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Published online before print January 25, 2008, 10.1148/radiol.2463062155
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(Radiology 2008;246:917-925.)
© RSNA, 2008

Technical Developments

Three-directional Myocardial Phase-Contrast Tissue Velocity MR Imaging with Navigator-Echo Gating: In Vivo and in Vitro Study1

Jana G. Delfino, PhD, Kevin R. Johnson, PhD, Robert L. Eisner, PhD, Susan Eder, RT, Angel R. Leon, MD, and John N. Oshinski, PhD

1 From the Department of Biomedical Engineering, Georgia Institute of Technology/Emory University, 101 Woodruff Cir, Suite 2001, Atlanta, GA 30322 (J.G.D., K.R.J., J.N.O.); and Departments of Radiology (R.L.E., J.N.O.) and Medicine, Division of Cardiology (S.E., A.R.L.), Emory University School of Medicine, Atlanta, Ga. Received December 18, 2006; revision requested February 16, 2007; revision received March 23; accepted April 25; final version accepted August 1. Supported by the Wallace H. Coulter Foundation. Address correspondence to J.G.D. (e-mail: jana.delfino{at}gatech.edu).


   ABSTRACT
TOP
 ABSTRACT
INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
DISCUSSION
 ADVANCES IN KNOWLEDGE
IMPLICATION FOR PATIENT CARE...
 References
 
The study protocol was HIPAA compliant and institutional review board approved. Informed consent was obtained from all participants. The purpose of the study was to prospectively validate the capability of navigator-echo–gated phase-contrast magnetic resonance (MR) imaging for measurement of myocardial velocities in a phantom and to prospectively use the phase-contrast MR sequence to measure three-directional velocity in the myocardium in vivo in volunteers and in patients scheduled for cardiac resynchronization therapy. An excellent correlation between the measured velocity and the true phantom motion (R = 0.90 for longitudinal velocity, R = 0.93 for circumferential velocity) was observed. Myocardial velocities were successfully measured in 17 healthy volunteers (11 male, six female; mean age, 27.5 years ± 6.5 [standard deviation]) and 28 patients with heart failure (18 male, 10 female; mean age, 63.9 years ± 15.0). Velocity values were significantly lower in the patients than in the volunteers. The time to peak velocity in the lateral wall of the patients, as compared with that in the volunteers, was delayed. Phase-contrast MR imaging can be combined with navigator-echo gating to measure three-directional myocardial tissue velocities in vivo.

Supplemental material: http://radiology.rsnajnls.org/cgi/content/full/2463062155/DC1

© RSNA, 2008


   INTRODUCTION
TOP
 ABSTRACT
 INTRODUCTION
MATERIALS AND METHODS
 RESULTS
DISCUSSION
 ADVANCES IN KNOWLEDGE
IMPLICATION FOR PATIENT CARE...
 References
 
Myocardial contraction is a complex three-directional motion involving longitudinal and radial shortening, torsion, and shear. Characterizing this motion is of interest because the myocardial contraction pattern is a direct measure of the function and viability of the heart (1). Multiple studies have revealed that alterations in this contraction pattern are predictive of cardiac disease and transplant rejection (2–4).

Tissue Doppler imaging is often used to study myocardial motion and velocity. It has been used specifically to identify mechanical delays and changes in the longitudinal velocity in patients with dyssynchrony (5–12). However, some study results suggest that the timing of radial and circumferential shortening may be more sensitive than the longitudinal velocity for the identification of dyssynchrony (13). Tissue Doppler imaging, however, is usually restricted to measurements of local long-axis velocities near the base of the left ventricle (LV). Furthermore, image quality depends on patient habitus, sonographer skill, and Doppler beam angle.

Phase-contrast magnetic resonance (MR) imaging also can be used to measure myocardial velocity, with the benefit of not being restricted to one direction. With use of cardiac and navigator-echo respiratory gating combined with phase-contrast MR imaging, one can measure velocity in the myocardium with high spatial and temporal resolution without patient breath holding. Since three-directional velocity information is available for each voxel, one can determine the location, timing, and magnitude of mechanical contraction in any direction.

