Technical Note

Inflow Quantification in Three-DimensionalCardiovascular 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 lackthereof) in three-dimensional (3D) cardiovascular MR for bothsingle phase whole-heart and cine biventricular functions.

Materials and Methods: A 3D imaging sequence is pro-posed in which radiofrequency excitation gradient ischanged without modifying image acquisition or phase/slice encoding. This imaging sequence enables direct inflowmeasurement while retaining static voxel signal-to-noiseratio. Inflow measurements were performed for bothspoiled gradient-echo (GRE) imaging and balanced steady-state free precession (SSFP) in 18 healthy subjects.

Results: For single phase imaging, increasing slab thick-ness from 3 to 10 cm lead to 73% and 59% reductions incontrast-to-noise ratio (CNR) with GRE and SSFP, respec-tively. For cine acquisitions, systolic CNR was reduced by85% and 50% for the GRE and SSFP acquisitions, respec-tively, while diastolic CNR was reduced by 64% and 42%.

Conclusion: There is significant loss of CNR between bloodand myocardium when using larger 3D slabs due to satu-ration of inflowing spins. The loss of contrast is less pro-nounced for SSFP than for GRE, though both acquisitiontechniques suffer.

Key Words: 3D cardiac imaging, 3D whole heart imaging,inflow quantificationJ. Magn. Reson. Imaging 2008;28:1273–1279.© 2008 Wiley-Liss, Inc.

IN TWO-DIMENSIONAL (2D) balanced steady-state freeprecession (SSFP) and spoiled gradient echo (GRE) ac-quisitions, fresh and unsaturated spins are continu-ously flowing into the imaging slice. These inflow spins

tend to improve the blood–myocardium contrast-to-noise ratio (CNR) as they yield higher blood signal tonoise ratio (SNR), particularly for SSFP sequences thatuse high imaging flip angles. The 3D cardiovascular MR(CMR) imaging is appealing due to its potentially higherSNR compared with 2D imaging (1). The 3D CMR imag-ing has recently gained importance due to the availabil-ity of phased arrays with a higher number of coil ele-ments that enable the use of 2D parallel imaging toreduce total acquisition time (2–6). Whole heart 3Dcoronary imaging, which applies a single large slab overthe entire heart in an axial orientation obviates the useof multiple thin slab 3D acquisitions targeted to theright and left coronary arteries (6–8). Similarly, singlebreath-hold (9) or self-navigated 3D cine acquisitionsare now becoming feasible. In addition to SNR gain in3D cine, slice misregistration is typically improved, per-mitting 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 imagingis more robust at higher magnetic fields (e.g., 3.0T),which display increased field inhomogeneity. Imagingof coronary vein anatomy for evaluation of patients un-dergoing cardiac resynchronization therapy has alsobeen shown to be more robust with GRE than SSFPeven at 1.5T, mainly due to increased B0 field inhomo-geneity surrounding the veins (12).

The transition from targeted thin 3D slab acquisi-tions to thick 3D slab whole heart should be advanta-geous with both GRE and SSFP due to higher spatialcoverage and potential SNR increase of �Nz, where Nz isthe number of slice partitions or encoding steps (1). Theeffects of slab size on inflow-based image contrast areyet to be explored in a systematic manner.

In 2D and small slab 3D imaging, fresh inflow resultsin blood signal enhancement thus blood–myocardiumCNR improvement. With large 3D slab acquisitions, in-flowing spins experience a higher number of radiofre-quency (RF) pulses within the imaging volume, drivingthe magnetization closer to its steady-state magnitude(13). Therefore, the effect of inflowing spins on imagecontrast is expected to be greatly reduced (13). In theabsence of inflow enhancement, GRE produces a T1weighting, leading to a reduction of blood–myocardiumCNR because both tissues have very similar T1s (1). In

1Departments of Medicine (Cardiovascular Division) and 2Radiology,Beth Israel Deaconess Medical Center and Harvard Medical School,Beth Israel Deaconess Medical Center, Boston, Massachusetts.3Clinical Sites Research Program, Philips Research North America,Briarcliff Manor, NY.4Tomographic 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 Cen-ter, 330 Brookline Avenue, Boston, MA, 02215.E-mail: rnezafat@bidmc.harvard.eduReceived February 19, 2008; Accepted June 6, 2008.Grant sponsor: American Heart Association; Grant number AHA SDG-0730339NDOI 10.1002/jmri.21493Published online in Wiley InterScience (www.interscience.wiley.com).

