At both public and private imaging centers, the majority of the expenditures represent fixed costs. As a consequence, patient throughput has an enormous effect on the bottom line. Beyond some threshold number of scans per day, additional studies reflect almost pure profit for the imaging center. Optimum patient utilization thus potentially can increase overall revenues, and thereby stabilize the amount of the fixed cost required for individual MR exams. Surely thisd explains some of our emphasis on reducing scan time. But there is a different spin we can put on the problem. With each substantive decrease in imaging time, we open a new range of applications for MRI.
Prior to the introduction, in 1982 [1] , of practical 2D multi-slice imaging, and shortly thereafter of multi-echo multi-slice imaging [2] , MRI scan times were unrealistically long for patient care. Once a brain examination could be completed in a time somewhat shorter than a patient's tolerance for the procedure, in a scan time comparable to that of CT, MR imaging's acceptance began to grow. While scan times decreased steadily with technological improvements, the next major gain in speed came with the introduction, in 1986 [3] , of the FLASH (Fast Low Angle SHot) gradient echo method. By reducing practical scan times to as little as 10 seconds, the FLASH method considerably increased MR applications in body regions, especially the abdomen, where suspended respiration could eliminate most motion related image distortions [4, 5, 6] . In many cases however, the burden of altered contrast (as compared to spin echo) in FLASH imaging has limited its acceptance [7] . Where contrast based on tissue relaxation time differences is not the key clinical issue, the speed gains with FLASH imaging can be used to tremendous advantage. MR angiography, for example, simply would not be practical without gradient echo techniques.
Cardiac MRI, though the subject of intense interest for years, has been plagued by the difficulty of reliably triggering MR scans to the cardiac cycle. Each error in triggering potentially results in tremendous image degradation. When scans could be completed in a time short compared to the human heartbeat, a major physiological threshold was reached. In the past few years, at least two major techniques have cleared the hurdle of sub-second imaging: practical echo planar imaging in 1987, and sub-second FLASH methods three years later. While most gross organ motion can be resolved with these scan times of a few tenths of a second, there remain essential physiological events that require still better temporal resolution. Precise scan synchronization allows visualization of temporal details in, for example, cardiac function, but many physiological processes, such as peristalsis, swallowing, saccadic eye motion, joint clicks, phonation, and so on, are not cyclically repetitive. The most recent threshold of scanner speed vs. physiology was therefore crossed when MRI scan times reached the tenth of a second range. MRI now can examine such details of function as mental processes, rendering brain physiology and pathophysiology for the first time subject to minimally invasive, in vivo imaging. Ultimately, the major impact of high speed imaging may be on a highly expanded range of applications, rather than a simple increase in the number of scans routinely obtained per site.
FIGURE 1. Relationship between physiological processes and imaging speeds of available methods. In order to avoid image artifacts from motion, scan times must be not be longer than the duration of the motion. In order to study that motion, scan times must be five to ten times faster.
One way to overcome this is to use an excitatory RF pulse which disturbs the magnetization to a lesser degree: a so-called shallow RF pulse. The flip angle of the RF pulse describes how much of the magnetization is converted to MR signal with each excitation. A 90 degree; pulse fully converts that magnetization; the signal is, in general, proportional to the sine of the RF flip angle (figure 2).
Figure 2. The strength of the MR signal is shown in [blue] and is equal to the projection of the magnetization on the transverse or "imaging" plane. Large flip angles yield larger MR signals.
Given that shallow flip angles yield less signal, of what use is a shallow flip technique? The answer lies in the realization that, with small flip angles which convert less magnetization, less time is required for the protons to re-magnetize after the excitation pulse (Figure 3).
Figure 3. When small flip angles are used for excitation, there is relatively little loss of longitudinal magnetization (shown vertically). As a result, the tissues will re-magnetize more rapidly following a small excitation pulse.
The first imaging technique to utilize this reduced flip angle advantage was the FLASH method of Frahm and Haase [3] . Though small flip angles cannot yield signals as strong as those of large flip angles, it is possible at any TR to specify a flip angle, less than 90 degree;, which will give the maximum signal possible at that TR. Generally speaking, with FLASH-type scanning it is possible to control the T1 contrast weighting with both flip angle and TR: as the TR is increased, or the flip is angle decreased, the T1 weighting in the image is reduced (figure 4).
Unfortunately, FLASH imaging does not admit the use of 180 degree; RF pulses to form spin echoes. Therefore it is not possible to use TE to control T2-contrast in the same fashion as conventional imaging. The contrast behavior of FLASH scans is therefore unlike that of spin echo. Where contrast based on T1 or proton density is sufficient, FLASH has proven a reliable alternative to spin echo scans [4, 5, 6] with a scan time short enough for suspended respiration. Omission of the 180 degree; pulse makes FLASH images more susceptible to the effects of iron-containing substances that distort the magnetic field. This too, has been used to clinical advantage in, for example, the evaluation of hemorrhage [8, 7, 9] . As we will see below, the relatively high sensitivity of FLASH and other gradient echo imaging techniques to magnetic field distortions has been helped the development of functional brain imaging.