Although previous studies have revealed excellent correlation between phase-contrast MR–derived and tissue Doppler imaging–derived myocardial velocities, differences in the magnitudes of velocity measured at phase-contrast MR and tissue Doppler imaging have been observed (14–16). Since the true myocardial velocity cannot be easily determined in vivo, assessing the accuracy of phase-contrast MR velocity measurements is difficult. The purpose of our study was to prospectively validate the capability of navigator-echo–gated phase-contrast MR imaging for measurement of myocardial velocities in a phantom and to prospectively use the phase-contrast MR sequence to measure three-directional velocity in the myocardium in vivo in volunteers and in patients scheduled for cardiac resynchronization therapy (CRT).


   MATERIALS AND METHODS
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
RESULTS
DISCUSSION
 ADVANCES IN KNOWLEDGE
IMPLICATION FOR PATIENT CARE...
 References
 
Phantom Study
Description of motion phantom.—An MR-compatible model (Fig 1) of the LV was constructed from two concentric cylinders—representing the epicardial and endocardial surfaces of the LV—whose dimensions were based on in vivo measurements reported in the literature (2). Polyvinyl alcohol cryogel, a material with T1 and T2 similar to those in the myocardium, simulated myocardial tissue between the two cylinders (17).


Figure 1
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  		Figure 1: Motion phantom. Top: Schematic illustration of entire phantom setup shows the linear and rotational motor controllers and the LV phantom in the MR magnet bore. Bottom: Close-up photograph of the phantom model shows a polyvinyl alcohol cryogel–filled cylinder simulating the LV; the motion of the LV is controlled by two independent piezoelectric motors that generate rotational and through-plane motion. The phantom experiments revealed the capability of phase-contrast MR imaging for measurement of three-directional velocities in the myocardium.

 
The three-directional motion values for the phantom were based on previous measurements of coronary artery motion in vivo: The maximal rotational displacement was 6°, and the maximal longitudinal displacement was 1.6 cm (18). Two computer-controlled piezoelectric high-force motors (Bayside Manufacturing, Port Washington, NY) controlled the phantom's movement: One motor moved the phantom linearly along the axis of the MR magnet bore, while the other rotated it in the transverse plane. A movie demonstrating the motion of the phantom model is available online as supplemental material (http://radiology.rsnajnls.org/cgi/content/full/2463062155/DC1). A 3-m-long, 2.5-cm-diameter acrylic rod connected the phantom to the motors, allowing the motors to remain outside the 5-G (0.0005-T) line. A series of support cradles ensured that the location of the phantom correctly simulated the location of the human heart within the MR magnet bore.

The motors were controlled by a computer program (Galil; Galil Motion Control, Rocklin, Calif), which used the input data in time-versus-position pairs. A 5-V pulse at the start of each cycle served as the electrocardiographic trigger input to the MR imaging unit (Intera CV; Philips Medical Systems, Best, the Netherlands). The program ran an infinite loop to simulate the multiple heartbeats needed for MR imaging. The system was equipped with a series of feedback sensors that measured the actual displacement versus time of both motors. The displacement information served as the reference standard for the position of the phantom over time. The LV phantom was placed inside a static chest wall phantom during imaging. We simulated the chest wall by filling the space between two concentric cylinders—which had an outer diameter of 26 cm and an inner diameter of 18 cm—with polyvinyl alcohol cryogel, as in the LV phantom. Velocity in the static chest wall was used for background phase offset correction (19).

MR imaging.—MR images were acquired by using a five-element phased-array cardiac coil (Philips Medical Systems) during one imaging session. Velocity images were acquired by using an electrocardiographically gated, segmented (three lines of k-space per shot), gradient-echo phase-contrast sequence. Velocity encoding was performed in a Hadamard fashion by using four-point velocity vector extraction with encoding for different directions performed in different heartbeats (20). Acquisition parameters were as follows: 7/4 (repetition time msec/echo time msec), 1.4-mm in-plane resolution, 8-mm section thickness, velocity-encoding value of 30 cm/sec, 15° flip angle, and 370-mm field of view. Raw data were saved, and a separate delayed reconstruction was used to extract velocity information for each direction.