JOURNAL OF MAGNETIC RESONANCE IMAGING 28:1273–1279 (2008)

© 2008 Wiley-Liss, Inc. 1273

contrast, SSFP yields images with T2/T1 contrast,which in the absence of any inflow enhancement, mightbe more advantageous in terms of blood–myocardiumCNR (1). Also, because the innate contrast produced bythese sequences may not be as good as that observedwith thin slab acquisitions, prepulses such as T2 mag-netization preparation (14,15) or magnetization trans-fer (MT) may be used to create additional contrast(12,16). However, during the long transit time of theblood through a thick 3D-slab, considerable accumu-lative phase errors may evolve due to flow encoding ofthe phase encoding gradients (17).

In this study, we sought to investigate inflow satura-tion in 3D single phase and cine CMR imaging by quan-tifying the contribution of inflowing spins to SNR andCNR with varying slab thickness. An imaging sequenceis designed such that RF excitation slab thickness isvaried in different acquisitions, thereby enabling directinflow measurement while retaining static voxel SNR.

MATERIALS AND METHODS

To study the effect of inflowing blood in a large slabprescription, a series of imaging studies were performedin each subject. The RF excitation to image encodingratio, that is, excitation slab thickness relative to imageslab size, was changed between different image acqui-sitions by changing the amplitude of the slab selectiongradient of Gz (i.e., reducing the excitation size) betweendifferent acquisitions, as shown in Figure 1. A sliceselective 3D acquisition is used with excited RF profilecovering the entire imaging volume. The excitation slabthickness was reduced manually in consecutive acqui-sitions without modifying the image acquisition vol-ume, that is, frequency and phase encoding steps. Bykeeping the imaging parameters of slice, phase andfrequency encoding the same between all acquisitions,the SNR and CNR changes between different acquisi-tions can be associated with the saturation of inflowing

spins. The flowing spins into the imaging plane willexperience different RF histories based on the excita-tion profile.

MR imaging studies were performed on two 1.5T Phil-ips Achieva (Philips Medical Systems, Best, NL). Onesystem was equipped with a 16-channel receiver and a5-element cardiac phased-array receiver coil for singlephase whole heart coronary acquisitions. The secondsystem used a 32-channel receiver and a 32-elementcardiac phased array for 3D cine scans.

In Vivo Studies

Three groups of healthy adult subjects were imaged. Ineach group, images were acquired in 6 subjects (total-ing 18 subjects, 10 females, mean age 27 years). For 3Dcine imaging, 6 healthy subjects were studied using a32-channel coil with both 3D SSFP and GRE cine. Forsingle phase imaging, the other 12 subjects were stud-ied using a 5-element coil using either 3D single phaseGRE (in 6 subjects) or 3D single phase SSFP (in 6 othersubjects). Written informed consent was obtained fromall participants and the protocols were approved byboth Institutional Review Boards.

3D GRE Single Phase Whole Heart CMR

An electrocardiogram triggered, segmented 2-chamberSSFP cine (repetition time [TR] � 3 ms, echo time [TE] �1.4 ms, � � 60°, temporal resolution of 29 ms, spatialresolution of 1.6 � 2.0 mm2 reconstructed to 1.21 �1.21 mm2) dataset was acquired at the level of themid-ventricle to identify the onset of the diastolic restperiod. This was followed by four whole heart CMRimaging studies that differed only in RF slab encodinggradient amplitude. A single phase sequence similar toone used for coronary artery and vein imaging was used(7,12). To improve blood–myocardium contrast, an MTpulse train consisting of eight 600° Sinc-Gaussian RF

Figure 1. Image acquisition set-up: a large 3D axial 10-cm slab (100 slices of 1 mm thickness) is prescribed to cover the entireheart. As usual, frequency (x), phase (y), and slice (z) encoding are performed. In the reference experiment, the imaging RF pulseexcites the entire 10 cm volume, as shown, with an extra 20% slice encoding to prevent fold-over from RF imperfections. Insubsequent 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 leftunchanged and cover the entire volume. This experiment maintains voxel size and scan time constant, and, therefore, changesin SNR and CNR can be attributed to inflowing spins.

