Cardiac MRI, Technical Aspects Primer
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Cardiac magnetic resonance imaging (MRI) has a wide range of clinical applications. Many of these applications are commonly employed in clinical practice—for example, in the evaluation of congenital heart disease, cardiac masses, the pericardium, right ventricular dysplasia, and hibernating myocardium. [1, 2]
Other applications, such as evaluation of myocardial perfusion and of valvular and ventricular function, are very accurately evaluated with MRI, but competing modalities, such as single-photon emission computed tomography (SPECT) scanning and echocardiography, are more commonly employed in clinical practice. Some applications, such as coronary artery imaging, are currently more accurately evaluated with other modalities. For examples of cardiac MRI scans, see the images below.
One of the main advantages of cardiac MRI is the lack of ionizing radiation, which is substantial with SPECT and computed tomography (CT) scanning. The strength of cardiac MRI, as compared with CT scanning, is its superior temporal and contrast resolution. However, the spatial resolution of CT scanning is superior.
While there are competing modalities for every clinical application of cardiac MRI, there is no one modality that can provide as comprehensive an evaluation as MRI. For this reason, cardiac MRI is often known as the “one-stop shop.”
The technical aspects of cardiac MRI are often more daunting for the novice than are the technical aspects of other modalities. The intent of this article is to serve as a primer on the technical aspects of cardiac MRI.
As the number of cardiac MRI applications is broad, the number of potential imaging techniques is correspondingly broad and cannot be covered in depth in 1 article. Therefore, this article will focus on providing an in-depth review of the most common cardiac MRI techniques. Several common clinical applications of cardiac MRI will also be addressed.
The main cardiac imaging planes are oblique to one another. As the cardiac imaging planes are also at arbitrary angles with respect to the scanner, they are called “double oblique” planes. The 3 main cardiac imaging planes are the short axis, as seen in the first image below; the horizontal long axis, as seen in the second image below; and the vertical long axis, as seen in the third image below (the long axis is the line from the center of the mitral valve orifice to the left ventricular apex).
The horizontal long-axis view is also known as the 4-chamber view, and the vertical long-axis view is also known as the 2-chamber view. Note that the initial vertical long-axis view that is prescribed from an axial image is only approximate; a true vertical long-axis view should be prescribed from the horizontal long-axis view. Methods to determine the correct location and orientation of the standard cardiac imaging planes have been well-described (see the images below). [3]
Other imaging planes that may be useful include a left ventricular outflow tract view (see the first image below), for ascending aortic pathology, and a 3-chamber view (see the second image below). [4] The 3-chamber view can be prescribed from the left ventricular outflow tract view of a short-axis view. This view displays the aortic and mitral valves immediately adjacent to one another. [5]
Unlike the pulmonic and tricuspid valves, which are separated by a muscular crista supraventricularis, the aortic and mitral valves are in close proximity and are often both affected by pathologic processes.
Electrocardiographic gating can be performed prospectively or retrospectively. Prospective gating is most common. In prospective gating, the MR acquisition is triggered by the R wave. Within an R-R interval, there may be a trigger delay, acquisition window, and trigger window. [6]
Diastolic imaging may be desirable with fast spin echo sequences, and a trigger delay can be used to delay image acquisition after the R wave trigger. A trigger window is an interval between the end of data acquisition and the next R wave. With a trigger window, earlier than expected heartbeats will still trigger acquisitions. The trigger window is typically 10-15% of the R-R window. The acquisition window is the duration of data acquisition. With a standard trigger window and no trigger delay, this would be 85-90% of the R-R window. Because of the trigger window, prospective-gating sequences will exclude late diastole.
Common problems with electrocardiogram (ECG)-triggered acquisitions include poor or inaccurate R wave detection (eg, triggering off a prominent T wave) and patient arrhythmias. R wave–detection problems can often be resolved by adjusting electrode position or by toggling the lead polarity.
Arrhythmias can result in inaccuracies in evaluation of cardiac function. Acquisition time can also be increased, as some heartbeats may not trigger data acquisition. The effect of arrhythmias can be mitigated with very fast sequences (eg, single-shot fast spin echo) or real-time sequences.
Retrospective gating is also useful in patients with arrhythmias, because data from irregular heartbeats can be rejected.
In retrospective gating, the data are acquired continuously, along with an ECG tracing. The data are retrospectively sorted using the ECG tracing after the acquisition. This is more computationally intensive.
Retrospective gating is helpful in patients with arrhythmias. In retrospective gating, there is no trigger window and the full cardiac cycle is imaged. Imaging of the full cardiac cycle may result in more accurate assessment of cardiac function. Retrospective gating is particularly helpful if peripheral pulse gating is used. Peripheral pulse gating is an option if central gating cannot be performed. Prospectively gated peripheral pulse triggered sequences will start after the onset of systole, as the systolic pulse must propagate to the finger before being detected.
Cardiac MRI imaging sequences may seem daunting to the novice. One way to approach the array of different sequences is to organize them by application. [7] Terms used in the following discussion will be explained in the sections on specific sequences. It is helpful to learn the generic names of the sequences rather than the trade names. [1]
It has been proposed that cardiovascular MRI pulse sequence terminology be simplified to increase the clarity of cardiovascular MRI reports, which might improve acceptance and application of cardiovascular MRI in clincal practice. [8] The following terms have been suggested for clinical reports:
Cardiac function is evaluated using cine gradient echo sequences often known as “bright blood” sequences (see the image below). Steady-state free precession (SSFP) gradient echo sequences have largely replaced spoiled gradient echo sequences for this purpose. Different trade names for these SSFP sequences are TrueFISP (True Fast Imaging with Steady-state Precession; Siemens), FIESTA (Fast Imaging Employing Steady-state Acquisition; GE), and b-FFE (Balanced Fast-Field Echo; Phillips). These sequences are typically used in conjunction with segmented k-space acquisition.
