 3D slab consisting of 25 slices (interpolated to 50 slices) was acquired using a FOV of 270 ⫻ 220 mm2, with a scan matrix of 128 ⫻ 102, yielding a voxel size of 2.1 ⫻ 2.1 ⫻ 4 mm3. A 2D acceleration of rate 2 ⫻ 2 in the phase and slice encoding directions and a partial Fourier acquisition (reduction factor 0.625) were used to reduce the acquisition time to fit within a single breath-hold (30 s). Although respiratory navigators can be used to reduce the time constraint on image acquisition, transition to steady-state artifacts that could present a complicating influence on the measurement of inflow effects. A 32element phased-array coil was used in all 3D cine experiments to enable highly accelerated imaging of rate 2 ⫻ 2. 3D GRE Cine CMR Although 3D cine GRE is not currently being used at 1.5T due to insufficient blood–myocardium CNR, it is commonly used in CMR at 3.0T because of SAR limita- Figure 2. a– h: Mid-diastolic 3D SSFP and GRE images acquired with different slab widths: 10 cm (a,e), 6.0 cm (b,f), 4.0 cm (c,g), and 3.0 cm (d,h). The images demonstrate the increased contrast found at thinner slab widths. Because all experiments are acquired with exactly the same parameters, the loss in contrast is known to be a direct result of saturation of inflowing spins. Note that the contrast in stationary tissues (e.g., the chest wall) is the same for all acquisitions. 1276 Data Analysis In whole heart 3D SSFP and GRE acquisitions, the SNR of arterial blood and myocardium were measured by drawing a region of interest (ROI) in the left ventricle and myocardium in the septum in the middle slice. The middle slice was chosen in all analyses to remove the effects of slice imperfection. Complex image data were not available due to use of partial Fourier acquisitions. Thus, the SNR analysis was based on the magnitude images. The standard deviation of the noise was measured using an ROI in the air across the chest wall. No corrections to the noise statistics were applied to account for the use of multi-channel magnitude images in the measurement process. SNR was calculated as the ratio of the mean signal to the standard deviation of the noise. CNR between the blood and myocardium was measured as the mean signal difference divided by the standard deviation of the noise. Regression analysis was used to calculate slope of SNR and CNR (i.e., rate of decline in SNR and CNR) in arbitrary unit. Percent change of SNR or CNR per cm increase in slab thickness was also calculated by dividing the calculated slope of each acquisition, calculated from the regression analysis, to the mean SNR or CNR of different subjects at 3 cm acquisition. Parallel imaging (SENSE) was used for 3D cine acquisitions making an absolute measure of SNR unfeasible. However, there was no difference in imaging parameters and identical coil sensitivity maps were used in all eight 3D cine acquisitions on each volunteer, keeping g-factors relatively constant. Therefore, “relative” SNR and CNR (i.e., SNR and CNR scaled with g-factor loss) were measured in 3D cine acquisitions and were directly comparable. The relative SNR was normalized to the maximum value observed in the Nezafat et al. respective measurement series (i.e., using the smallest slab thickness). Error bars were calculated as the standard deviation of the SNR measurements across different volunteers. RESULTS Figure 2 shows an example set of middle slices acquired using the 3D single phase whole heart SSFP and GRE sequences with RF excitation slab thicknesses of 10 cm, 6 cm, 4 cm, and 3 cm. These results show an improvement in contrast between ventricular blood and myocardium as the slab thickness decreases, which can only be associated with enhancement from inflowing spins. It is clear that the thicker slab thicknesses lead to saturation of spins flowing into the imaging volume. Compared to SSFP, there is more inflow contrast enhancement in the GRE images. Figure 3 shows the SNR and CNR measurements for single phase whole heart acquisitions. SNR measurements for 3D GRE acquisitions show that SNR is lost with a slope of 4.4 (5.6% loss per cm) and 1.5 (2.8% loss per cm) for arterial blood and myocardium, respectively (Fig. 3a). For SSFP acquisitions the slopes are 4.0 (5.3% loss per cm) and 1.2 (3.0% loss per cm) for arterial blood and myocardium, respectively (Fig. 3b). There is slight decline in myocardium SNR that could be associated with through plane motion of the myocardium. Blood– myocardium CNR shows losses associated with saturation of inflowing spins for both sequences with slopes of 2.4 (9.9% loss per cm) and 2.7 (8.2% loss per cm) for GRE and SSFP, respectively, with overall lower CNR in GRE (Fig. 3c,d). Blood–myocardium CNR was reduced by 73% and 59% by increasing the excitation slab Figure 3. a,b: SNRs for GRE (a) and SSFP (b) from 3D single phase whole heart acquisitions. Blood SNR declines with saturation for both sequences. c,d: Blood myocardium CNR measurements for GRE (c) and SSFP (d) imaging sequences show losses associated with saturation of inflowing spins for both sequences with overall lower CNR in the GRE. Error bars represent standard deviation. Inflow Quantification in 3D Cardiovascular MRI thickness from 3 cm to 10 cm in GRE and SSFP acquisitions, respectively. Figure 