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pulses, each with duration of 15 ms and frequencyoff-set of 500 Hz was used for all MT preparations (12).A spectrally selective fat saturation sequence was alsoused to suppress the fat signal. The navigator beam waspositioned at the dome of the right hemidiaphragm withan acceptance window of 7 mm, automatic respiratorydrift adaptation, and RF excitation angle of 25°. Anaxial volume covering the entire heart was prescribedoff 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 100one-mm-thick slices. For imaging, 20–25 RF excita-tions (asymmetric Gaussian-weighted-Sinc with fourand two lobes and duration of 672 �s) with spoiled GREreadouts: TE � 1.1 ms, TR � 3.6 ms, � � 30°, BW � 383Hz/Pixel, acquisition time of � 4:40 min for heart rateof 60 bpm. Partial echo (62.5%) was used in the acqui-sition. A field of view (FOV) of 270 � 270 mm2 wasimaged with a scan matrix of 140 � 140 yielding a voxelsize of 2 � 2 � 2 mm3 reconstructed to a 0.9 � 0.9 � 1mm3. A total of four studies with slab thickness of 10cm, 6 cm, 4 cm, and 3 cm were acquired. Due to thelong duration of these studies, the images were pre-scribed with lower spatial resolution to ensure the suc-cessful completion of all 4 studies in all subjects. Noparallel imaging was used to enable absolute SNR andCNR comparisons.

3D SSFP Single Phase Whole Heart CMR

For 3D SSFP inflow quantification, a similar study tothe one used for 3D GRE (previous section) was per-formed on six subjects. The following imaging parame-ters, if different from above, were used: TE � 1.8 ms,TR � 3.7 ms, � � 90°, BW � 1470 Hz/Pixel. An excita-tion Gaussian-weighted Sinc RF pulse with three lobesand duration of 1.49 ms was used for SSFP imaging.

3D SSFP Cine CMR

A 3D cardiac cine acquisition was prescribed to imagethe heart throughout the cardiac cycle. The imagingparameters were as follows: TR � 3.6 ms, TE � 1.8 ms,BW � 1562 Hz/Pixel, �� 60°. A 10-cm 3D slab consist-ing of 25 slices (interpolated to 50 slices) was acquiredusing a FOV of 270 � 220 mm2, with a scan matrix of128 � 102, yielding a voxel size of 2.1 � 2.1 � 4 mm3.A 2D acceleration of rate 2 � 2 in the phase and sliceencoding directions and a partial Fourier acquisition(reduction factor 0.625) were used to reduce the acqui-sition time to fit within a single breath-hold (30 s).Although respiratory navigators can be used to reducethe time constraint on image acquisition, transition tosteady-state artifacts that could present a complicatinginfluence on the measurement of inflow effects. A 32-element phased-array coil was used in all 3D cine ex-periments to enable highly accelerated imaging of rate2 � 2.

3D GRE Cine CMR

Although 3D cine GRE is not currently being used at1.5T due to insufficient blood–myocardium CNR, it iscommonly used in CMR at 3.0T because of SAR limita-

tions and imaging artifacts with SSFP from increasedfield inhomogeneity. However, for consistency, in thisstudy, 3D GRE cine images were acquired at 1.5T aswell. The imaging parameters were: TR � 3.6 ms, TE �1.3 ms, BW � 1562 Hz/Pixel, and � � 20°. The scangeometry and matrix size were identical to those usedfor the SSFP acquisition with the same acceleration rateand coil. Both 3D GRE and SSFP cine images wereacquired in one imaging session in one group of healthysubjects.

Figure 2. a–h: Mid-diastolic 3D SSFP and GRE images ac-quired with different slab widths: 10 cm (a,e), 6.0 cm (b,f), 4.0cm (c,g), and 3.0 cm (d,h). The images demonstrate the in-creased contrast found at thinner slab widths. Because allexperiments are acquired with exactly the same parameters,the loss in contrast is known to be a direct result of saturationof inflowing spins. Note that the contrast in stationary tissues(e.g., the chest wall) is the same for all acquisitions.

