MR image generation depends on the alignment of a charged particle with a net magnetic moment (hydrogen
nucleus or proton) along an externally applied magnetic field. The alignment and spin of this proton can be changed (excited) if it is exposed to a radiofrequency pulse at a certain frequency, called the Larmor frequency, which is specific to the strength of the magnetic field. After the radiofrequency pulse, the protons relax back toward their original alignment in the magnetic field, and the signal produced by this relaxation can be detected by receiver coils. The timing of the excitation pulses and the changing gradients of the magnetic field define the imaging sequence and are responsible for the different signal intensities produced by different tissues within each imaging sequence. Each repetition of the excitation-relaxation sequence produces a line of data that contributes to the image reconstruction. The higher the resolution that is desired, the more repetitions are needed.

Image quality depends not only on the size of the image matrix, but also on the magnitude of the signal produced. A strong external magnetic field will cause a larger proportion of hydrogen nuclei to align with the field, thereby producing greater signal during the imaging sequence. Most MR scanners currently used for cardiac imaging have a field strength of 1.5 Tesla. With low field strength (0.3 to 0.5 Tesla) in so-called “open” scanners, the signal-to-noise ratio is poor, and the imaging time is longer than in scanners that have higher field strengths. Therefore, scanners having low field strengths are unsuitable for cardiac imaging. MR scanners with a field strength of 3 Tesla are commercially available and have the potential to produce high-quality images with short imaging times. However, imaging at 3 Tesla also has disadvantages. Susceptibility artifacts are more prevalent than when imaging at lower field strengths. Also, routine cardiac sequences at 3 Tesla may deposit an amount of radiofrequency energy in the patient’s body that exceeds the specific absorption rate (SAR) limits imposed by the Food and Drug Administration (FDA).

To decrease cardiac motion artifacts, all cardiac MR studies are performed using electrocardiographic (ECG) gating. The R wave is used to the trigger the MR sequence, and an appropriate trigger delay is used to position the data acquisition window at the point in the cardiac cycle having least motion.

In MRI, attaining high spatial and high temporal resolution can be mutually exclusive. Data sufficient to reconstruct low-resolution images can be acquired over one or a few cardiac cycles. However, for MR coronary angiography, the data acquisition window should be less than 100 msec with normal heart rates and less than 60 msec in patients with tachycardia. With such short acquisition windows, data acquisition over many cardiac cycles is necessary to reconstruct a high-resolution image. For the submillimeter resolution required for coronary imaging, data acquisition can take several minutes. This approach assumes that the heart returns to the same shape and size at the same point in each cardiac cycle. Variations of the length of the cardiac cycle will result in image blurring.

Because of the long image acquisition time, respiratory motion also can be a cause of poor image quality in cardiac and coronary MRI. Various breath-hold and non-breathhold techniques have been used to address the problem of respiratory motion. Early approaches to MR coronary angiography relied on image acquisition during a sustained breath-hold. ECG-triggered k-space segmented gradient echo sequences allowed one single 2D slice to be acquired with each breath-hold. A series of contiguous, parallel slices are acquired, covering the coronary arterial tree in different imaging planes. This sequence can be performed on most commercially available 1.5T MRI scanners but requires experience and knowledge of cross-sectional cardiac anatomy on the part of the examiner. The breath-hold required for this technique (at least 25 to 45 seconds) is too long for many patients. Also, during suspended respiration, diaphragmatic drift of up to 1 cm can occur, causing blurring of the image. Numerous serial breath-holds usually are necessary to image the entire coronary tree. However, the diaphragm will not be always in the same position on successive breath-holds, and misregistration between the contiguous image slices will occur. Second-generation MR angiographic techniques using navigator echo to track the position of the diaphragm allow the acquisition of high-resolution images during free breathing.

Improvements in data acquisition speed now make it possible to acquire a 3D volume rather than a single slice during one breath-hold. These ECG-gated, segmented Turbo-FLASH (fast low-angle shot) or steady-state free procession (SSFP) sequences can be performed with or without gadolinium contrast enhancement. The 3D volume slabs are oriented, using scout images, along the axis of the vessel being imaged. These newer imaging sequences have a higher signal-to-noise ratio than gradient echoes and the 3D data acquisition approach minimizes misregistration artifacts.

ECG-monitoring during MRI often is limited to rhythm assessment because the ST-segments can be distorted by the magnetic field. Also, in the presence of a strong magnetic field, pulsating blood flow through the aorta produces a small electrical charge (magnetohydrodynamic effect) that causes artifacts in the ECG. The augmentation of T-waves by this effect can affect triggering of the scanner sequences if the scanner software mistakes them for Rwaves. Observation of and communication with the patient can be limited by the bulk of the MRI equipment and the layout of an imaging suite that is dictated by MRI safety issues. This can be of concern in patients who are suspected of having an acute coronary syndrome or patients undergoing pharmacologic stress testing. (See the section on MRI Stress Imaging.)