In Vivo Study
Study participants.—Seventeen healthy volunteers (11 male, six female; mean age, 27.5 years ± 6.5 [standard deviation]; age range, 20–45 years) and twenty-eight patients (18 male, 10 female; mean age, 63.9 years ± 15.0; age range, 43–86 years) participated in this study. The volunteers were recruited from August 2004 to May 2005, had no history of cardiovascular disease, and had a normal 12-lead electrocardiogram. The patients were recruited from May 2004 to January 2005 from the electrophysiology clinic where they were being evaluated for CRT. Inclusion criteria were New York Heart Association class III or IV heart failure, electrocardiographic evidence of dyssynchrony (QRS duration > 120 msec), and LV ejection fraction lower than 35%. The patients were receiving optimal medical therapy, which included β-blockers, angiotensin-converting enzyme inhibitors, angiotensinogen receptor blockers, diuretics, aldosterone antagonists, and digoxin. The study protocol was compliant with Health Insurance Portability and Accountability Act regulations and was approved by the institutional review board of Emory University. Informed consent was obtained from all participants.

MR imaging.—MR imaging was performed in the volunteers and patients by one trained technologist (S.E.) with 23 years MR experience by using the same MR unit that was used in the phantom experiments. After survey images were acquired, two-chamber vertical long-axis, four-chamber horizontal long-axis, and short-axis steady-state free precession cine images were obtained. The technologist measured the length of the LV from the apex to the mitral valve plane on an end-diastolic horizontal long-axis image, and a short-axis orientation at 70% of the distance from the apex to the base was chosen for the phase-contrast MR velocity image acquisition. The same velocity imaging protocol used in the phantom experiments was used, with the addition of two presaturation slabs—each 30 mm thick and placed 10 mm away from the imaging section—for suppression of the signal from the blood pool and navigator-echo gating for respiratory compensation (21,22).

A trailing navigator with an acceptance window of 6 mm was placed on the diaphragm at the lung-liver interface. The navigator pulse was applied during end diastole and took approximately 74 msec to execute. Navigator-echo gating enabled velocity data acquisition at high spatial and temporal resolution and ensured that all three velocity-encoding directions were properly registered for postprocessing (Fig 2). The temporal resolution between cardiac phases was 35 msec. The time for velocity-encoding image acquisition was 1 minute 30 seconds with a navigator efficiency of 100%; the actual imaging time depended on the navigator efficiency, which ranged from 30% to 80% (Fig 3). The mean time for the phase-contrast MR examination was approximately 5 minutes ± 3 (standard deviation). Phase-contrast MR tissue velocity maps were successfully obtained in all volunteers and patients. The quality of all acquired images was sufficient for interpretation.


Figure 2
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  		Figure 2: Sagittal ventricular short-axis phase-contrast MR images (7/4) acquired during diastolic relaxation with the velocity mapping technique. A, Magnitude image, and, B–D, phase velocity maps of left-to-right motion (B), anterior-to-posterior motion (C), and apex-to-base motion (D) are shown. All images are correctly registered for postprocessing. On B–D, moving objects are either light, indicating movement in the positive velocity direction, or dark, indicating movement in the negative velocity direction, while static objects are gray. Air in the lungs and outside the chest walls appears as noise.

 

Figure 3
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  		Figure 3: Schematic illustration of navigator-echo–gated three-directional phase-contrast MR sequence used to measure myocardial velocity in vivo. Velocity encoding was performed in a Hadamard fashion by using four-point velocity vector extraction with encoding for different velocity directions performed in different heartbeats. The bipolar flow-encoding gradient is shown in gray. A trailing navigator (NAV) during end diastole was used for respiratory compensation. k1, k2, and k3 refer to three different lines of k-space. P1, P2, P3, and PN refer to the different temporal phases in the acquired velocity data. Vx, Vy, and Vz are the three acquired velocity directions. X, Y, and Z are the three (section-select, phase-encode, and frequency-encode) gradient axes. RF = radiofrequency.