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Data Analysis

In whole heart 3D SSFP and GRE acquisitions, the SNRof arterial blood and myocardium were measured bydrawing a region of interest (ROI) in the left ventricleand myocardium in the septum in the middle slice. Themiddle slice was chosen in all analyses to remove theeffects of slice imperfection. Complex image data werenot available due to use of partial Fourier acquisitions.Thus, the SNR analysis was based on the magnitudeimages. The standard deviation of the noise was mea-sured using an ROI in the air across the chest wall. Nocorrections to the noise statistics were applied to ac-count for the use of multi-channel magnitude images inthe measurement process. SNR was calculated as theratio of the mean signal to the standard deviation of thenoise. CNR between the blood and myocardium wasmeasured as the mean signal difference divided by thestandard deviation of the noise. Regression analysiswas used to calculate slope of SNR and CNR (i.e., rate ofdecline in SNR and CNR) in arbitrary unit. Percentchange of SNR or CNR per cm increase in slab thicknesswas also calculated by dividing the calculated slope ofeach acquisition, calculated from the regression analy-sis, to the mean SNR or CNR of different subjects at 3cm acquisition. Parallel imaging (SENSE) was used for3D cine acquisitions making an absolute measure ofSNR unfeasible. However, there was no difference inimaging parameters and identical coil sensitivity mapswere used in all eight 3D cine acquisitions on eachvolunteer, keeping g-factors relatively constant. There-fore, “relative” SNR and CNR (i.e., SNR and CNR scaledwith g-factor loss) were measured in 3D cine acquisi-tions and were directly comparable. The relative SNRwas normalized to the maximum value observed in the

respective measurement series (i.e., using the smallestslab thickness). Error bars were calculated as the stan-dard deviation of the SNR measurements across differ-ent volunteers.

RESULTS

Figure 2 shows an example set of middle slices acquiredusing the 3D single phase whole heart SSFP and GREsequences with RF excitation slab thicknesses of 10cm, 6 cm, 4 cm, and 3 cm. These results show animprovement in contrast between ventricular blood andmyocardium as the slab thickness decreases, whichcan only be associated with enhancement from inflow-ing spins. It is clear that the thicker slab thicknesseslead to saturation of spins flowing into the imagingvolume. Compared to SSFP, there is more inflow con-trast enhancement in the GRE images.

Figure 3 shows the SNR and CNR measurements forsingle phase whole heart acquisitions. SNR measure-ments for 3D GRE acquisitions show that SNR is lostwith a slope of 4.4 (5.6% loss per cm) and 1.5 (2.8% lossper 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 bloodand myocardium, respectively (Fig. 3b). There is slightdecline in myocardium SNR that could be associatedwith through plane motion of the myocardium. Blood–myocardium CNR shows losses associated with satura-tion of inflowing spins for both sequences with slopes of2.4 (9.9% loss per cm) and 2.7 (8.2% loss per cm) forGRE and SSFP, respectively, with overall lower CNR inGRE (Fig. 3c,d). Blood–myocardium CNR was reducedby 73% and 59% by increasing the excitation slab

Figure 3. a,b: SNRs for GRE(a) and SSFP (b) from 3D singlephase whole heart acquisi-tions. Blood SNR declines withsaturation for both sequences.c,d: Blood myocardium CNRmeasurements for GRE (c) andSSFP (d) imaging sequencesshow losses associated withsaturation of inflowing spinsfor both sequences with overalllower CNR in the GRE. Errorbars represent standard devia-tion.

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thickness from 3 cm to 10 cm in GRE and SSFP acqui-sitions, respectively.

Figure 4 shows sample 3D cine images of middle sliceacquired with GRE (top two rows) and SSFP (bottom tworows) in mid systole and end diastole. As with the singlephase whole heart acquisitions, four slab sizes areshown: 10 cm, 6 cm, 4 cm, and 3 cm. An improvementin CNR with decreasing slab thickness can be seen inboth acquisitions with a more pronounced change inimages acquired with GRE. Comparing images acquiredin mid-systole versus diastole shows that there are con-siderable changes in blood signal in mid-systolic phase,in which there is significant inflow, compared with thediastolic phase. These images suggest that the variabil-ity of inflow throughout the cardiac cycle has an impacton CNR and SNR.

