MRI measures of signal decay without refocusing, with intermittent refocusing and with continuous refocusing reflected by time constants T2*, T2, and T1ρ can yield important clinical information in myocardial pathophysiology. T2* is used to characterize iron overload, although specificity is reduced by susceptibility effects. By incorporating refocusing of static inhomogeneity effects, T2 yields more specific characterization of signal changes associated with changes in blood oxygenation reflecting ischemia and changes in water mobility reflecting inflammation. Decreasing refocusing interval reduces dephasing due to diffusion through gradients and chemical exchange effects and has been used to increase T1ρ contrast between healthy and infarcted myocardium.
T2* is the decay time constant of magnetization following excitation by a radio-frequency (RF) pulse (i.e. tipped into the transverse plane) in the absence of refocusing pulses, and is essentially T2 relaxation (see below) plus an additional term describing dephasing associated with variations in proton resonance frequencies within a given volume (1,2). This reversible dephasing is superimposed with the irreversible dephasing generated by random magnetic field fluctuations which drive T2 decay in the sample to create a net decay T2*, which is shorter than T2. Any inhomogeneity in the local static field arising from imperfections in the magnetic field, susceptibility differences at air-tissue interfaces or from the presence of paramagnetic or ferromagnetic substances, will result in such additional reversible dephasing. Thus, T2* is particularly useful for assessing body iron levels, deoxyhemoglobin content (blood oxygen), and distribution of iron oxide contrast agents (3,4). T2* measurements are typically performed using a spoiled gradient echo sequence with multiple echoes (i.e. a multi-echo gradient-echo sequence) (5). For cardiac studies, following a 30 degree RF excitation, the signal is sampled at times of gradient echo formation generated by applying a series of alternating readout gradients. Typically, up to eight echoes are acquired after the excitation, with each echo corresponding to a different T2*-weighted image as determined by TE (TEs ranging from 1.4 to 15 ms). At each echo, the same line of k-space is acquired. In a segmented k-space acquisition, this is repeated around 8-12 times in a row (i.e. every 17 ms) to acquire 8-12 lines of k space within a time window gated to diastole in the heart cycle. This process is repeated at each heartbeat. Imaging is complete when the entire k space is filled giving eight T2*-weighted images from which T2* maps can be calculated. This is similar to an echo-planar-imaging (EPI) type readout; however, in an EPI sequence, the echoes acquired are assigned to different lines the k-space matrix of a single image where the effective TE is the time at which the center of k-space is read.
T2 is the effective transverse relaxation time constant describing MR signal decay that cannot be reversed by refocusing pulses (1,2). This effect arises from random temporal fluctuations in the proton spins’ resonant frequencies. Variation in precession frequencies across protons results in different phase accumulation and thereby loss of phase coherence and transverse magnetization decay. In biological tissue, protons bound to slowly tumbling macromolecules experience rapid T2 decay due to the presence of relatively slow fluctuating fields, which cause a large net phase dispersion. On the other hand, highly mobile protons (e.g. in water, edema) decay much more slowly since rapid fluctuations tend to average out yielding a smaller net phase dispersion over time. The gold standard for T2 mapping is a Carr-Purcell-Meiboom-Gill (CPMG) spin-echo sequence, that involves taking measurements at different echo times (TE) in an echo train to sample the T2 decay curve. Here, a 90 degree excitation pulse is followed by a series of 180 degree refocusing pulses, and signal is measured at the mid-points between refocusing pulses where the spin-echo is formed. In a conventional T2 spin echo sequence, signal contrast is primarily determined by placing the data acquisition block at the appropriate time (TE) following magnetization tip down by the RF pulse. However, for cardiac applications, magnetization prepared T2 contrast is a more desirable technique for efficient T2 mapping, particularly in the presence of blood flow and motion (6,7). Here, the longitudinal magnetization is prepared in advance with a 90 degree excitation pulse followed by a train of refocusing pulses over a duration TE after which magnetization is returned to the longitudinal axis (Mz) with a “tip-up” pulse (typically a composite -90 degree pulse). The resulting desired T2-weighted contrast stored in Mz is then sampled using any excitation and readout strategy. This approach is very efficient and especially useful for obtaining T2 measurements (using multiple T2 preparations with varying TEs) in organs, such as the heart, where flow and motion produce undesirable artifacts. A T2-prepared spiral readout sequence has been recently adopted to characterize myocardial edema by T2 quantification following acute myocardial infarction in both clinical and preclinical models (8,9). A T2-prepared SSFP sequence has also been developed for edema detection in myocardial infarction (10).