Hardware permitting, the FLASH method may be extended to the domain of extremely short TR's if sufficiently small flip angles are used. This has been the goal in the development of "snapshot" or "turbo" FLASH methods [10, 11, 12, 13] . A TR of 5 msec and flip angle of 6 degree; allows even moving structures, such as the heart, to be scanned in about one-half second (using a 128 line imaging matrix) with minimal motion artifact. Because small flip angles must be used, this can yield only one-twentieth of the signal available with a 90 degree; flip and a long TR. Many of today's commercial scanners now allow TR's of well under 10 msec and have incorporated "RF spoiling" to provide more precise control of image contrast in FLASH scans.
The generalized success of FLASH imaging has resulted in the proliferation of a variety of small flip angle pulse sequences known by the acronyms GRASS, FISP, FAST, CE-FAST, SPGR, SSFP, PFI, PSIF, FFE, and so on. This short article cannot possibly hope to summarize all of these methods, but the interested reader is invited to read a more comprehensive review in chapter 31A of Taveras' Radiology [14] .
Figure 5. Equipment block diagram showing hardware modifications required for the addition of Instascan echo planar imaging to existing hardware.
The development of positive MR contrast agents for use in imaging the human bowel has been hampered by prominent smearing artifacts produced by the agents in the presence of respiratory and peristaltic motion. Since the latter is involuntary, it is essential that the MR images be obtained rapidly compared to normal bowel motion. As demonstrated by Hahn at the MGH, standard CT contrast agents can be used to great advantage in MR imaging of the gastrointestinal system when combined with ultra-fast MRI [22] (figure 6).
Figure 6. T2-weighted Instascan (echo-planar) images of the gut after administration of the readily-available CT contrast agent, Readi-Cat 2 (E-Z-M corporation). The long T2 of the agent results in dramatic signal enhancement in these scans. Matrix 128 x 128, TR = ∞, TE= 75 msec, slice thickness 10 mm.
While the most obvious motion artifact challenges are in the thorax and abdomen, a surprising number of cranial MR images become technical failures due to patient motion. Ultra-fast imaging may be a useful alternative in such cases (Figure 7).
It has been recognized for some time that the NMR signal from skeletal muscle changes dramatically with exercise [29] . By combining ultra-fast T2-weighted imaging with active exercise - within the magnet - it has been possible to elucidate the time course of such signal changes and so to demonstrate recruitment patterns of muscles during complex exercise [30] (Figure 9). Such studies have clear implications for both clinical and basic scientific research.
Figure 9. Transverse sections of the thigh during active knee flexion. Following only thirty seconds of exercise, changes in the signal intensity of the hamstring muscles are clearly visible, as a result of large T2 changes. As exercise progresses, the signal intensity of the various synergists in this region increase (unequally) until fatigue. Such variations in signal changes may reflect both use and recruitment patterns during this active joint motion. TR = ∞, TE = 60 msec, slice thickness 10 mm, 64 x 128 matrix.
Another key application arises from the realization that diffusion is strongly temperature dependent. Our group, working with collaborators at the Brigham and Women's Hospital [37, 18] , and that of Le Bihan at the National Institutes of Health [38] , has shown that these temperature sensitive techniques may be applicable in the real-time monitoring of intraparenchymal surgery, potentially minimizing the surgical invasiveness and morbidity of even intracranial surgery. Since adequately diffusion-weighted imaging can now be performed using suitably modified gradient echo sequences, this is likely to become an area of increasing interest and growth.
Slightly coarser grained motions may also be measured with ultra-fast MRI. In a series of experiments inspired by the work of Feinberg, Poncelet has shown that when combined with velocity encoding gradients, echo planar imaging may be used to quantify precisely the pulsation-induced three-dimensional motion of the human brain within the calvarium [39] . Using similar imaging technology, but a different analysis strategy, Wedeen [40] has shown that velocity-dependent phase changes may be used to generate strain-rate images of the human heart, in which local deformations of the myocardium (as opposed to bulk tissue displacements) can be seen readily. Using this method it is possible to assess accurately the contractile activity of any voxel of myocardial muscle.
As commonly utilized in the CNS, injectable Gd-DTPA provides enhanced T1 relaxation when the blood brain barrier (BBB) is disrupted grossly. Since these T1 effects are local, and because the intravascular space within the brain is small, such agents have little impact in when the BBB is intact. Some tumors, acute infarcts, and most neurodegenerative diseases do not result in BBB destruction, while in others, such as multiple sclerosis, the disruption is variable [41] . As a result, conventional Gd-DTPA-enhanced imaging may not be adequate to characterize these disease states fully.