Fast spin echo sequences often known as “black blood” sequences are typically used (see the image below). Multiple options are available, but half-Fourier, single-shot, fast spin echo (SS-FSE) sequences are the fastest. Different trade names for these half-Fourier single shot sequences are HASTE (Half-Fourier Acquired Single-shot Turbo spin Echo; Siemens) and SS-FSE (GE, Phillips).
These sequences are typically used in conjunction with double inversion recovery prepulses. The “bright blood” SSFP sequence can also be used to assess cardiac morphology if it is altered to produce images of the entire heart (rather than a cine loop at a single location).
Magnetization-prepared gradient echo sequences are used to assess myocardial perfusion (see the image below). The magnetization preparation prepulse can be a saturation or inversion recovery pulse and is used to improve T1-weighted contrast. Different trade names for these sequences are TurboFLASH (Fast Imaging using Low Angle Shot; Siemens), Fast SPGR (Spoiled Grass [Gradient Recall Acquisition using Steady States]; GE), and TFE (Turbo Field Echo; Phillips). Echoplanar sequences can also be used.
Contrast-enhanced MR evaluation of myocardial viability utilizes inversion recovery gradient echo sequences, with the inversion time set to null viable myocardium. Either spoiled gradient echo or SSFP sequences can be used in conjunction with the inversion recovery prepulse. These sequences typically utilize segmented k-space acquisition.
Flow quantification utilizes cine phase contrast sequences (see the images below).
Many different sequences have been used to image the coronary arteries. These sequences are typically used in conjunction with segmented k-space acquisition. Two-dimensional (2D), segmented, gradient echo sequences can be used to evaluate coronary artery anomalies. Three-dimensional (3D) techniques are used to evaluate the arteries for stenosis. Images can be acquired during breath-holding or free breathing. Images can be obtained with or without intravenous contrast. A 3D, segmented SSFP sequence without intravenous contrast is well suited to evaluate the coronary arteries. [9] If intravenous contrast is employed, intravascular contrast agents are the most useful.
Standard 3D, spoiled gradient echo sequences with intravenous contrast are used to evaluate the aorta and great vessels.
The following discussion will cover the most commonly used sequences in the clinical practice of cardiac MRI: cine SSFP, fast spin echo with double inversion recovery, and inversion recovery gradient echo. A conceptual approach that addresses why specific techniques and sequences are used is employed.
Black blood MRI scans (see the image below) are produced with sequences designed to null the signal of flowing blood. These images allow for anatomic assessment of the heart and vascular structures without interference from a bright blood signal. While black blood sequences are standard in most imaging protocols, they are particularly important for assessment of cardiac masses, the myocardium (eg, in suspected arrhythmogenic right ventricular dysplasia), and the pericardium.
In clinical practice, there are 3 general options for black blood imaging [10] :
Half-Fourier, single-shot fast spin echo with double inversion recovery
Breath-hold, single-slice fast spin echo with double inversion recovery
Multislice fast spin echo
The first 2 options are the most commonly used.
In ECG-gated spin echo cardiac imaging, TR (repetition time) depends upon heart rate or R-R interval. Thus, the acquisition time can be calculated by substituting the R-R interval for TR in the standard equation:
Acquisition time = R-R interval × number of phase encoding steps × number of acquisitions/echo train length
Note that if the heart rate is 70 beats per minute, the R-R interval is 857 msec, which may not be adequate for T2-weighted imaging. In this case, triggering can be performed after every other R wave, and (2 x R-R interval) should be used in place of R-R interval in the above equation.
In many cases, the purpose of black blood imaging is to assess anatomy, and weighting is not important. In such cases, TR should be as short as possible to minimize imaging time; thus, black blood MRI scans are often T1-weighted. For certain applications, such as cardiac mass evaluation, specific T2-weighted sequences may be performed.
Protons must experience the 90° excitation pulse and the 180° refocusing pulse to generate a spin echo. If protons in flowing blood are not present in the slice long enough to experience both pulses, no spin echo is generated.
Thus, a way to minimize the signal from flowing blood is to decrease the chance that flowing blood will experience the 90° and 180° pulses. This can be done by minimizing the time the blood is in the slice, such as by decreasing the volume of the slice (thinner slices), creating the shortest path (slice positioning orthogonal to flowing blood), or increasing the speed of flowing blood (imaging during systole). Another method is to increase the time interval between the 90° and 180° pulses (increase TE, or echo time).
In standard spin echo imaging, acquisition during systole will result in more nulling of blood signal. However, as will be discussed, in fast spin echo imaging (see below), diastolic imaging is usually more optimal.
Standard spin echo black blood imaging has little utility in clinical practice, because acquisition times exceed patient breath-holding times. Although the resulting respiratory artifacts can be remedied to some extent with signal averaging (which further increases acquisition time), acquisition during free breathing is better performed with multislice fast spin echo imaging. The fastest fast spin echo sequences can be performed during a breath-hold.
A basic disadvantage of fast spin echo imaging relative to spin echo imaging is the image blurring that results from acquiring data at different effective echo times during the echo train. In cardiac imaging, this image blurring is exacerbated by the increased motion at systole. Thus, to minimize artifact, fast spin echo cardiac MRI is best performed in diastole.
However, as previously discussed, blood signal is optimally nulled at systole where blood flows fastest. Diastolic imaging may result in more blood signal than optimal. Fast spin echo cardiac MRI sequences are therefore typically performed with the addition of double inversion recovery (see below) pulses to achieve optimal nulling of blood signal.