 
Data Analyses
MATLAB programs (R2006a; Mathworks, Natick, Mass) to automatically process the phase-contrast MR velocity images obtained in both the phantom and the human subjects were developed in house (J.G.D.). Background phase errors were removed by using a least-squares plane fitted to static tissue (19). The acquired three-directional velocity measurements were converted to radial (positive toward center of LV blood pool), longitudinal (positive toward apex), and circumferential (positive for clockwise rotation when viewed from apex) velocities. The times of aortic valve opening and closing on the LV outflow tract images were used to define systole.

In each direction of motion (radial, circumferential, and longitudinal), the velocity was averaged throughout the imaging section to generate a velocity-versus-time curve. The peak and time to peak systolic and diastolic velocities were computed for each curve. The magnitudes and times of peak velocity in the three directions of motion during systole and diastole were compared. Normal systolic and diastolic values were compared, and patient values were compared with volunteer values.

Differences in time to peak velocity between the healthy volunteers and the patients were compared in the lateral and septal myocardial walls. According to the American Heart Association standard segmentation model (23), the lateral wall was defined as the average of the inferolateral and anterolateral segments, and the septal wall was defined as the average of the anteroseptal and inferoseptal segments.

Statistical Analyses
Statistical analyses were performed by using Excel 2000 software (Microsoft, Redmond, Wash). Values are reported as means ± standard deviations. The correlation between measured and true values in the phantom was determined by using linear correlation analysis. The MR-measured and true values were compared by using a modified Bland-Altman analysis, in which the difference between the two values was plotted against the known true value (24). In vivo peak velocity and time to peak velocity values were compared between the volunteers and the patients by using a two-sample, unequal variance, two-sided Student t test. The peak velocities at systole and diastole were compared by using a paired two-sided Student t test. Velocity magnitudes were compared between the different directions of motion by using multiple paired two-sided Student t tests with Bonferroni correction. Adjusted P < .05 was considered to indicate a significant difference.


   RESULTS
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
DISCUSSION
 ADVANCES IN KNOWLEDGE
IMPLICATION FOR PATIENT CARE...
 References
 
Phantom Study
There was an excellent correlation between the motion recorded by the feedback sensors and the velocity measured at phase-contrast MR imaging (R = 0.90 for longitudinal velocity, R = 0.93 for circumferential velocity) (Fig 4).


Figure 4
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  		Figure 4: Longitudinal velocity curves from phantom experiments. Excellent correlation (R = 0.90 for longitudinal velocity, R = 0.93 for circumferential velocity) between the known motion of the phantom and the velocity measured with phase-contrast MR imaging was observed. Phase-contrast MR imaging facilitated correct determinations of the times of peak velocity (mean differences: 9.4 msec ± 24.4 in longitudinal direction, 3.1 msec ± 3.6 in circumferential direction) but slight overestimations of the magnitudes of peak velocity (mean differences: 1.0 cm/sec ± 0.9 in longitudinal direction, 3.2 cm/sec ± 1.9 in circumferential direction).

 
Phase-contrast MR imaging facilitated accurate measurement of the velocities in the phantom: Mean differences between the measured velocity and the true velocity throughout the cardiac cycle, as computed at Bland-Altman analysis, were –0.15 cm/sec ± 2.8 in the longitudinal direction and 0.06 cm/sec ± 1.38 in the circumferential direction.

The peak systolic and diastolic velocities measured with phase-contrast MR imaging were a mean of 1.0 cm/sec ± 0.9 greater in magnitude than the true velocities in the longitudinal direction and a mean of 3.2 cm/sec ± 1.9 greater than the true velocities in the circumferential direction (Table 1).


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  		Table 1. Velocity Measurements in the Phantom

 
Phase-contrast MR imaging enabled accurate measurement of the timing of peak velocity: Mean differences between the measured and true values were 9.4 msec ± 24.4 in the longitudinal direction and 1.0 msec ± 20.3 in the circumferential direction. The temporal resolution of the acquired velocity data was 35 msec, so observed differences constituted less than one time frame.