Figure 5 shows the results from relative SNR andCNR measurements made on cine acquisitions, both inmid-systole and late diastole. Systolic blood SNR inGRE images decreased with a slope of 27.3 (11.4% lossper cm), while diastolic blood SNR in GRE images de-creased with a slope of 7.0 (7.0% loss per cm). A lowerdiastolic decline rate could be associated with a reduc-tion in fresh inflowing blood during the diastolic phase(Fig. 5a). Myocardial SNR did not show considerablechange with GRE (Fig. 5a). In SSFP images, both sys-tolic and diastolic blood SNRs decreased with a slope of19.2 (6.8% loss per cm) and 13.1 (5.4% loss per cm),respectively (Fig. 5b). Similar to the results observedwith GRE, myocardial SNR in SSFP images showedlittle change (Fig. 5b). Overall, SSFP showed higher SNRfor both diastolic and systolic phases compared withGRE. There was insufficient blood–myocardium CNR in

the GRE cine acquisitions, with slopes of 27.2 (12.2%loss per cm) and 6.9 (8.4% loss per cm) for systolic anddiastolic phases (Fig. 5c), respectively. The relative CNRof GRE acquisitions decreased when the excitation slabthickness was increased from 3 cm to 10 cm for bothsystolic (85%) and diastolic (64%) cardiac phases. The3D SSFP cine images showed higher CNR with a declineslope of 18.9 (7.0% loss per cm) and 12.8 (5.6% loss percm) for systole and diastole, respectively (Fig. 5d). Therelative CNR of SSFP acquisitions decreased when theexcitation slab thickness was increased from 3 cm to 10cm for both systolic (50%) and diastolic (42%) cardiacphases.

DISCUSSION

In this study, we sought to investigate the saturation ofinflowing spins in both 3D single phase whole heart andcine imaging. 3D cine is appealing due to ease of prescrip-tion, faster acquisition speed, and the reduction of slicemisregistration which could result in miscalculation ofleft ventricle (LV) end diastolic and end systolic volumes.Although recent advances in coil technology have alreadyenabled the acquisition of 3D cine images, image qualityhas not been as good as that observed with 2D cine im-aging in all slices mainly due to loss of contrast (e.g., suchas reduced CNR in LV apex (5,6)). Our results suggest thatblood signal saturation may be one of the main sources ofinsufficient image quality in 3D cine. Other sources ofpoor image quality could be imperfect 3D slice profile,field inhomogeneity, mixing of blood components with adifferent excitation history profile, or phase errors thatlead to flow artifacts due to undesired flow encoding by

Figure 4. Example 3D cine images in systoleand diastole for GRE (top rows) and SSFP (bot-toms rows) acquired with different slab sizes.The 3D SSFP images have higher CNR com-pared with GRE images. The CNR of SSFP im-ages is improved with thinner slabs.

Inflow Quantification in 3D Cardiovascular MRI 1277

phase encoding gradients. Further study is required toquantify the CNR loss resulting from a transition from 2D,the gold-standard in CMR LV function evaluation, to thickslab 3D imaging, as used here. The losses in SNR andCNR quantified in this work reduce the effective gainsexpected when increasing the number of z-partitions in3D CMR. New approaches should be tested and found tocounteract the loss of contrast due to inflowing spins.These may include the use of contrast agents as well asother volumetric approaches such as multi-slab acquisi-tions.

At 1.5T, 3D whole heart coronary artery imaging us-ing the SSFP sequence is an alternative to targetedsmall slab acquisition using either SSFP or GRE. How-ever, GRE acquisitions are more robust in the presenceof artifacts from spins flowing through field inhomoge-neities (18) and pericardial fluid. Nevertheless, wholeheart imaging using GRE at 1.5T can be problematicdue to insufficient contrast. At 3.0T, the GRE sequencehas emerged as the choice for coronary imaging due toits robustness to field inhomogeneity and lower specificabsorption rate. The methodology for inflow quantifica-tion proposed in this study could be used at 3.0T, withthe caveat that additional signal loss would result fromartifacts caused by flowing blood transiting through aninhomogeneous field. Additional experiments are re-quired to quantify inflow effects in CMR at 3.0 T.

Considering single phase whole heart imaging, weobserved 73% and 59% losses in blood SNR in GRE andSSFP acquisitions, respectively, when comparing thethinnest slabs (3 cm) and the thickest (10 cm) slabs.The methodology used in this work was designed not toaffect voxel SNR in the absence of inflow, keeping it

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 increasein SNR of 82% (��(10/3) without considering bloodsaturation due to a decrease in the noise. Hence, thetransition from thin slab to thick slab might still beadvantageous from an SNR perspective, albeit withmuch reduced benefits. This SNR increase would alsodirectly translate into a CNR increase of the same mag-nitude. This work shows that this theoretical CNR im-provement is reduced by considering inflow saturation.Hence, alternative approaches should be considered torestore contrast and image quality. Preparatory pulses(e.g., T2 Prep or MT) or use of exogenous contrast agentsmight further assist in maintaining contrast.