Double inversion recovery sequences are designed specifically to null the signal from flowing blood. There are 2 prepulses. A nonselective 180° RF (radiofrequency) pulse inverts all protons. This is followed by a slice-selective 180° pulse that reverts all protons in the imaging slice back to the original alignment. There is no effect on stationary protons in the imaging slice. However, the flowing blood in the imaging slice will have experienced only the nonselective pulse (the blood that experienced both pulses will no longer be in the slice at the time of imaging). Double inversion recovery sequences begin imaging when the magnetization vectors of the flowing blood crosses the null point — the inversion time.
Typical inversion times for double inversion recovery sequences are between 400 and 600 msec and depend on heart rate. Note that the inversion time is a substantial portion of a typical R-R window, which limits the amount of time available to acquire the echo train. Also note that images performed 400-600 msec after the R wave will conveniently be in diastole.
The fastest sequences are half-Fourier, single-shot fast spin echo with double inversion recovery in which the data needed to generate an image can be acquired during one heartbeat. However, while these images have the least cardiac and respiratory motion artifact, the half-Fourier single-shot acquisition decreases spatial resolution and signal-to-noise. For applications where optimal resolution and signal are useful (eg, evaluation of the right ventricular wall in suspected arrhythmogenic right ventricular dysplasia), breath-hold, single-shot fast spin echo with double inversion recovery (1 slice per breath-hold) may be more useful.
Another option is to use multislice fast spin echo imaging during free breathing. This technique is similar to basic spin echo imaging with the addition of a short echo train to decrease the imaging time. As in spin echo sequences, multiple signal averaging is used to decrease respiratory motion artifact. As blurring is minimal with a short echo train, systolic imaging is possible, and blood nulling is similar to spin echo sequences. Inversion recovery pulses may not be necessary with this technique.
It is also possible to use bright blood sequences to evaluate cardiac morphology.
Steady-state gradient echo imaging has largely replaced spoiled gradient echo imaging for bright blood cine cardiac MRI (see the image below).
In gradient echo (GRE) imaging, the TR is often shorter than the T2 of most tissues, and the transverse magnetization will not have fully decayed before the next RF pulse. Thus, there will be residual transverse magnetization that adds T2 contrast (in addition to T1 contrast) to the image. This additional T2 contrast is undesirable for many applications, as the T1 and T2 contrast may be competitive. For example, a liver lesion that is hypointense on T1 and hyperintense on T2 may be isointense with both T1 and T2 weighting. To achieve T1 weighting with a short TR GRE sequence, spoiling the residual transverse magnetization is necessary. This spoiling can be accomplished with an RF pulse or gradients. The majority of fast GRE sequences used in noncardiac clinical MRI are spoiled.
In steady-state GRE sequences, spoiling is not performed, and residual transverse magnetization is retained. The retained residual transverse magnetization increases the signal-to-noise ratio (SNR) of steady-state sequences relative to spoiled sequences. The image contrast will depend on the T2-to-T1 ratio. As stated previously, this is undesirable for many applications. In steady-state sequences, only fluid and fat will have a high signal (fluid and fat have comparable T1 and T2 times, while in most other tissues, T2 time is much shorter than T1 time). However, in bright blood cardiac MRI, hyperintense blood relative to other tissues is exactly what is needed; thus, steady-state GRE sequences are optimal for cine cardiac imaging (cMRI).
The sequences used in cardiac imaging are balanced SSFP sequences. Different trade names for these sequences are TrueFISP (Siemens), FIESTA (GE), and balanced FFE (Phillips). These sequences are very fast and have a high SNR, but the T2-to-T1 image contrast limits the role of these sequences for noncardiac applications.
SSFP cine MRI has largely replaced spoiled GRE cine MRI for evaluation of cardiac function. SSFP sequences do not depend on flow; they have a higher SNR; and they are faster. Spoiled GRE sequences are T1 weighted and depend on through plane flow enhancement (similar to time-of-flight MR angiography) to generate contrast. The blood may become saturated if the flow is slow or the TR is short. Thus, spoiled GRE cine MRI does not allow for the use of very low TRs, because there is not enough time for saturated blood to be replaced by unsaturated blood between excitation pulses.
With SSFP sequences, blood signal is dependent on intrinsic contrast rather than inflow effects, and TR can be as short as possible. SSFP cine MRI can be almost 3 times as fast as spoiled GRE cine MRI. In addition, the SSFP sequence has a higher SNR due to the residual transverse magnetization. This is particularly true at low TRs. With spoiled GRE sequences, SNR decreases with decreasing TR. With SSFP sequences, SNR is high even at low TRs, because residual transverse magnetization increases with shorter TRs.
High-quality SSFP imaging depends on a low TR, a high flip angle, and a uniform magnetic field. [9]
In SSFP imaging, residual transverse magnetization must be preserved. Field inhomogeneity and unbalanced gradients can disrupt the steady-state transverse magnetization. The sequences are implemented with balanced gradients to minimize gradient-induced dephasing. SSFP sequences are very sensitive to field inhomogeneities. In regions of high local magnetic-field variations, SSFP images often suffer from characteristic bands of signal loss (off-resonance banding artifact), which can disrupt the steady-state.
As TR is increased, any off-resonance banding artifact will become more pronounced because of the increased off-resonance precession per TR. Thus, the lowest TR possible is desirable for SSFP imaging. Typical TRs are less than 4 msec, with TEs being less than 2 msec. Banding artifact is a particular limitation for 3T MRI, as the banding artifact becomes more pronounced as the main magnetic-field strength (and any associated inhomogeneity) is increased.