Susceptibility (T2 or T2*) -based contrast occurs over greater distances. When paramagnetic substances such as Gadolinium or Dysprosium are confined within the vasculature, they induce magnetic field distortions which extend beyond the vascular boundaries and result in transverse relaxation enhancement of the surrounding tissue protons [42] . Since the intact BBB effectively compartmentalizes injected Gd-DTPA, it is possible to use this effect to track signal intensity changes following bolus contrast agent administration. The resulting relaxation enhancement shows a near linear relationship to the intravascular contrast agent concentration, at least within the physiological range of study [43] . When combined with ultra-fast imaging during the passage of Gd-DTPA through the cerebral vasculature, it is thus possible to derive concentration time curves of the cerebral transit of the injected contrast agent [44] and, by using the principles of tracer kinetics, to derive cerebral blood volume (CBV) maps [45, 46] . Interestingly, depending upon the MR imaging sequence used, the signal intensity decreases have a variable dependence on the size of the vessel carrying the paramagnetic agent. Images based on gradient echo contrast demonstrate total blood volume, while spin-echo based contrast, as achievable with Instascan imaging, is biased towards the microvascular volume [43] . At our center we utilize such spin echo scans for their relative insensitivity to artifacts as well as for their optimum tissue sensitivity.
An important application of CBV imaging is in the evaluation of stroke, as demonstrated by Rosen [47] . The absence of perfusion within in farcted tissues is conspicuous in the lack of signal change during contrast agent passage. In animal models the results compare favorably to those obtained with more conventional perfusion measures [45, 48] . Brain tumors are generally accompanied by neovascularization, which supports the increased metabolic requirements of active tumor. It is thus reasonable to expect that the MRI-derived cerebral blood volume maps might find application in the study of cerebral neoplasm. Because each echo-planar image is acquired in about 1/10th of a second, several slices can be acquired at once while still achieving good temporal resolution with this method. Figure 10, from the work of Aronen [49] , shows results from our multi-slice imaging protocol. The multi-slice technique affords clear advantage in that the volume extent of the lesion can be evaluated clearly without the need for repeated contrast agent administration. In a study of thirteen tumor patients, the MRI cerebral blood volume mapping technique was able to demonstrate features of tumor activity not seen by conventional MRI and, when used longitudinally, to provide information about microvascular changes that occurred secondarily to either radio- or chemotherapy [49] . Perhaps more surprising, the correspondence between the MRI CBV maps on the one hand and those derived using positron-emission tomography (PET) with 18F fluorodeoxyglucose (18FDG) on the other, was good to excellent in the majority of the tumor patients studied [44]. these results may imply a linkage between the metabolism, angiogenesis, microvascular density, and tumor grade.
Figure 10. Multi-slice cerebral blood volume (CBV) map of tumor patient. Using a single injection of Gadolinium, this multi-slice CBV map was calculated from the signal intensity time course derived from over 300 separate images. The grey scale represents relative cerebral blood volume.
The method developed by Belliveau in 1991 required the use of exogenous contrast agents. Some years before, Thulborn had shown that paramagnetic changes in hemoglobin result in MR-visible decreases in T2 as hemoglobin goes from an oxygenated to a deoxygenated states [51] . Ogawa, using conventional strategies [52] , and Turner using echo-planar methods [53] , have shown that signal intensity changes are apparent in apnea in animal models. Based on these results, Kwong suspected that echo-planar MRI could be used to follow changes in brain oxygenation in humans, first during breath-holding and ultimately during photic stimulation [54] . Similar results now have been achieved using a variety of high speed imaging methods and field strengths, and have resulted in an enormous wealth of data on the use of MRI in functional studies of the human brain [55, 56] . The oxygen contrast methods rely on the hyperperfusion of brain tissue which accompanies neuronal activity. The consequence of this increased blood supply is an increase in the oxygen content of the venous blood, implying that blood flow increases exceed the metabolic requirements of the cells. Kwong's study used both flow-dependent (T1) contrast and O2 dependent (T2*) contrast and showed that the onset of the O2 increases had a rise time of about 4.4 seconds. This suggests that the temporal resolution requirements of the functional MR imaging method must be of about the same order - that is, a few seconds. Because this is long compared to the speed capabilities of the echo-planar experiment, as in the CBV mapping experiments discussed earlier, EPI will provide the advantage of multi-slice or volume imaging for mapping of human brain function. Figure 11.
Figure 11 shows the activation of primary motor cortex during finger tapping moving of either the right or left hand (data from Chantal Stern, Massacusetts General Hospital).
In their book, NMR Imaging in Biomedicine [57] Mansfield and Morris wrote:
"Although reported only at an anecdotal level, it is known that some of the early pioneers of NMR had used themselves as specimens in some early biological experiments. For example, as long ago as thirty years, Edward Purcell is reputed to have placed his head in a suitable NMR coil and magnet in an effort to try to observe differences in NMR signal shape or line shape caused by his thinking intensively or concentrating hard on a specific task as opposed to, as far as possible, making the mind blank. Erwin Hahn also tried experiments of a similar nature on his own head. In all cases there appeared to be no difference between the observed signals when the brain was idling as opposed to when the brain was concentrating. No effects, adverse or otherwise, were reported. Although these experiments could in no way be interpreted as imaging experiments, or indeed, forerunners of NMR imaging experiments, it is clear that the early practitioners of NMR had strongly in their minds the possibility of applying NMR to the study of biological systems."They could hardly have been more prescient.
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