Perhaps the most significant reason for the new advances in magnetic resonance (MR) has been the rapid development of magnet and field gradient designs at higher field strengths of 3 tesla (T) and more recently, 7 T.1–5 From its beginning in the 1980s, MR as an imaging method suffered from inherently low signal intensity (or signal-to-noise ratio [SNR]), in which approximately 5 hydrogen protons of every 1 million within a voxel contributes to the MR signal at 1.5 T. At higher magnetic fields (B0) of 3 T for example, that number of protons contributing to the MR signal is 10 of every 1 million, which means that the all-crucial SNR in MRI is also approximately linear with the strength of the main magnetic field B0, such that the raw SNR from 1.5 T to 3 T is about a factor of 2, and a factor of 4.7 at 7 T. Why is this important? Two reasons: The only effective way to increase the SNR at a given field strength is to increase the number of signal averages, which at 1.5 T would be a fourfold increase in the number of averages, which typically translates into a fourfold increase in the scan time. A second important reason is that the extra SNR at higher fields can be used to trade off for other image enhancements, such as higher resolution, shorter scan times, and shorter echo times (TEs), to name only a few trade-offs. This second reason explains why the phenomenon of “parallel imaging” in MRI has accelerated the field so rapidly with the advent of higher magnetic field strengths; it facilitates these trade-offs.
The purpose of this article is to highlight recent advances in neuroimaging from two aspects: (1) those advances directly benefited by increases in field strength (increased T1, SNR, magnetic susceptibility-sensitivity, and chemical shift) and how the increased SNR can be used to trade off for other advantages and (2) those advances made in response to attempts to try to reduce the inherent artifacts encountered at higher field strengths (eg, reducing specific radiofrequency (RF) absorption in tissue and magnetic susceptibility).
Other game-changing inherent advantages of higher field strengths are the linear increase in the chemical shift, which, along with increased SNR at high field, has made MR spectroscopy more of a routine clinical use.6,7 In addition, longer proton relaxation times, T1s, are observed at higher fields, again typically a linear response. Because of the longer T1s, inflow methods such as MR angiography and postcontrast agent T1-shortening effects are better visualized at higher fields.2,8 The longer T1 values have also been important in the development of arterial spin labeling or non–contrast-enhanced perfusion MRI for clinical use in such diseases as stroke when used at high field.9,10
Looking back, perhaps the biggest reason why fields of 3 T and higher became important throughout the last decade was the improvement of the blood oxygenation level dependent (BOLD) approach used for functional MRI (fMRI). For fMRI, the sensitivity created by the nearly fourfold increase in T2* sensitivity at 3 T over 1.5 T (where T2* can be thought of as the relaxation time T2 measured in the presence of magnetic susceptibility effects such as blood, air, contrast agents, and so forth) has allowed for crucial fMRI advances, in single-event studies for example. Whether fMRI has made, or will make, a true clinical impact in neuroimaging is debatable; however, other blessings in disguise will likely overshadow the fMRI clinical advantages, to such an extent that 3 T is rapidly becoming, or has already become, the de facto standard field strength for clinical neuroimaging11 for the reasons discussed below.
Not all T2* effects lead to discussions regarding fMRI. Magnetic susceptibility effects from increased T2* sensitivity at high fields,12 although initially a curse of high-field MR (for causing image distortion artifacts, among other things) are today the means for cutting-edge MR applications such as susceptibility-weighted imaging (SWI) or T2*-sensitive imaging for detection of bleeds, calcium, oxygen, or hemosiderin. In other words, T2*-weighted MRI at 3 T is becoming a key sequence relevant for detection of microbleeds in vascular encephalopathies or for visualizing cerebral hemorrhage.13–16 Increased T2* sensitivity in high fields leads to more “dephasing” of the proton signal, depending on the proton microenvironment. The dephasing can be mapped using a phase-sensitive image reconstruction method. (The authors normally only map the magnitude of the proton signal).
Why the interest in high-field dephasing of the proton signal? It is simply that phase-sensitive methods can better visualize the microscopic effects of vascular blood oxygenation changes (eg, in stroke or in vascular disease) that might occur at resolutions below the threshold for MR angiography. SWI at high field has considerable potential13–18 in visualizing tissue oxygenation changes on a vascular level, ideal for venography. By mapping the proton phase changes due to the susceptibility T2*effects of iron, SWI becomes sensitive to small venous vessels. More broadly, phase-sensitive imaging will lead to an entirely new class of tissue contrast mechanisms based on the proton microenvironment and to the frequency shifts experienced during image acquisition, shifts that are sensitive to tissue oxygenation, structure, and heterogeneity. One may expect to see phase-sensitive MRI (and SWI) be used for amyloid deposits, calcifications, white matter heterogeneities, and other vascular abnormalities (Fig. 1).