In spoiled GRE sequences, optimal SNR is dependent on matching the flip angle to the TR (the lower the TR, the lower the flip angle). In SSFP sequences, the SNR does not change substantially with different flip angles, but the T2/T1 weighting will increase with an increasing flip angle. SSFP sequences, therefore, should use the largest flip angle achievable, because this will maximize the contrast-to-noise ratio. As RF pulses are continuously applied to maintain the steady state, specific absorption rate limits are often a factor in SSFP sequences and limit the use of very high flip angles. Flip angles in SSFP sequences are typically 40 to 70°.
SSFP sequences are prone to off-resonance banding artifacts. As these artifacts are caused by local field inhomogeneities, a very uniform magnetic field is required to avoid artifacts. [11]
Because SSFP sequences are typically performed with very low TRs and TEs, the low TE time may result in a chemical shift artifact of the second kind (India ink artifact).
SSFP sequences (see the first image below) may be less sensitive to turbulent flow (eg, in regurgitant valves) than spoiled GRE sequences (see the second image below), because SSFP sequences do not depend on time-of-flight effects.
Multiple images at the same slice position, corresponding to different time points in the cardiac cycle, are obtained during cine GRE imaging. Each image is called a frame. Typically, 12-18 frames are obtained during a cardiac cycle. The temporal resolution is the duration of the cardiac cycle that each frame represents. High temporal resolution is necessary to accurately assess cardiac motion, particularly during systole. [11]
The ideal temporal resolution should be 50-60 msec or less. With faster heart rates, greater temporal resolution is needed. [11]
The temporal resolution and the number of frames are directly related, but in general, the temporal resolution is more important than the number of frames obtained. For example, with a very fast heartbeat, functional evaluation may still be adequate with good temporal resolution, although the number of frames may be lower than that typically obtained with slower heartbeats. [11]
In cine MRI, the echoes are partitioned into k-spaces, with each k-space corresponding to a frame. If there are 12 frames, the echoes would be partitioned into 12 k-spaces. The amount of data (number of phase-encoding steps) needed to fill each of the k-spaces corresponds to the spatial resolution. In conventional cine MRI, each of the 12 k-spaces is filled with only 1 phase-encoding step of the necessary data during a single heartbeat. The total acquisition time is therefore the number of heartbeats necessary to fill a k-space.
A standard study with 128 phase-encoding steps will take 128 heartbeats to complete, which does not allow for breath-hold imaging. With segmented k-space cine MRI, multiple phase-encoding steps of data (per frame) are acquired after a single heartbeat.
The number of lines of k-space per frame acquired per heartbeat is referred to as the views per segment or lines per segment. For a study with 128 phase-encoding steps, 8 views per segment would reduce the imaging time from 128 heartbeats to 16 heartbeats. This allows for breath-hold cardiac cine imaging.
It is important to understand the relationship between temporal resolution, spatial resolution, and imaging time in cine cardiac MRI. [11]
The temporal resolution is directly related to the views per segment:
Temporal resolution = TR × views per segment
In this case, TR is used in the standard sense to refer to the time between consecutive RF pulses. Lee refers to this as “true TR,” because TR is also used to refer to the temporal resolution in cine MRI. [11]
There is a direct trade-off between imaging time (views per segment) and temporal resolution. Decreasing the imaging time by increasing the views per segment will decrease the temporal resolution. For example, if the number of views per segment is doubled, the overall imaging time will decrease by half, because twice as much data are acquired during every heartbeat. However, acquiring twice as much data per heartbeat takes twice as long per frame, which will halve the number of attainable frames per cardiac cycle and worsen the temporal resolution by a factor of 2.
Another way to decrease the imaging time is to decrease the resolution by decreasing the number of phase-encoding steps. An in-plane spatial resolution of 2-2.5 mm is adequate for most cardiac function studies, although higher spatial resolution can be helpful for evaluating structures such as cardiac valves. For patients who are poor breath-holders, temporal resolution, spatial resolution, or both must be compromised if the scan time needs to be decreased.
Heart rate can be helpful in determining the number of views per segment. With slow heart rates, more views per segment can be used. Because the R-R interval is longer, more views per segment can be added while maintaining an adequate number of frames, but temporal resolution will still be decreased. This will decrease the number of heartbeats necessary to complete the study, which is especially helpful if the heart rate is slow.
One way to increase temporal resolution with minimal effect on acquisition time is view sharing or echo sharing. In echo sharing, echoes are recycled over multiple images and can improve the perceived temporal resolution.
A discussion of parallel imaging techniques is beyond the scope of this article, but parallel imaging is especially useful in conjunction with SSFP cine cardiac imaging. Parallel imaging techniques can reduce the imaging time substantially by reducing the number of phase-encoding steps necessary to reconstruct an image severalfold. The drawback is decreased SNR, but SSFP cine MRI has inherently high SNR, which is therefore more tolerant of the decreased SNR resulting from parallel imaging.
Many methods of ventricular volume calculation are used. Rehr et al, [12] as well as Pearlman et al, found excellent correlation between findings at volumetric analysis with MRI and findings with ventricular casts (0.99 correlation, 4.9 mL standard error). Accuracy increases with the inclusion of long-axis measurements. Three-dimensional volumetric calculations are well correlated with ventriculographic findings and have low interstudy variability (< 5%), as compared with ventriculographic and echocardiographic results.
The first step in the calculation of ventricular volume is the selection of representative ED and ES cardiac-phase images. According to Semelka et al, either the phase images that depict the largest and smallest ventricular volumes or the phase images obtained immediately before mitral valve closure (ie, ED) and opening (ie, ES) are chosen. Next, the right ventricle (RV) and left ventricle (LV) are traced along the endocardial margin on each section obtained in the selected ED and ES phases from the cardiac apex to the section just prior to one that depicts the mitral and tricuspid valves. [13]
There are some choices that the operator can make when tracing the ventricular margins. [11] One choice is whether to include the trabeculations and papillary muscles. It is most important that the technique be consistent between studies. In patients with hypertension, hypertrophic cardiomyopathy, or storage disease, the papillary muscles may also hypertrophy; therefore, inclusion of the papillary muscles may yield the most accurate quantification of myocardial mass. For evaluation of function, the trabeculations and papillary muscles can be included or excluded, as long as the technique is consistent.