In Vivo Study
Peak velocity.—In both the volunteers and the patients, velocities were greatest in the longitudinal direction and lowest in the circumferential direction (Fig 5). In the volunteers, the mean peak systolic velocity (Table 2) was significantly greater in the longitudinal direction (5.7 cm/sec ± 2.1) than in the radial (3.8 cm/sec ± 1.2) or circumferential (3.2 cm/sec ± 1.2) direction (P < .05 for all after Bonferroni correction). The mean peak radial, longitudinal, and circumferential velocities during diastole for the volunteers were significantly different in magnitude (–6.5 cm/sec ± 2.2, –12.0 cm/sec ± 3.1, and –3.1 cm/sec ± 1.4, respectively; P < .05 for all after Bonferroni correction). The magnitude of peak diastolic velocities was significantly greater than the magnitude of peak systolic velocities in the longitudinal and radial directions (P < .05 for both) but not in the circumferential direction (P = .97).


Figure 5A
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  		Figure 5a: Three-directional lateral wall velocity curves for (a) a healthy volunteer and (b) a patient with dyssynchrony. The scales are different for the two subjects. The magnitude of peak velocities is greater during diastole than during systole for both the patient and the volunteer. Also, the magnitude of longitudinal (ie, through-plane) velocities is greater than the magnitudes of radial and circumferential velocities.

 

Figure 5B
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  		Figure 5b: Three-directional lateral wall velocity curves for (a) a healthy volunteer and (b) a patient with dyssynchrony. The scales are different for the two subjects. The magnitude of peak velocities is greater during diastole than during systole for both the patient and the volunteer. Also, the magnitude of longitudinal (ie, through-plane) velocities is greater than the magnitudes of radial and circumferential velocities.

 

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  		Table 2. Peak Velocities during Systole and Diastole in Healthy Volunteers and Patients with Heart Failure

 
The magnitudes of peak longitudinal and peak radial velocities were significantly greater in the volunteers than in the patients during both systole and diastole. The difference in longitudinal velocity during diastole was the most profound (–12.0 cm/sec ± 3.1 for volunteers, –3.8 cm/sec ± 2.0 for patients; P < .05) and potentially reflected diastolic dysfunction in the patients (Table 2). Peak circumferential velocity was significantly greater in the volunteers than in the patients during systole (P < .05), but no significant difference in the magnitude of peak circumferential velocity between the volunteers and the patients with heart failure was observed during diastole (P = .13).

Time to peak velocity.—In the healthy volunteers, the peak longitudinal velocity during systole occurred a mean of 38 msec ± 114 before the peak radial velocity and a mean of 57 msec ± 103 before the peak circumferential velocity. During diastole, the peak longitudinal velocity occurred first; the peak circumferential velocity (mean of 1.4 msec ± 89 later) and then peak radial velocity (mean of 3.1 msec ± 41 later) followed.

We observed a delay in the time to peak systolic velocity in the lateral wall of the patients with dyssynchrony compared with this parameter in the volunteers (Table 3). The difference was significant in the longitudinal (mean values: 99 msec after R wave ± 17 for volunteers, 198 msec after R wave ± 83 for patients; P < .05) and radial (mean values: 168 msec after R wave ± 51 for volunteers, 224 msec after R wave ± 66 for patients; P < .05) directions only. In the circumferential direction, no significant difference in the time to peak systolic velocity in the lateral wall was observed between the volunteers and patients (P = .21).


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  		Table 3. Time to Peak Systolic Velocity in Septal and Lateral Walls in Healthy Volunteers and Patients with Heart Failure

 
In the patients with dyssynchrony, we also observed a delay in the time to peak diastolic velocity in the lateral wall in the longitudinal (mean values: 507 msec after R wave ± 34 for volunteers, 613 msec after R wave ± 112 for patients; P < .05) and radial (mean values: 478 msec after R wave ± 34 for volunteers, 530 msec after R wave ± 50 for patients; P < .05) directions. No significant delays in time to peak diastolic velocity were observed in the septal wall in the radial and longitudinal directions. Delays in the time to peak diastolic velocity were not significant in the circumferential direction (P = .06).