Some types of cardiac MRI studies, such as viability and perfusion studies, require the intravenous administration of gadolinium chelates as a contrast agent. Gadolinium is a very safe contrast agent that has been used for a wide variety of MR examinations for many years. Unlike iodinated contrast agents, gadolinium has no significant nephrotoxicity at usual clinical doses. Side effects of and allergic reactions to gadolinium are rare.

Because the strong magnetic field of MRI scanners can interact with ferromagnetic materials, patients scheduled to undergo a cardiac MRI study should be carefully interviewed for the presence of metallic devices, implants, and foreign bodies. Reference resources should be consulted when it is unclear whether a specific device is safe in the MR environment (1). Pacemakers and implantable cardiac defibrillators generally are considered unsafe. However, it may be permissible and safe to scan patients with pacemakers under close monitoring if the clinical information needed cannot be otherwise obtained (2). Immediate follow-up is necessary, because devices or implants with electrical systems may malfunction after a clinical MRI examination. All prosthetic heart valves are considered safe at 1.5 Tesla. Coronary stents can safely be imaged by MRI using field strengths up to 1.5 Tesla beginning immediately after placement (3). Although most metallic surgical clips, wires, stents, and orthopedic implants are MR compatible, they cause local distortion in the magnetic field, and the resulting susceptibility artifact will obscure the representation of adjacent anatomic structures.

Most cardiac CT studies, with the exception of coronary artery calcium (CAC) scanning, require the intravenous administration of iodinated contrast medium. The amount of contrast medium needed to perform, for example, a CT angiogram, depends on the duration of the cardiac scan and the rate of injection (typically 2 to 4 mL/sec); this is approximately 100 to 140 mL for most CT scanners currently in clinical use (4). The short duration of cardiac scans using the newest MDCT scanners will allow CT angiograms to be performed using 60 to 80 mL or less.

The effective radiation dose received by the patient is higher for cardiac MDCT studies than for cardiac EBCT studies, even if dose-reducing scan acquisition techniques are used (for a CT coronary angiogram, approximately 8.1 mSievert versus 1.1 mSievert, respectively) (5).

CT imaging of the coronary arteries allows a quantification of the degree of coronary calcification. Common units for quantification are the Agatston score and the volume score. The prevalence and quantity of CAC is highly correlated with age. Women have a lower prevalence of coronary artery calcification than men until approximately 65 to 70 years of age. At any age, the quantity of CAC is less in women than in men (10).

The quantity of CAC is proportional to the coronary atherosclerotic plaque burden. For example, in a histopathologic analysis of excised coronary arteries, the square root sum of areas of calcification detected by CT imaging was highly correlated with, and represented approximately one-fifth of, the square root sum of the planimetry-derived total cross-sectional plaque area on an artery-by-artery basis and for all major coronary artery branches taken together (11). However, the location of coronary artery calcification does not predict the location of coronary artery stenoses.

Ongoing controversy centers on whether CAC scores measured by MDCT scanners are numerically equivalent to CAC scores measured by EBCT. Most studies of CAC quantification as a predictor of the presence of high-grade coronary artery stenoses or of the risk of future ACEs have been done using EBCT, and more studies are needed to confirm or refute that MDCT-derived CAC scores have the same implications as EBCT-derived CAC scores.

Because of the biologic process of coronary remodeling, in which arteries can increase their cross-sectional area to accommodate growing atherosclerotic plaque, not all CAC found on CT imaging translates into the presence of high-grade coronary stenoses. A large amount of plaque, represented by a large quantity of CAC, can be present before clinically significant luminal narrowing occurs. However,
when coronary plaque occupies approximately 40% of the enlarged coronary artery cross-section, the vascular remodeling capacity becomes exhausted and angiographically discernible—and eventually hemodynamically significant—encroachment upon the coronary artery lumen occurs with increasing plaque burden. Therefore, and because of the epidemiology of coronary artery calcification described in the previous section, appropriate ageand gender-adjusted thresholds must be considered if the quantity of CAC is to be used to predict the presence of high-grade coronary stenoses as the cause of chest pain in symptomatic patients.

In one study of 1,225 men and 539 women, thresholds of the Agatston score measured using EBCT differentiated between symptomatic patients who had a high or low probability of significant coronary artery stenoses detected by invasive angiography, and EBCT was highly sensitive and moderately specific in predicting high-grade coronary artery stenoses. In particular, the absence of any CAC on EBCT scanning was associated with a very low probability of high-grade coronary artery stenoses (Fig. 4.1) (12). In another study of 1,851 symptomatic patients, the power of coronary calcium scores to predict the severity and extent of angiographically significant coronary artery disease (CAD) was incremental to and independent from the clinical predictors “age” and “gender” in symptomatic patients, and most useful in an intermediate risk population (13).