4 shows sample 3D cine images of middle slice acquired with GRE (top two rows) and SSFP (bottom two rows) in mid systole and end diastole. As with the single phase whole heart acquisitions, four slab sizes are shown: 10 cm, 6 cm, 4 cm, and 3 cm. An improvement in CNR with decreasing slab thickness can be seen in both acquisitions with a more pronounced change in images acquired with GRE. Comparing images acquired in mid-systole versus diastole shows that there are considerable changes in blood signal in mid-systolic phase, in which there is significant inflow, compared with the diastolic phase. These images suggest that the variability of inflow throughout the cardiac cycle has an impact on CNR and SNR. Figure 5 shows the results from relative SNR and CNR measurements made on cine acquisitions, both in mid-systole and late diastole. Systolic blood SNR in GRE images decreased with a slope of 27.3 (11.4% loss per cm), while diastolic blood SNR in GRE images decreased with a slope of 7.0 (7.0% loss per cm). A lower diastolic decline rate could be associated with a reduction in fresh inflowing blood during the diastolic phase (Fig. 5a). Myocardial SNR did not show considerable change with GRE (Fig. 5a). In SSFP images, both systolic and diastolic blood SNRs decreased with a slope of 19.2 (6.8% loss per cm) and 13.1 (5.4% loss per cm), respectively (Fig. 5b). Similar to the results observed with GRE, myocardial SNR in SSFP images showed little change (Fig. 5b). Overall, SSFP showed higher SNR for both diastolic and systolic phases compared with GRE. There was insufficient blood–myocardium CNR in Figure 4. Example 3D cine images in systole and diastole for GRE (top rows) and SSFP (bottoms rows) acquired with different slab sizes. The 3D SSFP images have higher CNR compared with GRE images. The CNR of SSFP images is improved with thinner slabs. 1277 the GRE cine acquisitions, with slopes of 27.2 (12.2% loss per cm) and 6.9 (8.4% loss per cm) for systolic and diastolic phases (Fig. 5c), respectively. The relative CNR of GRE acquisitions decreased when the excitation slab thickness was increased from 3 cm to 10 cm for both systolic (85%) and diastolic (64%) cardiac phases. The 3D SSFP cine images showed higher CNR with a decline slope of 18.9 (7.0% loss per cm) and 12.8 (5.6% loss per cm) for systole and diastole, respectively (Fig. 5d). The relative CNR of SSFP acquisitions decreased when the excitation slab thickness was increased from 3 cm to 10 cm for both systolic (50%) and diastolic (42%) cardiac phases. DISCUSSION In this study, we sought to investigate the saturation of inflowing spins in both 3D single phase whole heart and cine imaging. 3D cine is appealing due to ease of prescription, faster acquisition speed, and the reduction of slice misregistration which could result in miscalculation of left ventricle (LV) end diastolic and end systolic volumes. Although recent advances in coil technology have already enabled the acquisition of 3D cine images, image quality has not been as good as that observed with 2D cine imaging in all slices mainly due to loss of contrast (e.g., such as reduced CNR in LV apex (5,6)). Our results suggest that blood signal saturation may be one of the main sources of insufficient image quality in 3D cine. Other sources of poor image quality could be imperfect 3D slice profile, field inhomogeneity, mixing of blood components with a different excitation history profile, or phase errors that lead to flow artifacts due to undesired flow encoding by 1278 phase encoding gradients. Further study is required to quantify the CNR loss resulting from a transition from 2D, the gold-standard in CMR LV function evaluation, to thick slab 3D imaging, as used here. The losses in SNR and CNR quantified in this work reduce the effective gains expected when increasing the number of z-partitions in 3D CMR. New approaches should be tested and found to counteract the loss of contrast due to inflowing spins. These may include the use of contrast agents as well as other volumetric approaches such as multi-slab acquisitions. At 1.5T, 3D whole heart coronary artery imaging using the SSFP sequence is an alternative to targeted small slab acquisition using either SSFP or GRE. However, GRE acquisitions are more robust in the presence of artifacts from spins flowing through field inhomogeneities (18) and pericardial fluid. Nevertheless, whole heart imaging using GRE at 1.5T can be problematic due to insufficient contrast. At 3.0T, the GRE sequence has emerged as the choice for coronary imaging due to its robustness to field inhomogeneity and lower specific absorption rate. The methodology for inflow quantification proposed in this study could be used at 3.0T, with the caveat that additional signal loss would result from artifacts caused by flowing blood transiting through an inhomogeneous field. Additional experiments are required to quantify inflow effects in CMR at 3.0 T. Considering single phase whole heart imaging, we observed 73% and 59% losses in blood SNR in GRE and SSFP acquisitions, respectively, when comparing the thinnest slabs (3 cm) and the thickest (10 cm) slabs. The methodology used in this work was designed not to affect voxel SNR in the absence of inflow, keeping it Nezafat et al. constant relative