   DISCUSSION
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
DISCUSSION
ADVANCES IN KNOWLEDGE
IMPLICATION FOR PATIENT CARE...
 References
 
Our study results demonstrate the feasibility of combining three-directional phase-contrast MR myocardial tissue velocity mapping with navigator-echo–gated respiratory compensation in healthy volunteers and in patients scheduled for CRT. This combination technique enables the acquisition of three-directional velocity data at high spatial and temporal resolution without the need for patient breath holding; thus, it can be used in patients with limited breath-hold capability with assurance that all three velocity directions are spatially registered for postprocessing.

The capability of phase-contrast MR imaging for correct measurement of blood flow velocity has been extensively demonstrated in vivo and in vitro (1,25–28). Although tissue velocity imaging involves the use of the same MR pulse sequence that is used for blood flow imaging, the lower velocities of myocardial tissue pose some unique problems. Background phase offset correction is especially important when measuring the lower velocities of myocardial tissue, because the background phase errors are the same order of magnitude as the desired signal. Lower velocities require higher gradients, which require longer echo times and higher bandwidths. Therefore, it is important that the feasibility of the described phase-contrast MR technique for myocardial velocity measurement be verified in a motion phantom before the procedure is applied in vivo.

During systole, the peak velocities measured in the patients scheduled for CRT were significantly lower than those measured in all three directions in the healthy volunteers. During diastole, however, no significant difference in the magnitude of peak circumferential velocities was observed. Decreased peak longitudinal and peak radial velocities have been documented in other patient populations, but circumferential velocity has been shown to be preserved. Young et al (2), in 1994, found that the longitudinal motion was depressed in a group of patients with hypertrophic cardiomyopathy, but in the hypertrophic cardiomyopathy group, the radial motion remained normal and rotation was actually greater; however, the results were not significant. Markl et al (29) found that the regional motion abnormalities in patients with a low ejection fraction after infarction could be detected by measuring the depressed radial—but not the depressed circumferential—velocity. However, because we averaged velocities over the entire imaging section, it is possible that some regional differences between the volunteers and patients were obscured.

The peak velocities in the volunteers reported in our study were greater than those previously reported for some phase-contrast examinations of the myocardium (29–35). In a basal myocardial section, peak radial velocities of 2–4 cm/sec during systole and –1 to –5 cm/sec during diastole have been reported. In the circumferential direction, peak systolic velocities of 0.5–1.5 rad/sec and peak diastolic velocities of 0.05–1.5 rad/sec also have been reported (29,31,36).

It is important that the temporal resolution in our study (35 msec) was significantly higher than those in previous phase-contrast velocity studies (90, 60, or 68 msec [29,31,32,36]). Poor temporal resolution may act as a low-pass filter and lead to underestimated velocities, which may account for some of the differences between the velocities reported in the literature and those measured in our study (35). In a recent study in which phase-contrast MR imaging of the myocardium was used at a higher temporal resolution (37–87 msec) with a breath-hold technique, peak velocities similar to those observed in our study were reported (37).

We observed a significant delay in the time to peak systolic velocity in the lateral wall of the patients with dyssynchrony, but only in the longitudinal and radial directions. Because the majority of patients had left bundle branch block, a delay in lateral wall contraction was expected. In patients with left bundle branch block, the onset of electrical depolarization is substantially delayed in the lateral free wall, so this area contracts much later than it normally would (38). This dyssynchronous contraction results in blood sloshing from early-activated to late-activated regions and then again to early-activated regions, ultimately causing a decreased ejection fraction (39). During CRT, a pacing lead is inserted and forces the delayed-activation region to contract in sync with the normally activated septal wall. Therefore, correctly identifying the time of late activation in this region is more important clinically than measuring the magnitude of tissue velocity.