T1ρ is the time constant for MRI signal decay in the presence of continuous radiofrequency excitation called a spin-lock pulse in the transverse plane with amplitude B1 corresponding to spin-lock frequency γB1. As demonstrated by Santyr et al (11), this is equivalent to signal decay in the presence of a refocusing pulse train (e.g. CPMG) with a short refocusing interval (τ180), where the spin-lock frequency is 1/((2τ180). On clinical MRI systems, development of T1ρ contrast is usually achieved using a magnetization preparation similar to a T2 preparation, with a 90 degree excitation followed by a long, low amplitude radiofrequency pulse oriented in the transverse plane with phase shifted 90 degrees to the excitation pulse, so that the magnetization “locks” or nutates around the axis of this pulse. The spin-lock pulse is then followed by a “tip-up” pulse. The duration of the spin lock pulse (TSP) is similar to the TE in a T2 measurement. The phase of the spin lock pulse is often shifted by 180 degrees half-way through the pulse to correct for field inhomogeneities (12). This preparation can then be followed by a small-tip acquisition with Cartesian or other k-space trajectories, again like T2 preparation. Originally, T1ρ was studied at high spin locking frequencies on the same order as protein dynamics in biological tissue, with the purpose of uncoupling dipolar interactions of adjacent protons. For instance, for water “bound” to collagen, the residence time is on the order of the 10us so one would need locking pulses on the order of 100 KHz. For hardware and heating considerations, typical spin-lock pulses in commercial clinical systems range from 0.1 to a few KHz. A locking pulse at 0.1KHz is equivalent to a refocusing train with a refocusing interval 5ms. Such pulses primarily affect decay associated with proton undergoing slower dynamic processes such as diffusion of water through local field inhomogeneities around susceptibility centres (e.g. deoxygenated red blood cells where time to sample the local field inhomogeneity is on the order of 10ms) and chemical or spin exchange which takes place on the same time order.
T2* – The utility of quantitative T2* has been best demonstrated in the management of patients suffering from cardiac iron overload (thalassemia, sickle cell and hereditary hemochromatosis). Anderson et al. were the first to describe the relationship between myocardial T2* and left ventricular ejection fraction in thalassemia major patients (13). They demonstrated that patients with myocardial T2*<20 ms exhibited a proportional decline in ejection fraction, thus establishing a link between iron concentration and cardiac dysfunction. Recent human autopsy studies (14) and preclinical studies (15) have revealed the direct relationship between iron and relaxation parameters. Recent preclinical studies have also demonstrated how relaxation measurements can help predict the clearance of iron based on type of iron chelator and acuity of iron loading in the heart (16). The utility of T2* mapping has also expanded to the detection of intramyocardial hemorrhage, a consequence of ischemia-reperfusion injury in acute myocardial infarction (8,17,18). Hemorrhagic infarcts have been shown to produce the worst clinical outcomes. In general, T2 shortening can also identify hemorrhage (19); however, T2* is more sensitive and specific to iron degradation products of hemorrhage.
T2 – The inflammatory state of cardiac tissue post injury is typically reflected by the presence of edema or swelling. T2-weighted signal has been found to increase under edematous conditions due to short molecular correlation times in free water; however it is disadvantaged by concerns with coil sensitivity correction and the inability to monitor serial progress. Alternatively, T2 mapping is growing in popularity as it allows monitoring of the inflammatory state regionally, serially as well as across subjects (8,20), a critical consideration while interrogating therapeutic strategies. Secondly, blood-oxygen-level-dependent (BOLD) imaging with T2 offers a novel way to probe oxygenation state by exploiting the paramagnetic properties of deoxyhemoglobin in blood and thereby myocardial perfusion reserve (21,22). A recent study has demonstrated that T2-based BOLD effects can identify impaired vasodilator function in infarcted and remote myocardium after an acute ischemic event (23).
T1ρ – Given the use of T1ρ to explore cartilage where proteoglycans and collagen play a major role, there has been significant recent interest in extending the use of this contrast mechanism to myocardial characterization in association with development of fibrosis after ischemic events (12). Signal intensity in T1ρ-weighted images (with typical spin-lock frequency around 500 Hz and duration around 50 ms) of myocardial infarction in a pig model 8 weeks post-infarction demonstrated hyperintensity in the region of infarct. Increasing the spin-lock frequency increased the contrast between infarct and healthy myocardium. It is not clear to what extent the contrast is associated with edema versus water interacting with collagen; one also sees significant elevation in T1ρ-weighted signal intensity in the infarct region in acute phases where tissue swelling is predominant. T1ρ is also diminished in the presence of microvascular obstruction (likely hemorrhage) (12) in a manner similar to T2* and T2. Recently, this technique has been translated to studies in patients (24,25).
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