to the acquisition with a 10 cm slab. Normally, the change from a 3 cm slab to a 10 cm slab (assuming 2 mm z-partitions) should yield an increase in SNR of 82% (⫽公(10/3) without considering blood saturation due to a decrease in the noise. Hence, the transition from thin slab to thick slab might still be advantageous from an SNR perspective, albeit with much reduced benefits. This SNR increase would also directly translate into a CNR increase of the same magnitude. This work shows that this theoretical CNR improvement is reduced by considering inflow saturation. Hence, alternative approaches should be considered to restore contrast and image quality. Preparatory pulses (e.g., T2 Prep or MT) or use of exogenous contrast agents might further assist in maintaining contrast. Study Limitations In the 3D cine studies, parallel imaging was used to reduce the acquisition time to one single breath-hold induration. Nevertheless, the breath-hold duration was significant (30 s), and not appropriate for use in a clinical examination. However, this study was designed to minimize confounding factors and imaging artifacts that could arise from nonbreath hold acquisitions, permitting accurate assessment of inflowing spin saturation. For example, respiratory navigators could interrupt the steadystate, while averaging could introduce blurring. The loss in blood SNR observed in this work is primarily due to saturation of the signal of inflowing spins. However, other second order effects could have also contributed to the loss. Though care was taken to minimize the differences between the 4 different Figure 5. a,b: Relative SNR for GRE (a) and SSFP (b) measured from 3D cine images for systolic blood, diastolic blood and myocardium. Relative CNRs for GRE (c) and SSFP (d) imaging sequences show that both systolic and diastolic CNRs decrease with increasing slab thickness. Error bars represent standard deviation. Inflow Quantification in 3D Cardiovascular MRI acquisitions (with four different slab thicknesses), the method of quantification used in this work does lead to one difference: the amplitude of the RF slab encoding gradient. Although the change in amplitude had negligible effects on the timing of the acquisition, it could have affected signal intensity. Intra-voxel dephasing is most prominent for stronger gradients, suggesting that experiments carried out with thinner slabs were more prone to signal loss in the blood pool. Hence, this mechanism could have biased the results by decreasing the SNR of blood observer in thinner slab acquisitions. Although the experiment could have been carried out using first moment nulling gradients which minimize intra-voxel dephasing, these gradients are too time consuming as they extend TE and TR and are not regularly used in cardiac imaging. Therefore, they were not used in this work. Although this might have introduced a bias into the resulting measurements, it is unlikely to be highly significant for the SNRs reported in this work, which were measured in the center of the blood pool. An accurate computer-controlled flow phantom or numerical simulation of inflow in the SSFP or GRE imaging sequences study could further validate this experimental imaging study. However, it is difficult to mimic in vivo conditions due to the complicated flow patterns observed in the cardiovascular system, especially in the ventricles. There are also additional in vivo conditions such as transition of blood flow through the inhomogeneous magnetic field (e.g., lung) or variations and position of the heart within the chest cavity that could further complicate such theoretical analysis. Although no flow phantom or numerical study is presented, this in vivo study is sufficient to demonstrate the loss of LV blood and myocardial CNR and SNR in 3D CMR. Furthermore, the static phantom study confirmed the accuracy of the imaging sequence by validating the slice profile and measuring the slice thickness. The SNR and CNR measurements were performed only in the middle slice to remove the effects of imperfections in the slice profile. Inflow measurements would differ if performed closer to base or apex of the heart, although it is likely the trends would have been similar to those observed in this work. Additionally, the measurements of inflow saturation were performed in a relatively young cohort of healthy subjects with relatively higher velocity of the LV blood flow (i.e., lower inflow saturation) compared with patients with cardiovascular disease. The changes in the SNR and CNR values presented in this study may differ for different imaging flip angles. For example, a higher imaging flip angle could increase saturation, yielding lower overall CNR. The flip angles chosen in this work reflect those used both in the literature and in our clinical practice, for both GRE and SSFP. Therefore, the results obtained within this work are likely relevant to the majority of cardiac MR scans performed using 3D imaging. Nevertheless, the effect of imaging flip angle on inflow saturation in the context of slab thickness needs further investigation. 1279 CONCLUSION In conclusion, in this study, we quantify the inflow enhancement in 3D single phase and 3D cine cardiac imaging. 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