Because the long-axis dimension of the left ventricle is shorter during systole, the basal short-axis slice may include the left atrium during systole and the left ventricle during diastole. If short axis images are used, the operator may need to decide which slices to include at the left ventricular base. Usually, only slices on which a complete circumferential left ventricular rim is visualized should be included.
ED and ES volumes for each section are totaled to yield the RV and LV end-diastolic volume (EDV) and end-systolic volume (ESV). If short-axis images are used, the volumes can be calculated using Simpson’s rule: the sum of the cross-sectional areas of each slice × distance between slices.
The stroke volume (SV) equals the EDV minus the ESV, or SV = EDV – ESV.
The ejection fraction (EF) equals the SV divided by the EDV times 100, or EF = (SV / EDV) x 100, to give a value reported as a percent.
Cardiac output equals SV multiplied by the heart rate.
For myocardial mass assessment, the RV and LV epicardial borders are traced in ED. The interventricular septum is assigned to the LV and excluded from the RV tracing of the myocardial mass (see the images below). The volumes of all sections are added, and the corresponding EDV is subtracted to determine the myocardial volume. This result is then multiplied by the specific gravity of the myocardium (ie, 1.05 g/mL) to calculate the mass. This measurement is useful in the assessment of hypertrophy and to follow up the ventricular response to therapy.
The AV and ventriculoarterial valves also can be assessed with cine GE sequences. [14] Valvular stenosis or regurgitation produces turbulent jets of signal void in the appropriate directions. Regarding the AV valves, regurgitation is graded according to echocardiographic criteria and is related to the distance to which the jet extends into the atrium. Grades of valvular stenosis are calculated more reproducibly. The valve orifice areas can be measured and graded according to current standards (see the image below).
It is important to note that the MRI sequence used may affect the calculated ventricular volume and mass, [15] which is dependent on accurate delineation of the endocardial and epicardial borders. Delineation of the endocardial border is dependent on contrast between the myocardium and the ventricular blood pool. SSFP images (see the image below) have greater contrast between the myocardium and the blood pool than do spoiled GRE images, primarily because the blood pool signal is not dependent upon flow.
On spoiled GRE images (see the image below), poor delineation of the endocardial border may result in an artifactual increase in apparent myocardial thickness. Using SSFP sequences, ventricular volumes are higher and myocardial mass is lower than those values derived from spoiled GRE sequences.
Normal LV volumes and masses (see Table 1) differ between sexes and among certain ethnic groups. [16] Men have significantly higher LV volumes and masses than do women. Asian Americans, regardless of sex, have lower LV volumes and mass than do other ethnic groups. African-American men have the largest LV volumes and mass. LV volumes do not differ between African-American women and white or Hispanic-American women. These findings are significant after normalization for body-surface area. LV mass is independent of age if indexed for body surface area.
Table 1. Left Ventricular Parameters [16] (Open Table in a new window)
Parameter
Men
Women
LV EDV (mL)
142 +/- 34
109 +/- 22
LV ESV (mL)
47 +/- 19
31 +/- 9
LV EF (%)
67 +/- 7
72 +/- 6
LV SV (mL)
95 +/- 21
78 +/- 17
Cardiac output (mL/min)
5.6 +/- 1.2
4.9 +/- 1.1
LV mass (g)
164 +/- 36
114 +/- 24
EDV: end-diastolic volume; ESV: end-systolic volume; EF: ejection fraction; SV: stroke volume
Typical areas for the aortic and mitral valves are 2.5-3.5 cm2 and 4-6 cm2, respectively; areas of less than 0.8 cm2 and less than 1 cm2, respectively, indicate severe stenosis. The values for aortic and mitral valve area apply to males and females.
Delayed enhancement of infarcts distinguishes infarct from viable myocardium. The difference in intensity between infarcted myocardium and viable myocardium may be difficult to detect on T1-weighted images. GRE sequences can be performed with an inversion recovery prepulse to optimize visualization of enhancing infarcted myocardium. The inversion time is selected to null signal from viable myocardium.
Typical correct inversion times are between 200 and 300 msec. [10] The correct inversion time can be determined empirically with an inversion-time mapping sequence (eg, a segmented cine GRE [typically, SSFP] with images generated at multiple inversion times). This process has been called “inversion-time surfing.” The SSFP sequence is optimal for this technique, because longitudinal magnetization is minimally disturbed during readout.
A complicating factor is that the nulling inversion time will slightly increase during the course of the exam. [17] As gadolinium washes out of viable myocardium, the inversion time increases and inversion time is longer.
If inversion time is optimal, normal myocardium should be very dark (see the image below), infarct should be the most intense structure on the image, and left ventricular blood pool should be of intermediate intensity. [18]
If the inversion time is too short, the blood pool will appear dark. If the inversion time is close to the optimal value but still short, the myocardium may have a speckled appearance, and the endocardial and epicardial borders may appear as hypointense lines around intermediate-intensity myocardium. If the inversion time is too long, the area of delayed enhancement will be only slightly more intense than the myocardium (see the image below).