Study Limitations

In the 3D cine studies, parallel imaging was used to re-duce the acquisition time to one single breath-hold indu-ration. Nevertheless, the breath-hold duration was signif-icant (30 s), and not appropriate for use in a clinicalexamination. However, this study was designed to mini-mize confounding factors and imaging artifacts that couldarise from nonbreath hold acquisitions, permitting accu-rate assessment of inflowing spin saturation. For exam-ple, respiratory navigators could interrupt the steady-state, while averaging could introduce blurring.

The loss in blood SNR observed in this work isprimarily due to saturation of the signal of inflowingspins. However, other second order effects could havealso contributed to the loss. Though care was takento minimize the differences between the 4 different

Figure 5. a,b: Relative SNR for GRE(a) and SSFP (b) measured from 3Dcine images for systolic blood, dia-stolic blood and myocardium. Rela-tive CNRs for GRE (c) and SSFP (d)imaging sequences show that bothsystolic and diastolic CNRs decreasewith increasing slab thickness. Errorbars represent standard deviation.

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acquisitions (with four different slab thicknesses),the method of quantification used in this work doeslead to one difference: the amplitude of the RF slabencoding gradient. Although the change in amplitudehad negligible effects on the timing of the acquisition,it could have affected signal intensity. Intra-voxeldephasing is most prominent for stronger gradients,suggesting that experiments carried out with thinnerslabs were more prone to signal loss in the blood pool.Hence, this mechanism could have biased the resultsby decreasing the SNR of blood observer in thinnerslab acquisitions. Although the experiment couldhave been carried out using first moment nullinggradients which minimize intra-voxel dephasing,these gradients are too time consuming as they ex-tend TE and TR and are not regularly used in cardiacimaging. Therefore, they were not used in this work.Although this might have introduced a bias into theresulting measurements, it is unlikely to be highlysignificant for the SNRs reported in this work, whichwere measured in the center of the blood pool.

An accurate computer-controlled flow phantom ornumerical simulation of inflow in the SSFP or GREimaging sequences study could further validate thisexperimental imaging study. However, it is difficult tomimic in vivo conditions due to the complicated flowpatterns observed in the cardiovascular system, es-pecially in the ventricles. There are also additional invivo conditions such as transition of blood flowthrough the inhomogeneous magnetic field (e.g.,lung) or variations and position of the heart withinthe chest cavity that could further complicate suchtheoretical analysis. Although no flow phantom ornumerical study is presented, this in vivo study issufficient to demonstrate the loss of LV blood andmyocardial CNR and SNR in 3D CMR. Furthermore,the static phantom study confirmed the accuracy ofthe imaging sequence by validating the slice profileand measuring the slice thickness.

The SNR and CNR measurements were performed onlyin the middle slice to remove the effects of imperfections inthe slice profile. Inflow measurements would differ if per-formed closer to base or apex of the heart, although it islikely the trends would have been similar to those ob-served in this work. Additionally, the measurements ofinflow saturation were performed in a relatively youngcohort of healthy subjects with relatively higher velocity ofthe LV blood flow (i.e., lower inflow saturation) comparedwith patients with cardiovascular disease.

The changes in the SNR and CNR values presentedin this study may differ for different imaging flip an-gles. For example, a higher imaging flip angle couldincrease saturation, yielding lower overall CNR. Theflip angles chosen in this work reflect those used bothin the literature and in our clinical practice, for bothGRE and SSFP. Therefore, the results obtained withinthis work are likely relevant to the majority of cardiacMR scans performed using 3D imaging. Nevertheless,the effect of imaging flip angle on inflow saturation inthe context of slab thickness needs further investiga-tion.

CONCLUSION

In conclusion, in this study, we quantify the inflowenhancement in 3D single phase and 3D cine cardiacimaging. The results show there is significant loss ofCNR between blood and myocardium when using larger3D slabs due to saturation of inflowing spins. The lossof contrast is less pronounced for SSFP than for GRE,though both acquisition techniques suffer.

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