Echocardiographic measurements of peak velocity timing parameters have previously revealed dyssynchrony (11,40). Furthermore, good agreement between phase-contrast MR imaging and tissue Doppler imaging for measurement of longitudinal myocardial velocity timing parameters and detection of dyssynchrony has been demonstrated previously (14,41). Our study results support previous findings (14,41) that phase-contrast MR imaging can be used to measure longitudinal timing parameters and extend the previous results to include measurements of in-plane velocities. Although radial and circumferential myocardial velocities have emerged as important indicators of dyssynchrony, they are difficult to measure with echocardiography (13,42). Therefore, phase-contrast MR imaging may yield additional information for identifying dyssynchrony and selecting patients for CRT.

Our study had limitations. Since the direction of twist changes along the length of the LV, the lack of a significant difference in peak diastolic circumferential velocity between the volunteers and the patients possibly was due to the imaging section location. We consistently placed the short-axis imaging section at a location that constituted 70% of the LV length, where tissue Doppler imaging velocity measurements are usually made on long-axis images. However, the clockwise-to-counterclockwise twist transition may have varied among the study participants overall or between the volunteers and the patients with dyssynchrony. A large discrepancy in age existed between the volunteers and patients. Although it is desirable to compare patient data with data from a group of age-matched control subjects, in this preliminary study, our purpose was to demonstrate the feasibility of the phase-contrast MR technique in the patient group. No attempt was made to demonstrate the diagnostic value of the measurements.

Although our study revealed excellent temporal resolution for the described navigator-echo–gated phase-contrast MR examination, the temporal resolution was still poor compared with that of Doppler examinations, in which data can be acquired at up to 5-msec intervals (15). This means that some aspects of the velocity curves (such as distinction of the E and A waves during early diastole) often cannot be seen with phase-contrast MR imaging. Since we measured velocity in one basal section, we were unable to measure torsion over the LV. However, now that the feasibility of phase-contrast MR imaging with navigator-echo gating has been demonstrated in patients with heart failure, examination of myocardial velocity at multiple locations is possible.

In conclusion, our study results demonstrate the feasibility of combining phase-contrast MR imaging with navigator-echo respiratory gating for the acquisition of three-directional (ie, longitudinal, radial, and circumferential) velocity data in myocardial tissue without patient breath holding. The capability of phase-contrast MR imaging for accurate measurement of myocardial tissue velocities was verified in a phantom, and the technique was used to study three-directional velocity in the myocardium of both healthy volunteers and patients with heart failure scheduled for CRT. Peak longitudinal and peak radial velocities were lower in magnitude in the patients than in the volunteers. In the patients, the time to peak systolic velocity was significantly delayed in the lateral wall in the radial and longitudinal directions. Navigator-echo–gated phase-contrast MR imaging may be a useful technique that yields a more complete description of the three-directional myocardial contraction pattern.


   ADVANCES IN KNOWLEDGE
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
DISCUSSION
 ADVANCES IN KNOWLEDGE
IMPLICATION FOR PATIENT CARE...
 References
 


   IMPLICATION FOR PATIENT CARE
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
DISCUSSION
 ADVANCES IN KNOWLEDGE
IMPLICATION FOR PATIENT CARE...
References
 


   FOOTNOTES
 

Abbreviations: CRT = cardiac resynchronization therapy • LV = left ventricle

Guarantors of integrity of entire study, J.G.D., J.N.O.; study concepts/study design or data acquisition or data analysis/interpretation, all authors; manuscript drafting or manuscript revision for important intellectual content, all authors; manuscript final version approval, all authors; literature research, J.G.D.; clinical studies, R.L.E., S.E., A.R.L., J.N.O.; experimental studies, J.G.D., K.R.J.; statistical analysis, J.G.D., K.R.J., J.N.O.; and manuscript editing, J.G.D., K.R.J., R.L.E., J.N.O.

Authors stated no financial relationship to disclose.


   References
TOP
 ABSTRACT
 INTRODUCTION
 MATERIALS AND METHODS
 RESULTS
DISCUSSION
 ADVANCES IN KNOWLEDGE
IMPLICATION FOR PATIENT CARE...
 References
 

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