One way to minimize the importance of choosing inversion time correctly is to use phase-sensitive reconstruction. [10] Unlike typical magnitude images, which do not distinguish between signal from protons above and those below the xy plane, signal intensity in phase-sensitive images varies with longitudinal magnetization across the full spectrum. Thus, it is much less important to accurately choose TI, because infarcted myocardium should always be higher in signal intensity than viable myocardium at all inversion times.
The disadvantage of phase-sensitive reconstruction is that background noise that has a random phase is pixilated. If inversion time is correctly chosen, magnitude images may be preferable.
The contrast-enhanced inversion recovery GRE sequence is performed with k-space segmentation. Either a spoiled GRE or SSFP sequence can be used. Typically, a breath-hold technique is used with 1 section per breath-hold and 10-12 sections total. If the patient can sustain a longer breath-hold, a single 3D acquisition can cover the entire heart. The sequence can also be performed with free-breathing and respiratory gating.
The following is a brief overview of several common clinical applications of cardiac MRI. With an understanding of the technical aspects specific to cardiac MRI, the imager should be able to optimize image quality to aid in the diagnosis and management of these and other cardiovascular diseases.
Ventricular function can be quantitatively assessed by reviewing the dynamic images. Chronic transmural infarcts exhibit a lack of wall thickening in systole and reduced myocardial thickness (< 6 mm). [18] Spin echo images may show an area of decreased signal intensity that corresponds to postinfarct scar formation.
Myocardial tagging is used to track segmental motion and can help in directly distinguishing impaired myocardium from myocardium that may move abnormally because of its proximity to a disease area (ie, tethering effect). Presaturation pulses are used to create a cross-hatched pattern over the myocardium. As systole proceeds, the pattern distorts in a direction corresponding with myocardial movement (see the images below). Thus, myocardial contraction can be assessed reproducibly and quantitatively.
Areas lacking pattern distortion indicate nonfunctioning myocardium. Experience with tagging has elucidated the characteristics of cardiac dynamics. [19] Ventricular contraction contains a twisting component in addition to short- and long-axis movement. A wringing effect occurs, whereby the base moves clockwise and the apex moves counterclockwise. This torsion reverses during isovolumic relaxation and before the mitral valve opens. This mechanism may generate ventricular suction, promoting early diastolic filling. In addition, long-axis shortening is pronounced in the lateral and posterior walls, as opposed to that of the anterior and septal regions.
The ultimate focus of MRI in this application is the development of techniques for identifying acute infarction and for differentiating viable myocardium from nonviable myocardium. The literature is filled with reports of MRI studies of acute myocardial infarction.
Many articles discuss the presence of increased signal intensity on T2-weighted images of regions with acute infarction; however, Filipchuk et al revealed that although sensitivity was adequate (88%), specificity was only 17% compared with controls. Myocardial thinning was the most specific finding in MI (88%), and sensitivity was only 67%. [20] Subendocardial signal intensity changes also can be difficult to distinguish from flow-related enhancement.
A more reliable indicator of acute myocardial infarction is delayed contrast enhancement after the IV administration of gadolinium-based contrast material. Studies in patients after they had suffered a myocardial infarction showed a good correlation between increased enhancement and infarcted tissue, compared with healthy myocardium.
After the contrast has washed out of other regions (5-10 min), it is retained in the altered cells of a myocardial infarction. In normal myocardium, gadolinium is excluded from the myocyte intracellular space. Sarcolemmal membrane integrity is lost in cell death, allowing gadolinium to extravasate into the myocyte and resulting in hyperenhancement.
In the chronic setting, scar tissue has increased extracellular collagen and a larger interstitial space than normal myocardium. A larger interstitial space would account for delayed hyperenhancement seen in scar. Thus, T1-sensitive inversion recovery imaging set to null (blacken) normal myocardium produces a “scar map.” The size and transmural depth are slightly larger shortly after a myocardial infarction, but from 1 week onward, they are stable indicators of scar extent.
Delayed hyperenhancement is associated with myocyte necrosis in the setting of acute and chronic infarction. “At risk” myocardium and severe reversible ischemic injury (even in the setting of stunning) do not exhibit hyperenhancement. A greater transmural extent of infarction (eg, hyperenhancement involving >50% of the wall thickness) can predict regions that are less likely to improve in function after revascularization or beta-blocker therapy (see the image below). The extent of dysfunctional, but nonhyperenhanced, myocardium can predict improvement in left ventricular ejection fraction after therapy. [21]
In clinical practice, imaging is typically performed 10-15 minutes after contrast injection to detect delayed hyperenhancement. An imaging window of 10-30 minutes is probably acceptable. In acute infarcts, hyperenhancement could potentially overestimate infarct size if a long delay is not used. [9]
Delayed hyperenhancement can be seen in diseases other than acute and chronic infarcts. [21] These include sarcoidosis, dilated and hypertrophic cardiomyopathy, myocarditis, amyloidosis, and arrhythmogenic right ventricular dysplasia. In many cases, the pattern of enhancement in these other disorders is different from the subendocardial or transmural hyperenhancement seen in infarct. For example, enhancement may have a midwall, epicardial, or global endocardial distribution.
Enhancement patterns have been explored with combined first-pass and delayed imaging to assess viability and predict functional recovery.
In a study, Rogers et al concluded that reperfused myocardium with the HYPER pattern (defined below) is viable or stunned and regains significant function, whereas those with a HYPO pattern (defined below) probably have irreversible myocardial damage. In the study, 17 patients with reperfused myocardial infarction underwent imaging with tagging and IV contrast enhancement at 1 and 7 weeks. Imaging was performed during the first pass of a bolus gadolinium-based contrast agent and after a 7-minute delay. [22]
The 3 patterns of enhancement that emerged in the study were as follows: (1) HYPO, that is, hypoenhancement on first-pass images and normal enhancement on delayed images; (2) HYPER, or normal enhancement on first-pass images and hyperenhancement on delayed images; and (3) COMB, or hypoenhancement on first-pass images and hyperenhancement on delayed images. Gadolinium-based contrast agents (gadopentetate dimeglumine [Magnevist], gadobenate dimeglumine [MultiHance], gadodiamide [Omniscan], gadoversetamide [OptiMARK], gadoteridol [ProHance]) have been linked to the development of nephrogenic systemic fibrosis (NSF) or nephrogenic fibrosing dermopathy (NFD).
NSF/NFD has occurred in patients with moderate to end-stage renal disease after being given a gadolinium-based contrast agent to enhance MRI or MRA scans. NSF/NFD is a debilitating and sometimes fatal disease. Characteristics include red or dark patches on the skin; burning, itching, swelling, hardening, and tightening of the skin; yellow spots on the whites of the eyes; joint stiffness with trouble moving or straightening the arms, hands, legs, or feet; pain deep in the hip bones or ribs; and muscle weakness.
Echocardiography is usually the initial step in the evaluation of cardiac masses. However, MRI is better in assessing the relationship of the mass to the cardiac structures, and MRI provides more reliable indications of the histologic diagnosis. A particular advantage of MRI is that it can be used to distinguish thrombus from tumor. On spin echo images, thrombus and tumors can have intermediate signal intensity. Unlike a tumor, a thrombus has low signal intensity on GE images because of the presence of deoxyhemoglobin. [23, 24, 25, 26]
A diagnostic dilemma may arise in the differentiation of thrombus (see the images below) from myxoma, which can appear hypointense because of its calcification and high hemoglobin content secondary to hemorrhage.
Myxoma usually originates from the interatrial septum, it can be pedunculated or sessile, and its contour is mostly smooth. Thrombus occupies the atrial appendage, is broad based, and has an irregular contour. Myxoma can prolapse through the mitral valve at cine GE imaging, whereas thrombus usually is associated with mitral valve disease. However, both can originate from the posterior atrial wall in a minority of patients. Another pitfall in thrombus detection is that a hyperacute clot has increased signal intensity on GE images, which can be confused with the appearance of a tumor.
Any clinical question regarding the differentiation of tumor and thrombus can be answered by the administration of contrast material, which causes enhancement of the tumor but not the thrombus. The presence of a clot in the right atrium is suggestive of a tumor thrombus and should prompt further exploration. The most likely source is renal, hepatic, or adrenal.
Other primary cardiac tumors exist, and their signal intensity characteristics are summarized in Table 2, below.
Table 2. Cardiac Tumor Characteristics (Open Table in a new window)
Mass
T1-weighted MRI
T2-weighted MRI
Enhancement
Distribution and Features
Thrombus
Isointensity
Hypointensity
None
Atrial appendage, signal void on GE images
Myxoma
Hypointensity or isointensity
Hyperintensity(often heterogeneous)
Mild or moderate
Atrial septum, in women aged 40-60 y, hypointense on GE images
Fibroma
Isointensity or hyperintensity
Isointensity or hypointensity
Rim
Most patients < 10 y, anterior wall of the left ventricle and/or septum, cystic or calcified, associated with Gorlin syndrome
Rhabdomyoma
Isointensity
Isointensity
Mild or none
Most patients < 1 y, multiple, tuberous sclerosis present in 50%
Hemangioma
Isointensity
Isointensity
High
Mostly intramural but can be exophytic and polypoid
Pheochromocytoma
Hypointensity
Extreme hyperintensity
High
Usually juxtacardiac, chromocytoma or pericardial mass
Lymphoma
Hypointensity
Hyperintensity
Heterogeneous
No data
Malignant fibrous histiocytoma (MFH)
Heterogeneous intensity
Hyperintensity
Moderate
Posterior part of the left atrium, multiple in two thirds of patients
Angiosarcoma
Heterogeneous intensity
Heterogeneous intensity
Heterogeneous or high
Frondular, in men aged 20-50 y, extension into the great vessels
Adapted from Martin DR, Merchant N, MacDonald C. MR imaging of cardiac masses: a review of current application and approach. Appl Radiol. 2000;Mar:10-20.
Another advantage of MRI is its ability to depict the characteristics of a pseudomass. Many normal structures can simulate a tumor or thrombus on echocardiographs, and MRIs can help in the diagnosis in many cases. A prominent crista terminalis in the right atrium (see the first image below) or moderator band in the right ventricle (see the second image below) can be misinterpreted as a lesion. These can be shown to be normal anatomic structures.
Large trabeculations and papillary muscles can be problematic, and asymmetrical ventricular hypertrophy can simulate tumor, as well. Cine GE imaging findings can confirm the normal contraction and function of these structures, which are not observed with a tumor.
Furthermore, obese, female, and older patients have a predilection for lipomatous hypertrophy of the myocardium. These can occur in any location, but a classic location is the interatrial septum (see the image below), typically sparing the fossa ovalis.
Pericardial disease is evaluated initially with echocardiography. The most common reason for imaging is to assess effusion and tamponade. Evaluation of the pericardium with echocardiography is limited if no effusion exists or if the effusion is complex. Echocardiography can cause the misdiagnosis of pericardial cysts, tumors, and diaphragmatic hernias as effusion. MRI does not have these limitations. [24]
At MRI, simple axial and coronal or sagittal imaging planes are used. Effusions complicated with adhesions or loculations are clearly shown. Transudative effusion with low signal intensity on spin echo images can be differentiated from an exudative or hemorrhagic effusion that has high signal intensity. In addition, cine images can depict diastolic collapse of the chambers, which indicates tamponade. Pericardial thickening can be evaluated reliably with MRI, unlike echocardiography. A pericardial thickness that is greater than the typical 2 mm may suggest an inflammatory process in association with effusion.
MRI has high sensitivity in the diagnosis of constrictive pericarditis (see the image below). The presence of pericardial thickening differentiates constrictive pericarditis from restrictive cardiomyopathy. Proper diagnosis is crucial; although their presentations can be identical, their treatments differ markedly. Constrictive pericarditis requires pericardiectomy; restrictive cardiomyopathy requires medical management. As a finding in constrictive pericarditis, diffuse pericardial thickening of 4 mm or greater has an accuracy of 93%.
Other findings, such as atrial dilation and ventricular or septal alterations, can be present, but these are less specific. The inability to image pericardial calcification is a limitation in the diagnosis because the specificity of pericardial calcification is high.
In the appropriate clinical setting, pericardial calcification on computed tomography (CT) scans or plain radiographs is diagnostic of constrictive pericarditis, and MRI is unnecessary. MRI is used to examine the minority of patients with symptoms of constrictive pericarditis in whom calcification is absent. MRI can be used to examine asymptomatic patients and those with atypical symptoms in whom pericardial calcifications are found incidentally to ascertain if constrictive pericarditis exists.
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Parameter
Men
Women
LV EDV (mL)
142 +/- 34
109 +/- 22
LV ESV (mL)
47 +/- 19
31 +/- 9
LV EF (%)
67 +/- 7
72 +/- 6
LV SV (mL)
95 +/- 21
78 +/- 17
Cardiac output (mL/min)
5.6 +/- 1.2
4.9 +/- 1.1
LV mass (g)
164 +/- 36
114 +/- 24
EDV: end-diastolic volume; ESV: end-systolic volume; EF: ejection fraction; SV: stroke volume
Mass
T1-weighted MRI
T2-weighted MRI
Enhancement
Distribution and Features
Thrombus
Isointensity
Hypointensity
None
Atrial appendage, signal void on GE images
Myxoma
Hypointensity or isointensity
Hyperintensity(often heterogeneous)
Mild or moderate
Atrial septum, in women aged 40-60 y, hypointense on GE images
Fibroma
Isointensity or hyperintensity
Isointensity or hypointensity
Rim
Most patients < 10 y, anterior wall of the left ventricle and/or septum, cystic or calcified, associated with Gorlin syndrome
Rhabdomyoma
Isointensity
Isointensity
Mild or none
Most patients < 1 y, multiple, tuberous sclerosis present in 50%
Hemangioma
Isointensity
Isointensity
High
Mostly intramural but can be exophytic and polypoid
Pheochromocytoma
Hypointensity
Extreme hyperintensity
High
Usually juxtacardiac, chromocytoma or pericardial mass
Lymphoma
Hypointensity
Hyperintensity
Heterogeneous
No data
Malignant fibrous histiocytoma (MFH)
Heterogeneous intensity
Hyperintensity
Moderate
Posterior part of the left atrium, multiple in two thirds of patients
Angiosarcoma
Heterogeneous intensity
Heterogeneous intensity
Heterogeneous or high
Frondular, in men aged 20-50 y, extension into the great vessels
Adapted from Martin DR, Merchant N, MacDonald C. MR imaging of cardiac masses: a review of current application and approach. Appl Radiol. 2000;Mar:10-20.
Eugene C Lin, MD Attending Radiologist, Teaching Coordinator for Cardiac Imaging, Radiology Residency Program, Virginia Mason Medical Center; Clinical Assistant Professor of Radiology, University of Washington School of Medicine
Eugene C Lin, MD is a member of the following medical societies: American College of Nuclear Medicine, American College of Radiology, Radiological Society of North America, Society of Nuclear Medicine and Molecular Imaging
Disclosure: Nothing to disclose.
Bernard D Coombs, MB, ChB, PhD Consulting Staff, Department of Specialist Rehabilitation Services, Hutt Valley District Health Board, New Zealand
Disclosure: Nothing to disclose.
Robert M Steiner, MD Professor of Radiology and Medicine, Temple University School of Medicine; Radiologist, Jeanes Hospital, Temple University Hospital
Robert M Steiner, MD is a member of the following medical societies: American College of Cardiology, American College of Chest Physicians, American College of Radiology, American Heart Association, Radiological Society of North America, Society of Thoracic Radiology, North American Society for Cardiac Imaging
Disclosure: Nothing to disclose.
Eugene C Lin, MD Attending Radiologist, Teaching Coordinator for Cardiac Imaging, Radiology Residency Program, Virginia Mason Medical Center; Clinical Assistant Professor of Radiology, University of Washington School of Medicine
Eugene C Lin, MD is a member of the following medical societies: American College of Nuclear Medicine, American College of Radiology, Radiological Society of North America, Society of Nuclear Medicine and Molecular Imaging
Disclosure: Nothing to disclose.
Evan J Samett, MD Interventional Radiology
Evan J Samett, MD is a member of the following medical societies: American College of Radiology, Radiological Society of North America
Disclosure: Nothing to disclose.
Justin D Pearlman, MD, ME, PhD, FACC, MA Chief, Division of Cardiology, Director of Cardiology Consultative Service, Director of Cardiology Clinic Service, Director of Cardiology Non-Invasive Laboratory, Chair of Institutional Review Board, University of California, Los Angeles, David Geffen School of Medicine
Justin D Pearlman, MD, ME, PhD, FACC, MA is a member of the following medical societies: American College of Cardiology, International Society for Magnetic Resonance in Medicine, American College of Physicians, American Federation for Medical Research, Radiological Society of North America
Disclosure: Nothing to disclose.
Steven R Klepac, MD Staff Radiologist, Teleradiology Solutions
Disclosure: Nothing to disclose.
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